Nanostructured biomaterials for regenerative medicine 9780081025949, 0081025947, 9780081025956

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Nanostructured Biomaterials for Regenerative Medicine

Woodhead Publishing Series in Biomaterials

Nanostructured Biomaterials for Regenerative Medicine

Edited by

Vincenzo Guarino Michele Iafisco Silvia Spriano

An imprint of Elsevier

Woodhead Publishing is an imprint of Elsevier The Officers’ Mess Business Centre, Royston Road, Duxford, CB22 4QH, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, OX5 1GB, United Kingdom © 2020 Elsevier Ltd. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/ permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such ­information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN: 978-0-08-102594-9 (print) ISBN: 978-0-08-102595-6 (online) For information on all Woodhead publications visit our website at https://www.elsevier.com/books-and-journals

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Contributors

Argelia Almaguer-Flores School of Dentistry, National Autonomous University of Mexico (UNAM), Mexico City, Mexico Marcela Arango-Ospina Institute of Biomaterials, Department of Materials Science and Engineering, University of Erlangen-Nuremberg, Erlangen, Germany Adriana Augurio  AO Research Institute Davos, Musculoskeletal Regeneration Program, Davos Platz, Switzerland Aldo R. Boccaccini Institute of Biomaterials, Department of Materials Science and Engineering, University of Erlangen-Nuremberg, Erlangen, Germany Martina Cazzola  Department of Applied Science and Technology (DISAT), Politecnico di Torino, Turin, Italy Christophe Drouet CIRIMAT, University of Toulouse, CNRS/INPT/UPS, Toulouse, France Sara Ferraris Department of Applied Science and Technology (DISAT), Politecnico di Torino, Turin, Italy J. Gaspar INL—International Iberian Nanotechnology Laboratory, Braga, Portugal Vincenzo Guarino  Institute of Polymers, Composites and Biomaterials (IPCB), National Research Council of Italy, Naples, Italy Jie Huang  Department of Mechanical Engineering, University College London, London, United Kingdom Michele Iafisco Institute of Science and Technology for Ceramics, National Research Council, Faenza, Italy Jonathan Massera Faculty of Medicine and Health Technology, Tampere University, Tampere, Finland Marco Morra Nobil Bio Ricerche S.r.l., Portacomaro (AT), Italy

xiiContributors

Roger Narayan UNC/NCSU Joint Department of Biomedical Engineering, Raleigh, NC, United States Qaisar Nawaz  Institute of Biomaterials, Department of Materials Science and Engineering, University of Erlangen-Nuremberg, Erlangen, Germany Ortensia Ilaria Parisi Department of Pharmacy; Macrofarm s.r.l., c/o Department of Pharmacy, Health and Nutritional Sciences, University of Calabria, Rende, Italy Silvia Pascale Sorin Group Italia S.r.l. (LivaNova), Saluggia (VC), Italy L.R. Pires INL—International Iberian Nanotechnology Laboratory, Braga, Portugal Francesco Puoci  Department of Pharmacy; Macrofarm s.r.l., c/o Department of Pharmacy, Health and Nutritional Sciences, University of Calabria, Rende, Italy Alice Ravizza DIMEAS Politecnico di Torino, Torino, Italy Christian Rey  CIRIMAT, University of Toulouse, CNRS/INPT/UPS, Toulouse, France Sandra E. Rodil Institute of Materials Research, National Autonomous University of Mexico (UNAM), Mexico City, Mexico Mariarosa Ruffo  Department of Pharmacy; Macrofarm s.r.l., c/o Department of Pharmacy, Health and Nutritional Sciences, University of Calabria, Rende, Italy Ahmed Salama  Cellulose and Paper Department, National Research Centre, Giza, Egypt Tiziano Serra AO Research Institute Davos, Musculoskeletal Regeneration Program, Davos Platz, Switzerland Nadia Shukry  Cellulose and Paper Department, National Research Centre, Giza, Egypt Phaedra Silva-Bermúdez  Department of Tissue Engineering, National Institute of Rehabilitation Luis Guillermo Ibarra Ibarra, Mexico City, Mexico Silvia Spriano Department of Applied Science and Technology (DISAT), Politecnico di Torino, Turin, Italy Pichaporn Sutthavas Department of Instructive Biomaterials Engineering, MERLN Institute for Technology-Inspired Regenerative Medicine, Maastricht University, Maastricht, The Netherlands

Contributorsxiii

Paulo Tambasco de Oliveira University of São Paulo, School of Dentistry of Ribeirão Preto, Ribeirão Preto, Brazil Riccardo Tognato  AO Research Institute Davos, Musculoskeletal Regeneration Program, Davos Platz, Switzerland Sabine van Rijt  Department of Instructive Biomaterials Engineering, MERLN Institute for Technology-Inspired Regenerative Medicine, Maastricht University, Maastricht, The Netherlands Fernando Warchomicka Institute of Materials Science, Joining and Forming, Graz University of Technology, Graz, Austria Seiji Yamaguchi  Department of Biomedical Sciences, Chubu University, Kasugai, Japan Aygul Zengin Department of Instructive Biomaterials Engineering, MERLN Institute for Technology-Inspired Regenerative Medicine, Maastricht University, Maastricht, The Netherlands Bin Zhang  Department of Mechanical Engineering, University College London, London, United Kingdom Leonardo Raphael Zuardi University of São Paulo, School of Dentistry of Ribeirão Preto, Ribeirão Preto, Brazil

Preface

The rational design of biomaterials plays a crucial role on the success of devices for different medical applications. The investigation of ex novo chemical synthesis as well as the use of unconventional technological strategies allows manipulating biomaterials into appropriate two-dimensional (2D) and three-dimensional (3D) forms, in order to tailor the main physical, chemical, structural, and biological properties requested to achieve desired clinical efficacy. As a function of their peculiar properties—that is, degradation, mechanical response, biocompatibility, fluid transport, and absorption—biomaterials from natural or synthetic source can be properly manipulated to (a) design innovative devices to support lost/dysfunctional tissues or organs, (b) mimic extracellular matrix (ECM) by temporary 3D substrates able to guide neo tissue formation and organization, or (c) fabricate nano-sized systems with controlled interfaces and/or molecular release for innovative therapeutic uses. This book offers a large but punctual overview of innovative platforms based on nanostructured biomaterials, focussing the attention on the peculiar chemistry of used biomaterials—by a well-consolidated classification including polymers, ceramics, and metals. For this purpose, the book includes 15 chapters, divided into four different subsections. An introductory section is mainly aimed at introducing the current state of art for the use of biomaterials in tissue repair and regeneration (Chapter 1), also taking into account basic regulatory aspects (Chapter 2) and future targets in clinical use (Chapter 3). The other three sections singularly addressed the recent discoveries on the use of polymers and their composites (Chapters 4–7), ceramics (Chapters 8–11), and metals (Chapters 12–15) for the design of devices for relevant medical needs. In each section, the peculiar role and the impact of nanotechnology and nanomaterials in helping to tackle today’s urgent challenges and to achieve more effective treatments as well as more successful diagnoses is emphasized.

Introducing biomaterials for tissue repair and regeneration

1

Vincenzo Guarino*, Michele Iafisco†, Silvia Spriano‡ Institute of Polymers, Composites and Biomaterials (IPCB), National Research Council of Italy, Naples, Italy, †Institute of Science and Technology for Ceramics, National Research Council, Faenza, Italy, ‡Department of Applied Science and Technology (DISAT), Politecnico di Torino, Turin, Italy *

1.1 Introduction Over the last decade, there has been a significant progress toward the development of biomaterials to be used for the fabrication of innovative devices for a wide variety of biomedical applications. The manipulation of material chemistry and processing technologies allows for the design of tailor-made systems with peculiar mechanical/ morphological/surface/functional properties which can be tailored on the specific application of interest. In particular, current state of the art in biomaterials design is continuously evolving to offer a portfolio of innovative devices to support the functionalities of natural tissues. In recent years, there has been increasing emphasis on materials that could be used in tissue repair and engineering areas. After an early empirical phase of biomaterials selection based on availability, some attempts were primarily focused on either achieving structural/mechanical performance or on rendering biomaterials inert and thus unrecognizable as foreign bodies by the immune system. Hence, biomaterials were used as implants in the form of sutures, bone plates, joint replacements, ligaments, vascular grafts, heart valves, intraocular lenses, dental implants, and medical devices like pacemakers and biosensors [1, 2]. Secondly, biomaterials have increased their relevance as elementary unit to design synthetic frameworks, namely, scaffolds, matrices, or foams—with features at micro-, submicro- and nanoscale—able to guide the in vitro and/or in vivo mechanisms during the regeneration process of natural tissues. More recently, particular interest has been addressed to biomaterials at the nanoscale for the enormous opportunities to exploit peculiar properties (i.e., extended surface area, high surface to volume ratios, high reactivity) to design a large variety of conventional or unconventional devices with improved interfaces with cells and macromolecules, toward the definition of innovative therapeutic/ diagnostic/theragnostic therapies. To date, several criteria have been used to select biomaterials, mainly based on their chemistry, molecular weight, solubility, shape and structure, hydrophilicity/hydrophobicity, lubricity, surface energy, water absorption degradation, and erosion mechanism. Commonly, materials for biomedical use can be divided into three principal classes: metals, ceramics, and polymers—natural or synthetic ones, respectively [3]. Herein, we would discuss the main features of biomaterials, basically remarking how Nanostructured Biomaterials for Regenerative Medicine. https://doi.org/10.1016/B978-0-08-102594-9.00001-2 © 2020 Elsevier Ltd. All rights reserved.

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the peculiar properties of materials influence specific functions, thus directing their use toward different fields of applications in tissue repair and regeneration.

1.2 Biomaterials: Basic concepts 1.2.1 Metals The applications of metals as biomaterials are varied and range from surgical equipment components to bones substitutions, artificial joints or valves, fracture fixation devices, stents, and dentures. Most are used in orthopedic and dental surgery, but cardiovascular applications are also widely diffused. The main causes of failure of metal implants were, and partially still are, wear, low fracture toughness or strength, stress shielding, fibrous encapsulation, inflammation, and infections. The evolution of metallic biomaterials was targeted as first to biocompatibility and mechanical issues (to face wear, low fracture toughness or strength, stress shielding), then to fast osseointegration (to face fibrous encapsulation), and recently to multifunctional properties (with a focus on modulation of inflammation and limitation of the risk of infection), as summarized in Fig. 1.1. Metallic materials have important mechanical characteristics, with respect to other materials, such as high elastic modulus (100–200 GPa) and high yield strength (300– 1000 MPa), thus make it possible to build structures capable of supporting high loads without large elastic deformations or permanent plastic deformations. As last, because of good ductility, when the applied stress exceeds the yield strength, the metal structure shows plastic deformation rather than brittle breaking, allowing the replacement of the deformed implant before it breaks. Adequate mechanical strength, under static or cyclic loading (fatigue), currently is an almost met demand by metal biomaterials (0.5% of people experiencing hip prosthesis breakage). On the other side, a low elastic modulus is of interest in bone contact applications where stress shielding effect has to be avoided and this is still an open issue. Concerning the biological requirements, biocompatibility (i.e., strictly related to high corrosion resistance), stable anchorage with tissues, and bioactive behavior (to

Failure of metal implants

Wear

Low fracture strength/ toughness

Biocompatibility

Stress shielding

Mechanical properties

Fibrous encapsulation

Fast osseointegration

Inflammation

Infections

Multi-functional properties

Evolution of biomaterials

Fig. 1.1  The main causes of failure of metal implants and evolution of metal biomaterials in order to face them.

Introducing biomaterials for tissue repair and regeneration3

avoid fibrous encapsulation) are the main needs. Even though currently used bulk metal materials usually have good biocompatibility and corrosion resistance, a still unmet issue is related to minimization of the friction forces and wear in the artificial joints in order to avoid toxic wear debris and to guarantee a long implant life. Stable anchorage and bioactive behavior were widely investigated and several surface treatments, coatings, and functionalization processes were developed and are in use, even if some un-met issues are still remaining. Any innovation concerning surfaces must avoid introducing further complications due to, for instance, excessive surface roughness, stress concentration factors, or material changing due to added thermal treatments. More recently, the clinical demand moved to multifunctional surfaces able to simultaneously give a specific response to colonization by different cells (osteoblasts, fibroblasts, macrophages) and infection agents (bacteria, viruses) [4]. This complex system of multiple biological stimuli and surface responses is the main focus of the current research on biomaterials. The development of antibacterial surfaces able to avoid biofilm formation are highly challenging for bone implants. It has been reported that deep infections typically occur in 1%–2% of patients with total hip arthroplasties [5] and dental peri-implant disease and infection have become a main focus in terms of prevention and treatment of oral implantology [6]. When deep infection occurs, removal and reimplantation of the implant is often necessary, with additional discomfort to the patients and costs for the health-care services. The main metals currently used as biomaterials are: titanium and titanium alloys, Co–Cr–Mo alloys, and shape memory or pseudo-elastic Ni–Ti alloys; stainless steel (AISI 316L) is mainly used in surgical equipment and only in a limited range of implants. Titanium and titanium alloys have excellent properties from the engineering standpoint [7]: low density, high strength (resistant as steel and twice as resistant as aluminum), low modulus of elasticity, low thermal conductivity, low thermal expansion, excellent corrosion resistance, biocompatibility, and extremely short radioactive halving period (which allows its use in nuclear systems). The combination of low density and high strength of Ti alloys produces particularly favorable strength/weight ratios, superior to almost all other metals. Titanium has a thermal expansion coefficient significantly lower than other metals, being much more compatible with ceramics or glasses, especially when coatings are involved. Titanium develops very stable surface oxides with high integrity, tenacity, and good adhesion, as last surface oxide on titanium, if scratched or damaged, can immediately be reconstructed in the presence of air or water. Usually, unalloyed titanium is used in applications where excellent corrosion resistance is desired and where high strength is not a determining factor. Beta titanium alloy containing nontoxic elements, such as Mo, Zr, Ta, Ta, and Nb, are of great interest because they show a lower modulus of elasticity and are under investigation to replace the TiAl4Va alloy, which is currently considered the most important biomedical alloy [8]. Cobalt-based alloys used in biomedical field comprise cast Stellite 21 (ASTM F75), wrought Stellite 25 (ASTM F90), forged alloys (ASTM F799 and F1537), and the multiphase alloy MP35N (ASTM F562). Young’s modulus is not affected

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by ­manufacturing processes because it is similar to all alloys varying among 200– 230 GPa. The wrought alloys have better mechanical behavior than the cast ones, especially compared to the tensile resistance, when subjected to hot and cold work. Wrought alloys F1537 and F799 (respectively, with high and low carbon content) have very similar tensile behavior, but not fatigue strength which is lower for F1537. Introduction of Ni in wrought F562 and several deformation workings improve the tensile strength, but decrease the fatigue stress. Due to their good wear resistance, Co– Cr–Mo alloys are used in artificial joints, but several biocompatibility issues are still unsolved. Metallic debris (mainly produced in metal-on-metal coupling) is smaller in volume and size than polyethylene debris, but they expose high surface area and are more prone to corrosion [6, 9]. Hence, even if Co and Cr are present in human body tissue, an increment of them can cause hypersensitivity and drastic inflammatory reactions. The potential toxic behavior of Co and Cr ions, such as Co2+, Cr3+, and Cr6+, is well recognized. They could cause chromosome breakage and DNA damage, cell apoptosis, and subsequently necrosis [9]. The nickel-titanium (Ni–Ti) alloys, especially the near-equiatomic ones, are characterized by excellent mechanical and functional properties, such as shape memory effect (SME), super-elasticity (SE), low elastic modulus, high corrosion resistance, and biocompatibility [10]. Pseudoelasticity refers to the recovery of shape from deformation which is much larger than conventional elastic deformation in metallic materials, in an isothermal condition. In general, the elastic strain of a metallic material is around 0.2%, and higher loading will result in plastic deformation, due to slip or twin boundary motion. On the other hand, the elastic strain could be up to 10% in the case of the pseudoelastic Ni–Ti alloys. The near-­ equiatomic Ni–Ti alloys are the only pseudoelastic materials nowadays available for industrial applications. The pseudoelastic behavior can be obtained at human body temperature, making the material an optimal choice for a self-expanding stent and other cardiovascular devices; the use of NiTi is limited as orthopedic implants and orthodontics wires.

1.2.2 Ceramics The term ceramic comes from the Greek word “keramos” which literally means “burnt stuff” alluding to the fact that the raw material must be heated at elevated temperature in order to form a device with a useful range of engineering properties. Ceramics can be defined in a very simplified way as inorganic nonmetallic materials [11]. More precisely, they are solid inorganic materials that combine metal and nonmetal atoms in which the chemical bonding ranges from very ionic to covalent [12]. They can be crystalline, partly crystalline (polycrystalline), or even amorphous and are usually formed by the action of heat and subsequent cooling. Ceramic materials have high melting temperatures, low conduction of electricity and heat, and relatively high hardness [12]. With regard to their mechanical behavior, ceramic materials exhibit great compression strengths and very low tensile strengths [12]. Moreover, they are stiff, with high Young’s modulus, and brittle because failure takes place without plastic deformation [12].

Introducing biomaterials for tissue repair and regeneration5

Ceramics as biomaterials, the so-called bioceramics, have been used for millennia especially to augment or replace various calcified parts of the human body such as bone and dental tissues [13]. Interestingly, Mayan skulls of >4000 years old in which missing teeth were replaced by nacre-based substitutes were discovered in 1972 by Amadeo Bobbio [14]. Nacre is a natural composite consisting of 95–98 wt% of calcium carbonate in the form of aragonite (the “ceramic” phase) and 2–5 wt% of organic matter (fibrous proteins, polysaccharides) [15]. In clinical practice, the routinely implantation of bioceramics started in the late 18th century in dentistry with the use of porcelain for crowns and in the 19th century in orthopedics with the use of plaster of Paris or gypsum (calcium sulfate dihydrate) for bone filling. Only during the 20th century ceramic materials specifically designed for hard tissue replacement become available thanks to the progress of chemistry and materials science, making possible the synthesis of fine and pure ceramic powders [16]. In particular, bioceramics appeared in the late 1960s as an effort to overcome some biocompatibility problems associated with the use of metallic implants in orthopedic surgery [17]. In fact, the biological environment is a humid and oxidative medium for many metals, and implant loosening, inflammation, and pain were common drawbacks in patients with metallic prostheses. Corrosion, metal particles debris, or loosening through wear still result in the failure of many metallic prostheses. Orthopedic implants (hip prostheses, etc.) are probably the most renowned application of bioceramics, even though they are also used for a large number of other applications such as dental, periodontal, middle ear, cranial, maxillofacial, otolaryngology, and spinal surgery [7]. Unfortunately, the high stiffness of a bioceramic may restrict its use in soft tissue applications; therefore, hybrid biomaterials have been developed with the combination of bioceramics and the polymers to produce composites that have the advantages of both materials (see next paragraph). The most common ceramics used in biomedical applications are silicates, oxides, carbides, sulfides, refractory hydrides, selenides, and carbon structures such as diamond and graphite [18]. The organization of bioceramics is usually carried out employing the “three generations” classification (Table  1.1). This classification is not based on the chemical-physical characteristics of the materials but rather on the performances of the ceramics when implanted in the organisms (i.e., in vivo reactivity). The first generation includes ceramics that are bioinert, namely they do not elicit any adverse reaction from the body and are “transparent” to the immune system maintaining their mechanical and physical properties after implantation [19]. When implanted for orthopedic or dental applications, no direct bone-ceramic interface is formed and the ceramic part replaces permanently the biogenic tissue. The study of the first generation of bioceramics started in the 1960s, when the goal was to maintain the reactivity as low as possible. The most common bioinert ceramics are alumina (aluminum oxide, Al2O3), zirconia (zirconium oxide, ZrO2), and their composite; pyrolytic carbon and silicon nitride are also used [20]. Bioinert ceramics are principally used for applications such as total hip and knee replacements, dental implants, dental crowns, etc., need to present sufficient high mechanical properties as compressive and tensile strengths, hardness, toughness, low wear, and good corrosion resistance against biological medium [20, 21]. Alumina or alumina-based composites are the most used ceramics for joint replacement and are often preferred to metals due to their

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Table 1.1  Classification of bioceramics employing the “three generations” classification Type of bioceramic

In vivo reactivity

Examples

Isolated by a nonadherent fibrous capsule

Alumina Zirconia Pyrolytic carbon Silicon nitride

Dissolved after a specific time Tightly bonded to living tissues

Calcium phosphates Silica-based bioactive glasses Calcium sulfate

Stimulating living tissue regeneration

Porous bioactive and biodegradable ceramics Advanced bioceramics: biomimetic ceramics; mesoporous materials, organic-inorganic hybrids

First generation Bioinert Nonresorbable

Second generation Bioresorbable Resorbable Bioactive Surface reactive

Third generation Scaffolds of biologically active molecules

better compatibility with human cells without cation release and their lower wear in friction stress conditions [22]. The use of alumina has been approved by US Food and Drug Administration (FDA) in 1982 for total hip arthroplasty. For dental applications, nowadays the most employed commercial products (i.e., ceramic crowns) are ­alumina-stabilized zirconia-based composites. Partially stabilized zirconia (with MgO or Y2O3) was proposed in 1986 as an alternative to alumina in ceramic femoral heads [22]. Since then, it has been gaining market share thanks to its enhanced mechanical properties in comparison to alumina. The first dental implants of zirconia were reported in 1993. Zirconia is more appreciated than titanium implants for aesthetical reason; in fact, zirconia being white is closer to the color of teeth. Moreover zirconia implants have a similar mechanical strength but higher fracture toughness and lower production cost than crystalline alumina. The limitations of zirconia as bioceramic are related to degradation and radiation. Improvements of these materials can occur through the preparation of zirconia/alumina composites. The second generation includes ceramics that are bioactive and bioresorbable. Bioactive means that they are able, when implanted, to chemically interact with the natural tissue and form bond at the interface. In this way, they can trigger a positive reaction from the body that is mainly an acceleration of the bone regrowth process. Bioresorbable means that material can be degraded over time by the organism and replaced by the endogenous tissue resulting in the possible regeneration and restoration of normal functional tissue. The most common bioactive and bioresorbable bioceramics clinically used for bone healing are calcium phosphates, silica-based bioactive glasses, and calcium sulfates [23, 24]. These materials were developed around

Introducing biomaterials for tissue repair and regeneration7

the 1980s when the principal objective changed to obtain favorable interactions with the living body, namely a bioactive response or degradation. Calcium phosphates being materials having chemical composition close to mineral phase of bone and tooth are probably the most important class of bioceramics thanks to their unrivalled biocompatibility [25]. Calcium phosphate bioceramics are widely used in medicine as bone substitutes, implants, and coatings on dental and orthopedic prostheses. Current biomedical applications include replacements for hips, knees, and teeth, as well as repair of periodontal disease, maxillofacial reconstruction, augmentation, and stabilization of the jawbone, spinal fusion, and bone fillers after tumor surgery. Chapter 8 is dedicated to synthetic calcium phosphates therefore the reader can find there more information on their chemical and physical properties and applications. Silica-based bioactive glasses were first prepared in 1969 by Larry Hench [26]. Since then, this research line has provided very interesting results in both academic and applied fields through the transformation of conventional glasses into glasses with biomedical added value. Bioactive glasses bond to and integrate with living bone in the body without forming fibrous tissue around them or promoting inflammation or toxicity. The high reactivity of these glasses is the main benefit for their application in periodontal repair and bone augmentation, since the products obtained from reactions between these types of glasses and physiological fluids lead to the crystallization of an apatite-like phase, similar to the inorganic component of bones in vertebrate species. Similar to calcium phosphates, Chapters  10 and 11 and focus on the silicate-based ceramics, where more details are reported. Calcium sulfate exists in two forms (alpha and beta), which differ greatly in physical properties [27]. Calcium sulfate has been used in bone regeneration as graft material and graft binder/extender and as barrier in guided tissue regeneration. It is an unusually biocompatible material and is completely resorbed following implantation [27]. It does not evoke a significant host response and creates a calcium-rich milieu in the area of implantation. Calcium ions may provide stimulation to osteoblasts, which may account for some of the positive results reported with this material. The raw material from which calcium sulfate is made is relatively inexpensive and abundant. Despite these advantages, calcium sulfate has never attracted the same degree of research interest as the other bioceramics such as calcium phosphates and silicate [27]. Recently, however, it has enjoyed a resurgence of sorts in the areas of periodontology, sinus augmentation, and orthopedic surgery. Third generation of bioceramics are materials designed and synthesized to stimulate specific cellular responses at the molecular level and to be able to release in vivo specific signals to improve the body response (e.g., porous ceramics, bioceramics functionalized by growth factors, adhesion peptides, or containing drug-­ releasing polymer capsules, etc.) (Fig. 1.2) [28]. They could be used as scaffolds for cells or other biologically active substances (i.e., ions, biomolecules, etc.) able to induce regeneration and repair of damaged tissues [29]. These materials were conceived at the end of the 20th century, when it was clear that the first-­generation bioceramics by themselves could not give a complete response to the clinical needs of biomaterials for implants. In fact, bioactive ceramics (second generation) are considered to be osteo-conductive (able to support bone formation in osseous defect sites), but not osteoinductive (able to induce de novo bone formation w ­ ithout

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Fig. 1.2  Third-generation bioceramics.

the presence of osteogenic factors). It is worth mentioning that the preparation of third-generation bioceramics does not aim to substitute plainly the materials from previous generations but to improve the initial physiological cues and bioactivity existing in the second-generation biomaterials. This can finally allow the biomaterials and surrounding environment to adapt, signal, and stimulate specific cellular activity and behavior in order to regenerate new tissues. The studies about the third-generation bioceramics are more based on biology and follow the purpose of substituting the “replacement” tissues by “regenerating” tissues [30]. This category includes bioceramics based on porous second-generation bioceramics, loaded with biologically active substances, and new advanced biomimetic ceramics. For example, tissue engineering calcium phosphates and silicate-based scaffolds with appropriate macro- and meso-porosity have been developed in order to achieve a desirable interaction with the surrounding environment and osteogenesis. The research has also focused on incorporating osteogenic inductive factors or osteogenic cells into bioactive scaffolds to be osteoinductive in tissue regeneration. Other examples of the most recent advancements on ceramic materials specifically designed for tissue regeneration are reported in the chapters of this book dedicated to calcium phosphates and silica-based materials.

Introducing biomaterials for tissue repair and regeneration9

1.2.3 Polymers In the last three decades, innovative devices for tissue repair and regeneration have been fabricated form various kinds of materials, including metals, ceramics, polymers, and their combinations. Due to the high mechanical strength and osteoinductivity, metals and ceramics have been preferentially used to face the bone tissue [31]. However, both metals and ceramics have some significant disadvantages, such as the fact that metals are nondegradable and the stress concentration may result from the stiffness of the metals. Although ceramics have limited biodegradability, their poor processability makes it difficult to transform the materials into scaffolds with specific shapes. In addition, most ceramics are brittle. In contrast, polymers have excellent processability, and their mechanical properties and degradation rates could be more easily tuned by adjusting the molecular designs. Therefore, polymers may be designed as suitable materials for use in scaffolds for tissue engineering [32]. Polymers for biomedical use can be classified into two types, namely natural polymers and synthetic polymers.

1.2.3.1 Natural polymers The growing attention toward the understanding of cell materials interaction has addressed the finding of new materials suitable to accurately reply the local biological microenvironment in order to improve cellular response and more efficiently modeling biological context. Hence, researchers have focused on the innate attitude of natural polymers to guide the cell behavior through biophysical and biochemical cues to mimic the native extracellular matrix (ECM). Indeed, natural polymers show an excellent biocompatibility and are able to send selective signals to cells to promote specific interactions. However, main constraints for the use of natural polymers include limited source to obtain the materials, difficulty to maintain the purity and quality of the product, poor mechanical properties, and uncontrollable biodegradation [33]. They can be further categorized into protein-type and polysaccharide-type. The commonly used natural protein-type polymers include collagen, gelatin, and silk proteins. Some common natural polysaccharide-type polymers are agarose, chitin, chitosan, hyaluronan, alginate, dextran, cellulose, and oxidized cellulose. The bacteria-produced polyester-type polymers, poly(hydroxyalkanoates) (PHAs), are also included in the category of natural biodegradable polymers.

Proteins Collagen is the most abundant protein in human body, whose fiber network can be found from the softest to the hardest tissue (such as nerves, blood vessels, skin, cartilages, and bones). Collagen fibers are formed from the assembly structure of multiple triple-helical collagen fibrils. Collagen (type) I, collagen II, and collagen III constitute >80% of all collagens in the human body [34]. There have been many reports regarding the fabrication of three-dimensional (3D) scaffolds consisting of collagen for the repair of tissue. Zhou et al. fabricated 3D scaffolds by freeze-drying collagen II, followed by cross-linking of 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide/N-­ hydroxysuccinimide (EDC/NHS) to enhance the mechanical strength. The freezedried scaffolds composed of type II collagen were found to enhance the proliferation

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of mesenchymal stem cells (MSCs) and promote MSCs to further differentiate into a nucleus pulposus-like phenotype [35]. A derivative of collagen is gelatin obtained by denaturation, that is, breaking down of triple-helix structure [36]. In particular, gelatin is simply the product of collagen hydrolysis, derived from sources rich in type I collagen from bovine, porcine, fish skin, bones, jellyfish, and bird feet. Based on the denaturation hydrolysis process, there are two types of gelatin, type A, which consist of an acid process where the collagen denaturation consists in the thermal hydrolysis of peptide bonds; type B in which the alkali breaks the cross-links, when collagen is heated the α-chains are released. Compared to collagen, gelatin is advantageous because it can be dissolved in aqueous solution during the fabrication process so that no ­acid-base neutralization is needed to obtain the final product. Nevertheless, cross-­linking of the materials was required to prevent rapid dissolution when the scaffolds were placed in the medium or body fluids. In particular, the use of specific cross-linking agents, that is, glyceraldehyde (GC), 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide hydrochloride (EDC), 1,4-butanediol diglycidyl ether (BDDGE), as cross-linker allows improving mechanical strength of protein-based scaffolds with relevant effects on the in vitro response [37]. Keratin is another fibrous protein, mainly synthetized in the cytoplasm of keratinocytes, being 90%–95% of cells in the epidermis. Keratin can be classified as α-helices or intermediate filament proteins (IFPs)—7–10 nm in size—and β-sheets, based on their secondary structure—3–4 nm in size [38]. The fundamental role of keratins is maintaining the architecture of cells, providing a scaffold for the cytoskeleton of cells and tissues and mechanical functions to maintain the structural integrity under mechanical stress, interacting on cell-cell junctions like desmosomes, hemidesmosomes, and focal adhesions [39]. Inside cells, it is involved in cell transport, in the regulation of protein synthesis, cell growth, in cell compartmentalization, and cell differentiation. According to previous studies, keratin is considered among the toughest biological material, making them suitable for the fabrication of platforms to be used as ECM analogs. Similarly, silk is a natural protein present in the form of fibers composed of a filament core protein—silk fibroin—and surrounded by sericin—a glue-like protein [40]. They form fiber filaments—10–15 μm in diameter—composed, from chemical point of view, of two major fibroin proteins, a heavy chain (~390 kDa) and a light chain (~26 kDa), at 1:1 ratio, linked by a single disulfide bond [41]. In this context, the formation of hydrogen bonds and hydrophobic interactions between hydrophobic blocks tend to form crystalline structures (β-sheets) which confer a high tensile strength to silk [42]. As a function of the local interactions of ordered hydrophobic blocks with the less ordered hydrophilic blocks, it is possible to modulate the elasticity and toughness features of the protein, with relevant benefits in the fabrication of structural scaffolds for tissue engineering applications.

Polysaccharides Polysaccharides are a family of carbohydrates playing fundamental roles in many biological contexts. Their structure is made of sugar rings linked by glycosidic bonds and various side functions [43]. Glycosidic bonds can be relatively easily biodegraded

Introducing biomaterials for tissue repair and regeneration11

via glycoside hydrolase enzymes while the side groups directly affect polymer charge density, hydration, and chemical reactivity, also being responsible for the formation of secondary structures. In the presence of charges, polysaccharides may behave like polyelectrolytes with the peculiar ability to ionize in aqueous media. Ionization aids in the solubilization of the polyelectrolytes, further being responsible for its unique properties. Noteworthy, polyelectrolyte dissolution is not comparable to the dissolution of a simple salt, because it does not produce ions, that is, cation/anion—with comparable size and independent mobility such as a salt in solution—but dissolving to yield a polyion and counter ions. In particular, polyions have mobility and hold a large number of charges in close proximity, so that individual charges are strongly attached to the macromolecular backbone, with short mobility within the domain of the macromolecular coil. This allows the self-assembly of differently charged polysaccharides by interactions between the negative and positive groups and by the entropic gain associated with these associations [44]. There are two main groups of polysaccharides: cellulose and all the derivatives from plant cells (i.e., alginates), or those from animal sources—that is, chitin from the shells of shrimp and other sea crustaceans (Fig. 1.2). Among the polyanionic polysaccharides, alginates are the more commonly used in biomedical field, because they are not toxic, biocompatible, highly hydrophilic, generally used as stabilizer, viscosifier, and gelling agent in food, textile, pharmaceutical, and biotechnological industries [45]. Alginic acid is obtained by acid extraction from algal tissue followed by neutralization with alkali and precipitation by the addition of calcium chloride or mineral acid. It is reconverted into sodium alginate through alkali treatment. Haug et al. have elucidated the structural constitution of alginic acid by partial acid hydrolysis, showing that chemically modified alginate is made from block of (1–4)-linked β-d-mannuronic acid (M) and α-l-guluronic acid (G) monomers, alternate into different forms of polymers. Customized formulations of these components, extracted from natural sources, by the main structural component of marine brown algae, confer them strength and flexibility [46] and assure a fully biocompatible response, making alginate a suitable biomaterial for tissue engineering and drug delivery applications. On the contrary, among polycationic polysaccharides, the most recognized one is the chitosan—a natural polymer obtained from chitin partially deacetylation, made from alternating monomers of 2-acetamido-2-deoxy-β-d-glucopyranose (GlcNAc; A) and 2-amino-2-deoxy-β-d-glucopyranose (GlcN; D). In this case, deacetylated amino groups are very electronegative and can take up a proton allowing chitosan to behave as a hydrophilic polycations, with different chemical, physical, and biological properties from chitin. In particular, deacetylation degree (DA) generally influences polarity, pH, ionic strength, and, ultimately, water-soluble behavior [47]. To improve thermal stability, swelling ratio and mechanical strength of chitosan, cross-linker agents like glutaraldehyde, oxalic acid, formaldehyde, glyoxal, and genipin can be used [48]. Moreover, positive charge of chitosan allows forming complexes with different ions, that is, transition metal and posttransition ions as a function of their selectivity Cu > Hg > Zn > Cd > Ni > Co = Ca [49]. Metallic anions, sulfate, citrate, and phosphate-bearing groups, such as β-glycerophosphate and tripolyphosphate (TPP), may react with chitosans to form derivatives well tolerated and biocompatible,

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useful in drug delivery. The main disadvantages of ionic cross-linking concern their lacks in mechanical and chemical stability, due to a highly pH-sensitive swelling [50]. Alternatively, γ-radiation and ultraviolet (UV) light have been used to prepare graft copolymers. Cai et al. used 60C-γ-radiation to graft N-isopropylacrylamide to fabricate pH-sensitive chitosan with good swelling properties [51]. Photo-induced grafting of polyethylenglicole (PEG) on carboxymethyl chitosan has been used to develop a pH responsive drug delivery system [52]. Yamada, et al. also proposed the use of enzymes to synthesize chitosan with unique pH-sensitive water solubility and adhesive properties [53]. Due to their cationic functions, chitosans may be successfully used to form complexes with oppositely charged polymers, that is, polyanions such as alginates or proteins, with enormous potential applications in cell encapsulation, drug delivery, and tissue engineering [54].

1.2.3.2 Synthetic polymers In the place of natural polymers, synthetic polymers, able to be synthesized or processed under controlled conditions, may offer an efficient strategy to exhibit highly predictable and reproducible mechanical and physical properties—that is, tensile strength, elastic modulus, and degradation rate. From biological point of view, interaction risks due to toxicity, immunogenicity, and infections delivery may be generally mitigated by the use of pure synthetic polymers, which are commonly composed of constituent monomeric units having a well known and simple structure [55]. For this reason, they have been selected long times as first choice to fabricate biodegradable porous scaffolds for tissue engineering applications. In particular, they include polyhydroxyesters, including polylactic acid (PLA) and poly(glycolic acid) (PGA), as well as poly(lactic-co-glycolide) (PLGA) copolymers [56]. PLA and PGA are degradable polymers approved by the US FDA. They can be processed easily and their degradation rates, and physical and mechanical properties can be finely modified into a wide range as a function of different chemical parameters such as molecular weights and copolymers ratios. Indeed, chemical properties of these polymers may deeply influence the hydrolytic degradation via de-esterification. Besides, the understanding of bulk erosion process of these polymers may be relevant to predict the in vitro fate of polymer systems, in order to evaluate any premature and uncertain failure. Indeed, the body contains highly regulated mechanisms for completely removing monomeric components of lactic and glycolic acids, which assure the complete removal of monomeric components by natural pathways. Moreover, the release of acidic degradation products can also occur, promoting a strong inflammatory response [57]. Polyester degradation occurs by uptake of water followed by the hydrolysis of esters. This mechanism is generally affected by several factors including chemical composition, processing history, molecular weight and polydispersity (Mw/Mn), environmental conditions, crystallinity, and porosity, especially in the case of 3D foams [58]. Hence, aliphatic polyesters can therefore exhibit quite distinct degradation kinetics. PGA, for example, is a stronger acid and is more hydrophilic than PLA, which is partially hydrophobic due to its methyl groups. Moreover, PLA can be dissolved in the

Introducing biomaterials for tissue repair and regeneration13

tricarboxylic acid cycle while PGA is converted to metabolites or eliminated by other mechanisms, further explaining why PGA degrades faster than PLA. Hence, PLGA, a copolymer of PLA and PGA, may show intermediate degradation rates that can be modulated as a function of the relative fraction of hydrophobic/hydrophilic phases, crystallinity, and composition of chains (i.e., contents in l-LA and d-LA and/or GA units) [59]. Indeed, the amount of d-LA or mesolactide present in the l-PLA polymer changes the properties significantly in terms of melting temperature, crystallization rate and therefore processibility and properties of foams. For example, the higher the d-isomer content in the polymer the lower the crystallization rate and the lower the melting point. In this context, the formation of debris or other waste products due to the reaction of acidic degradation products may provoke adverse tissue reactions [60]. This is the result of the heterogeneous degradation of polyesters, which occurs faster inside than at the exterior by the competition of the next two phenomena: the easier diffusion of soluble oligomers from the surface into the external medium than from inside and the neutralization of carboxylic end groups located at the surface by the external buffer solution (in vitro or in vivo). The combination of these events contributes to reduce the acidity at the surface, but enhances the degradation rate by autocatalysis due to carboxylic end groups in the bulk [61]. In this context, inorganic compounds, that is, calcium phosphates or bioactive glasses, may be further incorporated to stabilize the environment conditions surrounding the polymer in order to control its degradation. This counteracting effect of the acidic degradation of biodegradable polymers explains the high interest for the use of composite materials in tissue repair and engineering (Table 1.2).

1.3 From nanostructured materials to the devices 1.3.1 Prostheses for tissue repair Nanostructuring of metal biomaterials is a topic widely covered in the scientific literature and is mainly applied to the specific area of dental, spinal, and orthopedic implants following these approaches: modification of surface topography and/or chemistry as well as biological functionalization. The final aim is to get fast and effective osseointegration or guided tissue regeneration in order to avoid the late failure of implants linked to the advancement of bone resorption causing mobility after a few years. The problem is of interest both in the dental implant field, in particular in patients with a severely resorbed ridge, and for the orthopedic implants [68]. The term osseointegration, coined in the late 1960s by Per-Ingvar Brånemark, is used to define the intimate bond between the bone and artificial implants, without any interlayer of connective tissue. This intimate connection is defined to be reached when the space and relative movements between the bone and implant do not exceed 100 μm. Osseointegration is also a measure of implant stability in the implant site and it is mainly related to secondary stability. While primary stability derives from mechanical interconnection soon after implantation, secondary stability is biological stability achieved

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Table 1.2  Summary of main natural and synthetic polymers used for tissue engineering Material

Chemical attributes

Physical and biological properties

References

Natural polymers Collagen and gelatin

Keratin

Silk

Multilevel chemical organization, righthanded triple helix, composed of three α-chains (Gly-X-Y) RGD motifs RGD and LDV motifs cysteine content

Disulfide bonds

High biocompatibility adhesion, proliferation and differentiation signal

[62]

Cell adhesion, antibacterial properties, high mechanical strength in vitro High mechanical properties in vitro, biocompatibility

[63]

Deficiency of toxicity, FDA approved; high mechanical properties Nontoxic— noninflammatory— FDA approved

[65]

Nontoxic FDA approved Permeability to small molecules and drug

[67]

[64]

Synthetic polymers PCL

Hydrophobic, long degradation times (over 2 years)

PLA

Less hydrophobic, intermediate degradation times (6 months) Biodegradable Great resistance against organic solvents Impact resistance

Polybutylene terephthalate (PBT)

[66]

through bone regeneration and remodeling. Primary stability is a prerequisite for secondary stability and depends on implant shape and implant bed preparation. Excessive implant movement after implantation (absence of primary stability) leads to the formation of fibrous connective tissue between the bone and the implant and excludes osseointegration. Primary stability gradually decreases during the bone remodeling process. The transition from primary to secondary stability is dictated by the advancement of the healing process according to a process similar to healing of a bone fracture. The biological processes at the base of osseointegration derive from the hemorrhage from the bone and soft tissues surrounding the implant and formation of hematoma. The implant immediately comes in contact with blood and this triggers a series of biological processes such as protein deposition, platelet activation and c­ oagulation,

Introducing biomaterials for tissue repair and regeneration15

inflammation and signaling, and lastly tissue formation. These processes represent the host’s response to injury and implant introduction and are influenced by the ­chemical-physical and topographical characteristics of the implant surface. Fast healing and new bone formation are due to the cooperative action of several phenomena: bioactivity of the surface and apatite precipitation on it (mineralization), protein adsorption, fast adhesion, and proliferation of the osteoblastic cells, high degree of cell differentiation, as well as absence of infections. The implant failure factors, which can be hidden osseointegration, and are not related to the patient’s sex and age, are: inappropriate choice of implant type (shape and material); incorrect seat for the system, without observing the proper contraindications; incorrect surgical technique, with intraoperative and postoperative consequences; no immobilization or infection after the operation; excessive load or insufficient load, both not favorable to osseointegration; inappropriate amount or quality of bone (e.g., osteoporosis); health complications in the patient and poor oral hygiene. The main symptoms that appear as a result of implant failure are: mobility (already during the first year), infection formation, compression pain, bone destruction, osteitis, sensitivity disorders, uncovered implant parts, and maxillary sinusitis. Despite the wide scientific research, a fast and physiological bone integration of the metallic surfaces remains a challenge; the main consolidated results are here briefly resumed. A smooth surface is usually less suitable for osseointegration, so special treatments have been developed to obtain structured surfaces on the macro, micro, and/or nanoscale; the most common are acid etching or sandblasting. On the other hand, a highly rough surface is much more susceptible to bacterial colonization which can easily lead to implant loss and fatigue resistance can significantly be reduced, so a compromise solution, as typically occurs in a complex system, must be found. In the dental surgery the great success of micro-textured surfaces has been accompanied by more frequent peri-­implantitis mainly because rough superficial implants are more prone to it. As far as surface topography is concerned, the dimension of the surface features can affect different aspects of cell behavior. In particular, the surface macro-topography (features >100 μm) affects primary fixation and the surface micro-topography (between 1 and 100 μm) influences cell recruitment, adhesion, orientation, and morphology and even gene expression. The submicron and nano features influence the focal contacts and cytoskeletal arrangement, the orientation of the cells and their communication, as well as adsorption and conformation of the proteins and biomolecules on the surface [69]. The cellular response to the nanoscale features is currently under investigation, however, it can be inferred that the cells effectively sense nanostructures, because they interact in vivo with the ECM proteins, which are characterized by nanometric collagen fibrils; moreover, both the filopodia and lamellipodia are structures on the nanoscale, as well as the integrins [70]. The chemical characteristics of a surface significantly affect cell behavior, too. Titanium surfaces presenting low surface contamination can significantly increase fibronectin adsorption and osteoblast differentiation [71]. Treatments able to enhance the formation of hydroxyl groups on the metal surface are of interest in order to stimulate the cell differentiation and adhesion [72]. Different coatings and surface treatments have been proposed in order to improve the bone-bonding ability of titanium

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Nanostructured Biomaterials for Regenerative Medicine

and its alloys. Mechanical, chemical, physical, and even biological surface modifications have been considered in the scientific literature [1]. Biological functionalization allows to greatly expand the action of the implant surfaces by the delivery of specific signals directly from the surface to the biological environment and to obtain fully multifunctional surfaces. The biomolecules can be coupled to the artificial materials mainly in three ways: adsorption, covalent grafting, and inclusion into a resorbable carrier. Various biomolecules can be grafted to the material surface, proteins from the ECM, growth factors, and drugs can be cited as typical examples in order to mimic the physiological environment, to improve cellular adhesion, and to perform localized therapy. Moreover, natural molecules (pure or as a mixture of natural extracts) are attracting increasing interest in the scientific community [73, 74]. Over the past two decades, the osteotropic biomolecules (such as the growth factors) are increasingly considered as a source of inspiration for the bio-inspired design of the active implant surfaces. They have been shown to stimulate the formation of bone tissue around the implants with promising results in enhancing bone formation. Overdoses, as well as systemic therapies, must be avoided, that is why their grafting on the surface, in a small and controlled amount, is of great interest, in order to get a limited local action. Finally, metal implants are often in contact with soft tissues, too. Concerning dental implants, this is one of the key current matters, because a stable gum seal around the implant collar is needed in order to avoid peri-implantitis (bacteria infiltration) and down-growing of the junctional epithelium: a sealing gum tissue around the implant can prevent series of events often leading to bone resorption and the removal of the implant. Currently, collars are usually made as smooth as possible in order to avoid biofilm formation, but different engineered surfaces are under investigation. In fact, in the presence of a dental implant (without a specific 3D geometry), the fibers of the connective tissue are oriented parallel to the implant; the consequence is unwanted down migration of the junctional epithelium and the loss of the crystal bone. Moreover, some areas of loose connective tissue (1–3 μm wide) are often developed around a dental implant, with a poor fluid recirculation and difficulties for the macrophages to reach the tissues. The introduction of surface topographies with a predominant direction, such as grooves, both at the microscale (depth/width of the grooves 1–90 μm) and nanoscale (depth/width 10−6 (e.g., 10−5), considering the individual situation, including the risk assessment undertaken by the manufacturer of the device. Applicability of this clause can be sought through discussion with regulatory bodies. Usually, nanomaterials are not sterile and can be contaminated at any stage of their production with biological species, for example, endotoxins (which is a common contaminant coming from the Gram-negative bacterial cell membrane). For this reason, there are Good Laboratory Practice and protocols to take the necessary precautions in order to avoid/reduce endotoxin contamination, especially for some NPs that had shown high binding properties to endotoxin throughout synthesis and processing [5]. Moreover, nanomaterials could be susceptible to being damaged by sterilization techniques, particularly when biological materials are involved. For example, sterilization by autoclaving could origin chemical and physical changes caused by heating in NPs with “low melting point” or that contain proteins/antibodies. Otherwise, sterilization by gamma irradiation could induce the formation of free radicals that

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could cause chemical changes plus impact the stability of the product (especially in polymeric materials). “Aseptic manufacturing” is an option but could be quite complicated in particular for a multiple step process. Consequently, it is essential to test several methods in order to find the specific sterilization to be performed for each particular nanomaterial, and an appropriate series of tests must be utilized during the development phase in order to understand the impact of sterilization on the nanomaterial. Sterilization remains a critical step for the use of nanomaterials in medical devices, and the effects of sterilization on the integrity of the physicochemical properties of nanomaterials (immediately and after a long time) still require to be investigated.

2.4 What is to be demonstrated 2.4.1 Safety In the medical device field, a safe device can be described as a device with a favorable risk-benefit profile. In the Reg. EU 27017/745, this requirement is described as the fact that devices shall not compromise the clinical condition or the safety of patients. Patients are not the only stakeholders that shall be taken into consideration as the Regulation also requires that manufacturers assess the risk-benefit profile as regards of the safety and health of users or, where applicable, other persons. The Regulation also describes what a “favorable” risk-benefit profile is by explaining that any risks which may be associated with the medical device use shall be acceptable when weighed against the benefits to the patient. This also means that any risks are compatible with a high level of protection of health and safety, taking into account the generally acknowledged state of the art. This poses a special concern to manufacturers of medical devices containing nanomaterials, as the potential of exposure of patients and also of users shall be taken into consideration by assessing the state-of-the-art knowledge.

2.4.1.1 Role of the in vitro and in vivo testing Safety is typically demonstrated by compliance to safety standards, such as, for example, standards for biocompatibility and for sterility. Safety is a characteristic that is therefore proven by testing significant samples so that the results can adequately describe the safety profile of the commercial product that is actually placed on the market. The selection of a significant test sample can be especially challenging when testing medical devices based on nanomaterials. The applicable standard ISO 10993-22 provides guidance on the sample preparation and selection, as it requires that nanomaterials should be characterized in the form they will be delivered to the end-user, which is a clear requirement that the final device (or a significant portion) is tested. If only a portion of the final device is tested, the manufacturer shall describe representative samples from the final device in terms of materials processed in the same manner as the final device in order to directly evaluate a nanomaterial component.

Regulatory landscape of nanostructured biomaterials for tissue regeneration 39

The outcome of in vitro and in vivo testing is assessed, according to compliance to standards and to comparison to state of the art, in order to describe the potential device risk.

2.4.2 Clinical performance and clinical benefit According to the Regulation EU 2017/745,10 the term “clinical performance” of a device means the ability of that device, resulting from any direct or indirect medical effects which stem from its technical or functional characteristics, including diagnostic characteristics, to achieve its intended purpose as claimed by the manufacturer, thereby leading to a clinical benefit for patients, when used as intended by the manufacturer; clinical performance is best described by technical, chemical, and physical characteristics of the device itself. On the other hand, the Regulation describes the “clinical benefit” as the positive impact of a device on the health of an individual, expressed in terms of a meaningful, measurable, patient-relevant clinical outcome(s), including outcome(s) related to diagnosis, or a positive impact on patient management or public health. For this reason, there is a measurable and predictable cause-effect relationship between the performance (how the device acts on the human body) and the benefit (how the human body positively responds). Thus, the building up of the data required to show compliance to Regulations follows a sequence of steps of increasing complexity: in vitro data, often obtained by cell culture testing, provide ethical and scientific “green-light” for the preclinical animal, in vivo testing. Feedback from animal testing, evaluated through the requirements of safety, performance, and benefit, as just described, provide in turn green light for clinical evaluation as further described in the next section.

2.4.2.1 Role of the in vivo and clinical testing Clinical performance and clinical benefit are usually described not only by means of compliance with the standard, but also by assessing the device behavior and its impact on a living tissue or organism. Testing is therefore also performed on applicable animal models and, when technically necessary and ethically acceptable, clinical performance and benefit are studied with the involvement of patients. For clinical testing, there is a clear requirement in MDR 2017/745 that the endpoints of the clinical investigation shall address the intended purpose, clinical benefits, performance, and safety of the device. The endpoints shall be determined and assessed using scientifically valid methodologies. The primary endpoint shall be appropriate to the device and clinically relevant. For nanomaterials, the challenge in the assessment of the clinical performance and subsequent clinical benefit is, therefore, linked to the challenge of determining appropriate endpoints.

10

Medical Device Regulation 2017/745 Article 2.

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Nanostructured Biomaterials for Regenerative Medicine

2.5 From the project to the product Research Institutes, innovative Start-Ups/SMEs and Large Industries have individual plans to convert new “nano ideas” into products… but they have the same challenge: the commercialization of nanostructured devices for tissue regeneration is much more complicated than other products. When new products/processes are developed, a strong effort must be made to ensure that they are healthful and environmentally safe. The toxicity of NPs is mainly given by their physical/chemical characteristics (size, shape, surface charge, etc.) and the presence/ absence of active groups on the surface. Therefore, there are several complex regulatory and approval processes (depending on under which different product categories fall the new “nano idea”) to follow for commercialization, but they have often some “gray area” where neither the notify body nor the manufacturer has a clear idea of what to do. The development and commercialization of a “nano-product” (the term include both product containing NPs and produced using nanotechnologies) require collaboration among many different actors, due to the complexity of the matter and requires a long time (the time to market is often “close” to 20  years). Moreover, costs for translating an idea to a commercial product are extremely high (in part due to the enormous expenses requested for the clinical trials); therefore, a very careful screening of innovative ideas has to be performed as soon as possible (in order to not vast money).

2.5.1 Ensuring a consistent level of quality 2.5.1.1 Manufacturing Essential conditions for the development of an industrial scale “nano-manufacturing” are the capability to preserve nanoscale properties while incorporating into products, an online metrology in order to allow effective measurements/quality control and the availability of high throughput. The challenge is maintaining the properties and quality of “nano-system” during high-rate/volume production as well as during the lifetime of the product after production. Nano-manufacturing consists in a combination of process steps that often need specific tools (like the computer-aided design and manufacturing—CAD/CAM—­ especially if hybrid 3D printing is involved) and a precise quality control. The latter is a very important point because even small variations in processing of NPs can result in variations of toxicity. Sukhanova et al., in a recent paper, highlighted as the NP toxicity strongly depends on their physical and chemical properties, such as the shape, size, electric charge, and chemical compositions of the core and shell [6]. Moreover, nano-manufacturing must be safe for the workers, so any possible detection and measurement of the type and amount of NPs in the workplace have to be performed.

2.5.1.2 Quality control Nanomaterials possess exceptional unique nanoscale “size-dependent” physical, optical, and chemical properties that are completely different than their bulk counterpart. Therefore, what must be controlled and in which way?

Regulatory landscape of nanostructured biomaterials for tissue regeneration 41

The big challenge is not only the development of new analytical tools that make a possible measurement at nano-to-microscale (including dispersion of NPs within a matrix and at interfaces), but that these tools have to work in real-time embedded in the production line and above all, they must be “nonintrusively” or destructively. In the meantime, new methods of calibration and standardization have to be developed in order to ensure the accurate interpretation of results. A typical flow chart of a generic manufacturing process, with quality control steps, is shown in Fig.  2.1.11 Starting from the raw materials, an acceptance test must be performed in order to inspect materials and verify their conformity. Tests could be different for each material and could include, for example, Physicochemical characterizations (particle size, crystallinity, water solubility, surface characteristics, etc.). Each manufacturing phase must be subject to continuous analysis and during the manufacturing process, a statistical sampling (screening) could be performed in order to supervise the process and the final product. The controls that could be performed are, for example, an “in-process” monitoring of equipment, process parameters, and semifinished products, the control of dispersions, surface roughness measurements (to certify surface treatments), the control of the bonding and nonbonding interactions at the interface of the matrix and the transition region and last but not least the control of the toxicity of NPs. Critical control points (and procedures to monitor them) could be also determined in order to identify hazards, but it is even more important to set up corrective actions to be taken when a deviation is identified. At the end of the process, a final lot acceptance test and life test must be performed in order to confirm the quality of the final product.

2.5.2 Postmarket activities: surveillance and vigilance Postmarket activities involve actions to be performed after a medical device was firstly put on the market. They can be divided into two categories: ●

●

Proactive—Postmarket Surveillance Reactive—Vigilance

Postmarket surveillance is defined as: “the proactive collection of information on quality, safety or performance of medical devices after they have been placed on the market.” (According to the Global Harmonization Task Force.) Vigilance is defined as: reporting of incidents that can occur with medical devices, when they do not perform as intended, thereby leading in the worst case to injury or death. The purpose of medical device vigilance is to protect the health and safety of persons; evaluate incidents to prevent recurrence; determine the effectiveness of corrective actions and preventive actions, and monitor and learn from experience. Activities involved with postmarket surveillance typically involve seeking customer feedback 11

Modified from a Windbond website figure: http://www.winbond.com/hq/about-winbond/quality-policy/ management/management.html?__locale=en.

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Nanostructured Biomaterials for Regenerative Medicine

relating to whether the organization has met customer requirements. This may be accomplished by some of the following methods: Focus groups Reviewing customer complaint information

●

●

The quality manager will yearly monitor the performance and safety of company products already on the market by observing, collating, and reviewing the following: Vigilance results review Per defect Per product Customer complaints Medical practitioner feedback Independent research Trade journals

●

●

●

●

●

●

●

Vigilance has to do with Incidents. Examples of Incidents are: – Device malfunction or deterioration that occurs while the device is used as intended by the manufacturer. – Unanticipated adverse reaction or unanticipated side effect that is not stated in the information provided by the manufacturer. – Interactions with other substances or products. – Degradation/destruction of the device (e.g., fire). – Inappropriate therapy. – An inaccuracy, omission, or deficiency in the labeling, instructions for use and/or promotional materials, except information that is part of the general knowledge of the intended user.











Vigilance results can be defined using criteria such as: – – – –









number of incidents/field corrective actions in the year; number of incidents/field corrective actions per defect; number of incidents/field corrective actions per product; number of incidents/field corrective actions after complaints.

Results are compared with information from previous years. Occurrence of incidents should be notified to the relevant National Competent Authority.

2.6 Roles and responsibilities 2.6.1 Manufacturers According to the MDR 2017/745,3 a manufacturer is defined as: “A natural or legal person who manufactures or fully refurbishes a device or has a device designed, manufactured, or fully refurbished, and markets that device under its name or trademark.” The obligations of the manufacturers of medical devices (including the ones containing NPs) are, therefore, to demonstrate the conformity of their products with the legal requirements.

Regulatory landscape of nanostructured biomaterials for tissue regeneration 43

But it is also of extreme importance for manufacturers to be clearly aware of the supply chain and the impact of manufacturing activities. This is because in the global environment, nowadays, the use of contract manufacturing is often unavoidable in order to reduce costs and supply line issues. There are two documents prepared by the European Association of Notified Bodies for Medical Device that could help understanding the EU regulations on this matter: NB-MED/2.5.2/Rec1 “Subcontracting—QS related” and NB-MED/2.15/Rec1 “Voluntary Certification at an Intermediate Stage of Manufacture.”12 Moreover, manufacturers have to apply a postmarket surveillance system in order to constantly monitor the performance of their products and report every adverse incident (vigilance report) and field safety corrective actions (FSCAs) to EU Competent Authorities. An FSCA is an action taken by a manufacturer to report any technical or medical reason leading to a recall of devices and the new Medical Device Regulation addresses specific timelines for FSCA reports.

2.6.2 Competent authorities In Europe Competent Authorities, nominated by each government, are usually the Ministries of Health of each member State (instead, in the United States there is only one competent authority: the FDA). Therefore, each European country has its own Competent Authority in charge of market surveillance and designating and monitoring the independent Conformity Assessment Bodies (Notified Bodies). European National Competent Authorities for Medical Devices have created a group (CAMD) in order to enhance the level of collaborative work, communication, and surveillance of medical devices across Europe.

2.6.3 Notified bodies A notified body is an organization designated by an EU country to assess the conformity of certain products, including medical devices, before being placed on the market. These bodies carry out tasks related to conformity assessment procedures set out in the current Medical Device Directive and will from 2020 also carry out tasks related to conformity assessment procedures set out in the new upcoming regulation. The conformity assessment procedure for class I devices is being carried out, as a general rule, under the sole responsibility of manufacturers in view of the low level of vulnerability associated with such devices. For class IIa, class IIb, and class III devices, an appropriate level of involvement of a notified body is compulsory. Since devices containing nanomaterials are classified, in the Regulation 2017/745, at least in class IIa but more frequently in classes IIb and III, the intervention of a Notified Body in the marking procedure shall be expected. As part of their involvement in the certification process, notified bodies audit the technical documentation and the manufacturing, control, and distribution processes 12

http://www.team-nb.org/nb-med-documents/.

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Nanostructured Biomaterials for Regenerative Medicine

Process

Quality control Method Storage

Finished product quality confirmation

Reliability confirmation Life test and environmental test

Sampling inspection Lot acceptance test

Manufacturing

Screening

Equipment, submaterial environment, cleanliness discipline—Monitoring Process parameters, semifinished product inspection—Monitoring

Testing and visual inspection

Appearance and characteristics 100% inspection

Raw materials

Raw materials quality conformance

Inspection on raw materials Acceptance test

Fig. 2.1  Example of a quality control flowchart.

and have the right and duty to carry out both announced and unannounced on-site audits and to conduct physical or laboratory tests on devices to ensure continuous compliance by manufacturers.

2.7 Relationship between regulatory requirements and intellectual property The Intellectual Property of NPs (or of a product containing them) and the relationship between regulatory requirements is a “mine-laden” that will become of strategic importance in the next future owing to the increasing numbers of innovative products/ technologies involving NPs. There are concomitant reasons that brought to these very difficult circumstances… First of all, patent laws usually exclude the possibility to patent methods for the treatment of the human body using surgery and therapy but allow the patenting of products used in these procedures. But this could be subverted by the patenting of NPs because “at the end” the patent could cover also the way of their utilization that could be considered as “a method.” Moreover, NPs, and medical products containing them have to comply with strict regulations in order to enter in a commercial market. In vitro, in vivo tests, and clinical trials must be performed following standardized procedures for a given period of time,

Regulatory landscape of nanostructured biomaterials for tissue regeneration 45

but a medical device could be exempted from regulation if a similar product is already on the market. This allows shortening incredibly the time-to-market and the costs. On the contrary, the same MD is often recognized as “new” from the Intellectual Property Rights viewpoint and therefore it is protected by patent laws. These two “way of seeing the matter” could create several issues, because if there is a patent protecting the MD, this means that it is innovative/new and therefore, it must be validated following regulations and vice versa. Sure enough, if an MD containing NPs is considered “new” this could greatly impact the time and cost required for its commercialization, especially after the entering into force, on May 25, 2017, of two new MD regulations”13 that has led to longer time periods for obtaining marketing approval. As a result, MD manufacturers should seek compensation for their loss of patent protection by requesting a Supplementary Protection Certificate (SPC = extension of the patent term up to 5 years). But the SPC regulation is one of the most controversial areas of European IP law because SPC can be obtained only for products that have been authorized for the first time as medicinal or plant protection products in the European Union. That means in Europe, formulation and MD patents cannot be the subject of SPCs. However, several MDs often include active ingredients (combinations or borderline MD), thus, they may in principle be eligible for SPC protection, but it could generate a lot of problems the fact that a single borderline product may be classified as an MD or an MP depending on the approach of each different Member States. It is, therefore, possible to imagine a scenario where a particular product is not only authorized differently as an MD or an MP throughout Europe but also that the accessibility of asking for an SPC in those States where it has been authorized as an MD may be different among EU members. During the last years, several requests of SPC were evaluated by different EU member’s patent courts but only in The Netherlands and Germany some cases were accepted. Totally different is the situation in the United States where the US law allows obtaining patent term extensions (PTE) for Class III (high-risk) MD and furthermore gives the possibility to extend a patent in case of unreasonable delays during the examination at the US Patent and Trademark Office [the so-called patent term adjustments (PTAs)]. In order to avoid unequal treatments on “this site of the Ocean,” the hope is that also in Europe it could be recognized that MD Cass III are facing similar levels of regulation to pharmaceutical products because nowadays they are subject to high standards in order to obtain the required conformity assessment certificate, and therefore SPC could be requested. Actually, the European Commission is aware of this matter and is looking into potential reform of the SPC system (following the outcomes of consultation done in 2017). 13

https://eur-lex.europa.eu/legal-content/EN/TXT/PDF/?uri=CELEX:32017R0745 https://eur-lex.europa.eu/legal-content/EN/TXT/?uri=CELEX:32017R0746

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References [1] M. Racchi, S. Govoni, A. Lucchelli, L. Capone, E. Giovagnoni, Insights into the definition of terms in European medical device regulation. Expert Rev. Med. Devices 13 (10) (2016) 907–917, https://doi.org/10.1080/17434440.2016.1224644. [2] Regulation EU 2017/745, whereas (15) and Annex I Essential requirement 10.5 [3] IMDRF Definition and regulation in terms of mechanism of action and intended use IMDRF-3 /20 March 2013/. [4] MEDDEV 2.1/3 rev.3 Borderline products, drug-delivery products and medical devices incorporating, as integral part, an ancillary medicinal substance or an ancillary human blood derivative [5] R. Darkow, T. Groth, W. Albrecht, K. Lutzow, D. Paul, Functionalized nanoparticles for endotoxin binding in aqueous solutions, Biomaterials 20 (1999) 1277–1283. [6] A. Sukhanova, et al., Dependence of nanoparticle toxicity on their physical and chemical properties, Nanoscale Res. Lett. 13 (2018) 44.

Nanostructured biomaterials for regenerative medicine: Clinical perspectives

3

Bin Zhanga, Jie Huanga, Roger Narayanb a Department of Mechanical Engineering, University College London, London, United Kingdom, bUNC/NCSU Joint Department of Biomedical Engineering, Raleigh, NC, United States

3.1 Introduction Regenerative medicine holds great promise for developing biological substitutes, which can replace or regenerate human tissues [1]. Mimicking native tissue environment using bioengineered scaffolds, cells, and growth factors is a promising strategy to overcome the limitation of tissue volume supply [2, 3], the risk of immune reaction and disease transmission issue [4–6] in autografts and allografts clinical operations. Biomimetic tissue including hard tissue, mainly bone, and soft tissues, such as cartilage, blood vessels, skin, tendon, muscle, and heart have been successfully engineered and implanted in vivo [7–13]. Those bioengineered tissues serve as a template to guide cell organization and growth and allow diffusion of nutrients to transplanted cells, which ultimately leads to the generation of complex architectures which mimic native tissues. Recent advances in bioengineered tissue mimic the characteristics of natural tissue at the nanometre scale, have shown the benefits for cell attachment, proliferation, differentiation, and matrix deposition in vitro and induction in vivo [14]. Since cells dynamically interact with their local environment at the nanoscale [15], it is necessary to control the properties of engineered tissues at nanoscale lengths for more analogous to natural tissues. Some studies have focused on (a) use of nanostructured materials (e.g., incorporation of nanoparticles in a polymer matrix to mimic the nanocomposite architecture of natural tissue) and (b) manipulation of the mechanical properties (e.g., Young’s modulus and strength) of the scaffolds. Also, nanostructured biomaterials can decrease inflammatory response and increase wound healing in comparison to conventional biomaterials, possibly due to their high surface energy affecting protein adsorption and cell adhesion [16, 17]. However, despite the significance of nanoscale material on a cellular response, how nanostructure properties (e.g., mechanical property, surface energy, and surface roughness) regulate the behaviors of cells in vitro or in vivo are yet to be clearly understood. In this review, the role of nanostructured ceramics, metals, polymers, and their nanocomposites in tissue regeneration and regenerative medicine is considered. Also, the clinically approved nanostructured products for regenerative tissue applications are summarized. Nanostructured Biomaterials for Regenerative Medicine. https://doi.org/10.1016/B978-0-08-102594-9.00003-6 © 2020 Elsevier Ltd. All rights reserved.

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3.2 Ceramic nanostructures There is a high degree of demand for biomaterials to repair large defects of hard tissue (i.e., bone and tooth) [18]. Bone tissue has self-healing functions. However, it is difficult to restore a large bone defect size beyond 2–2.5 times of the affected bone diameter without using a graft or scaffold [19]. The bone tissue comprises an organic and inorganic compound. On the other hand, bone is mainly made up of the inorganic mineral crystallites hydroxyapatite (HA) at nano-sized scale [20, 21]. These crystallites are deposited and arranged in parallel to the collagen fibers thus providing the rigidity and stiffness to the hard tissue [22]. Fig. 3.1 shows the macrostructure, microstructure, and nanostructure of bone tissue. The ratio of organic to inorganic phases in bone tissue affects the mechanical property [23]. The collagen fibers retain similar characteristics to polymers, which give bone tissue the flexural resilience and toughness by decreasing the brittleness of the mineralized phase HA. Conversely, the mineral crystallites HA provides bone tissue the mechanical strength and stiffness.

3.2.1 Nano-hydroxyapatite Calcium phosphates, especially hydroxyapatites [HA, Ca10(PO4)6(OH)2] has already been successfully used as bone graft due to its similarity with the mineral crystal of bone. Robinson [24] reported an average crystal size in human bone tissue are roughly 50 nm long, 25 nm wide, and 2–5 nm, thick. These HA nanoparticles constitute approximately 70% of the native bone. Nano-sized hydroxyapatite (nano-HA) in the rod

Osteocytes

Osteon: 100–500 µm

Blood vessels

Collagen fibril

(A)

Macrostructure

(B)

Collagen fiber

Nano hydroxyapatite

Collagen molecule

g Microstructure

(C)

Nanostructure

Fig. 3.1  The macro, micro, and nanostructure of the bone tissue. (A) At the macrostructure level, bone consists of cortical bone and cancellous bone. (B) At the microstructure level, there are many repeated osteon units in cortical bone. In the osteons, the blood vessels and nerves are surrounded by concentric layers of collagen fibers. (C) Collagen fibers consist of repeating individual collagen fibrils (30–300 nm), and the collagen fibril is composed of a single collagen molecular species. Nanohydroxyapatite distribute in the collagen fibers and increase the stiffness of the bone. This figure is modified based on T. Gong, J. Xie, J. Liao, T. Zhang, S. Lin, Y. Lin. Nanomaterials and bone regeneration. Bone Res. 3 (2015) 15029.

Nanostructured biomaterials for regenerative medicine: Clinical perspectives49

or plate shape is commonly used in bioengineered tissue to mimic bone minerals, as shown in Fig. 3.2. Since Levitt et al. [26] developed apatite bioceramics and suggested the possibility of medical applications, HA has been widely used for clinical applications. HA material in dense or porous forms have been clinically used for alveolar ridge augmentation [27], the increment of spine fusion and repair of bone defects [28–30]. Many studies indicate that involving nano-sized HA could resemble bone minerals and better osteoconductivity would be achieved. Since engineered nano-architecture features a high surface area to volume ratio, it can systematically expose cells to multiple biological components with different functionalities [31]. Webster et al. [17] have shown the enhanced osteoblast adhesion and protein adsorption in vitro on the nano-size HA (less than 100 nm) compared to the traditional

Fig. 3.2  Transmission electron microscopy of (A) plate-shaped nano-HA in clusters; (B) singular plate-shaped nano-HA (both side lengths are 20 and 50 nm, separately); (C) roundshaped nano-HA in clusters; (D) singular round-shaped nano-HA (diameter is approximately 20 nm) [25].

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Nanostructured Biomaterials for Regenerative Medicine

micro-size HA. Liao et al. [32] indicated nano-sized HA influence the formation of adsorbed vitronectin which can subsequently mediate osteoblast adhesion in  vitro. Goené et  al. [33] investigated the influence of nano-HA treatment to the titanium implant surface with a clinical trial, which has shown that there was an increase in the extent of bone development in human maxillae after 4 weeks of healing. Although there is the ongoing discussion about the safety concern associated with nano-sized HA as it is likely that nanoparticles will transport across cell membranes to interact with DNA and RNA, it is not clear yet that the health consequences [34]. Scheel and Hermann [35] indicated there is no clear evidence that nano-sized HA is toxic. Furthermore, Kalia et al. [25] compared to the plate and round shapes of nano-HA in vitro, which demonstrated that the round shape nano-HA had relative higher alkaline phosphatase activity and matrix vesicle release than the plate-shaped nano-HA. The results indicated that the shape of nano-HA could be related to osteoblast viability and activity. Jang et al. [36] assembled a nanochannel network with nano-HA material, which provided both sufficient mechanical strength and efficient nutrient supply for bone cell growth and differentiation in  vitro. Further, in  vivo/clinical trials are required to establish the efficacy of the structure of nano-HA. Regardless of HA nanoparticles, another inorganic phase of human bone tissue is whitlockite nanocrystallites [WH: Ca18Mg2(HPO4)2(PO4)12]. These two crystals are distributed in different ratios depending on the regions of bone tissue, implying that HA and WH nanoparticles have distinguished biological roles [37]. Jang et al. [38] synthesized HA and WH nanoparticles, to mimic the inorganic composition of bone. The synthesis substance had the enhanced proliferation of bone cells and induced rapid regeneration of bone tissues in vitro and in vivo, compared with pure HA nanoparticles [39]. Thus, controlling HA and whitlockite nanoparticles composition distribution at the nanoscale is important for mimicking native bone tissue.

3.2.2 Nano-bioactive glass In addition to being osteoconductive and biocompatible, bioactive glass has also been investigated for use as a bone substitute. The 45S5 bioactive glass composition was invented by Hench et al. [40], which could release active biological ions and sequentially bond to living bone tissue. Merwin [41] induced the first application of 45S5 bioactive glass to replace the small bones in the middle ear for curing conductive hearing loss. Since then, 45S5 bioactive glass has the approval of the US Food and Drug Administration (FDA) and is successfully implanted in thousands of patients, mainly in bone tissue engineering application [42]. Bioactive glasses have been used as a coating in biomedical devices, dental fillers, tissue engineering scaffolds, and drug-delivery system [43–46]. Also, various bioactive glass compositions have been proposed, which contain no sodium or have additional elements incorporated in the network of silicate structure. For instance, the incorporation of silver [47] and zinc [48] in the silicate network, have been investigated to develop antibacterial materials. Waltimo et al. [49] have shown the stronger antimicrobial effect of nanoparticles of 45S5 (20–60 nm) than micro-sized particles by killing more Enterococcus faecalis. One

Nanostructured biomaterials for regenerative medicine: Clinical perspectives51

of the possible reasons is the reduction in size to nanometer scale can increase the surface area of nanoscale bioactive glasses, which allows a faster silica release and solution pH elevation. There are also other advantages of reducing the particles into nanosize, such as produce thin bioactive coating and reinforce polymeric nanofibers [16]. Bioactive glass exhibit potential benefit in comparison to HA regarding bioactivity. For instance, Wheeler et  al. [50] shown that the dissolution products from bioactive glasses upregulate the expression of genes that control osteogenesis in  vivo. However, the clinical application of bioactive glass is still lag than HA regarding commercial success. The possible reason can find from the literature that they generally performed poorly regarding mechanical strength and fracture toughness, therefore, not suitable for all grafting applications. Further, the development of apatite-wollastonite ­glass-ceramic was shown to possess a higher mechanical strength but lower fracture toughness than human cortical bone [51]. They could not involve cyclic loads with the host bone. Thus, the development of tougher scaffolds is required that still have all the bioactive properties of bioglass 45S5. The possible engineering solution is the development of composite materials [52], for example, incorporation of bioglass nanoparticles with a polymer, as discussed in detail further.

3.3 Polymeric nanostructures It is well known that native extracellular matrix (ECM) is the nanoscale dimensions of the physical structure, which is mainly composed of collagen fibers, between 50 and 200 nm in diameter [53]. Collagen fibers would regulate cell attachment, proliferation, and differentiation, and thus the design of engineered tissue scaffold should close mimic this structure. The polymeric nanofibers nonwoven scaffold is among the most promising biomaterials for the native fibrillar ECM. There are various polymeric materials have been nanofibrous, which mainly classified as natural materials, such as collagen, silk, alginate and chitosan, and synthetic polymers, such as poly (lactic acid) (PLA), poly (ethylene oxide) (PEO), and poly (ɛ-caprolactone) (PCL).

3.3.1 Fabrication of polymeric nanostructures Various methods have been applied to fabricate polymeric nanostructures, such as electrospinning, phase separation, and molecular self-assembly. The principle of electrospinning is applied an electrical field between the metallic nozzle and collector, to draw a thin liquid polymer solution. The fabricated polymer fibers can be produced in a range of nanometers to micrometers [54]. A variety of polymeric biomaterials, such as PEO, PLA, PCL, silk, chitosan, collagen, and alginate have been applied to form polymeric nanofibers using electrospinning [55–61]. In general, the minimum diameter of 9 nm fibers can be achieved using electrospinning, which can mimic the nanostructure of natural collagen fibers [59]. However, it is difficult to create 3D porous scaffolds using electrospinning. Moroni et al. [62] have applied electrospinning to combine with 3D fiber deposition to fabricate 3D scaffolds, but further investigation is required to optimize the scaffold structure.

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Phase separation, as another method, is frequently applied to fabricate 3D engineered tissue scaffolds. The porous polymer scaffolds can be manufactured by removing the solvent from the polymer solutions through free-drying [63]. Polymer scaffolds fabricated by the method of phase separation, have a sponge-like porous morphology. The micro or nanoscale polymer fibers can be obtained by controlling the solvent, polymer concentration, gelation time, and temperature. Yang et al. [64] fabricated poly(l-lactic acid) (PLLA) nanostructured porous scaffold using phase separation methods, with the minimum nanofibers diameter as 50 nm, which is comparable to those collagen fibers in the native ECM. The nanostructured fibers can also be fabricated by molecular self-assembly which is mediated by noncovalent bonds (e.g., hydrogen bonding as well as electrostatic, hydrophobic, and van der Waals interactions) [65, 66]. The fiber diameter can be reached less than 6 nm by molecular self-assembly, which is smaller than those fabricated using electrospinning and phase separation [67]. However, the mechanical properties of self-assembled scaffolds should be improved before using in tissue engineering application [68]. As Fig. 3.3 shown, cell spread as two-dimensional (2D) structure when seeding on microporous or microfibrous scaffolds. While, when cell cultured on nanostructured scaffolds, cells can make higher zone contacts which would lead to improved cell attachment and cellular interactions. The possible reason is that the higher surface area to volume ratio for the adsorption of proteins and binding of ligands [69]. Micropore scaffold

Microfiber scaffold

Cell binding

Scaffold architecture

Nanofiber scaffold

(A)

(B)

(C)

Fig. 3.3  Cells flatten and spread to microporous (A) or microfibrous scaffold (B), while cell exhibited more binding sites to nanofiber scaffold (C) [69].

Nanostructured biomaterials for regenerative medicine: Clinical perspectives53

3.3.2 Functionalized polymeric nanostructures 3.3.2.1 Controlling structures of polymeric nanostructure Cells are regularly oriented in native tissues, which is vital for tissue function. Thus, it is essential to control the orientation of the cell in tissue engineered scaffolds to mimic native tissue. In general, nanofibers fabricated on a static platform using electrospinning have a randomly oriented nonwoven fiber matrix. However, if the nanofibers are collected with a rotating collector, the aligned nanofiber matrix can be produced [70]. As shown in Fig.  3.4, Teixeira et  al. [71] found that human corneal epithelial cells could align and elongate along with the nanostructure substrate. Thus, the application of aligned polymer nanofibers can control cell orientation. Table 3.1 selectively summarized the functional nanopolymers for the application of tissue regeneration in vitro and in vivo. Those functional nanomaterials have successfully generated similar or even better tissue functions to stimulate cells to repair tissues. For instance, Dalby et al. [72] used nanoscale polymethylmethacrylate (PMMA) to stimulate stem cells toward osteogenic differentiation in  vitro without osteogenic supplements; these cells showed similar levels of the bone mineral to those of cells cultured with osteogenic media. However, compared with cells treated with osteogenic media, the topographically treated stem cells have a distinct osteogenic differentiation. McMurray et  al. [84] indicated that nanoscale surface topographies could determine cell fate and functions. Nanoscale square lattice symmetry patterns can promote the growth of stem cells and the retention of multipotency in vitro. Similarly, Badrossamay et al. [76] successfully controlled heart tissue constructs in vivo to mimic the ECM structure of myocardial tissue and the induced alignment of rat v­ entricular

(A)

(B)

10 µm 500 nm

Fig. 3.4  SEM images of cells cultured on the nanostructured substrate. (A) Human corneal epithelial cells aligned along the patterns of nanotopography. (B) Filopodia extend along the patterns of grooves and ridges with nano feature dimensions. This figure is modified based on A.I. Teixeira, G.A., Abrams, P.J. Bertics, C.J. Murphy, P.F. Nealey. Epithelial contact guidance on well-defined micro-and nanostructured substrates. J. Cell Sci. 116 (2003) 1881–1892.

54

Table 3.1  Functional nanopolymers for use in tissue regeneration Nanopolymers

Functionality

Results obtained

Bone

Nanofibrous PMMA with the various surface pattern

Modulate stem cell differentiation

Biodegradable poly(lactideco-glycolide) nanoparticles with osteogenesis related growth factors Peptide amphiphilic nanofibers functionalized with glycosaminoglycans (GAG) molecules Glycopeptide nanofibers selfassembled supramolecular GAGs

Nanoparticles could be as delivery carriers for growth factors

A topographically treated stem cell has a distinct osteogenesis differentiation The hMSCs transfection of PLGA nanoparticles significantly enhanced osteogenesis Enhanced aggregation of stem cells and deposition of cartilage-specific matrix elements Induced chondrogenic differentiation of MSCs and enhanced formation of hyaline-like cartilage Induced alignment of rat ventricular myocytes along with the nanofibers Improved vascularization of the scaffold with upregulation of gene expression related to ECM remodeling, after implanted

Cartilage

Heart

Highly aligned hybrid polymer-protein nanofiber Nanofibrous collagen scaffold

Mimicking composition, structure moreover, the function of the ECM, and chondrogenic differentiation Mimicking bioactive functions of natural cartilage tissue Controlled nanoscale surface topographies mimicking the function of myocardial tissue Mimicking composition of myocardial connective stroma and delivery of cardiomyocytes

Experimental types

References

In vitro

[72]

In vitro and in vivo

[73]

In vitro

[74]

In vitro and in vivo

[75]

In vivo

[76]

In vitro and in vivo

[8]

Nanostructured Biomaterials for Regenerative Medicine

Tissue

Tendon Skin

Vessel

Mimicking composition and structure of skeletal muscle basal lamina

Tenascin-C was incorporated into self-assembling nanofibers formed by peptide amphiphiles (PAs) Laminin mimetic peptide nanofibers

Facilitating the regeneration of skeletal muscle

Electrospun aligned poly-llactic acid nanofibers 3D PCL/collagen multilayered nanofibrous scaffold Multilayer nanofilm composed of hyaluronic acid and poly-l-lysine

Mimicking composition and structure of tendon tissue Mimicking composition of skin tissue with multiple types of cells Mimicking epidermal-dermal composition and structure of skin at the nanometre scale

Ar plasma-treated nanostructured surface of polymers Nanopores in the vessel wall mimicking a vascular bed

The surface roughness of the polymers changes with the plasma treatment Mimicking composition of vessel tissue

Enhanced cellular gene expression related to the skeletal muscle-specific marker Promoted satellite cell activation, accelerated myofibrillar regeneration, and reduced the time for muscle tissue repairing Upregulated tendon-specific genes Produced skin tissues with bilayer-epidermal and dermal layers Promoted adhesion of adhesion of keratinocytes, enhancing epidermal protective barrier function of the skin The increase of vascular smooth muscle cells cell adhesion Enhanced permeability and intercellular interference

In vitro

[77]

In vivo

[78]

In vitro and in vivo In vivo

[79] [80]

In vitro

[81]

In vitro

[82]

In vitro and in vivo

[83]

Nanostructured biomaterials for regenerative medicine: Clinical perspectives55

Muscle

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Nanostructured Biomaterials for Regenerative Medicine

myocytes along with the nanofiber. Thus, nanoscale structural cues can further control the function of tissue constructs. Aligned nanofibers are especially useful in guiding cellular orientation to mimic the anisotropy of natural tissues, including heart, nerve, tendon, skin, and blood vessels. Yin et al. [79] seeded human tendon progenitor cells on aligned PLLA nanofibers that recapitulated parallel collagen fibers in the tendon. These cells expressed a higher level of tendon specific genes compared with cells grown on random fibers in  vitro. Also, in vivo experiments, there were spindle-shaped cells formed on the aligned nanofibers and tendon-like tissue. Monteiro et al. [81] developed multilayer nanofilm composed of hyaluronic acid and poly-l-lysine on top of a hyaluronic acid scaffold by using layer-by-layer assembly for mimicking epidermal-dermal composition and structure of the skin at the nanometre scale. The results showed the promoting adhesion of keratinocytes, enhancing the epidermal protective barrier function of the skin. Further in vivo experiments are needed to establish an effective combination of hierarchical structures for the multifunction for tissue regeneration. Nanochannels in natural tissues are also vital for maintaining the activity of cells, as they provide transport paths for oxygen and nutrients [85]. Zhang et al. [83] incorporated nanochannel in vascularized cardiac tissue constructs and bone scaffolds in vivo. They used self-assembled and porogen methods to enhance permeability and permit cellular crosstalk while maintaining mechanical properties. Core-shell nanofiber structure can be fabricated using modified electrospinning technology. This technique uses a designed nozzle containing a core opening and a surrounding annular opening. This technique is commonly applied to embed growth factors [86] into the core of biodegradable polymer nanofibers, producing polymer nanofibers able to release growth factor. Also, the core-shell structure can be used to form nanofibers with a natural polymer as the shell material and a synthetic polymer as the core material, which would improve the mechanical strength [87]. He et al. [88] fabricated PCL/zein core/shell nanofibers, and the results indicated the encapsulation of natural zein resulted in enhanced cell adhesion and proliferation in  vitro. Thus, the core-shell structure approach also solves the problem of poor biocompatibility of synthetic polymer nanofibers, as the biocompatibility of synthetic polymer is not as good as the natural polymer. More in vivo experiments are necessary to investigate the advantages of mechanical strength and biocompatibility in the core-shell nanofiber structures.

3.3.2.2 Surface modification of polymeric nanostructures Various synthetic biodegradable polymers have been utilized as tissue engineered scaffolds; however, one of the disadvantages is lacking biological recognition on their surface [89]. Thus, it is possible to promote cell-scaffold interaction by modifying the scaffold surface to obtain the desired characteristics. Common techniques for surface modification are plasma treatment [82, 90], laser treatment [91, 92], nanoparticle modification [93–95], and self-assembly [96]. Plasma treatment leads to changes in surface roughness of the polymers, which could affect the biological response. Reznickova et al. [82] applied Ar plasma-treated

Nanostructured biomaterials for regenerative medicine: Clinical perspectives57

low-density polyethylene, high-density polyethylene, and ultra-high-molecular-weight polyethylene surfaces with different nanomorphologies. The in  vitro results proved there was a significant increase of vascular smooth muscle cells cell adhesion, compared with the untreated polymers. Laser treatment can generate periodic surface nanostructures. Rebollar et al. [92] modified the surface of polystyrene foils by laser, and the results indicate that the presence of nanostructures on the surface can guide cell alignment. Also, the changes in surface chemistry and morphology have a positive influence on human embryonic kidney cells proliferation, compared with the untreated one. The chemical bonds are broken by laser radiation, which leads to highly reactive radicals reacting with the surrounding atmosphere on the surface. This approach leads to forming new functional groups, such as the ­oxygen-containing group, and the amino group [91]. The presence of amino acid and oxygen functional groups has been proved directly proportional to the spreading and adhesion of seeded cells [97]. However, plasma and laser treatment mostly focuses on 2D film surfaces, and it is difficult for the modification of 3D scaffold surface. Liu et al. [96] developed electrostatic layer-by-layer self-assembly technique to modify 3D nanofibrous PLLA scaffolds with gelatin. Cell proliferation was distributed more effectively and uniformly throughout the surface-modified PLLA nanofibers scaffolds, compare to the control scaffold. Another method for modifying the polymer surface is the incorporation of ceramic [94], noble metal nanoparticles [95, 98], and carbon-based nanotubes [93] on top of the scaffold surface. The electrospinning techniques could be applied to modify the scaffold surface, but it does not suit for 3D scaffold surface [99]. To obtain a uniform apatite layer on the 3D scaffold surface for bonding with living tissue, the polymer scaffold can be soaked in simulated body fluid (SBF) to allow apatite crystals to grow onto the scaffold surface [100].

3.3.2.3 Bulk modification of polymeric nanostructures Regardless of the surface modification of polymeric nanostructures, the ceramic nanoparticles can be directly mixed into polymers during processing. Thus, the incorporated ceramic nano-sized particles remain locked inside the polymeric scaffold structure. By bulk modification of polymeric nanostructures, the mechanical strength of nanocomposite could improve compared with the pure polymers. However, most of the ceramic particle is within the polymer scaffold, rather than on top of the surface. The cell-scaffold interaction should occur at the scaffold surface. Thus, the no exposed ceramic nanoparticles is a waste [68]. Moreover, the polymer can be functionalized by mixing with growth factors, and the biocompatibility of polymer can be improved [101, 102]. Although the presence of growth factors is of crucial importance to trigger healing for tissue regeneration, one of the limitations to growth factor therapy is insufficient local retention and require large quantities due to the fast inactivation of growth factors [103]. Nanoscale delivery systems have attracted a great deal of attention by researchers in the field of regenerative medicine based on their unique features, such as high s­ urface

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area and easiness of surface functionalization, which can promote the adsorption of growth factors or relevant gene materials [104–106]. For instance, Kong et  al. [107] incorporated chitosan nanospheres into mineralized collagen coatings in vivo to improve rhBMP-2 loading and morphogen releasing based on the good affinity of chitosan for proteins and the large surface area of nanospheres. Also, Kumar et al. [108] developed an injectable, self-assembled peptide-based nanofibrous hydrogel that contains peptides for proangiogenic moieties, which can rapidly form mature vascular networks and induce tissue integration after subcutaneous delivery in vivo via a syringe needle. Also, Park et al. [73] used biodegradable PLGA nanoparticles as carriers loaded with runt-related transcription factor 2 (RUNX2) protein and coated with bone morphogenetic protein 2 (BMP2) plasmid DNA (pDNA), delivery to hMSCs in vitro and in vivo. The hMSCs transfection of PLGA nanoparticles significantly enhanced osteogenesis. These examples proved that nanoparticles can be exploited as delivery carriers for growth factors.

3.4 Metal nanostructures Since Faraday [109] firstly identified metallic nanoparticles in an aqueous solution, noble metal nanoparticles have been extensively studied as there is a significant difference with its bulk counterparts in the respects of physical, chemical, and biological properties [110]. The unique characteristics are high energy atoms located on the particle surface area [111], a high ratio of surface area to volume, high surface energy, and electron storage capacity [112]. Silver (Ag) and gold (Au) represent pure metallic nanoparticles, while iron oxide nanoparticles (IONPs), titanium oxide (TiO2), and zinc oxide (ZnO) are metal oxide nanoparticles, which have been used for tissue engineering [113–119]. These applications can be achieved by directly adding nanoparticles into culture media, coating or incorporating with other materials as composites. For instance, silver nanoparticles show antimicrobial activity while gold nanoparticles show good biocompatibility and rarely induce an allergic response. They can be seen as good candidates for various biomedical applications. The nanoparticle size, shape, stiffness, and surface property are essential for internalization into the cells. Ko et al. [117] indicated that smaller size gold nanoparticles (30–50 nm) have higher uptake and more cellular internalization, compared to bigger size nanoparticles (50–200 nm). Zhang et al. [119] indicated that nanoparticles possessing a positive charge could avoid lysosomal degradation, compared with the negative or neutral charges. Thus, the nanoparticle surface charge is related to cellular internalization rates. Zhang et al. [120] also mentioned that spherical-shaped nanoparticles had higher uptake and more cellular internalization, compared to 2D disk-shaped nanoparticles. Further, Li et al. [118] in vitro compared spherical-shaped and rod-shaped gold nanoparticles, and the spherical-shaped nanoparticles showed higher uptake rate than that of rod-shaped nanoparticles. There is evidence in the literature that modifying surface approaches, such as coating, can increase the bioactivity of metal or polymer surface. The commonly applied approach is ceramic coating using plasma spray method, for example, nano-HA [121–123]. However, HA coatings are prone to degradation over a short period [124]. Thus the metal nanoparticle coating can be an alternative solution.

Nanostructured biomaterials for regenerative medicine: Clinical perspectives59

Augustine et  al. [114] shown ZnO nanoparticles could generate reactive oxygen species which enhanced angiogenesis through growth factor-mediated mechanisms. The results showed fibroblast growth factor and vascular endothelial growth factor upregulated due to the incorporation of ZnO nanoparticles. IONPs has the potential for use in tissue engineering due to its specific physical properties and excellent biocompatibility, epically IONPs have been widely investigated for magnetic hyperthermia [125]. Also, IONPs can be possibly used for targeting growth factors (e.g., nerve growth factor) to the desired location within the body using an external magnetic field [126]. TiO2 nanoparticles have been clinically applied for the coating material since it has good bactericidal activity but has the nontoxic property [127, 128]. The membrane of microorganisms can be damaged since the formation of superoxide anions (O2−), hydroxyl radicals (OH−), and hydrogen peroxide (H2O2) [129, 130]. For instance, Shiraishi et  al. [131] shown TiO2 film had a strong bactericidal effect on Staphylococcus aureus which is one of the primary bacteria causing pin (stainless steel) site infection. Koseki et al. [132] in vivo evaluated the plasma spray TiO2 coating in inhibition of infection when using percutaneous external fixation pins. There was less infection in the TiO2-coated pin group than the control pin group. Dr. Tadashi Kokubo’s group successfully used the sol-gel method to form thin uniform titanium oxide layers on polymer materials (e.g., polyethylene terephthalate [133] and polyetheretherketone (PEEK) [134]). They also demonstrated that the temperature used in the sol-gel coating method is significantly lower than the plasma spray technique and thus does not exceed the glass transition temperature of most polymers. The in vitro experimental results show that adding a sol-gel-derived TiO2 layer can positively affect the formation of apatite in SBF. Shimizu et al. [135] pretreated PEEK with dilute HCl acid before the sol-gel coating. Both in vitro and in vivo results posttreatment with acid was proved to confer the apatite-forming ability to the surface, compare to the untreated one. Moreover, chemical and heat treatments were used with metal implants; this is a commonly used method to modify the implant surface of Ti metal. For instance, NaOH and heat treatment was applied to the Ti metal implant to form sodium titanate with nanoscale needle-like network structure on the surface. The treated Ti metal formed an apatite layer on their surfaces in SBF and tightly bonded to living bone. These treatments have been successfully applied in the clinic since 2007 [136]. More recently, Kokubo and Yamaguchi [51] directly modified the titanium surface to form nanometer-scale roughness with heat treatment after exposure to acid (HCl) and alkali (NaOH) solution, which conferred good bone bonding ability. The formed rutile induces a gel layer on the titanium surface. They also demonstrated that the titania layer is vital for forming bone-like apatite on the implant surface, which would bond to living bone. This treatment method has already been applied to porous Ti metal for spinal fusion device in clinical trials [137, 138]. Divakarla et  al. [139] modified the surface of the commercial titanium alloy GUMMETAL (Ti59Nb36Ta2Zr3O0.3) with alkali and heat treatments. The metabolic activity of bone marrow stem cells on the surface treated commercial alloy was increased in comparison to control one.

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3.5 Composite nanostructures 3.5.1 The applications of polymer/hydroxyapatite nanocomposites Nanocomposite structures are used widely, as they can enhance the mechanical strength of hybrid organic/inorganic composites, and thus influence cellular proliferation and differentiation. Table 3.2 shows functional polymer/ceramic nanocomposites for the application of tissue regeneration. To mimic the organization of bone tissue that is composed of inorganic minerals and organic collagen matrix, silicate nanoparticles were incorporated into organic materials, enhancing mechanical properties, and further promoting cellular proliferation. Chae et al. [94] successfully fabricated HA/alginate nanocomposite fibrous scaffolds using electrospinning for mimicking mineralized collagen fibers in bone tissue, as shown in Fig.  3.5. The results showed that osteoblasts were a more stable attachment on HA nanoparticles/alginate scaffolds than pure alginate scaffold. The osteoblasts were round-shape on the pure alginate scaffold, whereas they presented spindle-shape on the HAp/alginate scaffolds. Stiffness is another key parameter for altering cell growth and differentiation. Alakpa et al. [150] fabricated supramolecular nanofiber hydrogels and controlled their stiffness to direct the differentiation of stem cells in  vitro without any biochemical functionalization. In developing regenerative biomaterials for hard tissue engineering (i.e., engineering of bone tissue), one material alone (either an organic polymer or an inorganic bioceramic) cannot meet the requirement. For closely mimic the characteristics of bone tissue at the nanometre scale, HA nanoparticles have been combined with natural and synthetic polymers [151–153]. Wang et al. [153] have shown the HA/ polyamide nanocomposite scaffolds had better biocompatibility and extensive osteoconductivity with host bone, comparing the pure polyamide scaffolds, at the preliminary period after implantation in vivo. Frohbergh et al. [151] in vitro indicated there are a higher expression and enzymatic activity of alkaline phosphatase (osteogenic marker) on the HA/chitosan nanocomposite scaffolds than the pure chitosan scaffolds. Also, there was an enhanced osteoinductivity of the HA/chitosan nanocomposite scaffolds due to the higher rate of osteonectin mRNA expression in composite scaffolds. Ribeiro et al. [152] have shown that the compression modulus of nano-HA/silk fibroin composite scaffolds was improved with an increasing nano-HA concentration. The alkaline phosphatase activities of osteoblastic cells in vitro were enhanced by the incorporation of nano-HA in the silk fibroin matrix.

3.5.2 The applications of polymer/bioactive glass nanocomposites Bioactive glass/biodegradable polymer composite materials have emerged with main applications of composite coating for the implant [154] and tissue engineering scaffolds [146, 155, 156]. In particular, recent studies indicate the application of

Tissue Bone tissue

Skull tissue Dental tissue

Polymer/ceramic nanocomposites

Experimental type

Functionality

Results obtained

PCL/nano-HA composite

Mimicking bone graft structure and composition

In vitro and in vivo

[140]

PCL/bioactive glass nanocomposite

Mimicking bone graft structure and composition

In vitro and in vivo

[141]

The nanocomposite of poly(ethylene oxide) and laponite nanoparticles A nano-HA—pullulan/dextran polysaccharide composite

Mimicking bone mineralization

Better attachment and proliferation, and higher alkaline phosphatase activity and calcium content on the PCL/nano-HA than the PCL/micro-HA The bioactivity and mechanical stability were higher in PCL/bioactive glass nanocomposite than PCL/bioglass microparticles composite Induced increasing mechanical strength, and enhancing cellular activities and mineralization The nano-HA matrix induced a higher mineralized tissue than the one without nano-HA

In vitro

[142]

In vitro and in vivo

[143]

Improved alkaline phosphatase activity and human osteoblasts proliferation with increasing nano-bioglass content Improved mechanical properties of the scaffold and increased osteoconductivity The scaffolds presented better cell viability and increased the formation of hydroxyapatite once immersion in SBF Extensive mineralization in SBF in the PCL/ nano-HA and the content of nano-HA affect antimicrobial activity

In vitro

[144]

In vitro and in vivo In vitro

[145]

In vitro

[147]

Dextran hydrogels incorporated with bioactive glass nanoparticles Chitosan/nano-HA scaffolds Nanobioglass into the chitosan gelatin to fabricate composite scaffolds PCL/nano-HA and loaded with amoxicillin

Mimicking bone tissue stimulating bone cell differentiation Mimicking bone

Mimicking skull tissue Mimicking alveolar bone tissue Mimicking dental tissue

References

[146]

Continued

Nanostructured biomaterials for regenerative medicine: Clinical perspectives61

Table 3.2  Functional polymer/ceramic nanocomposites for the application of tissue regeneration

62

Table 3.2  Functional polymer/ceramic nanocomposites for the application of tissue regeneration—Cont’d Tissue Calcified cartilage

Polymer/ceramic nanocomposites Self-assembled peptide amphiphile nanofibrous matrices to induce biomimetic nucleation of hydroxyapatite crystals

HA/alginate nanocomposite fibrous scaffolds

References

Promoted new bone formation in a rat femoral defect

In vivo

[148]

Induced continuous deposition of lamellar bone tissue while maintaining osteoblast activity

In vivo

[149]

Osteoblasts attached on HA/alginate scaffolds are more stable than on pure alginate

In vitro

[94]

Results obtained

Mimicking bone mineralization with collagen fiber structure and nucleation of hydroxyapatite crystals Mimicking natural bone tissue based on the inorganic materials and natural polymers Mimicking mineralized collagen fibers

Nanostructured Biomaterials for Regenerative Medicine

HA composites sponge with concentrated collagen nanofibres

Experimental type

Functionality

Nanostructured biomaterials for regenerative medicine: Clinical perspectives63

Fig. 3.5  Microscopy images of synthesized HA nanoparticles/alginate scaffolds. This figure is modified based on T. Chae, H. Yang, V. Leung, F. Ko, T. Troczynski. Novel biomimetic hydroxyapatite/alginate nanocomposite fibrous scaffolds for bone tissue regeneration. J. Mater. Sci.: Mater. Med. 24, (2013) 1885-1894.

n­ ano-sized bioactive glass particles in composites could improve the performance for biomedical applications in tissue engineering [17, 157]. Couto et al. [154] developed chitosan and bioactive glass nanoparticle multilayer coatings for the application on prosthetic devices. The bioglass improved bioactivity for the organic-inorganic structure, and the chitosan provided viscoelastic properties of the coating. The in vitro results showed that the multilayer coating induced apatite formation. Apart from the use of polymer/bioactive glass nanoparticle for surface coatings, it has been utilized as a tissue scaffold material. Peter et al. [146] have shown protein adsorption was increased with the addition of bioactive glass nanoparticle into the ­chitosan-gelatin

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to fabricate composite scaffolds fabricated with the sol-gel method. de Oliveira et al. [155] developed polyurethane/bioactive glass nanoparticle scaffolds, and the scaffolds present good cell viability and hydroxyapatite layer formation upon immersion in SBF. Ji et al. [156] indicated that the elastic modulus of the bioactive glass/PCL nanocomposite was improved by an increase of the bioactive glass nanoparticle content. Comparing with pure PCL polymer, the hydrophilic property, degradation behavior, and mineralization behavior were all improved with the addition of the bioactive glass nanoparticle. The polymer/bioactive glass nanocomposites also have the improved roughness, wettability, and surface area, which could promote bone regeneration through the increased nutrient exchange, protein adsorption, and porosity [16, 158, 159]. In summary, nanostructured ceramics, especially nano-HA, are popular as coatings, filler materials, and incorporation with other materials as nanocomposites to be used as bone substitutes. The high surface fraction in ceramic nanoparticles would increase osteoblast functions, such as adhesion, proliferation, and differentiation. Nanocomposite scaffolds also provide a support structure for cells, and thus if changing the tissue at the nanoscale level would affect cell-scaffold adhesion, interaction, and cellular function.

3.5.3 The application of polymer/metal nanocomposites There are increasing studies about the use of metal nanoparticles and polymer as the composites to apply in the field of regenerative medicine. With the optimized combination of polymer and metal nanoparticles, the desired property of nanocomposite can be obtained. Silver nanoparticles have been commonly used in tissue engineering due to its capability to release silver ions which in turn leads to an antibacterial activity [95, 98]. However, silver nanoparticles are easily aggregated due to the high surface free energy, and also they can be oxidized or contaminated in the air. Thus, these difficulties in processing have restricted its application [160]. The valid solution to solve the problem is the incorporation of silver nanoparticles into the biodegradable polymer [113, 161]. Also, by varying the concentration of silver nanoparticles in a certain range, the surface morphology of polymer/silver nanocomposite would change, and further ­affect the roughness and wettability on the nanocomposite surface. These characteristics can influence the bacterial adhesion on the nanocomposite surface [162, 163]. Moreover, the incorporation of gene materials at the local site can sustainably produce growth factor with gene transfection. For instance, Tandon et al. [164] demonstrated that polyethylenimine (PEI)-conjugated gold nanoparticles was an efficient carrier for mediating BMP7 gene delivery in vivo, which also modulated wound healing and inhibits fibrosis. Also, the mechanical properties of the polymer can be improved by incorporating of metal nanoparticles, due to the characteristics of nanostructured metal (e.g., high modulus and large surface area) [165]. Chatterjee et al. [166] indicated that the mechanical property of nanocomposites can be achieved by increasing the interaction area between the polymer matrix and nanoparticles. However, some of these interactions also lead to toxicity, which can be a serious problem for tissue regeneration. Thus, in-depth investigations of nanomaterials in vivo are required.

Nanostructured biomaterials for regenerative medicine: Clinical perspectives65

3.5.4 Other nanostructured composite biomaterials With adding nanoparticles, many studies have proved that a nanostructured composite can promote osteogenesis in the absence of growth factors. Gaharwar et al. [167] indicated the adhesion, spreading, and proliferation of MC3T3-E1 mouse preosteoblast cells can be modified by varying the concentration of laponite nanoparticles in laponite-PEO nanocomposites. The average diameter and thickness of the laponite nanoparticles were 25–30 nm, and 1 nm, respectively. Laponite nanoparticles also influence the differentiation of preosteoblast cells as there was increased mineralized phosphate producing on the nanocomposite surfaces. Wu et al. [168] indicated that adding laponite, the composite hydrogel had higher elongation and improved toughness comparing with pure hydrogel. Xavier et al. [169] studied the cellular response to the nano-laponite/collagen-based hydrogel film. The nanocomposite hydrogels could promote osteogenesis in vitro without the involvement of osteoinductive growth factor. There were an increased in alkaline phosphatase activity and the formation of a mineralized matrix by adding the nano-laponite to the collagen-based hydrogels. Moreover, the Kalpana S. Katti’s group modified nanoclay with amino acids to mineralize HA mimicking biomineralization in bone [170]. In Ambre et al. [171] work, the modified nanoclay was incorporated into chitosan/polygalacturonic acid films; hMSCs were seeded on the films to investigate the cellular response. The in  vitro results indicated there is the formation of mineralized nodules on the chitosan/polygalacturonic acid films without adding osteogenic supplements used for hMSCs differentiation. Viability and differentiation assays results show that the composite scaffolds were favorable for the viability and differentiation of hMSCs. Further, Ambre et al. [172] incorporated modified nanoclay in polycaprolactone to form PCL/nano-HA clay films; the cellular response of the composite film was examined. The results indicated that hMSCs composed mineralized ECM without the use of the osteogenic supplement. Also, PCL/nano-HA clay films had significantly increased in elastic moduli nanomechanical properties. Liao et al. [173] developed polyethylene glycol/graphene oxide nanocomposite scaffold for mimicking cartilage engineering, and the results show the improvement of mechanical properties and electrical conductivity of scaffold by graphene oxide, which leads to enhanced regeneration of cartilage tissue. Ahadian et al. [93] incorporated aligned carbon nanotubes into hybrid hydrogel scaffold, and there are tunable and anisotropic mechanical and electrical characteristics, which lead to enhanced cardiac differentiation of embryoid bodies with increased beating activity.

3.6 The clinical products for tissue regeneration based on nanotechnologies In 2014, the US FDA defined nanotechnology products as those which have at least one dimension between 1 and 100 nm in size [85]. The development progress of regenerative medicine nanostructured products for tissue regeneration in the clinic, starting with the basic concept product, then in vitro and in vivo studies, and culminating with clinical investigations and commercialization. Table 3.3 summarizes the selected nanostructured medicine products that have obtained from the FDA and applied in clinics.

Product/company Nanostructured polymers Regranex/Smith & Nephew, Inc.

HemCom/Medical Technologies, Inc. Nanostructured metal Acticoat/Smith and Nephew, Inc.

NanOss/Angstrom Medica, Inc.

Indications for use

Nanoparticle advantage and clinical outcomes

Year approved

References

Sodium carboxymethylcellulose gel, containing growth factor Chitosan acetate

Diabetic foot ulcers

Low concentrations of growth factors in nanoparticles minimize side defect

1997

([174], [175])

External wound healing

More effective in reducing bacterial luminescence

2008

[176]

Nanocrystalline silver

External wound healing

Prevents bacterial infection and improves wound healing

2009

[177]

Hydroxyapatite nanocrystalline

The filler is injected into a bone void or defect

2004

([178], [179])

Hydroxyapatite nanocrystals

Filler, osteoconductive, resorbable bone graft for osseous defects

This filler facilitates bone regeneration, based on its bone mimetic chemical composition and crystalline structures HA nanoparticles mimic the microstructure and the composition of bone and have higher mechanical properties and osteoconductive effects

2005

([180], [181])

The filler can control timed the release of calcium sulfate that supports bone augmentation

2006

[182]

Nanostructured composites BoneGen TR/BioLok Calcium sulfate-based International, Inc. nanocomposite

Filler, oral surgery, periodontics, endodontics, implantology

Nanostructured Biomaterials for Regenerative Medicine

Nanostructured ceramics Ostim/Osartis, Inc.

Nanostructured material descriptions

66

Table 3.3  Selective FDA approved nanostructured products for the application of tissue regeneration in the clinic

2008

[183]

2009

[184]

2010

[184]

This filler is controlled to be degraded over the time, stimulating bone regeneration There are increased alkaline activity and collagen production

2011

[185]

2013

([186], [178])

Implant, spinal fusion procedures

This implant has a nanotubeenhanced surface which can promote bone regeneration around the implant

2014

[85]

Porous 3D composite construct as bone boid filler

This filler has interconnected porosity mimic human cancellous bone to promote tissue interaction and regeneration

2014

[183]

Type I collagen fibers with nano-HA particles

Spinal surgery/bone void filler

EquivaBone/ETEX Corporation

Paste composed of a bone matrix and nanocrystalline hydroxyapatite

Bone void filler in spinal and trauma surgery

Beta-BSM/ETEX Corporation

Synthetic calcium phosphate bone graft material in a nanocrystalline matrix Calcium sulfate hemihydrate based nanocomposite Type I collagen fibers with Mg-HA nanoparticles Implant composed of a highly porous titanium a scaffold that is integrated with a PEEK core Hydroxyapatite nanogranules suspended in a collagen-based foam matrix

Injectable bone substitute material for orthopedic trauma and bone void filler Filler, osseous defects

NanoGen/Orthogen, LLC

RegenOss/JRI orthopaedics FortiCore/Nanovis, Inc.

nanOss Bioactive 3D bone void filler/Pioneer Surgical Technology, Inc.

Filler acetabular defects and spinal fusion

Nanostructured biomaterials for regenerative medicine: Clinical perspectives67

The filler is resorbed and remodeled into the new bone as part of the natural healing process This scaffold has osteoconductive effect by providing hydroxyapatite nanocrystalline and osteoinductive growth factors This filler has osteoconductive properties based on the bone mimetic chemical structure

Healos/DePuy synthes spine

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Nanostructured Biomaterials for Regenerative Medicine

Polymer nanoparticles can be used as carriers loaded with drugs or biological molecules, which is a promising tool for skin regeneration and wound healing. Regranex, becaplermin gel, was first approved by the FDA for the healing of diabetic neuropathic foot ulcers. It is sodium carboxymethylcellulose-based gel containing recombinant platelet-derived growth factor which is beneficial in wound healing [187]. However, a side effect has been reported (e.g., carcinogenic effects), which could be minimized by the application of low concentrations of growth factors in nanoparticles [174]. HemCom, a chitosan-based hemostatic agent, has been widely used for the control of severely bleeding wounds [176]. Biological molecules, such as corticosteroids [188] and antioxidants [189] have been loaded into chitosan nanoparticles in  vitro and in vivo for promoting wound healing; however, further clinical study is required. Silver nanoparticles are the only metal nanoparticles used for wound healing in the clinical market due to their low systemic toxicity, antibacterial activity, biocompatibility, and low cost [190]. Acticoat, nanocrystalline silver wound dressing, has been clinically used for topical treatment of infected burns, open wounds, and chronic ulcers [174, 177]. Also, silver nanoparticles-based drug-delivery systems can improve antibacterial efficacy by directly targeting antimicrobial agents to the site of infection [191]. Most of the current nanotechnology-based regenerative medicine products are made for bone tissue regeneration, fillers, and osseous defects. Bioceramic, mainly nano-sized HA has been used in commercial products (e.g., nanOss, Ostim, Cerabone, and BoneSave) [179, 181] and have gained the requisite regulatory approval. Recently, some nanoparticle-based composite biomaterials also have been obtained from the FDA. For instance, Healos is nanocomposite of type I collagen fibers and nonsintered calcium phosphate mineral nanoparticle. However, Kraiwattanapong et al. [192] indicated Healos was ineffective as a bone graft substitute when combined with autogenous bone marrow in rabbit’s body. The composition of RegenOss is the type I collagen fibers coated with Mg-HA nanoparticles. It has been shown to be biocompatible and was resorbed 90 days after implantation in sheep [193].

3.7 Conclusion and future perspective In this review, we highlighted leading edge nanostructured materials that mimic the composition and structure of the hard and soft tissue. Based on recent advances in nanotechnologies, bioengineered scaffolds are becoming more similar to natural tissue, thus enabling the recovery of damaged tissue. To mimic the organization of hard tissue (e.g., bone, calcified cartilage, and dental tissue) that is composed of inorganic minerals and an organic collagen matrix, ceramic nanoparticles were incorporated into polymer materials, enhancing the mechanical properties of the polymer and further promoting cellular proliferation. The increased roughness and the ratio of area to volume can be achieved, which are beneficial for the enhancement of protein adsorption. In addition, HA and WH are mainly two bone crystal nanoparticles, which are distributed in different ratios depending on the regions of bone tissue. By controlling their composition distribution to mimic native bone

Nanostructured biomaterials for regenerative medicine: Clinical perspectives69

tissue at the nanoscale can enhance the proliferation of bone cells and induced rapid regeneration of bone tissues. Various nanotechnologies, such as electrospinning, nanolithography, self-assembly, phase separation method, have been developed, to fabricate nanocomposite scaffolds. The nanocomposites have better mechanical and biological property, compared with either sole using polymer or nanostructured ceramic materials. However, the most common method for a bone substitute is mainly ceramic filler/paste form. This could be due to the absence of understanding of biological components in bioengineered tissue fabricated using various nanotechnologies. Moreover, some studies focus on the application of nanostructured material to closely mimic the soft tissue. By the surface modification of vascular polymer graft surface, the nanometer surface roughness could improve endothelial cell functions. Laser treatment, one of the surface modification methods, can be used to develop nanometer surface morphology. The modification process also involves the change of surface chemistry, such as the formation of new oxygen-containing groups and amino groups, which can facilitate better cell adhesion. Nanopatterns also play an essential role in directing various cellular behaviors, and control cell functions. The alignment of nanofibers in bioengineered tissue can guide cell movement and orientation, which can mimic the heart, tendon, and blood vessels tissues. Furthermore, the growth factor can regulate cellular migration, differentiation, and proliferation, and it is vital for the growth factor located in the site of injury for the repairing process. Due to the size and surface chemistries, the nanostructured biomaterials can be the delivery carriers for growth factors, peptides, and gene materials. Some studies demonstrated there were improved transport properties and more efficient delivery of growth factors to target sites. Also, the incorporation of laponite nanoparticles in nanocomposite can promote osteogenesis without growth factors. Although there are many positive results about the application of nanostructured biomaterials in soft tissue, the majority of the developed nanomaterials have not been utilized for soft tissue regeneration in a clinical setting. The main challenge is to scale up the formation of soft tissue from the nanoscale to macroscale for the tissue repair. The use of nanotechnology to create bioengineered scaffolds to closely mimic natural tissues in regenerative medicine has received increased attention over the years. It is ideally able to design the bioengineered scaffold to match with the natural tissue regarding integral structure, material composition, and surface morphology. However, the biological environment in vivo is dynamic, and the bioengineered scaffold is unlikely to sustain tissue growth with time. Although many nanoparticles have a positive performance effect on cellular internalization rates and have been utilized in regenerative medicine, there are still safety concerns about the use of these nanomaterials. The mechanisms have not yet been clarified about if the nanoparticles can cause toxic effects by crossing cell barriers, which will need to be explored further. These problems could be addressed thorough investigating the physicochemical characterization of nanomaterials. Based on the understanding of the effectiveness and safety of nanomaterials, proper in vivo studies should be pursued for clinical translation of nanomaterials.

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Nanostructured Biomaterials for Regenerative Medicine

Despite these shortcomings, the use of nanostructured materials can be used to precisely control bioengineered scaffold structures, and positively affect the release of bioactive factor(s) that are vital for tissue regeneration. There will be a continuing trend in (a) the use of nanostructured biomaterials and (b) the approval of nanostructured products by the FDA. The study of nanostructured material in a clinical stage has more growth potential and overcome current challenges in regenerative medicine to heal damaged tissues.

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Nanostructured biomaterials with antimicrobial activity for tissue engineering

4

Argelia Almaguer-Floresa, Phaedra Silva-Bermúdezb, Sandra E. Rodilc a School of Dentistry, National Autonomous University of Mexico (UNAM), Mexico City, Mexico, bDepartment of Tissue Engineering, National Institute of Rehabilitation Luis Guillermo Ibarra Ibarra, Mexico City, Mexico, cInstitute of Materials Research, National Autonomous University of Mexico (UNAM), Mexico City, Mexico

4.1 Biomaterial-related infections During the last 60  years, implantable medical devices have gained growing importance, mainly due to the rising median age of world population and the scientific and technological advances that have allowed the development of implantable medical devices for replacing or augmenting different tissues and organs. Unfortunately, medical devices are an important source of health-care-related infections, being biomaterialsrelated bacterial infections one of the major health risks associated to the use of either permanent or temporary implants. Throughout the three stages of implant surgical intervention, microorganisms can be encountered: (a) preoperatively, from contaminated wounds, hospital environment, or the normal skin flora, (b) perioperatively, from contaminated implants or implants that become contaminated upon contact with hospital environment, wounds, physiological fluids, etc., and (c) postoperatively, during the lifetime of the implant from hematogenous or lymphogenic routes [1, 2]. Nevertheless, the most common source of infection occurs during the perioperative stage due to contamination from microorganisms that adhere to the biomaterial surface from infected wounds, normal skin flora, etc. According to Leaper et al. [3], half of the two million cases of health-care-related infections occurring annually in the United States might be associated with implantable devices. Gastmeier et al. [4] reported contaminated medical devices as the source of 21.1% of health-care-associated infections. For example, in the trauma environment, infections following primary implant surgery occur in 0.5%–5% of the patients, particularly, open fractures are at higher risk for implant-associated infections (up to 30%); in the cardiovascular field, implant-associated infections occur in around 4% of the patients receiving heart valves, vascular grafts, or pacemaker-defibrillators; while, in neurological or cosmetic surgery, implant-related infection constitutes around 2%– 6% of primary implant surgeries [1, 5–7]. Even if average infection rate is below 6% for implanted devices, biomaterials-­ associated infections represent a significant burden for national health systems from both economical and public health perspectives. The treatment of these infections Nanostructured Biomaterials for Regenerative Medicine. https://doi.org/10.1016/B978-0-08-102594-9.00004-8 © 2020 Elsevier Ltd. All rights reserved.

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often requires long-period antibiotic therapies, combination of different antibiotics, several surgical interventions, and even removal and replacement of the implanted biomaterial. Biomaterials-associated infections are difficult to treat because the host immune defense system is depressed due to the foreign-body reaction (the number of bacteria required to cause infection is significantly lower in this case that when no foreign-body reaction is active) and mainly, because bacteria can easily adhere to biomaterials surfaces, replicate, and form a biofilm. The clinical relevance of preventing bacterial adhesion and biofilm formation on biomaterials used for tissue engineering (and it also applies for nanostructured biomaterials) relies on the fact that once the biofilm is formed, bacterial cells become highly resistant to the antibiotic therapy and the host immune defenses, making the treatment of this kind of infections very difficult to succeed. Clinical experience has shown that regarding biofilm-related infections, the microorganisms must be removed physically from the biomaterial; an attempt to disrupt the complex structure that protects bacterial cells from the antibiotics and the surrounding environment. Unfortunately, mechanical treatment for removing the biofilms often requires a second surgical intervention (revision surgeries) which can cause discomfort of the patients, additional pain, and also extra monetary cost, plus the fact that revision surgeries are associated with even higher frequencies of infections than primary implant surgeries [1, 2]. In the worst cases, when mechanical cleaning and antibiotic therapy used for eradicating the biofilms are not sufficiently effective, the loss of the implanted biomaterial is imminent and with this, the failure of the desired treatment. Thus, biofilm formation is quite relevant for biomaterials-associated infections pathogenesis.

4.1.1 Bacterial adhesion and biofilm formation on biomaterial surfaces Biofilms are fascinating and complex structures that constitute the predominant mode of growth of many bacterial species. A biofilm is an assemblage of microbial cells irreversibly associated with a surface and enclosed in a matrix of primarily polysaccharide material; this structure protects the microbial communities from predation, toxic substances such as antibiotics and physical perturbation [8]. The variety of microniches provided by the biofilm also promotes a considerable diversity of microbial life and metabolic potential [9]. It is interesting that biofilms can be observed in virtually every natural environment like streams, lakes, oceans, and also in human tissues like teeth, heart valves, and vaginal surfaces [10]. In addition, biofilms have the potential to cause device-related and other chronic bacterial infections that have gradually come to predominate in modern medicine causing harm to millions of humans annually. In the human body, biofilms can cause a significant amount of microbial infections including periodontitis, device-related infections, osteomyelitis, otitis media, infective endocarditis, and others [11, 12]. The difficulty of eradicating biofilm bacteria with standard systemic antibiotic treatments is a prime concern of medicine, because the bacteria in a biofilm can be up to a thousand times less susceptible to antimicrobial stress than their freely suspended counterparts [13, 14].

Nanostructured biomaterials with antimicrobial activity for tissue engineering83

The first step in the formation of a biofilm is the adhesion of planktonic cells to the surface of the biomaterial. The primary adhesion of a planktonic microorganism is reversible and is dictated by a variety of physicochemical properties that regulate the interactions between the bacterial cell surface and the biomaterial. The mechanisms of bacterial adhesion can be divided into physicochemical (phase 1) and molecular-­ cellular interactions (phase 2). Physicochemical interactions occurs via the long-range van der Waals forces and electrical double-layer forces, which form the basis for the well-known DLVO theory [15, 16]. DLVO theory describes the force between charged surfaces (bacteria-membrane and biomaterial) in liquid media. The net result of the different interactions led to attractive force between alike-charged surfaces and repulsive force for oppositely charged surfaces [17]. Van der Waals attraction tends to dominate at larger distances, but as the organism comes into a close approximation of the biomaterial surface, usually  90% Released of Ag + ions up to 7 days. No bacteria growth on the surface but fibroblasts viability was reduced with Ag coating Improved biocompatibility, no antibacterial tests

Conclusion

[170]

[169]

[168]

[167]

[166]

[165]

[164]

[163]

[162]

References

Polyamide

Polyaniline (PANI)

TiAlV-Cu alloys

5, 10%

0.2–1.0 ppm

5 wt%

S. aureus

E. coli, S. aureus, and C. albicans

E. coli

Methicillinresistant S. aureus, E. coli, S. epidermidis S. aureus, E. coli

Mg-Cu alloys

Cu-doped wollastonite (CaSiO3)

S. aureus (CFU’s 24 h)

Bacterial strain

Coatings/titanium

Matrix-support/ substrate

0.7 ppm

Nitrided TiCu coatings. 1.6 Cu at% 0.04, 0.09, 0.23 wt%

Cu load

Table 4.3  Copper (Cu) antimicrobial applications

NA

Blood cells

Mouse mesenchymal stem cells, human osteoblast cells NA

In vivo tests

NA

Mammalian cells

Addition of copper to the wollastonite enhances the antibacterial potentiality and osteoblastic differentiation. No cytotoxicity below 0.05 mg/mL 10 mg/mL inhibits 100% bacterial growth. Activity retained up to 35 days following the Cu2 + ion release MIC values of 1 ppm for the three species. Cu-PANI nanocomposite at investigated concentrations (0.7–20 ppm) displayed cytotoxic effect The annealing temperature of 740°C was the most promising heat treatment for Ti6Al4V5Cu. Biofilm inhibition was associated to the presence of globule Ti2Cu phase and/or blocky β-Ti phases, which contained high content of Cu

The coatings presented antibacterial effect and also wear and corrosion resistance The alloy containing 0.25 wt% Cu exhibited the highest antibacterial activity, with favorable biocompatibility

Conclusion

[180]

[179]

[178]

[177]

[176]

[175]

References

94 Nanostructured Biomaterials for Regenerative Medicine

Nanostructured biomaterials with antimicrobial activity for tissue engineering95

Laurenti and Cauda presented a comprehensive review about the use of ZnO for tissue engineering applications, where based on the collective analysis of different works, they conclude that besides the well-known antibacterial properties of ZnO, there is also clear evidence that it could promote both osteogenesis and angiogenesis [182]. Such results have been confirmed for different nanostructures, such as nanoflowers, nanowires, NPs, and thin films, as well as for ZnO containing nanocomposites using both ceramic and polymeric matrix. Many of the examples summarized in Table  4.4 have a composite structure including Zn or ZnO and have been deposited on the surfaces of materials designed for biomedical implants.

4.2.2.5 Polymeric molecules Surface coating of biomedical implants to prevent bacterial adhesion and biofilm formation based on known intrinsic antibacterial organic molecules is another mechanism to prevent infections due to the increasing prevalence of antibiotic-resistant strains. Such molecules could act to prevent bacterial adhesion or through a bactericidal effect to nonadherent or adherent bacteria. Recent results concerning the use of selective bacterial killing using coatings of organic or polymeric molecules, such as chitosan and curcumin [193] are summarized in this section (Table 4.5). Chitosan is a biodegradable and biocompatible natural polysaccharide obtained from different sources, being the exoskeleton of crustacean shells and squids, the most common. It has been widely studied due to its antimicrobial activity, probably conferred due to its polycationic character which allows binding to the negatively charged bacterial surface, leading to alteration in the membrane permeability and cell death [143, 194, 195]. However, this antimicrobial activity of chitosan is limited to acidic pH conditions since at neutral pH and above, the positive charges on the amine groups are lost, and moreover, the solubility of chitosan is drastically reduced. This phenomenon is the main reason for the different approaches investigated to improve the properties of chitosan by forming nano-microparticles, functionalization, or nanocomposites adding other antibacterial substances, metal, or oxides. Perinelli et al. have published a recent review paper to present the different physicochemical properties that affect the antibacterial properties of chitosan, as well as a summary of the recent results concerning chitosan composites formulations [195]. The main conclusion is that despite the thousands of papers devoted to the research on the antimicrobial properties of chitosan, the results are still controversial regarding the active mechanism and the comparison between the antimicrobial activity of chitosan and chitosan NPs.

4.2.2.6 Others Different studies have reported the addition of antibacterial metals or oxide NPs into titania nanotubes (TNTs) produced on Ti surfaces (as can be observed in Tables 4.2– 4.4). These TNTs can be produced by different means, such as electrochemical anodization, and their promising biomedical properties have been described before [204, 205]. Significant reduction of the bacterial attachment has been observed adding antibacterial metals/oxides into the TNTs, but the final effect might be a synergy between

PLGA-Ag/ZnO coatings/ titanium

Sr-ZnO-anodized Ti/Ti

Ag-ZnO/sulfonated PEEK

Zn2 + release up to 1000 μg/L in 15 days

Zn2 + release of 9000 μg/L in 5 days

ZnO-polydopamine (PDA)-arginine-glycineaspartic acid-cysteine (RGDC) nanorod (NR) arrays/titanium N-Halamine-immobilized silica-coated PSA ZnO nanoparticles/titanium

Matrix-support/ substrate

ZnO nanowire coating

~ 5 wt%

Hybrid

Zn load

Table 4.4  Zinc (Zn) antimicrobial applications

E. coli, S. aureus

E. coli, S. aureus

E. coli, S. aureus

P. aeruginosa, E. coli, S. aureus

S. aureus, E. coli

Bacterial strain

Osteoblast-like MG-63 cells

MC3T3E1 cells up to 7 days

Mouse calvarial cells

MC3T3-E1 preosteoblasts from mice

Osteoblasts

Mammalian cells In vitro and in vivo results reveal that the hybrid NRs possess contact-killing antibacterial activity against adherent S. aureus and E. coli as well as outstanding osteogenic properties Significant antibacterial properties for S. aureus and E. coli due to synergy effect of the N-halamine compound. Biocompatibility was similar with or without surface modification, but ALP expression was enhanced Enhanced Zn2 + release by the addition of PLGA/Ag. The ZnO nanorods-Ti samples coated with PLGA containing 3 wt% AgNPs had sufficient antibacterial activity, minimal toxicity and induced cell proliferation Ti nanotubes doped with Sr-ZnO with or without antifouling chemical agent. Antibacterial activity increased to nearly 90% for both strains while cell viability and ALP activity was comparable to bare Ti, except for the hydrophobic surface Bacterial adhesion reduction about 90% for Ag/SPEEK and AgZn/SPEEK; but no for ZnO/SPEEK coated surfaces. Zn containing surfaces show low toxicity and significantly higher ALP activity and genetic expression (RUNX2, OCN, Col1)

Conclusion

[187]

[186]

[185]

[184]

[183]

References

Electro deposition of ZnO NPs into titania nanotubes/Ti

S. aureus

Folic acid-ZnO quantum dots sealed drug-filled titania nanotubes/Ti

2.09, 2.46, 5.54 at%

E. coli, S. epidermidis

ZnO-hydroxiapatite (HAp) coating/Ti

Zn2 + release between 0.5 to 1.0 μg/L in 1 h ZnO quantum dots completely dissolved

S. aureus

E. coli

ZnO NPs embedded in poly(lactic acid)/ magnesium alloy

Nominal 5, 10 wt% Real: 4.36, 6.18 at%

E. coli, S. aureus

Hydrothermal treatment to Zn containing Microarc oxidation coatings/Ti

Zn2 + release of 250–350 μg/L in 14 days

Macrophage-like RAW cells

Osteoblasts MC3T3-E1 cells

Osteoblasts MC3T3-E1 cells

Slight increment in the antibacterial rate by the introduction of Zn, but the largest effect is due to the porous structure generated by the micro-arc oxidation of the Ti surface The results indicate that the incorporation of about 5 wt% ZnO NPs within the PLA coating increased the antibacterial properties without negatively impacting their cytocompatibility Antibacterial activity of the HAp coating is significantly reduced (~ 100%) with the addition of Zn Cell viability is increased by the formation of titania nanotubes and steeply decreases with the antibacterial drug and the addition of ZnO, effect more evident after 7 days. Meanwhile the antibacterial effect is increased 60%–80% in comparison to bare Ti The bacteria attachment was reduced above 90% by the addition of Zn in comparison to the titania nanotubes. Macrophage attached cells show poor activity, suggesting a decrease in the immune response with the addition of Zn [192]

[191]

[190]

[189]

[188]

Modified surface

Polyurethane

Ag NPs

Chitosan/nano hydroxyapatite (CS/nHA) scaffold containing zoledronic acid

Chitosan grafted to PTFE samples using different spacer molecules (GA, PEGb, PA)

Antibacterial molecule

Chitosan

Chitosan

Chitosan

Chitosan Xylella fastidiosa

E. coli, S. aureus

E. coli, S. aureus, C. albicans. (inhibition zone, MICs, MBCs)

S. aureus, P. aeruginosa

Bacterial strain

NA

Bone giant cell tumor cells and human bone marrow mesenchymal stem cells

NA

NA

Mammalian cells

Table 4.5  Organic or polymeric molecules with antimicrobial applications

Chitosan immobilized films had strong antibacterial activity for both species, demonstrating higher efficiency parallel to the chitosan concentration MICs and MBCs concentrations were lower for the colloidal Chitosan-Ag NPs in comparison to Ag colloids, penicillin, chitosan and AgNO3 solutions for the three species. However, inhibition ratios were comparable or slightly lower to penicillin Scaffolds with and without zoledronic acid showed almost 100% reduction of bacterial adhesion suggesting the activity of the chitosan. The zoledronic content promote tumor cells apoptosis, without toxicity to hBMSCs cells, nor blood cells Chitosan linked using PA exhibited better antibacterial response compared to GA and PEGb, which could be correlated to a denser coating promoting the interaction with the bacteria cell wall

Conclusion

[199]

[198]

[197]

[196]

References

Chitosan-heparin/ alkali treated Titanium

Furanones, FC/polymers

PVP-I/anodized Ti6Al4V

poly(ε-caprolactone) (PCL)/gum tragacanth (GT) curcumin loaded nanofibers

Chitosan

3-(10-Bromohexyl)5-dibromomethylene2(5H)-furanone. Br at% 0.11–8.10

Povidone-iodine (PVP-I)

Curcumin

Methicillin resistant Staphylococcus aureus (MRSA), extended spectrum β-lactamase (ESBL)

S. aureus

S. epidermidis

E. coli, S. aureus

Mesenchymal stem cells

NA

Murine cells

Blood and endothelial cells

The antibacterial rates against E. coli was between 46%–74% and against S. aureus was 55%–65%, increasing with the chitosan layer. Heparin induced excellent antithrombotic activity, and the biocompatibility was not compromised The bacterial load on control and FC biomaterials were similar at 1 h indicating that initial adhesion was not altered by furanone coating. However, at 24 h, the furanone coating significantly inhibited bacterial load on all the polymers PVP-I can provide long-lasting antibacterial effect even under mechanical damage of the surface The results suggested that the nanofibers demonstrated remarkable antibacterial activity against MRSA (99.9%) and ESBL (85.14%) [203]

[202]

[201]

[200]

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the micro-nanostructure introduced by the TNTs, the release metal ions, and a possible intrinsic antibacterial effect of the TiO2 nanotubes [206]. The antibacterial and self-cleaning properties of TiO2 have been usually associated with UV-illuminated conditions, since the activity is enhanced through the formation of electron-hole pairs that catalyzed chemical reactions. Such an approach is useful to sterilize samples [207], but no for tissue engineering applications. However, Jing et al. [208] showed the antibacterial activity of TNTs even without exposure to UV light. They produced TNTs by anodization of Ti samples having different diameters between 50 and 150 nm and tested the antibacterial activity against Staphylococcus aureus for 1, 3, and 7 days. Their findings show that the antibacterial activity was about 60% higher than on the smooth Ti surfaces, while good in vivo skin integration was observed. Recently, there has been an interest in the evaluation of the antibacterial activity of other metal oxides, such as magnesium oxide [209], hafnium oxide [210], and silica [131]. Emerging antibacterial carbon-based materials are graphene [211, 212] and graphene oxide (GO) [213], which are commonly combined with other materials (metals, metal oxides, or polymers), either to reduce toxicity or to enhance antimicrobial properties. GO is the oxidized form of the 2D single layer of carbon atoms, which is one of the allotropes of carbon presenting sp2 hybrids. According to the recent review paper of Yousefi et al., the antibacterial efficiency of GO is due to the damage of cell membranes via generation of ROS and the physical damage caused to the cell membrane due to the GO sharp edges [213]. GO on polymer (silicone rubber) substrates as an antibacterial coating was recently studied by Liu et al. [214]. In this study, cell viability and live/dead experiments using S. aureus and E. coli were done. Results showed that only the 14.2% of E. coli cells and 27.6% of S. aureus could survive on the surface of GO coatings in comparison to the 100% growth on the silicon rubber. Even more, scanning electron microscopy (SEM) analysis of the attached dead bacteria indicated that the sharp edges of the GO that was placed perpendicular to the silicone surface, created severe damage to the thinner wall of the E. coli, leading to cell death. More details concerning the biomedical applications of GO can be found in two recent reviews [215, 216]. Sulfur and metal sulfide NPs have also been considered as antibacterial agents. Argueta-Figueroa et al. published a review presenting the synthesis, biocompatibility, and antibacterial activity of silver, copper and ion sulfides, as well as other biomedical applications [217]. Concerning Ag2S, there was not a clear agreement between different publications concerning the biocidal effect to typical microbes; probably associated with the different sizes, synthesis methods and capping agents used. It was shown that Ag2S is photoactive in the visible, so it can produce ROS, which might explain the antibacterial effect observed in some cases. Copper sulfides comprise different compositions and structures, although the more common are CuS and Cu2S. For CuS, a research group has shown that sulfidation reduces the toxicity of Cu NPs while giving a good antibacterial effect [218–221]. A recent paper [222] shows that pyrrhotite (Fex − 1S) nanoplates synthesized by hydrothermal method drastically reduced bacterial growth (S. aureus, E. coli, and Enterococcus faecalis) at concentrations about 0.15  mg/mL. Moreover, complete ­inhibition for the three species was obtained at 0.625 mg/mL. The cytotoxicity was

Nanostructured biomaterials with antimicrobial activity for tissue engineering101

evaluated at this large concentration for three human cell lines (fibroblasts, pulp, and osteoblast), finding slight toxicity, not superior to 50% reduction in cell viability.

4.2.3 Grafting As it has already been mentioned in this chapter, a vast number of surface modification strategies have been developed to confer biomaterials surfaces with antiinfective properties, considering that requirements to which antiinfective biomaterials have to respond are very broad and use-specific, but also that they have to be biocompatible and exert appropriate tissue interactions. Among all the different strategies, the next paragraphs will attempt to rapidly enlist some of the latest grafting antibacterial strategies to achieve surfaces with antibacterial properties; from 2016 and backwards there are various comprehensive reviews [1, 2, 223–226]. Antibacterial agents can be grafted to surfaces with leaching or nonleaching purposes; nevertheless, in this section, we will talk about grafting strategies with nonleaching purposes. In this perspective, grafting antibacterial strategies can be roughly divided into (a) prevent bacteria attachment and (b) produce contact-killing or biocides surfaces.

4.2.3.1 Polymer brushes One of the first approaches of grafting antibacterial strategies was to passively prevent bacteria adhesion by surface immobilization of molecules or compounds that disrupt appropriate physical-chemical interactions between surface and bacteria; that is, mainly attachment of polymer brushes. In this approach, where surfaces are modified to prevent bacteria adhesion rather than to kill bacteria, grafting aims to tune surface energy, surface charge, mechanical properties, or water interactions (hydrophobic/hydrophilic nature) of the surface to prevent appropriate electrostatic bacteria-surface interactions. Thus, bacteria adhesion is minimized by reducing the bacteria adhesion forces to the solid surface, diminishing electrostatic interactions and Van der Waals forces in combination to creating mechanical modifications (soft grafting-modified low adhesion surfaces mainly consist of hydrophilic, highly hydrated, noncharged soft polymer brushes, where polyethylene glycol-based polymers are considered the gold standard for polymer brushes to prevent bacterial adhesion). However, this ­grafting-functionalization passive technique has different disadvantages in the field of biomaterials for tissue engineering. Polymer brushes have to be in their fully hydrated state to prevent bacterial adhesion, and typically, polymer brushes themselves do not possess any antibacterial activity; hence, microorganisms adhered to biomaterials surface in a dry environment (before implantation) can freely attach and grow upon implantation causing infections [2, 23, 146]. On the other hand, this low surface energy, low modulus surface modifications also prevent nonspecifically the adhesion of other molecules such as proteins, as well as the adhesion of mammalian cells, preventing appropriate implant-tissue interactions and consequently decreasing implant-tissue integration and tissue regeneration. Then, this strategy alone will not be further discussed in this chapter. Nevertheless, there are exciting approaches to

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combine the n­ onadhesiveness of polymer brushes with functionalization by grafting of antimicrobial agents with inherent antibacterial properties such as quaternary ammonium compounds (QAs), antibacterial peptides, antimicrobial proteins, or antibiotics [2, 23, 146].

4.2.3.2 Quaternary ammonium compounds Regarding nonleaching surface grafting-functionalization to add intrinsic, direct contact bactericidal activity to biomaterials, one of the most commonly used antibacterial agents are QAs, which are positively charged organic molecules containing four alkyl groups covalently attached to a central nitrogen atom with general formula R4N+. Different mechanisms have been proposed for QAs antibacterial properties; one of the most reported mechanisms is the penetration of the hydrophobic lipophilic alkyl chains in QAs on the bacterial cell membrane, which leads to denaturation of structural proteins and enzymes and cell content leakage, and eventually, to bacteria death [227–230]. Consequently, QAs antibacterial efficacy highly depends on the length of the substituting alkyl chains bound to the N central atom. This effect has been extensively studied in solution; however, it has also been corroborated that alkyl chain length also drives the balance of hydrophilic/hydrophobic properties of surfaces modified by QAs and consequently their antibacterial efficacy [231]. General trends indicate that longer alkyl chains display larger hydrophobic properties exerting higher antibacterial effects [232]. It has been observed that for chain length below 4 or above 18 carbon atoms, the antibacterial properties are almost entirely diminished and that QAs with a chain length of 8–18 carbon atoms display the best bactericidal activity for QAs bound to surfaces. Another antibacterial mechanism reported for QAs involves the effect of surface charge, which induces an ion exchange phenomenon that destabilizes the bacterial cell membrane homeostasis, mainly by exchange of divalent cations such as Ca2 + and Mg2 + [233–235]. Positively charged quaternary nitrogen in QAs (polycationic nature) is also strongly electrostatically attracted to the negatively charged phospholipids in the bacterial cell membrane, with the capacity of extracting these anionic lipids from bacterial cell membrane and the consequent membrane destabilization and leakage of bacterial cell material [236]. It has been reported that the surface charge density threshold to exert antibacterial effects is in the range of 1015–1013 N+ cm− 2, depending on bacteria strain and bacterial division activity [237, 238]. In an attempt to improve the local density of active QAs groups in close vicinity, the development of different QAs functionalized dendrimers has been reported. It has been hypothesized that amphiphilic dendrimers may exhibit antibacterial activity with minimal cytotoxicity [239]. There are already a considerable number of reviews that summarize the different bonding and synthesis strategies of QAs. Poverenov et al. described and analyzed different strategies for bounding QAs to biomaterials surfaces such as covalent linkage, sonication, hydrogen bonding, hydrophobic interactions, electrostatic interactions, and metal-coordination chemistry [229]. They mentioned covalent linkage as one of the most employed techniques for surface bounding of QAs due to the strength of the

Nanostructured biomaterials with antimicrobial activity for tissue engineering103

bonding, which allows resistant and stable surface functionalization. In addition, they categorized covalent linkage strategies in two general classes: ex situ and in situ. The first strategy, the most common approach for QAs covalent binding to surfaces, was denominated as “linking tail-activated head” considering that in this approach QAs already synthesized (ex situ synthesis of QAs) are bounded to surfaces. In this case, QAs must contain two main moieties, one linking moiety or tail which is used for covalent binding to the surface and one bioactive moiety or head which remains unaltered and exposes toward the environment to act as the main bactericidal contact point. In the in situ strategy, quaternization of tertiary amines with alkyl halides occur in situ on the surface after binding of tertiary amines on the biomaterial surface. Jiao et  al. also grouped QAs covalent linkage approach into ex situ and in situ strategies, but they named them as “grafting onto” and “grafting from” strategies, respectively [228]. Mentioning that the “grafting onto” approach, where QAs polymer molecules from solution are directly immobilized on suitably prepared surfaces, is an experimentally simpler strategy, but it produces functionalized surfaces with low grafting density; while the “grafting from” strategy, where surface-initiated polymerization occurs, allows better control of functionality, density, and thickness of the grafted QAs polymers. Addition of polymeric spacers to the surface before QAs grafting or grafting of long-chain polymers with QAs functionalities (synthesized either by direct copolymerization of monomers containing QAs functionalities or by postpolymerization/quaternization of reactive precursor polymers) are also widely used approaches, because they result in functionalized surfaces with improved QAs spatial flexibility. This enhances QAs bacterial cell penetration and the number of QAs-bacterium contact points, resulting in enhanced bactericide properties [240]. Then, QAs surface functionalization contact-killing approach provides an improved broad range and long-lasting antimicrobial activity, killing Gram-positive and Gram-negative bacteria by mechanisms of action mainly involving bacterial cell penetration and thus, these compounds are unlikely to develop antibiotic resistance [228]. Recent reviews [1, 2, 223] have extensively analyzed the different mechanisms of QAs surfaces. Thus, we aim in the following paragraphs to present a general overview of the latest 2  years works, that we believe report somehow innovative or different approaches for QAs surface functionalization of biomaterials. One of the disadvantages of grafting antibacterial compounds to biomaterial surfaces is that these coatings might be ultimately inactivated by protein adsorption (fouling) upon implantation [241]. Nevertheless, some works have demonstrated that the strong surface charge of QAs and their capacity to bound negatively charged lipids from bacterial cell membrane made them less susceptible to protein adsorption inactivation, in comparison to other grafting antibacterial compounds [23]. In 2016, Cavallaro et al. reported the development of a controlled surface gradient of QAs density (using glycidyltrimethylammonium chloride as QAs salt precursor) immobilized on plasma-polymerized allylamine surfaces to corroborate the threshold of surface density and minimum positive surface potential required to endow biomaterials with antiinfective properties [237]. Simultaneously, they analyzed the cytotoxicity of the QAs antibacterial coatings, determining that antibacterial surface density threshold was 4.18% NR4+ bonded nitrogen with a surface potential of + 120.4  mV and showing that adhesion of the

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culture medium constituents eliminated any cytotoxicity toward primary human fibroblasts. They stated that these findings do not necessarily imply QAs-based coatings as ineffective for indwelling biomaterials, because these coatings can still protect implant surface against viable bacteria attachment before implantation, and upon implantation they would become more biocompatible. Another possible disadvantage of materials endowed antiinfective properties by QAs grafting relays on the contact necessary to kill bacteria requisite because contact is usually restricted by Brownian movement in large amounts of volume. In this respect, Druvari et al. [242] reported an attractive strategy for improving antibacterial efficacy of QAs-based coatings. They developed dual-action (contact-killing and released-based antibacterial action) QAs-based biocidal surface modifications by producing environmentally friendly hexadecyltrimethylammonium styrenesulfonate and 4-vinylbenzyl dimethylhexadecylammonium chloride block copolymers, which exhibited control release and bactericidal contact action. A different approach to improve bacteria attraction to contact-killing surfaces overcoming the limitations due to restrictive Brownian motion of planktonic bacteria was reported by Jain et al. [243]. They proved the design of active/dynamic contact-killing surfaces with enhanced antimicrobial activities by using two soluble natural nonbiocidal chemoattractants (aspartate and glucose) to lure bacteria toward QAs-modified surfaces. Different works have functionalized QAs aiming to reduce their toxicity. Ergene and Palermo [244] developed interesting self-immolative polymers based on poly(benzyl ether)s end-capped with a silyl ether group and bearing pendant allyl side chains that were converted to polycations by photo-initiated thiol-ene radical addition using cysteamine HCl. Upon exposure to fluoride ions, these polymers spontaneously underwent depolymerization into their monomer components, which showed good antibacterial properties and low hemolytic toxicity [244]. Recently, some strategies to produce antimicrobials QAs-cellulose dressings for wound antiinfective protection and enhanced tissue regeneration have also been published. Żywicka et  al. [245] reported bacterial cellulose membranes modified with bioactive compounds, based on a long-chain dimer of C18 linoleic acid and ethylene diamine (EDA) that were gradually released from the bacterial cellulose membranes; the bioactive compound was synthesized by a simple coupling reaction. These membranes displayed both biocompatibility and antimicrobial activity [245]. Synthesis of QAs-modified cellulose-based materials was also reported by Littunen et  al. [246]. In that work, two types of QAs-modified nanofibrillated cellulose were prepared by aqueous synthesis using two different approaches, redox-initiated graft copolymerization and etherification with QAs. Materials synthesized by etherification exhibited more substitution and charge density and a broader antimicrobial spectrum in comparison to the copolymerized material, and neither of the two synthesized materials showed cytotoxicity against human cells. One of the requisites for surface grafting-functionalization should be mechanical strength to sustain surgical management. Contributing to this challenge, abrasion-­ resistant QAs-based coatings were developed by Gao et al. using dodecyl-alkylated quaternary ammonium and a benzophenone moiety that, under mild UV irradiation,

Nanostructured biomaterials with antimicrobial activity for tissue engineering105

generated within 1 min a densely cross-linked polymeric thin film which covalently attached to a variety of substrates containing CH bonds [247]. Coatings sustained high shear forces and abrasion and retained biocidal properties after exposure to mechanical stress and abrasion. To develop high durability and low toxicity QAs-based antimicrobial coatings, Li et al. synthesized a quaternary ammonium silane antimicrobial copolymer coating via a simple potentially scalable thermal-curing coating process. Developed copolymers were less toxic to different human cell lines and more durable than commercial antimicrobial QAs monomeric agent, dimethyloctadecyl[3-(trimethoxysilyl)propyl]ammonium chloride [248]. Hua and Odelius [249] addressed the environmental concerns regarding precursors and waste products from QAs synthesis processes. They reported the development of a UV-cross-linked polyhydroxyurethane film with antibacterial properties from the hydroxyurethane precursor (synthesized by aminolysis condensation) avoiding the use of isocyanates. Polyhydroxyurethane films were cross-linked under solvent-free processes (UV-triggered thiol-ene mechanism) proving to form hydrophilic, highly flexible hydrogels coatings with dual (contact and release) bactericidal properties. These coatings were achieved by controlled quaternization of the network’s tertiary-amine and methylation of thiol-ether functionality, resulting in QAs and sulfonium compounds [249]. Interesting combinatorial approaches have also been developed to produce enhanced antiinfective properties by synergistically combining QAs contact biocidal activity with other antibacterial compounds. Palantoken et al. [250] presented a simple method to fabricate dual antibacterial action (contact and release antibacterial activity) hydrogels by combining different concentrations of QAs and silver nitrate (a longknown antibacterial compound) into UV cross-linkable polyethyleneimine hydrogels that showed significant bacterial inhibition of adherent and planktonic bacteria. Cytotoxicity was dependent on QAs and AgNO3 concentration with some hydrogels showing to be biocompatible [250]. Recently, Pant et al. [251] generated an enhanced antimicrobial coating for potential medical devices by combining a nitric oxide releasing agent (S-nitroso-N-acetylpenicillamine) in CarboSil polymer coated with a QA compound. Antiinfective material developed showed significant contact and leaching antibacterial activity against both Gram-positive and Gram-negative bacteria [251].

4.2.3.3  N-halamines, N-chloramines After QAs, organic N-halamines are amongst the currently most studied contact-­killing antimicrobial agents [23, 252]. Similarly, to QAs, N-halamines are also low cytotoxicity, contact-killing compounds with broad-spectrum antibacterial activity. N-halamines are organic compounds containing one or more nitrogen-halogen covalent bonds, normally formed by imide, amide, or amine groups halogenation. Upon bacteria contact with N-halamines, halogen exchange reaction occurs causing the bacteria death. Organic N-halamines such as 1,3-dichloro-5,5-dimethylhydantoin and 3-­bromo-1chloro-5,5-dimethylhydantoin have been widely used as disinfectants [154]; however, N-chloramines are the most studied biocides grafting agents to endow antiinfective properties to biomaterials surfaces.

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N-Chloramines can be reduced to their precursor amines/amides/imides groups after reacting with biological components such as target proteins and thiol-containing enzymes. As a result, chlorine bleaching occurs, which can again convert precursor forms to N-chloramines. This mechanism has been described as the cause of one of the most interesting N-chloramines properties; that is, the rechargeable antibacterial activity of N-chloramine compounds grafted on surfaces. The antibacterial action of surface-bound N-chloramines has been speculated to involve the release of free chlorine (hypochlorous acid or hypochlorite anions) and contact-enabled transfer of active chlorine directly to biological receptors on bacterial cells, being contact-enabled chlorine transfer to bacterial cells the main bactericidal mechanism [253]. In 2017, Chen et al. reported the antiinfective modification of polymers by combining the antibacterial activity of QAs and N-chloramine compounds to produce QAs/N-chloramine polysiloxanes that were further interpenetrated into polyethylene terephthalate (PET) using, as an ecologically friendly solvent, supercritical carbon dioxide to form a nanometric biocidal layer [254]. Biocidal layer exhibited the synergistic effects of QAs and N-chloramines providing broad-spectrum antibacterial activity and rechargeability of lost chlorine. Recently, Ghanbar et al. reported the development of an enhanced biocide composite, based on N-chloramine and QAs, capable of reducing the risk of potential bacterial resistance sometimes associated with QAs [255]. In order to reduce the potential fouling of the biocide agents in biological media, the amide-based N-chloramine moiety of the composite was substituted with a secondary amine-based N-chloramine to improve its biocidal efficacy in highly protein media. New composite showed higher antibacterial efficacy than benzyldodecyldimethylammonium chloride in highly protein media. Li et al. also developed novel N-chloramine composites with improved antibacterial properties by synergistically combining the antibacterial activity of N-chloramine with other antibacterial compounds such as QAs, phosphonium, or pyridinium compounds [256, 257]. N-Chloramines have been also investigated in combination with other antibacterial compounds such as chitosan derivatives. Hoque et  al. synthesized N-(2hydroxypropyl)-3-trimethylammonium chitosan chlorides polymers and evaluated their antibacterial activity against multidrug-resistant bacteria and their in  vivo and in vitro cytotoxicity, showing that polymers displayed rapid microbicidal activity and very low in vitro cytotoxicity. In vivo, no skin inflammation was observed, and significant reduction of methicillin-resistant S. aureus load was corroborated [258].

4.2.3.4 Antimicrobial peptides One of the most actively studied surface functionalization strategies nowadays in the field of biomaterials is surface attachment of biomolecules to create biologically active biomaterials. Biomolecules can endow materials with different favorable properties for tissue regeneration such as selective protein and cell binding, enhanced cell proliferation, differentiation, and increasing extracellular matrix deposition. Nevertheless, biomolecules can also endow materials with antibacterial properties, while allowing cell attachment. Thus, grafting of antimicrobial proteins and peptides has been found as an

Nanostructured biomaterials with antimicrobial activity for tissue engineering107

important and growing field of research to develop antiinfective biomaterials surfaces. Antibacterial peptides are relatively easier to obtain in comparison with full-length proteins and can also present the specific cell- or protein-binding moiety of more complex macromolecules such as proteins; besides that, peptides (Mw ≈ 300 Da) are smaller and simpler structures than proteins (Mw > 150 kDa) [259]. Antimicrobial peptides are short (typically composed of 12–50 amino acids) cationic (average net charge of + 3.2), gen-encoded peptide antibiotics that are expressed in almost all living organisms including humans [260, 261]. Therefore, many antimicrobial peptides studied nowadays have been identified through screening of physiological fluids with antibacterial properties or from different animal and plants displaying natural antimicrobial mechanisms such as frogs, reptiles, marsupials, crustaceans, and many other marine organisms, insects, or fungi. Nevertheless, antimicrobial peptides have also been designed (synthetic or semisynthetic antimicrobial peptides) from phage, ribosome, and bacterial-displayed peptide libraries. Recently, in silico or knowledge-based approaches to design antibacterial peptides have been increasingly studied as simpler, faster, and lower-cost approaches than long-time consuming, screening, and synthesis process from combinatorial approaches [262–264]. Likewise QAs, antimicrobial peptides act on the electrostatic forces of bacterial cell membranes, inducing destabilization and physical damage to the membrane and the consequent bacterial cell material leakage and bacteria death. Several mechanisms have been proposed for antibacterial peptides action, explaining that the hydrophilic cationic moiety on antimicrobial peptides might initially bind to the negatively charged bacterial cell membrane, and after that, the hydrophobic residues might contribute for further antimicrobial peptide insertion into bacterial cell [259]. Other reports suggest that antimicrobial peptides also exert DNA damage, protein and cell wall synthesis inhibition, and prevention of the activity of DnaK molecular chaperone [260]. This combination of antimicrobial mechanisms makes the development of bacterial resistance more difficult. The ability of antimicrobial peptides to prevent biofilm formation has also been reported [265]. Some antimicrobial peptides of human origin seem to induce immunostimulatory and immune-suppressive effects, in addition to selectivity toward bacterial cells over human cells, mainly due to their action mechanisms and the different composition and structures of the bacterial cell and human cell membranes [261]. Antimicrobial peptides can be produced biologically using in vivo or in vitro expression platforms such as E. coli, or through chemical methods such as solid-phase peptide synthesis. Attachment of antimicrobial peptides to surfaces can be achieved through physical adsorption, direct covalent linking to properly activated surfaces, or through covalent tethering to polymer layers or brushes attached to surfaces. Upon attachment the small molecular size of antimicrobial peptides allows high-density layers; however, it is essential to control the orientation at which peptides are immobilized since their antibacterial mechanisms act through specific active motifs that have to be exposed toward the environment to effectively protect biomaterials surfaces from potential infection. Thus, chemical conjugation or click chemistry becomes very relevant for proper attachment of antimicrobial peptides on surfaces [1, 266]. In the last years, studies of antimicrobial peptides have acquired growing importance and excellent comprehensive reviews about their mechanisms of action, synthesis

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routes, and approaches for attachment on biomaterials have been published [267–269]. Diverse works have been recently published on wound dressings endowed with antiinfective properties by functionalization with antibacterial peptides coatings. Chouhan et al. synthesized a scaffold made of silkworm as bulk material and top-coated with the recombinant spider silk protein with a cell-binding motif, antimicrobial peptides, and a growth factor; these scaffolds demonstrated the formation of a keratinized epidermal layer in  vitro [270]. Costa et  al. showed a very interesting approach to genetically engineered protein-based polymers with multiple functionalities [271]. Synthesized polymers displayed high broad-spectrum antibacterial activity and no detectable cytotoxicity. Another silk fibroin-based membrane with antibacterial properties endowed through an antimicrobial peptide motif (Cys-KR12) originated from human cathelicidin peptide (LL37) was reported by Song et al., showing decreased expression of proinflammatory factors in  vitro and significant antibacterial activity [272]. Cassin et al. have recently reported a promising strategy to design antibacterial biomaterials for wound healing by fabricating detachable collagen/hyaluronic acid ­polyelectrolyte multilayers modified with the naturally occurring human antimicrobial peptide, LL-37. Multilayers presented no observable cytotoxicity and displayed  90%); meanwhile, an enhanced osteoblast proliferation and differentiation were also observed [298]. Laser surface texturing of Ti surfaces has also been used to alter the bacterial adhesion. The effect is complex, since lasers can simultaneously modify the surface topography, hydrophilicity, and composition. Cunha et  al. showed that adhesion of S. aureus was inhibited on laser-treated Ti surfaces because high-aspect-ratio nano features were obtained [299]. Plasma methods can also be applied to nonmetallic surfaces, such as it has been shown in the work of Rezaei et al. [300] that used an atmospheric-pressure oxygen plasma to modify the surface properties of PMMA. Depending on the specific plasma parameters, the ability to change the hydrophilicity and surface roughness was demonstrated. Therefore, by selecting appropriate conditions, an enhanced antibacterial activity of 50% above the control using E. coli was obtained. The authors showed that the optimum antibacterial activity and cellular (fibroblasts) response was obtained at the highest hydrophilicity, while an increase of surface roughness deteriorated both the antibacterial performance and cell adhesion. It was also shown that these responses are also mediated by modification in the functional groups of the surface, such as hydroxyl (OH) and ­carboxyl (COOH) groups, and their introduction might favor the biocompatibility [300].

4.3 Challenges to evaluate the antimicrobial properties of nanostructured biomaterials for tissue engineering In the previous sections, we have done a description of the information that has been gathered by different authors aiming to prevent implant device-related infections. However, not much has been said about the selection of the bacterial strains studied or the methods/techniques used to reach those conclusions.

4.3.1 Nature of the microbial infection The human microbiome project has identify more than 8  ×  106 microbial genes detected in the body habitats of healthy individuals, such as skin, urogenital tract,

Nanostructured biomaterials with antimicrobial activity for tissue engineering111

gastrointestinal system, and the oral cavity, each one forming unique and complex communities [75, 301, 302]. Such diversity demands to contemplate the surrounding microbial environment of the implant when designing strategies to prevent the ­biomaterial-related infections, and for this it is necessary to be aware of the ecological diversity and abundance of the human-associated microbial communities. Microorganisms use different strategies to colonize biomaterial surfaces; however, certain strains have been detected more frequently depending on the anatomic site. For example, catheter-related bloodstream infections are usually related to the presence of Gram-positive S. aureus and Staphylococcus epidermidis and Gram-negative P. aeruginosa [303, 304], while on infected urinary catheters higher quantities of­ E. coli, Enterococci, Klebsiella pneumoniae, Proteus mirabilis, and fungal species such as C. albicans have been detected [305, 306]. In the case of prosthetic joint infections, S. aureus and Propionibacterium acnes have been identified as the most prevalent microorganisms isolated from early and late infections [307, 308]. The oral cavity harbors the second most abundant microbiota after the gastrointestinal tract, and in dental implant infections a large number of Gram-negative anaerobic bacteria such as Prevotella sp., Fusobacterium sp., Capnocytophaga sp., Porphyromonas gingivalis, Aggregatibacter actinomycetemcomitans, and Treponema denticola have been associated to this condition [309, 310]. In addition, it has been reported an important presence of saccharolytic anaerobic Gram-positive rods such as Eubacterium sp., Filifactor alocis, and Slackia exigua as nonconventional periodontopathic bacteria associated to the active peri-implantitis lesions [311]. This complex ecosystem also harbors fungal species that have been associated to the infections in acrylic dentures [102]. Another aspect that should be considered is the route and the timing that microorganisms use to colonize the nanostructured biomaterial, causing perioperative, hematogenous, or contiguous early or late infections, since different microorganisms have been detected depending on the route of infection and the onset of symptoms after the implantation of the biomaterial [312, 313].

4.3.2 Limitations of the methods and techniques used to evaluate the growth of biofilms on biomaterials Different methods and techniques have been proposed to test the antimicrobial activity on nanostructured biomaterials, from traditional culture methods to those with a molecular genomic approach. Table  4.6 presents a summary of the methods and techniques employed to study the adhesion and biofilm formation process onto nanostructured biomaterials. These methods have the goal to achieve routine quality control and screening to evaluate the antimicrobial effect of biomaterials designed with this purpose. However, depending on the method used, different efficiencies can be obtained in the evaluation of the antimicrobial potential, and often results of different evaluation methods are conflicting [339].

Description: Inhibition zone diameters are obtained from this test when biomaterial samples with antibacterial activity are placed on agar plates and inoculated with test microorganisms. Interpretation should be made regarding the appropriate standards or controls Limitation: Zones of inhibition may depend on the rate of diffusion of the material through the agar. Interpretation criteria may vary, and potential clinical significance of the inhibitory halo size is made based on product literature This method is restricted for in vitro cultivable microorganisms and nanostructured biomaterials with antimicrobial release properties Associated techniques. Optical microscopy Description: Different concentrations of the nanoparticles could be tested preparing agar dilution plates to determine minimum inhibitory concentrations (MICs) Nanoparticles are added to molten agar, mixed, poured into a petri dish and allowed to solidify. Standardized microbial suspensions are inoculated onto the surface of each agar plate. After incubation colony forming units (CFUs) are visually counted Limitation. The inhibition may depend on the homogeneous dispersion of the nanoparticles through the agar This method is restricted for in vitro cultivable microorganisms and nanostructured biomaterials with an antimicrobial release or contact killing properties Associated Techniques. Optical microscopy Description: A stock solution with the nanoparticles is prepared, and serial microdilutions are performed usually in 96-well-plates. The same standardized microbial suspension is inoculated to each well, and after incubation, MICs can be obtained by different methods (i.e., direct culture on agar plates, absorbance at 600 nm, Live&dead XTT or MTT kits) Limitation: The antimicrobial effect may depend on the homogeneous dispersion of the nanoparticles through the broth during the incubation time. The risk of unseen contamination is high (no colonies could be observed) This method is restricted for in vitro cultivable microorganisms and nanostructured biomaterials with an antimicrobial release or contact killing properties Associated Techniques. Optical microscope, Dyes, Spectrophotometry

Agar zone inhibition (based on the antibiogram or KirbyBauer method [314])

Broth microdilution (for nanoparticles dispersed liquid media) (based on the CLSI methods for the evaluation of antimicrobial susceptibility tests [318])

Agar dispersion (for dry nanoparticles) (based on the CLSI methods for the evaluation of antimicrobial susceptibility tests [318])

Culture methods

Description and limitations

Name

Table 4.6  Methods to evaluate the antimicrobial activity of biomaterials

[140, 321]

[319, 320]

[315–317]

Related studies

Microscopy

Associated techniques

Direct culture on the nanostructured biomaterial

Optical microscopy (bright and dark field)

Description: It is used for qualitative observations. Microorganisms are usually stained, and some special staining methods allow the observation of bacterial surface structures such as capsules or appendages. Dark-field optical microscopy allowed the direct observation of living microorganisms without fixation and staining. The advances in image analysis make qualitative and semiquantitative analysis fast and more efficient Use to measure the radius of the inhibition zone and to count CFUs Limitation: Not adequate for direct culture on opaque biomaterials, the biomaterial has to be translucent

Description: This technique allowed to study bacterial adhesion and/or biofilm formation on biomaterial surfaces Biomaterial samples are exposed to a droplet of a bacterial suspension or are placed in cultured well plates under mild shaking conditions. After incubation, samples are carefully washing to remove all nonadhering bacteria. Then adhering bacteria is collected by sonication and serial dilutions of the sonicated samples are cultured on agar plates. Quantification is performed by direct counting of the CFUs on the agar plates. Variations of this technique have been proposed using different dyes or kits to directly measure the adhering bacteria Limitation: The washing of nonadherent bacteria should be done very carefully in order to avoid the unwanted removal of the attached bacteria. The nutrient concentrations in the culture broth should be carefully watched since exceeding the nutrients eliminates the need for planktonic bacteria to adhere to a surface and limited quantity of nutrients could interfere with the normal process of biofilm formation This method is restricted for in vitro cultivable microorganisms and can be used in nanostructured biomaterials with an antimicrobial release, contact killing, and nonadhesive properties Associated techniques. Optical, electron, confocal and atomic force microscopy, Dyes, Molecular techniques

(Continued)

[323, 324]

[295, 317, 322]

Name

Table 4.6  Continued

Scanning electron microscopy (SEM)

[326–328]

[325, 326]

Confocal scanning laser microscopy (CSLM)

Description: Is a three-dimensional technique to determine the microbial identity using fluorescent molecular probes (oligonucleotide probes). It allows the direct observation and quantification of microorganisms on both translucent and opaque surfaces. The examination of living fully hydrated biofilms in real time is possible and may be used to accurately assess the antibacterial properties of biofilm-resistant biomaterials since the physiological state (live vs. dead) can be distinguished by using specific dyes This technique offers several advantages, including the ability to control the depth of the field, elimination or reduction of background information away from the focal plane (that leads to image degradation), and the capability to collect serial optical sections from thick specimens Limitation: The bacteria need to be colored or labeled with oligonucleotide probes or specific fluorescent dyes for visualization. The technical requirements to obtain a requested image quality are expensive. Description: This is a well-established basic technique to observe the morphology of the microorganisms adhered on a material surface, the material surface morphology, and the relationships between the two. It is also used to observe the morphology of bacterial biofilms on surfaces Environmental SEM or low vacuum SEM do not require metal or carbon sputtering and is less prone to damaging the bacteria adhered on a surface or alter the surface characteristics of the specimen, therefore overcoming the drawbacks The chemical composition of samples can be determined by using energydispersive X-ray (EDX) for elements Limitation: It requires the specimen to be conductive (essentially “metal sputtered”). Cannot differentiate between live and dead bacterial cells. During sample preparation, the drying step is considered to cause noticeable cell shrinkage, and it exacerbates other undesirable outcomes, like damage and distortion of the biofilm

Related studies

Description and limitations

Dyes

XTT or MTT viability kits

Crystal violet

Atomic force microscopy (AFM)

Description: The AFM has proved to be useful in imaging the morphology of individual microbial cells and bacterial biofilm on solid surfaces, both in dry and hydrated states. It is used for mapping interaction forces at microbial surfaces. AFM is a noninvasive microscopic technique capable of imaging surfaces at nanometer resolutions and three-dimensional images at high resolution. Furthermore, as no stains or coatings are needed in this method, biofilms may be observed in situ. Preparation of sample surface is not required The sample does not need to be electrically conductive, no metallic coating of the specimen is required. Unlike the case with the SEM, no dehydration of the sample is required, and biofilms may be viewed in their hydrated state. The resolution of AFM is higher than that of the environmental SEM, where images can also be obtained with hydrated samples, and extracellular polymeric substances may not be imaged with clarity Limitation: The observation area is limited as compared with SEM. It cannot differentiate between live and dead bacterial cells. Imaging bacterial cells can be a time-consuming task Description: This assay is used for quantification of biofilm biomass. This dye stains both living and dead cells, by linking to negatively charged surface molecules and polysaccharides in the extracellular matrix It can be used with any culture method Limitation: This method provides no information about viability since this dye cannot differentiate between live and dead bacterial cells Description: These assays are used to test cell viability using any culture method. These colorimetric and rapid methods are based on cell metabolic activity. These assays were designed for use in eukaryotic cells lines and later have been applied for bacteria and fungi Limitation: An important risk of inaccuracies and misinterpretation of the test results could be obtained since the absence of metabolic activity does not necessarily mean bacterial death. This method is restricted for in vitro cultivable microorganisms. Care must be taken for chemical interactions with NPs

(Continued)

[326, 332–334]

[331]

[327, 329, 330]

Molecular techniques

Name

Table 4.6  Continued

PCR

DNA-DNA hybridization [336]

[335, 338]

[119, 337]

[327, 335]

Live/dead fluorescent stain

Description: This is a two-color fluorescence assay that allows distinguishing live and dead bacteria. A mixed population containing a range of bacterial type can be observed. This kit is composed of two nucleic acid-binding stains: SYTO 9 (green-fluorescent) and propidium iodide (red-fluorescent). These stains differ both in their spectral characteristics and in their ability to penetrate healthy bacterial cells. SYTO 9 stain generally labels all bacteria (those with intact membranes and those with damaged membranes); in contrast, propidium iodide penetrates only bacteria with damaged membranes, causing a reduction in the SYTO 9 stain fluorescence when both dyes are present It can be used with direct culture method Limitation: Depending on the experiment, the proportions of the two dyes must be adjusted for optimal live or dead discrimination; otherwise inaccuracies by direct observation and quantitation could be obtained Description: This is a microbial identification technique that allows to identify up to 40 different bacterial species at the same time and. Is useful to detect quantitative changes in the microbial composition of even noncultivable species that can colonize different biomaterial surfaces It might not be suitable for broth microdilution culture method using NPs Limitation: The microbial detection and quantification are limited to the available DNA probes Description: This is a very sensitive method for detection of specific species. Is also useful to detect changes in the microbial composition of even noncultivable species that can colonize different biomaterial surfaces. Quantification is also possible using RT-PCR It might not be suitable for broth microdilution culture method using NPs Limitation: The microbial detection is limited to the primer’s specificity.

Related studies

Description and limitations

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The sensitivity and reproducibility of a technique are crucial to determine a reliable in vitro evaluation method for biomaterials with antimicrobial properties. A remarkable number of international organizations (i.e., AATCC: American Association of Textile Chemists and Colorists, ASTM: American Society for Testing and Materials, EN: European Standard, ISO: International Standard Organization, JIS: Japanese Industrial Standard, SN: Schweizerische Normen Vereinigung, among others) have developed a variety of standard test, with the purpose of decreasing the variations in the results when materials with antimicrobial properties are tested [340]. Four critical factors have been identified influencing the outcome of antibacterial testing material; (1) the incubation time, (2) the bacterial starting concentration, (3) the physiological state of bacteria (stationary or exponential phase of growth), and (4) the nutrient concentration [341]. In addition, topics such as the definition of the antimicrobial activity and efficacy of a biomaterial are still debatable. According to ISO 20743, the antibacterial activity is based in the materials capacity to prevent or mitigate the bacterial growth, reduce their number, or directly kill the bacteria [342]. While the JIS Z 2801 defines as antimicrobial “the condition inhibiting the growth of bacteria on the surface of products” [343], making no distinction between killing and growth inhibition in their definition. Another critical aspect to consider is the choice of the correct or representative microorganism to test the antimicrobial properties of a nanostructured biomaterial. As previously mentioned, microorganisms exhibited specific tropism to human tissues or affinities for materials with specific chemical composition; it would be useless to test the sensitivity of a microorganism that will be never in contact with the nanomaterial once placed in a certain human tissue.

4.4 Conclusions and future trends in the development of antibacterial strategies Tissue engineering is an evolving field with the clear purpose to repair tissue functions to achieve positive health outcomes in the population. As an interdisciplinary new subject, it demands the close collaboration between biology, materials, and medical sciences, which is a challenge itself. The emergence of nanotechnology has provided this community with new strategies to solve the many demanding challenges associated with the wide range applications of tissue engineering such as soft or hard tissue replacement. In this chapter, we have briefly discussed the antimicrobial strategies required to prevent the device-­ associated infections using nanotechnology approaches but preserving the functional activity of the implant, device, or scaffold. The overview presented shows that the current trend is the development of complex strategies which combine diverse antimicrobial mechanisms aiming to optimize long-lasting and broad-range antimicrobial activity. This is necessary since the colonization of microorganisms and their capacity to form biofilms on biomaterial surfaces could affect the desirable host-cell interactions with the biomaterial, causing lack of integration, failure, and severe health damage in the worst cases.

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From the different factors influencing the development of biofilms, the p­ hysicochemical-surface properties have been the most studied and, therefore, a wide range of surface modification strategies have been tested. However, a profound examination of the complex interaction of the microorganisms with biomaterials makes evident that some important factors have to be further considered for designing the antimicrobial stratagems. Some of these considerations are listed here: ●

●

●

●

●

The most common approach is to evaluate the effectiveness of the antimicrobial strategy to prevent bacterial adhesion (short-time culture studies using planktonic cells), but less has been done to evaluate the biofilm growth (days of culture). This particularly important since it has been clearly stated that different phenotypes are displayed by bacteria when they are in the planktonic state or growing in biofilms. Therefore, the antimicrobial features of a biomaterial should be examined using both phenotypes. One of the most critical issues of antibacterial surface modifications is that modifications have to resist the surgical manipulation (mechanical stress) to which biomaterials are subjected to upon surgery to be introduced and implanted in the body. This can be a challenge to face, mainly in the case of grafting-modifications since the strength of the bonding of the antibacterial compounds or molecules to the biomaterial surface might not be enough to resist mechanical manipulation and those the surface modification might fail during surgery. It is important to remark that throughout the different stages (pre-, peri-, and postoperatively) of the surgical intervention, microorganisms can be encountered. However, the most common source of infection comes from the peri-operatively stage when sterile (preoperatively stage) biomaterials become into direct contact with the hospital environment, wounds, physiological fluids, etc. The selection of the bacterial strains to test the antimicrobial properties of a biomaterial has not usually been made considering the specific environment of their implantation, nor the stage of implant surgical intervention. Another one of the biggest challenges facing the development of biomaterials with antibacterial properties is the cell target of the antibacterial strategies. Most modifications performed to endow biomaterials with antibacterial properties aim to affect bacteria through unspecific mechanisms that can affect either eukaryotic or prokaryotic cells. For example, (a) binding compounds to cells through the charge of the cellular wall (which might be similar for eukaryotic and bacterial cells); (b) mechanically disturbing the cell wall (e.g., abrasive effect of NPs); and (c) entering the cell wall or membrane through ionic transport channels. Then, more specific antibacterial strategies have to be developed which target specific mechanisms of bacterial cells but not of mammalian cells; thus, biocompatibility of biomaterials can be fully preserved upon antibacterial modifications. Sometimes, there is a lack of rationality in the use of the methods and techniques, for example, the inhibition zone method cannot be used when the antibacterial agent/feature does not diffuse into the agar. It is only valid if release and diffusion of ionic species is the action mechanism; therefore, a careful selection of the techniques used to evaluate the antibacterial properties of a biomaterial should be done.

The novel antimicrobial strategies used to prevent device-associated infections using nanotechnology approaches, attempt to resolve a very complex problem that can occur once a biomaterial is placed in some part of the human body. The development and improvement of the nanostructured biomaterials with antimicrobial activity should be based on the rationale of the biological, medical, and materials sciences. With the appropriate knowledge and combination of the best of the three worlds, the success of the biomaterial used for tissue engineering will be more predictable and achievable.

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Molecularly imprinted polymers for selective recognition in regenerative medicine

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Ortensia Ilaria Parisia,b, Mariarosa Ruffoa,b, Francesco Puocia,b a Department of Pharmacy, Health and Nutritional Sciences, University of Calabria, Rende, Italy, bMacrofarm s.r.l., c/o Department of Pharmacy, Health and Nutrition Sciences, University of Calabria, Rende, Italy

5.1 Introduction Molecular imprinting (MI) represents a very interesting and powerful technology for the synthesis of polymeric materials, in the form of monoliths, micro- and nanoparticles, membranes, or gels, able to recognize in a selective way a predetermined analyte called template. This versatile technology is based on the introduction of the template during the polymerization process and its subsequent removal after the reaction. In a typical imprinting process, indeed, the template is reversibly bound to the chosen functional monomers, via covalent or noncovalent way, which is then polymerized in the presence of a cross-linking agent and an appropriate porogenic solvent. After the polymerization has taken place, the target analyte is removed from the obtained polymeric matrix leaving selective binding cavities complementary to the template in terms of chemical functionalities orientation but also of dimension and shape. Therefore, the resulting molecularly imprinted polymer (MIP) is a cross-linked three-dimensional network with a significant recognition function due to a chemical memory for the analyte of interest, which is selectively bind also in the presence of structural homologous and enantiomers. MIPs are attracting considerable interest not only for their recognition properties but also for other characteristics including a significant stability in an extensive variety of conditions, such as temperature, pressure, pH, and organic solvents. MIPs, indeed, can be used in aggressive media. Moreover, MIPs are cost effective and easily prepared materials for molecular recognition and can be regenerated and reused without any loss in their activity allowing a long-term employment. However, the optimization of the synthetic procedure requires relevant efforts in the aim to obtain a polymer with the desired properties. Due to their high sensitivity, stability, reduced costs of production, ease of preparation, and reproducibility, MIPs find potential applications in a wide range of fields including analytical and separation sciences [1–3], catalysis [4], chemo/biosensors [5], and artificial antibodies [6] development and drug delivery [7]. In the beginning, MIPs were designed for the separation and quantification of small molecules; then, their use in environmental cleanup; toxin sequestration from systems, Nanostructured Biomaterials for Regenerative Medicine. https://doi.org/10.1016/B978-0-08-102594-9.00005-X © 2020 Elsevier Ltd. All rights reserved.

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such as the gastrointestinal tract and for the controlled release of therapeutic agents was also investigated. Some soil bacteria are able to collect iron from the soil through specific enzymes, called siderophores, which chelate and release the metal ion for its internalization at the cell wall. In a research study [8], MIPs able to bind macrocyclic metal complexes were synthesized using Ni(II), N,N′,N″,N‴-tetra(2-carbamoylethyl) cyclam and acrylamide as a model environmental metal ion, ligand, and functional monomer, respectively. Due to the formation of noncovalent interactions, such as hydrogen bonds with acrylamide, N,N′,N″,N‴-tetra(2-carbamoylethyl) cyclam-Ni(II) complex was bound into the imprinted cavities highlighting the good recognition properties of the synthesized polymers in both acetonitrile and water. MIPs thermal stability, which allows sterilization processes, and their insolubility, which results in the impossibility of being absorbed, contributes to the safety of these materials that can find application also in toxin sequestration. Deoxycholic acid (DCA) is a toxin produced by bacterial flora and it is involved in heart disease, colon and esophagus cancer, liver cholestasis, and gallstones [9]. DCA imprinted polymers were prepared starting from N,N′-diethyl(4-vinylphenyl)amidine (DEVPA) and successfully used in clinical studies for the toxin selective removal from the gastrointestinal tract [10]. The use of MIPs as chiral stationary phases in high-performance liquid chromatography, capillary electrochromatography, and supercritical fluid chromatography for separation procedures have been also explored [11]. Enantiomers can be characterized by different properties and this represents a crucial point in the development of new drugs, which need to be enantiomerically pure. Therefore, MI represents a promising technology for the synthesis of polymeric matrices, of research and commercial interest, able to resolve a racemic compound as they are prepared in the presence of one of the enantiomers [9]. In the last decades, catalysis based on the use of MIPs has also attracted significant attention. MIPs, indeed, as synthetic materials present some advantages, including higher stability and low-cost and easy preparation compared to the biological counterpart, which make these polymers more suitable for industrial application. An example of MIP-assisted equilibrium shifting in favor of a product is the synthesis of aspartame starting from l-phenylalanine methyl ester and Z-l-aspartic acid. In this reaction, the aspartame imprinted polymer removes the obtained product pushing the reaction equilibrium to the right side [12]. Immunoassays, such as ELISA, are usually employed for detection and quantification of a wide range of compounds. However, due to their biological nature, they exhibit a restricted stability and an expensive production procedure. The MIPs ability to selectively bind a target analyte allows them to be used for the development of high sensitivity assays overcoming these drawbacks [13, 14]. This ability was also addressed to the production of label-free molecular sensors including luminescence MIPs-based sensors consisting of polymers containing fluorescent functionalities able to respond to the binding process throughout fluorescence changes [15], thermometric MIPs sensors based on the quantification of the heat developed during the template binding [16], and electrochemical sensors obtained by

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combining MIPs with conductive components such as metals, graphene or carbon nanotubes [17–19]. MI technology was also used to produce plastic antibodies that specifically recognize and bind targets, such as proteins, peptides, or their fragments, with a high affinity comparable to that of biological antibodies. The first efforts in this field were carried out by Mosbach and coworkers that synthesized MIPs for theophylline and diazepam detection in human serum developing a novel radiolabeled ligand-binding assay as an alternative to conventional ones [20]. Currently, the use of MIPs as drug delivery systems (DDSs) is attracting a growing interest due to their capability to control the delivery of the therapeutic agent used as an imprint molecule during the polymerization procedure. This results in a prolonged drug release, thus, reducing side effects and improving patient compliance [21, 22]. Moreover, MIPs can be used as excipients to develop targeted DDSs employing cell receptors or molecules expressed at the pathological site as templates. The targeting of MIPs particles can be also reached using magnetic composite systems. A research study describes the preparation of a novel drug device for targeted cancer therapy combining the controlled release ability of MIPs with magnetic responding properties of magnetite [23]. However, an open challenge consists of the optimization of polymeric materials for this application increasing their biocompatibility and biodegradability. Another big challenge is the development of smart scaffolds for application in regenerative medicine and, therefore, able to guide cell proliferation, differentiation, and adhesion. The main purpose of tissue engineering (TE) is the regeneration and/ or replacement of cells, tissues, and organs to enable them to perform their normal functions. In this context, MIPs can act in two different ways. First, MIPs are able to release therapeutic agents for tissue repair and regeneration in a prolonged manner, which is important as these processes take time. Moreover, these polymers can be used for the development of scaffolds able to recognize and bind biomolecules present in the bloodstream or in situ promoting cell recruitment, adhesion, growth, and/or differentiation to the site of injury [24]. Based on this context, the present chapter is devoted to MI technology focusing particular attention on MIPs design and synthesis and their application in regenerative medicine.

5.2 Natural body response to injury and regenerative medicine Regenerative medicine represents a promising and relatively new branch of medicine aimed to the development of therapies and treatments for regeneration and/or replacement of damaged cells, tissues, and organs in order to restore their normal structure and, therefore, function. Conventional clinical approaches often act by treating the symptoms, while regenerative medicine provides different strategies including cellular therapies, the use of medical devices, the implantation of artificial organs and TE. This last one involves biocompatible smart scaffolds able to attract cells to the

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injured site affecting their proliferation and differentiation in the aim to promote tissue regeneration. Regenerative medicine is based on the natural healing process and the amplification of the body’s innate response. Tissue repair and regeneration processes, indeed, occur after the onset of a lesion, which can result from a trauma or a pathological condition, and involve signaling molecular and cellular sequential events. The wound repair process can be divided into three main phases such as inflammatory, proliferative, and remodeling stages (Fig. 5.1) [25, 26]. In particular, inflammation plays a crucial role in effective wound healing, Successful tissue repair after injury, indeed, requires the resolution of the inflammatory process but, at the same time, inflammation represents a prerequisite to scarring. After the injury, the development of tissue edema is observed and, during the inflammatory stage, thrombocytes, platelets, and leukocytes aggregate in a fibrin network, which forms a barrier against microorganisms and a temporary matrix for cell migration. Then, a quick activation of immune cells, secreting chemokines, and cytokines may also occur. In response to the activation of the complement system, platelet degranulation, and bacterial degradation products, neutrophils are also activated and recruited. These cells play a key role in the cleanup of the injured tissue due to the expression of several pro-inflammatory cytokines and antimicrobial agents. After 48 h from the onset of the lesion, monocytes are differentiated into macrophages, which are activated through chemokine signaling and support neutrophils in phagocytosis. Furthermore, macrophages release cytokines and pro-angiogenic, inflammatory, and fibrogenic factors attracting other inflammatory cells to the injured site. These cells also produce potent vasodilators such as prostaglandins with the activation of endothelial cells and the production of other cytokines and growth factors able to stimulate the formation of granulation tissue [27–29]. The proliferative phase of a healing process involves fibroplasia, contraction, angiogenesis, and reepithelialization by keratinocytes [25]. Fibroplasia starts with the formation of granulation tissue, which is due to three main events such as the increase in fibroblasts proliferation, the collagenous biosynthesis, and the production of chemotactic factors by fibroblasts that are responsible for the formation of a

Fig. 5.1  The three stages involved in the wound healing process.

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three-­dimensional extracellular network of connective tissue. Wound’s contraction is caused by the action of myofibroblasts that contract the lesion’s borders toward the inner portion [30] while a suitable supply of nutrients and oxygen, which is essential in wound healing, is guaranteed by angiogenesis. This process consists of the formation of new blood vessels and occurs in the extracellular matrix (ECM) of the wound through migration and mitogenic stimulation of the endothelial cells [29]. Finally, the reepithelialization occurs by keratinocytes, which proliferate and migrate from the borders of the wound in order to close it. The remodeling stage represents the third and last phase of the wound healing process. During this stage, the granulation tissue is remodeled and the majority of inflammatory cells, fibroblasts, and blood vessels are removed by apoptosis or other mechanisms leading to the formation of scar tissue characterized by a limited number of cells. Moreover, type III collagen undergoes degradation, while the synthesis of type I collagen is intensified. Therefore, the two main body responses to injury are scarring and regeneration. The first one is the most frequent response and consists of a production of a fibrous tissue in the aim to stop the bleeding and avoid microbial infection at the site of injury. In this case, the overall cell, tissue, or organ function is preserved. The second one is less frequent and takes place at embryonic stages and below a critical size injury [31]. Regeneration, indeed, allows tissue and organs to completely recover their structure and thus their function. Wound healing involves cell-cell and cell-matrix interactions, which occur through both mechanical and biochemical signals that affect cell migration, adhesion, proliferation, and differentiation. The ECM plays a key role in this process providing mechanical support to cells and modulating cell behavior thanks to the growth factors action. In this context, the development of smart scaffolds based on MIPs, thus characterized by a high recognition ability toward target molecules or cells attract significant interest in the field of regenerative medicine and TE.

5.3 MI technology As reported above, MI represents a broadly explored technology that finds application in several fields including chromatography, immunoassays, chemical sensing, antibody mimicking, artificial enzymes, catalysis processes, and drug delivery. Natural processes, such as the antigen-antibody interaction and the enzyme catalysis, inspired this technology, which is based on giving a molecular memory to polymeric materials [24]. This synthetic strategy, indeed, leads to the production of MIPs characterized by selective recognition properties and, therefore, by an improved specificity toward a target analyte, such as a molecule, an ion, a complex or a cell. The first scientific remark concerning MI is attributed to M.V. Polyakov in 1931 [32]. The scientist reported a new synthetic procedure for silica polymers, which were prepared in the presence of an additive, and observed that the obtained silica particles were able to selectively absorb such additive molecule. The selectivity was described as a template effect involving changes in the silica pore structure that reflect the nature of the additive compound.

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In 1949, Dickey described the polymerization of sodium silicate in the presence of four different alkyl orange dyes, which were used as guidable templates [33]. In the performed rebinding studies, the prepared silica gels exhibited a good selectivity for the dye template present during the polymerization, which was bind in preference to the other dyes. Based on these results, Dickey understood that imprinted cavities, which fit the template molecule, were formed during the polymerization process and left by the dye when removed from the silica material. Afterward, different research groups adopted Dickey’s method in the aim to prepare selective silica-based adsorbents. In 1972, MI technology was applied for the synthesis of organic polymers by Wulff and Sarhan [34]. In this work, the monosaccharide phenyl-α-d-mannopyranoside was employed as a template and esterified with 4-vinylphenylboronic acid. Then, it was copolymerized with styrene/divinylbenzene or ethylene glycol dimethacrylate (EGDMA) obtaining the imprinted polymeric material. This study represents a milestone in the history of MI and reports on porous and swellable polymeric systems characterized by an enzyme-like behavior and able to resolve racemic mixtures due to the presence of selective binding sites. The described method involved the formation of covalent interactions between templates and monomers, thus, it represents the starting point for the development of the covalent approach for MIPs synthesis. In 1981, Arshady and Mosbach introduced the noncovalent approach, which does not require the template chemical attachment to the monomeric units to produce an imprint of the target analyte within the polymeric network [35]. The scientists called the developed strategy “host-guest polymerization” and described it as a simple mixing procedure in which a monomer mixture containing cross-linking units is polymerized in the presence of the target substrate. During the process, a host-guest relationship is formed leading to the development of an imprint of the guest molecule within the polymeric material. After the polymerization has taken place, a simple washing step allows to extract the template preserving the recognition cavities. This approach is based on the noncovalent self-assembly between template and monomers and, at the present, it represents the most widely employed method. Most of the interactions in biological systems, indeed, involve noncovalent forces.

5.4 Design and synthesis of MIPs MI technology allows to insert a permanent molecular memory for a given analyte into a polymeric material improving its specificity, loading capacity, and release control [24]. MIPs are synthesized in the presence of the chosen template by polymerization of suitable functional monomers and cross-linking agents to introduce the selective binding cavities, which are template-complementary in size, shape, and chemical functionalities, into the obtained three-dimensional polymeric matrix. The synthetic reaction involves four main steps. The selected target analyte and functional monomers are dissolved into a suitable porogenic solvent. Then, template and monomers self-assemble by covalent or noncovalent interactions leading to the

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Fig. 5.2  A schematic representation of the molecular imprinting process.

formation of a prepolymerization complex, which is stabilized by cross-linking within the resulting polymer. Finally, the template is extracted by performing several washing steps, using organic or inorganic solvents, or by enzymatic digestion obtaining a highly cross-linked polymeric material characterized by the presence of template-­ specific cavities (Fig. 5.2). The imprint analyte directs the monomers molecular orientation, while the cross-linking guarantees polymer rigidity that “freezes” the three-dimensional architecture of the recognition cavities after the template removal. The achieved MIPs exhibit relevant molecular recognition capabilities for the template, also in the presence of a mixture of other closely related molecules, durability, robustness, and endure reuse without significant loss in performance. The imprinting efficacy of a MIP can be investigated by performing binding studies in which amounts of the imprinted and the correspondent nonimprinted material (nonimprinted polymer, NIP), used as a control are incubated with template standard solutions. Then, three key parameters can be evaluated. The first one is the binding capacity (Qe, mol/g), which represents the amount of adsorbed template per weight of polymer at the equilibrium and it is calculated according to the following equation:

( Ci − C e ) × V

(1) m where Ci and Ce (mol/L) are the initial and equilibrium template concentrations in solution, respectively, V (L) is the volume of the solution and m (g) is the weight of the polymer. The second parameter is the imprinting efficiency (α), which corresponds to the ratio between MIP and NIP adsorption capacities for the target analyte. Finally, the selectivity coefficient (ɛ) consists of the ratio between the adsorption capacities of the template and a structural analog observed for the imprinted polymer. Qe =

5.4.1 Covalent, noncovalent and semicovalent approaches Three main approaches can be adopted for the synthesis of MIPs according to the nature of the interactions between the template and the functional monomers during the prepolymerization and the rebinding steps (Fig. 5.3). In all the cases, the chosen functional monomers are polymerized in the presence of the target molecule.

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Fig. 5.3  Template-monomers interactions involved in covalent, noncovalent, and semicovalent imprint approaches.

In the covalent approach, the template and one or more functional monomers are covalently bound during the prepolymerization step and the same reversible covalent interactions are reformed also during the rebinding phase. After the polymerization process in the presence of a cross-linker, the target analyte is extracted from the synthesized polymer by chemical cleavage of the covalent bonds through a simple procedure. Wulff and coworkers introduced this imprinting strategy synthesizing an imprinted polymer using specific sugar or amino acid derivatives containing a polymerizable function, such as vinylphenylboronate [36, 37]. After the polymerization reaction, the sugar moiety was hydrolyzed leading to a polymeric matrix characterized by a significant selectivity toward the template. This imprinting approach allows to obtain a more stable prepolymerization complex due to the covalent nature of the formed template-monomer stoichiometric interactions. This results in a more homogeneous binding cavities distribution and in a higher molecular recognition, which reduces the presence of nonspecific binding sites improving the polymer performance. On the other hand, this strategy presents some limitations such as the slow kinetics of covalent bonds formation and cleavage, which are processes that should take place under mild reaction conditions. Furthermore, the number of potential suitable templates and functional monomers is restricted. Given these characteristics, the covalent approach represents an impractical method for different applications, such as chromatography, that require faster dissociation and rebinding kinetics. The noncovalent approach, which was introduced by Arshady and Mosbach [35], involves the formation of relatively weak noncovalent interactions, such as hydrogen bonds, van der Waals forces, ion pair, and dipole-dipole interactions, between the template and the functional monomers during both the polymerization process and the rebinding step. This imprinting technique represents the most extensively used method for the production of MIPs due to several reasons such as the simple experimental protocol for the prepolymerization complex formation, the easy template removal and the great variety of potential template molecules, including biological compounds characterized by the presence of carboxyl, keto, or amino groups, and suitable functional monomers. These last ones include methacrylic acid, which is the most widely used, acrylic acid, 4-vinylbenzoic acid, 2-acrylamido-2-methyl-1-propanesulfonic acid, 4-vinylpyridine, acrylamide, and 2-hydroxyethyl methacrylate.

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In order to increase MIPs selectivity, a mixture of several monomers can be also used. However, templates with a single interacting group lead to poor recognition abilities compared to target analytes characterized by more points of interaction. In the noncovalent approach, the formation of the prepolymerization complex involves an equilibrium process in which a high amount of functional monomer results in a shift of the equilibrium toward the formation of the complex itself. On the contrary, an excess of monomer leads to the formation of nonselective binding sites within the polymeric matrix. Furthermore, the formation of template-monomer interactions is stabilized under hydrophobic environments, while polar environments interfere with this process. Due to its characteristics, the noncovalent approach is the most promising and widely used method for the preparation of MIPs. Finally, the semicovalent approach, introduced by Whitcombe et al. [38], allows to combine the advantages of the two imprinting methods described above due to the strict control of functional groups location and their uniform distribution, the mild operation conditions, and the reduced kinetic restriction during rebinding. This strategy involves covalent template-monomer interactions during the prepolymerization step, while noncovalent interactions are formed in the rebinding phase between the target molecule and the synthesized polymer.

5.4.2 Synthetic strategies and polymerization methods In the aim to prepare performant MIPs, the imprinting conditions have to be carefully selected to move the equilibrium toward the formation of the prepolymerization complex. Several factors, indeed, affect MIPs synthesis, such as nature, type, and concentration of template, functional monomers, cross-linking agents, initiators, and reaction solvent. The preparation of an effective MIP requires the optimization of the experimental conditions and, thus, of all these parameters to reach the best combination of rigidity and flexibility. Rigidity, indeed, allows to preserve the architecture of the recognition cavities, while flexibility provides fast release and rebinding kinetics. The template directs the orientation of the monomer functionalities and has to be chemically inert and stable under the adopted polymerization conditions. Small organic molecules are easily used as templates; on the contrary, the imprinting of larger compounds, including proteins and cells, represents a challenge in this field due to their poor stability under the reaction conditions and the difficult formation of well-defined binding cavities. The choice of suitable functional monomers (Fig. 5.4) plays a key role in the synthesis of effective MIPs. Monomers should be biocompatible and nontoxic and have to provide the functionalities involved in the formation of the interactions with the target analyte. Interaction strength, indeed, affects the selective recognition properties of the binding cavities. Moreover, the template to the functional monomer molar ratio can also influence the imprinting efficiency. In the covalent approach, the template determines the number of functional monomers that can be attached in a stoichiometric way; in the noncovalent approach, monomers are employed in excess compared to the number of template moles in the aim to shift the equilibrium to the formation of the prepolymerization complex. However, a surplus in the amount of monomer can lead to nonspecific binding sites (Fig. 5.4).

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Functional monomers O

O

O

OH

OH

NH2

Acrylic acid (AA)

Methacrylic acid (MAA)

O

O

Acrylamide (AAm)

OH

O

O N

2-Hydroxyethyl methacrylate (HEMA)

Methyl methacrylate (MMA)

4-Vinylpyridine (4-VP)

Crosslinking agents O

O N H

O

N H

N.N-Methylenebisacrylamide (MBAA)

Ethylene glycol dimethacrylate (EGDMA) O

O O

Initiator CN

O

N NC

O O Divinylbenzene (DVB)

N

2,2′-Azobisisobutyronitrile (AIBN)

Trimethylolpropane trimethacrylate (TRIM)

Fig. 5.4  Chemical structures of some functional monomers, cross-linking agents, and initiators.

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O

O

O

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Another important variable in MIPs synthesis is represented by the cross-linker, which carries out three main tasks. The chosen cross-linker, indeed, controls the polymer morphology, provides mechanical stability to the matrix and freezes monomer functionalities around the template molecule stabilizing the structure of the recognition cavities after the target removal. Therefore, a higher cross-linking promotes higher selectivity; on the other hand, an excessive degree of cross-linking leads to a difficult template extraction and negatively affects the rebinding properties. The most usually used cross-linkers include EGDMA, divinylbenzene (DVB), trimethylolpropane trimethacrylate (TRIM), and N,N′-methylenebisacrylamide (MBAA) (Fig. 5.4). Several initiators can be used as a source of radicals in free radical polymerizations for the synthesis of MIPs. Depending on the initiator nature, its decomposition can occur by heat, light or chemical/electrochemical means. Among the radical initiators, 2,2′-azobisisobutyronitrile (AIBN) (Fig.  5.4) represents one of the most commonly employed. This compound is decomposed by thermolysis or photolysis (UV) producing carbon-centered radicals [39]. Other examples of thermal initiators are water-­soluble inorganic compounds and organic peroxides. The first class includes ammonium and potassium persulfate, which can be used either alone or in combination with N,N,N′,N′-tetramethylenediamine (TEMED), while the second class includes benzoyl peroxide. In the presence of thermo-sensitive molecules, the photochemical decomposition of initiators, such as benzophenone and 2,2′-dimethoxy2-­phenylacetophenone, is preferred to avoid template degradation and to improve the prepolymerization complex stability [40]. High temperatures, indeed, have a negative impact on the prepolymerization complex stability. In the imprinting process, the porogen is responsible for the dissolution of all the reaction components, such as template, monomers, cross-linker, and initiator, and at the same time, affects the morphology and the porous structure of the synthesized polymer. Moreover, the nature and level of porogenic solvent determine the strength of the interactions in the noncovalent approach. In the aim to favor the formation of the template-functional monomer complex, indeed, aprotic and low polar organic solvents, such as chloroform, acetonitrile, and toluene, are preferred to stabilize hydrogen bonds. Several methods can be used for the synthesis of MIPs according to the final application of the prepared polymeric material. Due to the fast and simple experimental procedure, bulk polymerization represents the most widely used method. However, this synthetic strategy is characterized by several limitations. The obtained monolith, indeed, has to be crushed, ground, and sieved resulting in highly irregular particles in shape and size with diameters usually in the micrometer range. Moreover, these operations lead to the destruction of the recognition sites and to a decrease of the polymer yield [41]. Precipitation polymerization allows to prepare uniform and spherical imprinted polymeric particles in a one-pot reaction. Moreover, this synthetic technique provides an excellent control over the particle size. This methodology requires a higher amount of both template and porogen compared to the bulk polymerization and the polymer chains continue to grow individually by capturing newly formed oligomers and monomers from surrounding solution reaching a certain critical mass. Then, polymeric

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microgels precipitate due to their low solubility in the used solvent and, finally, the polymeric beads are recovered by filtration or centrifugation. This method is used to synthesize both micro- and nanospheres. In the suspension polymerization, droplets of the prepolymerization mixture are suspended in a continuous phase, such as water, mineral oil or perfluorocarbon, in the presence of a surfactant [42]. All the reaction components, indeed, are dissolved in a suitable organic solvent and the obtained solution is added to a larger volume of an immiscible solvent. The system is vigorously stirred in the aim to form droplets and, then, the polymerization is induced. This synthetic technique is characterized by some drawbacks. The presence of water and surfactants, indeed, interferes with the noncovalent interactions between template and functional monomers. Furthermore, the obtained polymeric particles are polydisperse in size. Emulsion polymerization allows to synthesize high-yield monodispersed polymeric particles, but it is also characterized by the presence of surfactants. This synthetic approach is effectively used for protein imprinting [43]. Finally, the multistep swelling method allows to prepare spherical imprinted particles monodisperse in size and shape by the stepwise swelling of seed particles, which are suspended in water [44]. Due to the addition of suitable organic solvents, the initial particles swell to the desired size and, then, all the components involved in the reaction are added to the solution inducing the polymerization. However, this method requires the use of an aqueous medium as a continuous phase, which could interfere with the formation of the template-monomer interactions resulting in a decrease in the selective recognition abilities.

5.5 MI and TE TE is an emerging field in which different scientists and research groups collaborate to allow an innovation in regenerative medicine. TE studies concern the investigation of macrolevel structures and of the combination of biomaterials, cells or tissues in the aim to improve or restore the damaged anatomic structure and function, or missing tissue. The purpose of this emerging field is to engineer the functional units of tissue and to control cellular environment, cell-molecular interactions and cell interactions by the construction of supercellular, cellular, subcellular, and nanoscale structures [45]. In the body, the ECM supports and maintains cell microenvironment while cellular processes, such as adhesion, migration, growth, and gene expression, are controlled by molecular mechanisms in which the substrates suitable for adhesion are recognized by cells through specific membrane receptors [46]. For example, some proteins (collagen, fibronectin, laminin), which are components of ECM, present particular peptide sequences that are recognized only by integrins, a family of trans-membrane proteins linked to the cytoskeleton on the cytoplasmatic side of the membrane. If these proteins were included in artificial substrates, they would become poorly effective because of the degradation phenomena following conformational changes in the protein structure. In the aim to overcome this problem

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and to improve cell adhesion and proliferation, it is possible to include short chains of oligopeptides on the artificial surface. A possible solution to this problem is represented by the formulation of biomaterial scaffolds that can maintain and regulate cell behavior [47]. In the TE applications, the development of new biomaterials plays a key role and the most used biomimetic materials must be able to direct new tissue formation, induce cellular responses, [48] mimic the interactions among cells and ECM proteins and allow specific interactions between surrounding tissues. To obtain the biomolecular recognition of materials by cells, there are two approach strategies: in the first design, biomaterials are equipped with a specific bioactivity by the incorporation of bioactive molecules (growth factors and plasmids DNA) and by subsequent release of these bioactive molecules, which modulate new tissue formation [49]. In the second approach, to allow specific interactions with cell receptors, biomaterials are modified in the surface and in the bulk by chemical or physical methods with ECM proteins or with short peptide sequences derived from them [50]. MI technology refers to a new approach for the formulation of synthetic biomaterials useful for cell adhesion and proliferation. This technology, in fact, allows a topographical patterning of a polymeric material with the introduction of recognition precise structural patterns with complementary binding sites that will allow recognizing specific receptors on the substrate [51]. MI technology is a valid strategy in the biomaterial field in general and in TE in particular. Indeed, the macromolecular matrices, obtained with this technology, have several advantages [51]: ●

●

●

stability in critical chemical and physical conditions; a duration of several years during which there is no reduction in their performance; high repeatability without any alteration to the memory.

The ability of scaffolds to recognize specific molecules or cells is fundamental in TE applications and, for this reason, MI technology has the potential to become a tool utilized in this field.

5.5.1 MIP-based scaffolds TE strategies provide to replace damaged tissue with a functional and regenerated tissue by mimicking and/or modulating the natural events of the wound healing process. TE strategies include several methods, which can be divided into either cell- or scaffold-based approaches. The first TE strategy allows the formation of new functional biomaterials with the implantation of progenitor cells [24]. This approach has several limits such as the difficulty to obtain in vitro cell expansion and differentiation. The second approach concerns the production of scaffolds with degradable biomaterials, which should be able to simulate the ECM function, maintain homeostasis, and provide signals to cells. To mime the natural cell environment in several varieties of tissue, such as skin and heart valves, decellularized extracellular matrices were widely used [52]. These matrices present several limits including the difficulty of process method and the possibility of contamination.

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Fig. 5.5  MIP-based scaffolds.

For this reason, it is necessary to develop an intelligent biomaterial/scaffold (Fig. 5.5), which is: ●

●

●

able to promote cell adhesion, proliferation, or migration in a biocompatible way; nontoxic; characterized by a biological activity, which depends on the inclusion of proteins or chemical functional groups able to interact with these molecules [53].

For the development of biocompatible, biodegradable, and intelligent scaffolds that are, moreover, able to improve cell adhesion and proliferation, MI technology is a valid new approach. With this technology, it is possible to guide and control the adhesion process by using polymeric materials, which will operate as “smart” systems for recognizing specific peptide sequences in ECM. In MI field, peptides are the most used template molecules because the use of entire proteins showed to be often unsuccessful. The large molecular dimensions of proteins complicate the polymer molecular recognition capacity and selectivity and, for this reason, stable short peptide sequences are used. These imprinted sequences are representative of an accessible fragment of several proteins of interest (i.e., collagen, fibronectin, laminin, vitronectin, etc.) and they are often being located in the receptor domains or in other parts directly involved in the molecular recognition process (epitopes) [54]. Below, different examples of MIPs useful for the preparation of scaffolds to be employed in TE are listed (Fig. 5.6). GRGDS peptide imprinted scaffold: In the research group of Rosellini, to realize MIPs able to recognize a pentapeptide segment of an exposed part of a fibronectin functional domain, the GRGDS (H-Gly-Arg-Gly-Asp-Ser-OH) peptide was used as template molecule [54]. The choice of the molecule to be recognized was justified in that fibronectin is the main ECM adhesive glycoprotein and it plays a fundamental role in cell adhesion, migration, and repair processes. Therefore, the potential application of the synthesized MIPs as functionalization structures in the development of bioactive scaffolds being able to guide and control cell behavior was also investigated. For this purpose, MIPs were employed to functionalize synthetic polymeric films by deposition on their surface and, then, the ability of the obtained MIP-modified films to improve cell adhesion and proliferation was confirmed by in vitro cell culture tests with C2C12 myoblasts.

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Fig. 5.6  Some examples of MIPs for the preparation of scaffolds to be used in tissue engineering.

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Peptide imprinted polymers for effective scaffolds development: In addition to the previously reported study, also the research group of Rechichi realized a polymeric material able to recognize the tripeptide Z-Thr-Ala-Ala-OMe segment of an exposed part of a fibronectin functional domain. This realized imprinted polymer allows to guide and control the adhesion process [47]. Therefore, the synthesized MIP could find a potential application in TE for the preparation of biocompatible, biodegradable, and effective scaffolds. MMP-9 imprinted scaffold: Cristallini and coworkers proposed MI as new nanotechnology for the preparation of advanced synthetic structures supporting cell adhesion and proliferation to prevent the expansion of postinfarct left ventricular remodeling. The cardiac ECM plays an important role in wound repair postmyocardial infarction and its degradation and remodeling is controlled by matrix metalloproteinase (MMPs) proteolytic system. The cardiac tissue repair processes are divided into three phases: inflammation, granulation tissue formation, and remodeling phase that can lead to a decline of heart contractile function. The MMP-9 enzyme is particularly present during the first and second phase in infarcted myocardium [55]. For this reason, in this study, the MI technology was used to obtain a novel biodegradable, biomimetic, and functionalized polymer scaffold that is able to selectively recognize a specific MMP at a particular time step. This study provided a new biodegradable multifunctional scaffold replicating cardiac ECM structure also reducing or preventing the left ventricle remodeling [56]. Proteins imprinted PMMA scaffolds: Imprinted polymethylmethacrylate (PMMA) scaffolds able to mimic the in vivo cellular microenvironment were developed for TE applications by combining MI with soft-lithography technology [57]. The scaffolds were prepared by using poly(dimethylsiloxane) (PDMS) microstructure molds functionalized with fluorescein isothiocyanate albumin (FITC-albumin) and TRITC-lectin as model proteins. Then, the functionalized PDMS molds were used to imprint PMMA scaffolds with macromolecules in the aim to form selective binding sites able to host adhesion proteins improving and guiding cell adhesion, growth, and proliferation.

5.5.2 Cell imprinting MI technology is also used for imprinting of macromolecules and of larger templates such as cells. In the aim to recognize a cell, MIP-mediated cell imprinting could be developed with a specific recognition site toward specific cell membrane molecules or toward the ensemble of the cell membrane and, so, it could be divided into two categories: the cell-membrane-MI and whole-cell-imprinting strategy [58].

5.5.2.1 Cell-membrane-MI strategy The cell-membrane-MI method refers to the imprinting of specific cell membrane molecules such as proteins, lipids, and glycans [59]. This advanced MI method could be used for the development of MIP-mediated cell recognition in cancer therapy and regenerative medicine.

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Regarding cell imaging, MIPs are very promising materials for tumor cell imaging because easy to functionalize with fluorescence or other photoactive molecules. According to this, nano-MIPs for cell membrane molecules were used for targeted cell imaging according to the recognition of specific glycans on the cell surfaces, such as glucuronic acid, sialic acid, mannose, and fucose [58]. In addition to cell imaging, nano-MIPs are also used for cancer cell targeting and for targeting surface antigens of red blood cells. In the first case, MIPs are used to target the epitope of the exposed proteins on the cell membranes [60], while, in the second example, MIPs with binding affinity to specific surface antigens offer the possibility of discriminating blood type [61]. Finally, as natural receptor mimics, MIP mediated cell adhesion was realized by imprinting a cell adhesive peptides or proteins that can bind specifically to the cell membrane receptors and, at the same time, allows indirect cell recognition and subsequent cell adhesion [58].

5.5.2.2 Whole cell imprinting strategy The whole cell imprinting strategy concerns the recognition of the cell membrane and it is suitable for the imprinting of bacteria and viruses but, at the same time, it could be used to imprint a mammalian cell, which commonly needs a pretreatment (Fig. 5.7). An example of imprinting mammalian cells was demonstrated, for the first time, in a study of Hayden and coworkers, in which they developed a QCM (quartz crystal microbalance) transducer surface imprinted by red blood cells for recognition of ABO blood grouping. Thanks to its selectivity, MIP technology offers a promising application for blood-group typing and, moreover, it could replace the expensive and traditional techniques [62]. In another study, Eersels and coworkers developed a novel biosensor able to recognize cancer cells and macrophages. In this novel device, a Heat transfer resistance (HTR) instrument is combined with a cell-imprinted polyurethane layer on aluminum chip and the detection depends on the change of HTR induced by binding the cells to the cell-imprinted surface [63]. The whole cell imprinting strategy, moreover, allows to direct stem cell differentiation and to regulate the growth of cells. In a recent research, Mahmoudi et al. developed cell imprinted substrates by using mature and dedifferentiated chondrocytes as templates in the aim to culture rabbit adipose-derived mesenchymal stem cells (ADSCs)

Fig. 5.7  Schematic representation of the whole cell imprinting strategy.

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for 1 week. The morphology of the cultured ADSCs on the different substrates was different and, in particular, ADSCs cultured on mature chondrocyte-­imprinted substrates expressed specific chondrocyte gene markers such as collagen type II and aggrecan, while ADSCs cultured on dedifferentiated chondrocytes-­imprinted substrates did not express collagen II and aggrecan. This result indicated that the type of cell-imprinted substrates modulated the morphology [64]. Bonakdar and coworkers developed cell-imprinted substrates by using different templates such as chondrocytes, tenocytes, and semifibroblasts and demonstrated that the cell-imprinted substrates were able to modulate the differentiation, redifferentiation, and trans-differentiation of various cells underline the potential application in regenerative medicine and cell-based diagnosis [65]. To program mammalian cell adhesion and growth, De Porter and coworkers proposed a cell-imprinting approach. The aim of this research was to develop a culture system as an alternative to the conventional high cost and multistep fabrication processes [66]. In the biomedical and TE technologies, the development of a defined pattern of mammalian cell adhesion and growth is an important component. The required approach in cell adhesion and growth is characterized by the conjugation of cell-­adhesive affinity reagent to a surface using expensive and/or specialized techniques such as photolithography, photoablation, 3D printing, and microfluidics [66]. To overcome the limits of these technologies, it is necessary to use a technique that allows to prepare substrates for programmed cell adhesion and growth that does not need specialized or expensive equipment. Several studies have demonstrated that the use of bacterial or virus cell-imprinted features supports the recognition of those cells from solution [67] and, similarly, erythrocyte stamps could be used to prepare polyurethane surfaces that recognize erythrocytes from solution [62]. The research group of De Porter successfully developed HeLa, HEK-293T (­epithelial-like cells) and MRC-9 (fibroblast-like cells) cells imprinted polyacrylamide hydrogels as substrates for mammalian cell adhesion and growth. This research demonstrated that cell-imprinting technology allows to obtain hydrogels or elastomer surfaces that support cell adhesion and growth. In another study, Jeon used MG63 osteoblast-like cells as a template in a poly (dimethylsiloxane) PDMS surface. With this study, this research group developed a biomimetic surface pattern model that can be used to evaluate the interface between various cells and substrates [68]. Cell imprinting is an emerging field in the early stages and more studies need to be performed to further explore its applications and influences on various cell functions.

5.6 Conclusions and future perspectives MI technology refers to a new approach for the formulation of synthetic biomaterials useful for cell adhesion and proliferation. This technology, in fact, allows a topographical patterning of a polymeric material with the introduction of precise structural patterns with complementary binding sites that will allow recognizing specific receptors on the substrate.

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MI technology is fundamental in TE applications because it allows to develop intelligent scaffolds, which are able to recognize specific peptide sequences in ECM and, thus, to guide and control the adhesion process. In MI technology, the most used template molecules are peptides because the use of entire proteins showed to be often unsuccessful. In this paragraph, all the studied peptide scaffolds aim to recognize specific peptide sequences and facilitate the cell adhesion, migration, and repair processes. MI technology is also used for imprinting of macromolecules and larger templates such as cells. This technology allows to overcome the limits of the most used technologies such as photolithography, photoablation, 3D printing, and microfluidics, which are expensive and require highly specialized techniques. MI technology, indeed, allows to prepare substrates for programmed cell adhesion and growth that does not need specialized or expensive equipment. Currently, cell imprinting is an emerging field and, to understand if this technology can replace the existing technologies, further investigations are needed.

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Further reading [69] G.  Pan, S.  Shinde, S.Y.  Yeung, M.  Jakštaitė, Q.  Li, A.G.  Wingren, B.  Sellergren, An ­epitope-imprinted biointerface with dynamic bioactivity for modulating cell-biomaterial interactions, Angew. Chem. Int. Ed. 56 (50) (2017) 15959–15963.

Polysaccharide-based hybrid materials for molecular release applications

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Ahmed Salamaa, Nadia Shukrya, Vincenzo Guarinob a Cellulose and Paper Department, National Research Centre, Giza, Egypt, bInstitute of Polymers, Composites and Biomaterials (IPCB), National Research Council of Italy, Naples, Italy

6.1 Introduction The term drug delivery system refers to a system that monitors the drug release and location in body where it should be released. Drug delivery systems can include excipients, such as inorganic material, lipid, or a polymer or their combination, used for the preparation of vesicles, micelles, nanocapsules, nanospheres, microspheres, microemulsions, films, or hydrogels [1]. The term excipient refers to any component other than the active substance intentionally added to support the chemical and physical properties as well as robustness of the pharmaceutical dosage form [2]. The therapeutic effects of any drug can be improved by constructing an effective target delivery system, that is, carriers and drugs, in which the carrier plays the role of target delivery. The carrier must possess, among other properties, target effect, strong adsorptive capability for drugs, enable drug transport to the effect-relevant sites and efficiency to release drugs at the proper site in a controlled manner. Drug delivery technologies can modify drug-release profile and improve drug efficacy and safety. The drug release can be by diffusion, degradation, swelling, and affinity-based mechanisms [3]. The common drug administration pathways include peroral, topical, transmucosal, and inhalation. However, peptides, protein, antibody, vaccine, and gene-based drugs cannot be delivered by these methods. These drugs are administrated through injection or microneedle arrays (Medical Dictionary). Controlled drug delivery is one which delivers the drug at a predetermined rate for locally or systemically for a specified period of time (Medical Dictionary). Since 1980 the progress of controlled release drug delivery systems focused on self-­regulated drug delivery systems, long-term depot formulations, and nanomaterials-based delivery systems [4]. The most important strategies of controlled release drug delivery systems are the maintenance of drug level within a desired range and to achieve more effective therapies. Controlled drug delivery occurs when a drug or an active substance, embedded within or combined with the excipient, is released in a predesigned manner, whether constant or cyclic, over a long period. Polymers are commonly used to modify the drug-release characteristics not usually offered by conventional dosage forms [5]. Nanostructured Biomaterials for Regenerative Medicine. https://doi.org/10.1016/B978-0-08-102594-9.00006-1 © 2020 Elsevier Ltd. All rights reserved.

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Sustained drug delivery is designed to slowly release a drug in the body over an extended period of time especially to sustain therapeutic level. In this context, sustained release dosage forms are designed to liberate a drug at a predetermined rate in order to maintain a constant drug concentration for a specific period of time with minimum side effects (Medical Dictionary). Recently, the improvement of sustained drug delivery is ascribed to the use of novel inorganic materials-based composite hydrogel systems of polysaccharides [6]. Extended release (ER) is a type of modified release dosages and can perform ~twofolds reduction in dosing frequency. The terms controlled release, prolonged release, sustained, or slow release (SR), long-acting release are synonymies with ER [5]. The therapeutic effects can be improved when the drug is delivered with high selectivity to the effect-relevant sites. Such selective delivery is called target treatment. Hence, in order to realize the target treatment and to reduce side effects in nontargeted tissues, it is necessary to construct an effective target delivery system, that is, carriers and drugs, in which the carriers play the role of target delivery. An ideal carrier for target drug delivery system should fulfill the following prerequisites for its function: the carrier itself has target effect, has strong adsorptive capability for drugs, enable drug transport to the effect-relevant sites, as well as the efficiency to release the drug at the proper site [7]. Moreover, the carrier must be able to bind, encapsulate, or adhere to drugs and other active materials, such as genes [1]. Nontoxicity, high drug upload and retention capacities, physicochemical stability, ability to protect drugs from degradation, preservation of the drug form at the target site, permeation, and suitable mechanical properties are essential requirements for an efficient drug delivery system. Versatility, biocompatibility, biodegradability, and low cost are also key factors for a successful delivery system [8]. It is also recommended that when the drug reaches the target site, it must maintain the adequate concentration level for a long period of time without going to its toxic and subtherapeutic concentration levels. Improved circulation half time, chemical stability in in  vivo conditions, and inertness toward enzymatic degradation are also of great importance. All these properties can help developing new techniques in drug delivery systems and alter the traditional method of use of drugs from injection to oral therapies [9]. Hydrogels represent excellent drug matrices owing to their outstanding water absorption and retention capacities, capability of protecting sensitive, and unstable drugs in harsh environments, for example, stomach [10]. Similarly, nanohydrogels exhibit high water content and are widely used in drug delivery systems because of their tunable, chemical, physical, and three-dimensional porous structure [11]; encapsulation of biomolecules in nanogels offer ideal candidates in nanomedicine [12]. Few ­nanomaterials-based products are promising in treatment of cancers, such as nanoparticles, liposomes, and polymer-drug conjugates and many have been applied in clinical research. These nanoformulations show several advantages over traditional therapeutics with respect to controlled release, targeted delivery, and therapeutic effects [13]. Moreover, nano-sized drug delivery systems have been recently applied for codelivery of antitumor drugs to overcome multidrug resistance in cancer chemotherapy. Recently, different drug resistance inhibitors have been applied as a promising technique to achieve the best therapeutic efficacy. Inspired by these findings, pronounced efforts

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have been carried out to prepare drug delivery systems that selectively target the cancerous organ with minimal damage to the other organs. However, most of these efforts are still in the preclinical stage and there are still trials that need further investigations. Additionally, polysaccharides can be functionalized by adding positive or negative charged groups, needed to form stable gels, by cross-linking the polymer network via physical or covalent bonds [14]. They can originate a large variety of organic/inorganic hybrids very attracting in biomedical applications due to their stability in in vivo conditions and their capability to entrap and release drugs, by the fabrication of tailored porosities at micro-, submicro-nanometric size scale [15].

6.2 Classification of polysaccharides Biopolymers are attractive raw materials for the generation of advanced controlled drug delivery systems. Moreover, polysaccharides represent excellent hosting materials for nanoparticles used in the field of drug delivery. In this chapter, we will therefore discuss the structure and properties of the most popular polysaccharides used for preparing hybrid materials for drug delivery applications. The chapter also discusses approaches to new polysaccharides-based hybrid materials using polysaccharides and inorganics followed by a discussion of the different kinds of polysaccharides/inorganic hybrids and subsequently the advantages appeared with polysaccharides-based hybrid materials.

6.2.1 Cellulose and carboxymethyl cellulose Cellulose is an abundant, renewable, and an inexhaustible source of raw material that find various applications for biocompatible products and materials [16]. The most commonly reported sources of cellulose include soft and hard wood from higher plants, certain bacteria, algae, fungi, and several marine animals. Cellulose consists of long-chain, nonbranched, homopolymer [17]. The chains are constructed of two anhydroglucose units linked together through an oxygen covalently bonded to C1 of one glucose ring and C4 of the adjoining ring and so-called the β-1,4-glycosidic bond [18] and every other monomer is rotated 180 degree with respect to its neighboring unit [19]. The inter- and intramolecular hydrogen bonds (Fig.  6.1), arising from OH groups, not only provide cellulose with excellent mechanical properties but also render it insoluble in most solvents. The degree of polymerization, is the number of anhydro-d-glucopyranose units, may vary from several hundreds to 20,000 depending on the source, extraction, and treatment method [19]. The reactivity of the three hydroxyl groups offers a variety of possibilities for making useful cellulose derivatives. The properties of the derivatives depend mainly on the type, distribution, and uniformity of the substituting groups. The average number of hydroxyl groups replaced by the substituents is the degree of substitution (DS), the maximum being three. Several chemical modifications can be used to obtain cellulose derivatives like etherification, esterification, cross-linking, or graft-copolymerization reactions [20]. Water-soluble cellulose ethers, prepared by the reaction of cellulose with aqueous

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CH2OH

CH2OH

CH2OH

Cellulose

O

O O

OH

4

H OH

CH2OH O

H O

H H

H

O n

NH2

H

COONa

HO

OH

G

H

OH

1 H OH

H

O H

O

1

4 NaOOC

4

H OH

O

G

O

H

H 1

H H

Sodium alginate

Chitosan H

H

H

OH

n–2

O

H OH

H

NH2

OH

1

3

OH

OH

CH2OH

OH 2

HO

H

O

5

O

OH

HO H

OH NaOOC

H H

COONa

O

1 O 4

H

OH

O

H

HO

O

H

M

H

M

H

Fig. 6.1  Chemical structure of the main polysaccharides.

sodium hydroxide and then with an alkyl halide, have a great interest especially in the field of biomaterials. Carboxymethyl cellulose (CMC), which is considered as one of the most water-soluble cellulose derivatives, can be prepared from the reaction of cellulose with sodium monochloroacetic acid. CMC with a DS ranging from 0.4 to 1.3 has become the largest industrial cellulose ethers because of its versatile applications in, for example, detergents and oil drilling. CMC degrades completely at low rates in the environment forming nontoxic intermediate [21]. In addition to the various superior characteristics of CMC, the chemical modifications, especially through grafting, can offer additional desired properties that enhance its overall characteristics and increase its potential applications in many fields. Many efforts have been made to synthesize CMC-based superabsorbents and to improve the swelling capacity. Cellulose and its derivatives have been studied recently in the field of drug delivery [22]. CMC has good packing ability for insoluble drugs and bioactive materials.

6.2.2 Chitosan Chitosan (CS, deacetylated chitin) is the second most abundant polymer next to cellulose. CS is a linear aminopolysaccharide derived from chitin through deacetylation process. It is a copolymer composed of N-acetyl-d-glucosamine and d-glucosamine units with one amino group and two hydroxyl groups in each repeating glucosidic unit [23]. The molecular weight of CS varies from 300 to 1000 kDa with a degree of deacetylation from 30% to 95%. CS comprises numerous primary amino (NH2) and hydroxyl (OH) groups; the amino groups are reactive sites for new group attachments.

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CS is insoluble in water and alkaline pH, but soluble in dilute acid solutions [24]. CS received a tremendous attention and has been extensively used during the last decades in drug delivery systems, tissue engineering, and wound-healing dressing, etc., owing to its excellent properties [25]. CS is nontoxic, biocompatible, and biodegradable and possesses antimicrobial activity. Besides biodegradability, CS is hemostatic, anticholesteremic, anticarcinogenic, and fungistatic. Pure CS is insoluble under physiological conditions which hinder its biomedical applications. However, unlike chitin, which is essentially insoluble in aqueous media, CS dissolves in dilute and concentrated aqueous acid. CS also poses significant solubility problems and better solvents for CS are still wanted. Moreover, CS is a pH sensitive polymer, exhibits gel-forming properties, with nonimmunogenicity and mucus adhesion properties. The pH sensitivity of CS is attributed to protonation/deprotonation of the NH2 groups, which enhances interaction with polyanions such as sulfate, citrate, tripolyphosphate ions leading to formation of strong cross-linked matrix. Polyanionic macromolecules such as alginate, gelatin, and CMC can also interact with protonated NH2 groups leading to formation to polyelectrolyte complexes. Also, protonation of the NH2 groups in CS results in enhanced mucoadhesion properties which led to prolonged residence time of CS at the site of administration, permeation enhancing properties as well as controlled release properties. CS is also degraded by body enzymes, allowing for safe excretion. It is also worth mentioning that modification of CS gives rise to water-soluble derivatives and improves its compatibility. Such derivatives are covalently or ionically cross-linked and through cross-linking irreversible chemical links are formed, which facilitate absorption of water and/or bioactive compounds. In this respect, modification of CS is a promising avenue to surpass its weak mechanical properties, which led to drug burst release in controlled drug delivery system. In the last decades, many successful trials were carried out to elucidate the potential of CS-based materials in the delivery of vaccine, proteins, and genes, as well as in drug delivery systems. New water-soluble CS derivatives covalently or ionically cross-linked has been investigated as promising materials for medical and pharmaceutical applications. Thiolated CS showed enhanced mucoadhesiveness properties due to the formation of disulfide bridges between polymer chains and glycoproteins of the mucosa. Novel cross-linked thiolated CS prepared by green chemistry as aqueous dispersion demonstrated to be a useful excipient in topical controlled delivery system [24]. Many polymers, other than CS, can also adhere to mucosal tissues via hydrogen bonding or nonspecific and noncovalent electrostatic attractions [13]. Polymeric micelles are considered as nanoparticulates able to incorporate drugs by physical entrapment and characterized by improved absorption and stability in the biological fluids. CS and CS derivatives micelles are mainly used as drug delivery targeting to the brain tumors and can host peptides and genes. However, the weak electrostatic attractions between CS and the negatively charged mucosal surfaces cause insufficient localization of the drug carrier at the target mucosal site. Introducing thiol groups to CS micelles causes prolonged residence time due to the formation of covalent bonds between thiochitosan and the mucosal surface. Also the small size of the micelles allows deeper penetration of the thiol groups and prevents back diffusion leading to good adhesion. Cross-linked thiolated CS as aqueous dispersions are considered useful excipient in

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topical controlled drug delivery systems [24]. Highly mucoadhesive thiolated CS micelles based on amphiphilic CS-stearic acid-conjugate was prepared. The amphiphilic micelles were prepared via stearic acid linkage to CS and later, thioglycolic acid was covalently attached to CS stearic acid. The newly developed thiolated CS micelles based on amphiphilic CS-stearic acid conjugate is promising candidate for prolonged delivery of hydrophobic drugs to mucosal membranes [26].

6.2.3 Sodium alginate Alginate is a water-soluble linear polysaccharide extracted from brown seaweed and is composed of alternating blocks of 1-4 linked α-l-guluronic and β-d-mannuronic acid residues. The molecular weight can vary between 10 and 1000 kDa depending on the source and production process. Fig. 6.1 shows the structures of mannuronic (M) and guluronic (G) acid residues and the binding between these residues in alginate. Because of the particular shapes of the monomers and their modes of linkage in the polymer, the geometries of the G-block regions, M-block regions, and alternating regions are substantially different. If two G-block regions are aligned side by side, a diamond-shaped hole results. This hole has dimensions that are ideal for the cooperative binding of calcium ions. The homopolymeric regions of β-d-mannuronic acid blocks and α-l-guluronic acid blocks are interdispersed with regions of alternating structure (β-d-mannuronic ­acid-α-l-guluronic acid blocks) [27]. The composition, the extent of the sequences, and the molecular weight determine the physical properties of the alginates [28]. The gelation of alginate can be carried out under an extremely mild environment by using nontoxic reactants. Upon adding multivalent cations, an alginate solution rapidly forms an ionotropic gel that makes it extremely interesting to be applied in the biomedical field. Alginate beads can be prepared by extruding a solution of sodium alginate containing the desired active materials, as droplets, into a divalent cross-linking solution such as Ca2+, Sr2+, or Ba2+. Monovalent cations and Mg2+ ions do not induce gelation [29, 30]. The gelation and cross-linking of the polymers are mainly achieved by the exchange of sodium ions from the guluronic acids with the divalent cations, and the stacking of these guluronic groups to form the characteristic egg-box structure as shown in Fig. 6.1. The divalent cations bind to the α-l-guluronic acid blocks in a highly cooperative manner and the size of the cooperative unit is >20 monomers. Each alginate chain dimerizes to form junctions with many other chains and as a result, gel networks are formed [31]. Numerous studies have shown that the chemical structure, molecular size as well as the gel-forming kinetics and the cation has a significant impact on several of its functional properties, including porosity, swelling behavior, stability, biodegradability, gel strength, and the gel’s immunological characteristics and biocompatibility. Alginate has been widely employed to fabricate gel beads for the drug delivery of biomolecules such as drugs, proteins, and peptides [32]. It was reported that the biological activity of drugs can be retained in the calcium-cross-linked alginate encapsulation process [33]. The entrapment of several proteins from alginate has been reported including melatonin [34] and heparin [35]. Sodium alginate cross-linked with divalent cations are particularly appropriate for the encapsulation of sensitive drugs because of the mild preparation conditions [36].

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Several trials have been carried out to overcome the problems associated with its applications for oral drug administration such as burst release, lower encapsulation efficiency, and quick dissolution in intestinal tract [37]. Consequently, combination of sodium alginate and inorganic nanoparticles can provide the ideal properties necessary for the sustained drug delivery.

6.3 Composites and hybrid materials for molecular release Hybrid materials and composites may be obtained by mixing at least two components, commonly inorganic and organic materials at different size scale—from nanometric to molecular scale. Biohybrids are produced by organisms for specific purposes including structural support, sensing, ion storage, and toxic waste removal. Composite materials, such as bones, teeth, spines, and shells show a wide range of morphologies and yet are made from a relatively limited number of chemical components. These components are assembled under genetic control, as a result they have intriguing structures and specific functions [16]. The use of hybrid materials in pharmaceutical formulations is highly beneficial owing to the combined properties of both the organic moiety, such as functionality, lightness, pliability, and flexibility with heat resistance and stability of the inorganic material [38, 39]. Hybrids can be synthesized via different approaches: sol-gel process, intercalation, exchange, or grafting. The interaction between the organic and inorganic parts in the hybrid can be via weak bonds (hydrogen, van der Waalʼs or ionic bonds) or strong chemical bonds (covalent or iono-covalent bonds) [40]. Organic/inorganic hybrids represent creative field for research activities, in which enormous number of new and multifunctional materials can be developed and applied. In biomedical area, such hybrids are stable in in vivo conditions and it has been noticed that porous hybrid materials are very effective for the encapsulation of hydrophobic drugs [15]. Also, increased drug upload and decreased release rate could be achieved as the drug composite was coated with a porous, nontoxic filler. Moreover, polysaccharides/inorganic hybrid materials represent generic importance in the formulation of biocompatible microcapsules [41]. Organic/inorganic hybrid materials are also qualified for the development of advanced nanocomposites, which offer good candidates for drug delivery systems [42]. Nanocomposites are polymers containing dispersed nano-inorganic filler with average particle size 40% of cell mortality [142]. To allow a repeated or continuous administration of molecules the use of nanostraws was found to be very promising. A biomimetic platform was developed that establishes continuous fluid access into the intracellular space for the delivery of small molecules (Fig. 14.5). The authors report 40%–70% efficiency in the delivery of small molecules and expression of delivered DNA plasmids [103]. Other studies report that nanostraw penetration was found to be dependent on adhesion time and although a penetration rate ranging from 6% to 12% was described, this was found to be stable and suitable

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Fig. 14.5  (A) Scanning electron microscopy image and (B) the proposed device for using nanostraws for continuous delivery of molecules. Adapted with permission from J.J. VanDersarl, A.M. Xu, and N.A. Melosh, Nanostraws for direct fluidic intracellular access, Nano Lett. 12(8) (2012) 3881–3886. Copyright (2012) American Chemical Society.

for a practical application namely in bio-manipulation [109]. The application of a short electrical pulse to compromise cell membrane integrity and provide access to the cell milieu improves the penetration and delivery mediated by nanostraws [143].

14.2.6 Nanoneedles for detecting circulating tumor cells An alternative application in which the use of vertically aligned nanostructures showed promising results concerns the capture of CTCs. These cells are released from a primary tumor to the circulation being associated with the occurrence of metastasis. Moreover, the study of CTCs is considered relevant to monitor cancer therapies [144]. The first device to capture CTCs—Nanovelcro—is based on vertically aligned silicon nanowires functionalized with a specific cell adhesion antibody— EpCAM—that identifies tumor cells, but not other blood cells. The assay provides for the detection of 70% of the cells from whole blood samples [145]. In the device, local topographic interactions between the nanoscale cellular components and the nanowire array are promoted. In order to increase cell-substrate contact frequency, an overlaid polymeric serpentine microfluidic chaotic mixer was integrated with the nanopatterned substrate to improve capture efficiency [146]. However, in these devices to recover the cells after capturing is a harsh process that requires enzymatic treatment and does not guarantee viable cells retrieval. To assure cells viability after capture and recovering, nanowires were coated with a thermoresponsive polymer poly(N-isopropylacrylamide), (PNIPAAm). When the temperature goes below 37°C, PNIPAAm turns more hydrophobic and supports the release of the cells [147,148]. More recently, fractal nano-bio interfaces based on 3D hierarchical nanomaterials were developed to further boost CTC capture based on hierarchically assembled indium tin oxide nanowire arrays with both horizontal and vertical nanowire branches. The branches provide more efficiency comparing to nanowires without branches, in a shorter period of time [102].

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14.3 Concluding remarks If for the one hand the use of microneedle patches faces the challenge of upscaling, regulation, and clinical trials, in the context of drug release, the integration of these microstructures in wearable and biosensing devices opens new opportunities for fully integrated, smart, and responsive devices. These are still in the significant growth stage. As the development of fourth-generation microneedle-based devices is really interdisciplinary, the field of microneedles faces now a very challenging period. For the other hand, the field of nanoneedles is still in its late infancy. Some of the concepts from the microneedle field are to be applied to the nanoscale too. Strict control of the architecture of these structures, the manufacturing, and upscaling issues are to be addressed and time will come to understand whether the current promises will become a solution.

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Metal nanoscale systems functionalized with organic compounds

15

Sara Ferraris*, Martina Cazzola*, Leonardo Raphael Zuardi†, Paulo Tambasco de Oliveira† * Department of Applied Science and Technology (DISAT), Politecnico di Torino, Turin, Italy, † University of São Paulo, School of Dentistry of Ribeirão Preto, Ribeirão Preto, Brazil

15.1 Proteins, peptides, and growth factors Nanostructuring of metallic surfaces and their functionalization with bioactive organic compounds represents the current generation of development of metal implant devices. Despite the ability of nanotopographic surfaces per se to differentially control cellular activities during tissue repair, functionalization with organic compounds adds specific cell functions with beneficial synergistic effects [1–5]. Importantly, the selection of one or more molecules to functionalize metals should take into consideration the biological effects that they exert on the various cell types and differentiation stages and whether their endogenous counterparts have their production modulated by the nanotopography itself. In that case, aggregates of endogenous extracellular matrix proteins laid down on the material surface would then act as a physiological surface coating, with potential impact on cell signaling and function [6–8]. The organic compounds that naturally occur at the tissue-biomaterial interfacial region have been selected to coat metals, including the major protein of connective tissues, type I collagen [9–12], noncollagenous matrix proteins or growth factors [13–18], keratin [18a], and peptides with known biological functions [3, 19–23]. As blood and its plasma constituents are the first elements to interact with biomaterials at the time of implantation, key proteins of the blood clot that is formed, for example, fibrinogen and fibrin, have also been considered in surface functionalization strategies aiming to provide more physiologically relevant interfacial areas, especially for bone application, as reported in [24–26, 28] and reviewed in [27]. The biological effects that proteins, peptides, and/or growth factors exert when functionalized on various metallic surfaces are summarized in Table 1. It is well established that type I collagen, the main structural protein of bone, promotes osteoblast cell adhesion and activities [50]. Based on that, type I collagen has been coated on titanium aiming to enhance osseointegration. Despite the beneficial cellular and tissue responses observed in in vitro and in vivo studies [2, 9–12, 29–32], the main source of collagen, from animal tissues, raises concerns about its quality, purity, lack of stability, immunogenicity, and the risk of transmission of infectious diseases, Nanostructured Biomaterials for Regenerative Medicine. https://doi.org/10.1016/B978-0-08-102594-9.00015-2 © 2020 Elsevier Ltd. All rights reserved.

Grafted protein, peptide or growth factor

In vitro and/or in vivo testing

Titanium alloy

Type I collagen

In vitro

Titanium

Acrylic acid surface graftingtype I collagen coupling

In vitro/in vivo

Titanium pins

Type I collagen

In vivo

Porous titanium surface topography

Acrylic acid surface graftingtype I collagen coupling

In vitro/in vivo

Titanium

Acrylic acid surface graftingtype I collagen coupling

In vitro

Acid etched titanium implants Plasma-sprayed titanium coating with rough surfaces

Type I collagen

In vivo

Type I collagen

In vitro/in vivo

Titanium screws

Type I collagen

In vivo

Screw-type titanium implants

Type I collagen nanofibers

In vivo

Main results

Reference(s)

Accelerated initial adhesion of rat calvarial osteoblasts In vitro: Reduction of the growth rate of SaOS-2 osteoblastic cells. In vivo: increase of bone growth and bone-to-implant contact in rabbit femur Increased early bone remodeling around titanium pins and a tendency towards increased bone formation around them Enhanced human mesenchymal cell growth and improvement of bone-to-implant contact in rabbit trabecular bone Altered mRNA expression profile of human alveolar bone-derived cells during the mineralization phase of cultures Promoted the upregulation of osteoblast markers and bone formation (mainly at 3 weeks) Promoted human mesenchymal stem cell migration and more active new-bone formation around implants (greater bone-to-implant contact; more abundant bone trabeculae) Ensured a greater mechanical stability and increased bone-to-implant contact in compromised bone Better contact of the gingival connective tissue; some collagen fiber bundles oriented perpendicularly to the implant surface

[29] [9]

[11]

[10]

[30]

[12] [31]

[32] [33]

Nanostructured Biomaterials for Regenerative Medicine

Metallic substrate

408

Table 1  Surface functionalization of metals with proteins, peptides, and growth factors

Type I collagen

In vitro

Titanium and titanium nitride alloys

Type I collagen

In vitro

Titanium

Collagen-mimetic GFOGER peptide

In vitro/in vivo

Hydroxyapatitecoated titanium implants Titanium alloy

Type I collagen-derived P-15 peptide

In vivo

Type I collagen-derived P-15 peptide Type I collagen-derived P-15 peptide

In vitro

Type I collagen-derived P-15 peptide

Type I collagen-derived P-15 peptide

Calciumphosphate-coated microtopographic titanium Calciumphosphate-coated microtopographic titanium Hydroxyapatitecoated titanium implants

Increased proliferation, ALP activity, matrix mineralization, and Coll1a, OSX, OC, and BAX expression in rat calvarial osteogenic cells Increased rate of cell proliferation in primary human gingival fibroblasts and higher expression levels of genes related to cell adhesion, differentiation and matrix production Triggered osteoblastic differentiation and mineral deposition in bone marrow stromal cells and improved peri-implant bone regeneration and osseointegration in a rat cortical bone model Increased peri-implant bone density

[2]

Increased cell attachment, spreading, and osteoblast differentiation of a mesenchymal cell line Promoted limited synergistic effects to the enhanced osteogenic potential of cultures grown on calciumphosphate coating

[36]

In vivo

Higher bone-to-implant contact at an early time point (1 week) in the dog tibia

[37]

In vivo

Improved peri-implant bone formation and remodeling at 6 months, but not at the early phase/ first 7 days of osseointegration

[22]

In vitro

[34]

[35]

[19]

[20]

Continued

Metal nanoscale systems functionalized with organic compounds409

Nanotopographic titanium

410

Table 1  Continued Grafted protein, peptide or growth factor

In vitro and/or in vivo testing

Titanium

Modified P17-peptide

In vitro

Hydrophilic gold

Osteopontin

In vitro

Titanium implants Chemically nanostructured titanium surface Titanium

Bone sialoprotein Bone sialoprotein

In vivo In vitro

Decorin

In vitro

Titanium implants

Collagen and chondroitin sulfate

In vivo

Titanium implants

Type I collagen and low sulfated hyaluronan derivatives BMP-2

In vivo

BMP-2 and multilayered coatings of gelatin/chitosan via layer-by-layer assembly Osteogenic growth peptide

In vitro

Titania nanotubes Titania nanotube arrays Titania nanotubes

In vitro

In vitro

Main results

Reference(s)

MP17-coatings reduced the effect of TGF-β1 on MC3T3-E1 cells Exhibited the highest level of hydration, and supported cell adhesion and spreading Osteoinduction under mechanical loading Increased calcium deposition and stimulated cell differentiation of primary human osteoblasts

[38]

Reduced fibroblastic cell proliferation, migration and collagen synthesis and enhanced osteoblastic cell functions Enhanced bone remodeling and de novo bone formation; improved both the quantity and quality of bone adjacent to implants in a model of ovariectomized animals Increased early peri-implant bone formation

[41]

Proliferation and osteogenic differentiation of mesenchymal stem cells Retained BMP-2 bioactivity and release properties, and improved the motogenic and osteogenic potentials of mesenchymal stem cells Enhanced osteogenic potential of osteoblastic cell cultures

[45]

[14,39] [15] [40]

[42,43]

[44]

[46]

[47]

Nanostructured Biomaterials for Regenerative Medicine

Metallic substrate

Serum albumin

In vitro

Hydrophilic microroughened titanium with nanostructures

Whole human blood

In vitro

Fiber network plates/444 ferritic stainless steel Micro/nano-rough titanium discs

Fibrinogen and thrombin

In vitro

Fibronectin/vitronectin exudate and leucocyte- and platelet-rich fibrin Fibrinogen and/or fibrin

In vitro

Acid-etched titanium

In vitro

Promoted early human gingival fibroblast adhesion, while suppressing late proliferation and type I collagen secretion Formation of a dense fibrin network; increased attachment, expression of key osteogenic markers, and mineralization of human bone-derived cell cultures Higher human osteoblast attachment to fibrincontaining fiber networks; lower cell differentiation on supraphysiological concentrations Increased thickness of a dense fibrin network in direct contact with the surface

[48, 49]

Higher cell proliferation, adherence, gene expression, and mineralization in human osteoblastic MG63 cells

[28]

[24]

[25]

[26]

Metal nanoscale systems functionalized with organic compounds411

Titania nanotube arrays

412

Nanostructured Biomaterials for Regenerative Medicine

thus limiting its use [51–54]. To circumvent these problems, the development of recombinant sources of human collagen provides a reliable, predictable and chemically defined source of purified human collagens [53, 54]. Interestingly, high-scale production of recombinant human type I collagen in a tobacco plant expression platform has been shown to provide thermally stable helical structures, fine homogenous fibrils displaying intact-binding sites, and bioactivity that resembles that of native collagen [55]. As a strategy of presenting specific biofunctional domains within the native ligand, another alternative is to use synthetic peptides mimicking amino acid sequences of collagen [35, 56]. For instance, P-15, a collagen-derived synthetic peptide analogue of the cell-binding domain of type I collagen, has been demonstrated to promote osteogenic differentiation and to enhance bone formation in distinct applications [20, 22, 36, 37, 53, 54, 57, 58]. The coating of titanium surfaces with the glycine-phenylalanine-­ hydroxyproline-glycine-glutamate-arginine (GFOGER) collagen-mimetic peptide, that selectively promotes alpha2beta1 integrin binding, resulted in an enhancement in osteoblast cell functions in vitro and in significantly improved in vivo peri-implant bone regeneration and osseointegration [35]. In addition to supporting bone formation adjacent to metal implants, the functionalization of titanium with type I collagen has also been applied with the aim at enhancing gingival fibroblast cell response, thus favoring the interactions of the gingival connective tissue and the metal surface in the transmucosal area [33, 34]. Similarly, titania nanotube arrays fabricated by anodizing titanium sheets have been loaded with serum albumin to modulate gingival fibroblast activities [48, 49]. Despite the benefits of supporting the repair of gingival fibrous connective tissue in the transmucosal area, attempts have been made to minimize fibrotic encapsulation of implants during the process of osseointegration. In this context, the loading of the noncollagenous protein decorin, a small proteoglycan, on a titanium surface via polydopamine film attenuates fibroblastic cell proliferation, migration, and collagen synthesis while enhancing the osteogenic potential of osteoblastic cells [41]. As TGF-β1 signaling has been associated with collagen production and the formation of fibrous tissues, the covalent immobilization of a TGF-β1 inhibitor modified P17-peptide on titanium surfaces is expected to inhibit TGF-β1 activity that will potentially minimize fibrotic encapsulation of implants during the process of osseointegration [38]. The multifunctionality of proteins can be exploited in surface functionalization strategies of nanoscale metal systems, as it might impact simultaneously the activities of diverse cell types at the interfacial region. For instance, osteopontin (OPN), a multifunctional, matricellular protein of the SIBLING (small integrin-binding ligand, N-linked glycoprotein) family, which is expressed by many tissues and cell types, has been considered a promising coating for biomaterials based on its importance in both formation and remodeling of calcified tissues and in immunological responses. In the case of bone applications, the choice of OPN is based on its preferential accumulation at bone interfaces, including that with biomaterial [59]. Depending on the type of coating, whether adsorbed at hydrophilic or hydrophobic surface chemistries or presented as nanopatterned patches, OPN has a significant impact on the initial events of cell

Metal nanoscale systems functionalized with organic compounds413

a­ dhesion and spreading, a finding that could be explained by the different conformations and/or orientations it adopts at the different material surfaces [14, 39]. Another multifunctional SIBLING glycoprotein, bone sialoprotein (BSP), has been coated onto femoral titanium implants with gelatin as a carrier, showing osteoinduction under mechanical loading conditions [15]. In another study [40], a chemically nanostructured titanium surface was used to functionalize BSP via an aminosilane linker or by simple adsorption (physisorption). The results showed that both methods were effective in promoting osteogenic differentiation in vitro, especially with the higher BSP concentration used (280 μg/mL). It is worth noting that any strategy to coat metals with either OPN or BSP (or even with other noncollagenous matrix proteins) should first consider the impact that the nanostructured surface exerts on the production and secretion of these glycoproteins by host cells at the interface [6, 20, 60]. Nonetheless, despite the possibility of controlling the endogenous accumulation of proteins at the titanium surface by tuning the nanoscale topographical features and/or surface chemistry, the coating with an exogenous protein offers the advantages of predictability of more homogeneous layer over the material surface and of knowing its concentration. Given the variety of matrix proteins that are secreted at the interface with metal implants during tissue repair, more complex coatings using two or more matrix components have been proposed to mimicking the physiological 3D environment and ultimately enhancing bone healing, especially for compromised bone. Coating titanium implants with an artificial matrix based on collagen and chondroitin sulfate have been shown to enhance bone remodeling and de novo bone formation in vivo [42], and appears to improve both the quantity and quality of bone formed around implants in a model of ovariectomized animals [43]. An artificial matrix consisting of type I collagen and either one of two regioselectively low sulfated hyaluronan derivatives has been coated on titanium implants and shown to increase early peri-implant bone formation [44]. In a systematic review and metaanalysis study, bone morphogenetic proteins (BMPs), a group of growth factors belonging to the transforming growth factor superfamily, were among the most commonly used bioactive chemical compounds for delivering to the bone-implant interface, and among them, the only organic compound that promoted a significant increase in the bone-to-implant contact in animal models of osseointegration [61]. Among BMPs, BMP-2 and BMP-7 are the two growth factors approved by the Food and Drug Administration (FDA) for treatments of spinal fusion and long-bone fractures with collagen carriers [62]. Concerning nanostructured metals, the conjugation of BMP-2 onto titania nanotubes via polydopamine showed beneficial effects on mesenchymal stem cells in vitro in terms of cell proliferation and osteogenic differentiation [45]. A similar strategy has been applied to functionalize osteogenic growth peptide, a growth factor peptide, with benefits for the osteogenic potential of osteoblastic cells [47]. In another study, titania nanotube arrays for loading of BMP-2 have been constructed on titanium substrates and then covered with multilayered coatings of gelatin/chitosan via layer-by-layer assembly. The multilayered coating retained the BMP-2 bioactivity and release properties, and was able to stimulate the motogenic and osteogenic potentials of mesenchymal stem cells [46].

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15.2 Antibiotics Bacterial infections at surgical sites have been identified as the main cause of dental and orthopedic metal implant failures. The formation of a pathogenic multispecies biofilm on the metal surface triggers an immunoinflammatory reaction, which hinders bone tissue repair and might progress to osteomyelitis. Based on that, various approaches have been applied to make the metal implant surface antibacterial by impeding the biofilm formation, including adhesion-­resistant coatings and coatings containing or releasing antimicrobial agents (reviewed in Refs. [63, 64]). Despite nanotopographic and chemical patterns of metallic surfaces per se have been shown to successfully reduce bacterial cell adhesion, that is critical for biofilm formation [65–72], such strategies of surface modifications have also been used to functionalize specific antibiotics, including gentamicin, vancomycin, and tetracycline, or a combination of antibiotics and/or antimicrobial agents. Ideally, in addition to exert bacteriostatic and/or bacteriocidal effects in peri-­implantitisassociated bacterial species in a sustained release/optimal bioavailability over the healing process and longer periods, antibiotic coatings on metal implants should also support biocompatibility, that is, bone cell survival, proliferation, differentiation, and bone repair at the implantation site, which has been found to exert a major impact on the performance, stability, and longevity of the implant [73]. The biological effects of functionalized antibiotics and/or antimicrobial agents on metallic surfaces are summarized in Table 2. Various surface nanostructuring methods have been applied to successfully functionalize titanium surfaces with gentamicin. Titania nanotubes loaded with gentamicin significantly reduced Staphylococcus epidermidis adhesion on the surface, while promoting enhanced osteoblast differentiation [74]. Similarly, nanotubular anodized titanium coated with gentamicin implant showed less bacterial growth in  vitro, reduced implant-related osteomyelitis, and enhanced bone biocompatibility in vivo [67, 68]. The coating of 3D-nanopillared titanium substrates with an ultrathin tannic acid/gentamicin layer-by-layer film resulted in a 10-fold reduction of the number of surface-attached Staphylococcus aureus [75]. High gentamicin delivery efficiency was achieved with gentamicin sulfate-loaded titania nanotubes covered by a thin film comprising a mixture of gentamicin sulfate and chitosan [76]. In a rat model, biodegradable, gentamicin-loaded poly(d,l-lactide) coating of titanium Kirschner wires significantly reduced bacterial growth and implant-related osteomyelitis [77]. A bioactive, macroporous titanium-oxide surface modified with gentamicin-sodiumdodecylsulfate and gentamicin-tannic acid coatings prevented implant contamination, with infection prophylaxis rates of 100% and 90%, respectively [78]. In a multicenter prospective study, the use of gentamicin-coated tibia nail has been effective in the prevention of implant-related infection or osteomyelitis in patients with a high infection risk, including open fractures or infected nonunions. Any local or systemic toxic effects related to gentamicin were noted, and deep surgical site infections occurred in approximately 5% of patients [79].

Metallic substrate

Grafted antibiotic and/or antimicrobial agent

In vitro and/or in vivo testing

Titania nanotubes

Gentamicin

In vitro

Nanotubular anodized titanium

Gentamicin

In vitro/in vivo

3D-nanopillared titanium

In vitro

Plasma chemical oxidized titanium alloy Titanium nails

Ultrathin tannic acid/gentamicin layer-by-layer film Mixture of gentamicin sulfate and chitosan Biodegradable, gentamicin-loaded poly(d,l-lactide) Gentamicin-sodiumdodecylsulfate and gentamicin-tannic acid Gentamicin

Titanium alloy

Vancomycin

In vitro

Titanium locking compression plates Titanium

Vancomycin

In vivo

Vancomycin encapsulated in a poly(ethylene glycol)-based hydrogel film

In vivo

Titania nanotubes Titanium Kirschner wires

In vitro In vivo In vivo In vivo

Main results

Reference

Reduced Staphylococcus epidermidis adhesion and enhanced osteoblast differentiation Less bacterial growth, reduced implantrelated osteomyelitis, and enhanced bone biocompatibility Reduction of the number of surfaceattached Staphylococcus aureus High gentamicin delivery efficiency

[74]

Reduced bacterial growth and implantrelated osteomyelitis Prevented implant contamination

[77]

Effective in the prevention of implantrelated infection or osteomyelitis in patients with a high infection risk Inhibited Staphylococcus epidermidis biofilm formation Prevented implant biofilm formation and supported bone repair Reduction of the inflammatory reaction and an antibacterial capacity

[79]

[67, 68]

[75] [76]

[78]

[80] [81] [82]

Continued

Metal nanoscale systems functionalized with organic compounds415

Table 2  Surface functionalization of metals with antibiotics and/or antimicrobial agents

416

Table 2  Continued Grafted antibiotic and/or antimicrobial agent

In vitro and/or in vivo testing

Chitosan-coated titanium

Tetracycline

In vitro/in vivo

Titanium

Tetracycline

In vitro

Ca-P-coated titanium

Metronidazole and simvastatin

In vitro

F-doped nanotubular anodic oxide layers fabricated on titanium alloy Polymeric nanofibercoated titanium Kirschner wires Indwelling titanium devices

Gentamicin and vancomycin

In vitro

Vancomycin, rifampin, linezolid and daptomycin Silver nanoparticles and/or ampicillin

Main results

Reference

Inhibited Aggregatibacter actinomycetemcomitans and Staphylococcus epidermidis growth, with limited cytotoxic effects and typical acute inflammatory response Prevented Escherichia coli colonization and supported osteoblastic cell adhesion and proliferation Antibacterial activity against Porphyromonas gingivalis and improved osteogenic potential of undifferentiated cells Limited growth of Staphylococcus aureus, Staphylococcus epidermidis and Pseudomonas aeruginosa

[83]

In vivo

Prevented biofilm formation and promoted enhanced osseointegration

[86]

In vitro

Decreased Staphylococcus aureus viability; bacteriostatic activity maintained over time

[87]

[84]

[48, 49]

[85]

Nanostructured Biomaterials for Regenerative Medicine

Metallic substrate

Metal nanoscale systems functionalized with organic compounds417

Vancomycin covalently attached to a titanium alloy surface was shown to inhibit S. epidermidis biofilm formation that has been prevalent in orthopedic infections [80]. In a sheep model of tibial osteotomy and S. aureus infection, a vancomycin-modified titanium locking compression plate effectively prevented implant biofilm formation and supported bone repair [81]. Vancomycin has also been loaded in polymeric coatings on orthopedic implants [82]. Briefly, vancomycin encapsulated in a poly(ethylene glycol) (PEG)-based hydrogel film that was covalently bound to titanium implants was then covered by a PEG-poly(lactic-co-caprolactone) membrane. Additionally, cross-linked starch was mixed with the hydrogel to inhibit its swelling and therefore slow down antibiotic release. In a rabbit model of S. aureus infection, the ­vancomycin-loaded implants showed a significant reduction of the inflammatory reaction and exhibited a good antibacterial capacity. Tetracycline has also been used in titanium coatings. The local delivery of tetracycline using chitosan coatings bonded to titanium via silane inhibited Aggregatibacter actinomycetemcomitans and S. epidermidis growth, with limited in  vitro cytotoxic effects and typical acute inflammatory response in an animal model for tissue healing [83]. The covalent coupling of tetracycline to titanium surfaces (­tetracycline-tethered titanium surfaces) has been applied to prevent Escherichia coli colonization, while supporting osteoblastic cell adhesion and proliferation in short-term cultures [84]. The strategy to combine two or more antibiotics and/or antimicrobial agents has been proved to be more effective in the control of bacterial survival and growth. For instance, metronidazole and simvastatin were integrated into biomimetic Ca-P coatings on a titanium surface, promoting antibacterial activity against Porphyromonas gingivalis and improving the osteogenic potential of undifferentiated cells derived from bone marrow and adipose tissues [48, 49]. Gentamicin and vancomycin have been loaded into F-doped nanotubular anodic oxide layers fabricated on titanium alloy, and the combination of both antibiotics was effective in limiting the growth of S. aureus, S. epidermidis, and Pseudomonas aeruginosa [85]. The delivery of combinatorial antimicrobial agents from titanium Kirschner-wire implants has been achieved with an electrospun composite coating comprised of poly(lacticco-­glycolic acid) (PLGA) nanofibers embedded in a poly(ε-caprolactone) (PCL) film. PLGA and PCL polymer matrices were used to contain and control the release rates of vancomycin, rifampin, linezolid, and daptomycin. Three sets of composite PLGA/PCL coatings loaded with different antibiotic combinations were generated as examples of local combinatorial therapy, and these were effective in preventing biofilm formation on the titanium surface and enhancing its osseointegration [86]. Multifunctionalization of titanium surfaces with silver nanoparticles and/or ampicillin has been developed to control bacterial adhesion and biofilm formation on indwelling devices. Such nanoscale system decreased S. aureus viability in more than 80%, and its bacteriostatic activity was maintained over time with the relatively low but steady release of silver [87]. The challenge of infections caused by multiple antibiotic-resistant bacteria has been addressed with the use of antimicrobial peptides as a promising strategy to eluting coatings on functionalized metal implant surfaces [88–92].

418

Nanostructured Biomaterials for Regenerative Medicine

15.3 Polyphenols and other natural molecules of vegetal origin Natural molecules of vegetal origin are gaining increasing interest in the scientific community due to their potential health benefit effects, low toxicity, safety, and sustainability. Among them, polyphenols stand out for their antioxidant, anticancer, antiinflammatory, antibacterial, vasculoprotective, and bone-stimulating properties, as well as for the possibility to be obtained from the byproducts of the food and beverages industries. The main limitations to effective clinical application of polyphenols are related to their poor stability and bioavailability. In fact, these molecules are sensitive to light, pH (>7.4), heat, oxidants and, upon oral administration they are subjected to rapid metabolism and clearance. The stabilization of polyphenols by means of emulsions [93] and nanoencapsulation [94] has been reported in the literature to increase bioavailability. Despite of the increasing interest in natural molecules and of the promising stabilizing effects of their coupling with materials, few research works consider functionalization of biomaterials and in particular of metal nanoscale systems with these compounds. Some attempts of polyphenols and essential oils grafting to metallic substrates are summarized in Table 3. Briefly, polyphenols have been coupled with metallic substrates in order to confer them specific properties, peculiar of the grafted molecules: thin and colorless polyphenols coating have been obtained on gold and steel starting from pure substances as well as from crude extracts [96], quercetin has been introduced onto TiO2 nanotubes on Ti6Al7Nb in order to induce antiinflammatory, antibacterial, and bone-stimulating properties [98] and pectin have been grafted onto titanium in order to improve osteoblast response [99]. Polyphenols have also been used for the improvement of coating properties on titanium surfaces, in particular the addition of lignin to HAp/HAp-Ag coatings obtained by electrophoretic deposition increases adhesion and stability [97]. Finally, polyphenols have been employed for reduction and stabilization of metallic nanoparticles, as an example, the synthesis and stabilization of gold nanoparticles with resveratrol has been proposed [95]. As far as essential oils are concerned the possibility to obtain cinnamon oil and chitosan coatings on stainless steel has been described [100]. Moreover, the use of essential oils as corrosion inhibitors has been reported for stainless steel [101, 102] and titanium [103]. In these cases, however, essential oils are introduced in the acidic medium used for corrosion evaluation and not directly on the metal surface.

15.4 Stability and durability of metal nanoscale systems functionalized with organic compounds 15.4.1 Effects of sterilization Sterility is a fundamental requisite for biomaterials intended for biological characterization (cells and bacteria culture) and clinical application.

Metallic substrate

Grafted natural vegetal compound

Main results

Reference

Gold nanoparticles

Resveratrol

[95]

Gold and steel

Tannic acid, pyrogallol and crude extracts of tea, wine, and chocolate Lignin

Reduction and stabilization of Au nanoparticles (35 nm). Effective binding of doxorubicin to resveratrol stabilized Au NPs Development of thin (few tens of nm) and colorless coatings of polyphenols Lignin addition to hydroxyapatite (HAp) and HAp/Ag coatings obtained by electrophoretic deposition improve their adhesion and protect HAp from decomposition during sintering Quercetin release kinetics can be tailored varying the thickness of chitosan coating Pectins increase surface wettability but do not alter morphology. Osteoblast response is affected by enzymatic modifications and pectins with high content of galactose induce the highest cell mineralization Coatings with at least 1% cinnamon oil are effectively antibacterial against S. epidermidis

[100]

Titanium

Ti6Al7Nb anodized (TiO2 nanotubes) Aminated (plasma polymerization with Allylamine) commercially pure titanium Blasted 316L stainless steel

Quercetin and chitosan Rhamnogalacturonan-I (RG-I) region of pectins from potato and apple, optionally enzymatically modified Cinnamon oil and chitosan

[96] [97]

[98] [99]

Metal nanoscale systems functionalized with organic compounds419

Table 3  Surface functionalization of metals with polyphenols and essential oils

420

Nanostructured Biomaterials for Regenerative Medicine

Sterilization means the destruction of all living forms (bacteria, fungi, virus, and spores); according to the World Health Organization (WHO), a material or device can be considered sterile if the probability to find a living organism is lower than 10−6 (sterility assurance level—SAL) [104]. This result can be obtained by the application of steam, dry heat, chemical agents, and ionizing or nonionizing radiations. The most commonly sterilization techniques used in the medical field are summarized in Table 4. The interaction between heat, steam, chemicals, or radiations with biomaterials can affect material properties. This phenomenon can be emphasized in case of presence of sensitive and degradable molecules such as organic compounds employed for surface functionalization and development of multifunctional materials (Fig. 1). Despite an intensive scientific research on biomaterials and functional materials, even coupled with organic active agents, few works consider the effects of sterilization on their properties. The impact of sterilization on surface-treated metals should also been considered, in fact, significant variation of materials properties has been reported even for nonfunctionalized substrates. A wide debate on the biocompatibility of titanium oxide nanotubes has been reported. Some authors suggest that the reason of this data dispersion can be associated with the different sterilization techniques used for the preparation of samples for cell cultures. Steam sterilization has been frequently associated with an increase in surface carbon contaminants, strictly related to an increase in the water contact angle, as well as to an alteration of nanotube structure [109, 110]. Moreover, it has been evidenced that autoclave sterilization (of dry or wet samples) can affect protein absorption and consequently the biological response to TiO2 nanotubes [111]. Dual acid-etched titanium surfaces (SLA-type, widely employed in dental applications) have also been shown as sensitive to sterilization. Analogously to TiO2 nanotubes, an increase in carbon contaminations has been reported after autoclave sterilization [112, 113]. On the other hand, oxygen plasma sterilization, gamma irradiation, and hydrothermal sterilization (autoclave sterilization in water in a sealed glass bottle) reduced the surface carbon content improving both wettability and biological response [112, 113]. Hydrothermal sterilization has also been shown as suitable for restoring good biological response of SLA surface after long storage [113] and for sterilization of alkaline-treated titanium surfaces without hampering bioactivity and cell adhesion, contrary to conventional steam sterilization [114]. This last point has been associated to an increase of surface crystallinity and a reduction in carbon contamination that compensate a certain morphological alteration at the nanoscale and a significant sodium loss from the surface [114]. An increase in surface carbon contamination after conventional steam sterilization has also been reported for mirror polished and blasted titanium surfaces [115, 116]. A similar behavior has also been evidenced for ethylene oxide sterilization [116]. On the other hand, a decrease in the carbon content has been observed after gamma irradiation [116] or 70% ethanol soaking [115]. However in this last case, the osteoblast adhesion to ethanol-treated titanium was lower than on steam or dry heat sterilized ones and was restored after sterile water rinsing. This phenomenon has been explained by

Temperature (°C)

Pressure (bar)

Humidity (%)

Time (min)

Mechanism

Autoclave

121–134

1.1–2.1

100

4–20

Dry heat Ethylene oxide

150–170 30–60

Atmospheric Vacuum

– 40–50

60–150 120–960

Pressure and temperature action Temperature Chemical

Gamma irradiation Electron beam

Room temperature

Atmospheric

–

240–360

Ionizing radiation

Room temperature

Atmospheric

–

1–15

Ionizing radiation

Notes

Degassing required for the removal of toxic compounds (days) 25-kGy recommended dose for medical devices 25-kGy recommended dose for medical devices Time depends on number of items to be processed

Metal nanoscale systems functionalized with organic compounds421

Table 4  Main parameters of the most common sterilization techniques used in the medical field [104–108]

422

Nanostructured Biomaterials for Regenerative Medicine

Fig. 1  Effect of sterilization, storage, and working conditions on metal nanoscale systems functionalized with organic compounds.

a strict correlation between the amount of surface OH groups and cell adhesion [115]; sterilization and water rinsing alter the density of OH groups on the surface, consequently the presence of these groups affected cell adhesion on sterilized materials. Finally, the impact of sterilization on the properties of other biomedical metals and alloys should be taken into consideration. As an example, it can be cited that significant morphological alterations as well as effects on cell adhesion have been evidence on Mg and MgCa after ethylene oxide, autoclave, glutaraldehyde, dry heat, and gamma (25 kGy) irradiation. The most significant impact was obtained with ethylene oxide while the most suitable sterilization technique was gamma irradiation [117]. Dry heat, autoclave, and steam sterilizations have been associated with morphological alterations, increase in the surface oxide layer and Ni content on Nitinol samples. A negligible effect, on the same material, has been reported for plasma sterilization [118]. The impact of sterilization on organic compounds can be more critical. Except for silk fibroin, which has been reported as highly resistant to autoclave (in dry, humid, or water soaking conditions), dry heat (180°C), ethylene oxide, 70% ethanol or antibiotic, and antimycotic solutions [119, 120], most of organic molecules/materials is sensitive to strong parameters involved in many sterilization techniques. Substantial alterations after steam sterilization have been reported for polypropylene meshes coated with maleic anhydride by plasma polymerization [121] and PLA films [122]. The same substrates were better resistant to electron beam, gamma irradiation, ethylene oxide, and plasma sterilizations [121, 122]. Gamma irradiation (25 kGy) has been reported as suitable technique for the sterilization of salmon gelatin scaffolds [123], of materials for guided tissue/bone regeneration (based on tricalcium phosphate, bioglass, equine bone, PLGA, PLA, and collagen) [124] and of freeze-dried chitosan beads functionalized with enzymes (papain and bromelain) [125]. Limited alteration of the material structure has been evidenced for

Metal nanoscale systems functionalized with organic compounds423

salmon gelatin scaffolds [123], however mechanical properties and tissue integration ability were maintained. On the other hand, significant alteration of materials properties has been observed after gamma irradiation of collagen sponges (degradation, shrinkage, and cell response) [126] and carboxymethyl-chitin sponges (molar weight reduction, high degradability by PBS and lysozyme) [127]. Supercritical CO2, with the eventual addition of H2O2, has been suggested as alternative sterilization technique for sensitive biomaterials (e.g., biopolymers) as well as for surface decontamination of titanium [128]. Properly adjusting the H2O2 content acceptable alterations of molecular weight, stability, mechanical properties, and biocompatibility have been obtained on polysaccharidic membranes (alginate and hyaluronan) [129]. As far as metal nanoscale systems are concerned, the effect of sterilization after functionalization with organic compounds is investigated in few research works; a summary is reported in Table 5. UV sterilization has been demonstrated a suitable technique for the sterilization of PEG-grafted gold nanoparticles [130] and COOH-grafted titanium [134]. A negligible or, at least acceptable, alteration of surface properties has been observed after gamma sterilization (25 kGy, widely employed in the sterilization of commercial medical devices) of Ti6Al4V grafted with lysozyme [133], BMP-2 [135], and ALP [137]. ALP-grafted Ti6Al4V surfaces resulted quite stable also to electron beam irradiation at the same intensity [137]. Variable results are reported for ethylene oxide sterilization, while it can maintain acceptable activity level of ALP grafted to Ti6Al4V [137], a complete inactivation of TGF-β1 grafted to Ti6Al4V was induced by the same sterilization [136]. Ethanol vapor and ethanol (70%) soaking resulted suitable for the sterilization of Au nanorods functionalized with polystyrene sulfonate (PSS) and methyl-polyethyleneglycol-thiol (m-PEG-SH) [131] and titanium grafted with dopamine and polydopamine [132], respectively. This last material was also stable to autoclave sterilization [132]. A crucial point is that it is extremely hard to find a sterilization technique which leaves completely unaltered the properties of biologically based materials; the critical step is to define proper acceptability level in order to preserve biomaterials peculiar properties and guarantee effective sterility. Similar considerations can be done for stability over time and durability in physiological conditions, discussed in the following.

15.4.2 Stability over time and durability in physiological conditions As discussed in the previous sections organic compounds are extremely sensitive molecules and their exposition to humidity, liquid media, pH change, light, or temperature can alter their properties and impair their function (Fig. 1). The physiological environment foresees the continuous contact with physiological fluids (pH 7.4 in the majority of cases, acidic in the case of the gastrointestinal tract) at 37°C in the presence of oxygen, chlorides, and proteins. These working conditions are quite harsh for traditional materials and can become extremely critical for biomaterials grafted with sensitive organic compounds.

424

Table 5  Effect of sterilization on metal nanoscale system functionalized with organic compounds Metallic substrate

Grafted organic compound

Sterilization

Comments

Reference

Plasmonic hollow gold nanoparticles

PEG

UV-B lamp 15 min

[130]

Au nanorods

Cetyltrimethyl ammonium chloride (CTABr), polystyrenesulfonate (PSS), and methyl-polyethyleneglycol-thiol (m-PEG-SH) Carboxymethyl chitosan (CMCS, grafting mediated by (3-aminopropyl) triethoxysilane—Silane, dopamine— DA and polydopamine—PDA) and bone morphogenetic protein 2 (BMP-2) Lysozyme, via plasma surface activation and first surface treatment with phase-transited lysozyme and citric acid

Freeze-drying and ethanol vapor sterilization

Effective sterilization of infected particles without alteration of morphology and SPR response PEG-SH and PSS resulted stable after freeze-drying and sterilization

Ethanol 70% soaking (1 h) Autoclave (121°C, 20 min)

Functionalization performed with DA and PDA resulted stable to sterilization

[132]

Freeze-drying and gamma irradiation (25–42 kGy)

Progressive reduction of enzymatic activity after several freeze-drying cycles 8 > 3) but complete maintenance after gamma irradiation

[133]

Atmospheric Pressure Plasma functionalization with NH2 (aminopropyltriethoxysilane as precursor) or COOH (methylmethacrylate as precursor)

UV 35 min

Moderate variation in surface characteristics (FTIR peaks of amide, carboxylic acids and ketons) for NH2grafted surfaces, higher stability for COOH functionalized ones

[134]

Ti6Al4V, polished or surface treated (biomimetic advanced surface, Avinent Implant System) Titanium

Nanostructured Biomaterials for Regenerative Medicine

Titanium

[131]

Bone Morphogenetic Protein—2 (BMP-2) grafted via polyelectrolyte multilayer

Gamma irradiation 25 and 50 kGy

Ti6Al4V

Transforming growth factor β1 (TGF-β1) by absorption on untreated or collagen-I coated titanium Alkaline Phosphatase (ALP)

Ethylene oxide at 42°C for 12 h

Ti6Al4V chemically treated (HF etching and H2O2 controlled oxidation)

Gamma irradiation (2, 6, 10, 25 kGy) Electron beam irradiation (2, 6, 10, 25 kGy) Ethylene oxyde

Gamma irradiation increases PEM crosslinking and induces a 50% (25 kGy) reduction of ALP activity induced on cells (Murine C2C12 skeletal myoblasts) or 80% for higher intensity (50 kGy). In vivo (rat ectopic model) sterilized materials maintain osteogenic activity Complete inactivation of TGF-β1 after ethylene oxide sterilization independently from the presence of collagen Limited reduction of enzymatic activity after all the considered conventional sterilization methods

[135]

[136]

[137]

Metal nanoscale systems functionalized with organic compounds425

Commercially pure titanium

426

Nanostructured Biomaterials for Regenerative Medicine

Despite of a wide research in the field of surface functionalization of biomaterials with organic compounds and a certain awareness of their sensitivity to storage and activity in physiological conditions, few research works consider the evaluation of their stability and durability. A selection of papers that reports data of stability over time and durability in physiological conditions are reported in Table 6. The majority of the reported tests simulate the stability in physiological conditions by soaking in phosphate buffered saline (PBS) at various temperatures for times up to 15 days [132, 133, 138] or by the evaluation of stability by cyclic voltammetry and step-potential chronoamperometry, directly in PBS [139]. Dopamine- and polydopamine-mediated chitosan grafting to titanium surfaces and lysozyme grafting to Ti6Al4V [133] resulted stable to PBS soaking while a significant reduction of NH2 density and activity was observed on Ti6Al4V grafted with 3-­aminopropyltrimethoxysilane (APTMS) [138]. A dependence of grafting stability on the site of anchorage has been highlighted for 3-mercaptopropionic acid and dl-thioctic acid on Au electrodes [139]. Two works consider the stability of functionalized titanium surfaces with ALP [137] and vitamin-D precursor [140] respectively, after storage in different conditions. Both works evidenced the beneficial effect of 4°C storage while a complete inactivation of ALP was observed after freeze drying [137] and a reduction of vitamin D precursor after 23°C storage [140].

15.5 Conclusions Grafting of organic compounds to biomaterials, and in particular to metal nanoscale system, is a promising strategy to improve their properties (e.g., bone bonding ability, antiinflammatory, and antibacterial activities) opening the opportunity to produce highly tailorable surfaces able to match complex medical needs. Surface functionalization allows the local administration of active principles to the site of interest with a significant dose reduction and together with molecular stabilization at the surface. Surface grafting of organic compounds to metal nanoscale systems is gaining increasing interest in the scientific literature, as demonstrated by the growing number of studies in the field discussed in this chapter. From the functionalization point of view, results are encouraging and worthy of further investigation for customized applications. On the other hand, the presence of sensitive organic compound on inorganic surfaces makes their sterilization, storage, and certification more difficult. This point is fundamental for the scaling up of innovative biomaterials from the laboratory to the clinical application, although still poorly studied. It is indeed a major challenge to find sterilization, packaging, and storage conditions that leave completely unaltered the properties of functionalized surfaces. Such aspects should then be taken into account in the design of innovative biomaterials grafted with organic compounds aiming to define proper acceptability level that ultimately preserves biomaterials peculiar properties and ensures effective sterility. Thus, further studies on biomaterials grafted with organic compounds should include analyses to verify the impact of sterilization/storage procedures on their physicochemical and biological properties.

Table 6  Stability over time and durability in physiological conditions of metal nanoscale systems functionalized with organic compounds Grafted organic compound

Stability/durability test

Comments

Reference

Titanium

Carboxymethyl chitosan (CMCS, grafting mediated by (3-aminopropyl) triethoxysilane— Silane, dopamine—DA and polydopamine—PDA) and bone morphogenetic protein 2 (BMP-2) Lysozyme, via plasma surface activation and first surface treatment with phase-transited lysozyme and citric acid Alkaline phosphatase (ALP)

14 days soaking in PBS

DA and PDA grafting are stable after PBS soaking

[132]

15 days soaking in PBS at 4°C

Negligible reduction of enzymatic activity

[133]

50 days storage in different conditions: Temperature: 4°C, −20°C Packaging: vacuum bag, paper-PE sterilization bag Optional treatment: freeze-drying

[137]

Ti6Al4V polished and piranha activated

3-aminopropyltrimethoxysilane (APTMS) and transcinnamaldheyde

PBS soaking at room temperature under stirring (700 rpm) up to 360 h

Au electrodes (mono/ poly-crystalline)

3-Mercaptopropionic acid, 2-mercaptoethanol, 1,4-dithiothreitol and dl-thioctic acid

Commercially pure titanium blasted with TiO2

UV irradiated vitamin D precursor (7-dehydrocholesterol) and Vitamin E

Evaluation of stability in PBS by cyclic voltammetry and step-potential chronoamperometry 3, 6, and 12 weeks storage in nitrogen packaging in the dark at −20°C, 4°C, or 23°C

Freeze drying extremely reduces enzymatic activity with all temperature and packaging conditions. Both packaging and storage condition (without freezedrying) induce acceptable ALP activity reduction PBS soaking decrease the NH2 density on the surface (70% reduction) and reduction of the protection ability against bacteria 3-Mercaptopropionic acid and dl-thioctic acid resulted more stable when bonded to (100) sites compared to (111) ones 23°C storage induces loss of activity after 3 weeks while 4°C preserves molecular activity up to 12 weeks

Ti6Al4V, polished or surface treated (Biomimetic Advanced Surface, Avinent Implant System) Ti6Al4V chemically treated (HF etching and H2O2 controlled oxidation)

[138]

[139]

[140]

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Metallic substrate

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Index Note: Page numbers followed by f indicate figures and t indicate tables.

A Acid etching, 320 Adipose-derived mesenchymal stem cells (ADSCs), 157–158 Advanced therapy medicinal products (ATMPs), 30 Agarose hydrogel, 207, 208f Ag NCs. See Silver nanoclusters (Ag NCs) AIBN. See 2,2′-Azobisisobutyronitrile (AIBN) Alginic acid, 11 Alumina/alumina-based composites, 5–6 Alumina-stabilized zirconia-based composites, 5–6 Amino group, 57 3-Amino-propyltriethoxysilane (APTES), 207, 265 Anisogels fabrication process, 211, 214f Antimicrobial activity biomaterial-related infections antibiotic therapy, 82 average infection rate, 81–82 bacterial adhesins, 83 biofilms, 82–83 (see also Biofilm formation) clinical relevance, 82 foreign-body reaction, 81–82 fungal colonization, 85–86 health-care-related infections, 81 implant surgical intervention, 81 device-associated infections, 118 nanostructured biomaterials, 110 limitations, 111–117, 112–116t microbial infection, 110–111 physicochemical-surface properties, 118 Antimicrobial peptides electrostatic forces, 107 gen-encoded peptide antibiotics, 107 in vivo/in vitro expression, 107 silk fibroin-based membrane, 107–108 surface functionalization strategies, 106–107 titanium alloys, 108

Anti-programmed death-1 (aPD-1) immunotherapy, 387 APTES. See 3-Amino-propyltriethoxysilane (APTES) Aseptic manufacturing process, 38 Atmospheric-pressure oxygen plasma, 110 ATMPs. See Advanced therapy medicinal products (ATMPs) Atomic force microscopy (AFM), 347–348 2,2′-Azobisisobutyronitrile (AIBN), 151 B Bacterial adhesion, 83, 88, 355, 358t Bacteria-produced polyester-type polymers, 9 Beam techniques, patterning/texturing, 318 electron beam, 322 laser beam, 321–322 Beta-cyclodextrin (beta-CD), 295 BIC. See Bone-implant contact (BIC) Bioactive glasses, 50 apatite-layer formation, 281 biocompatibility, 278 (see also Biocompatibility) biodegradability, 278 cells, 278 crystallization, 278–279, 281 gel-casting, 278 hot stage microscopy, 279, 279f mechanical stability, 278 nanocomposites, 60–64 nanofibers, 260 reproducibility, 278 scaffold design, 278 scanning electron microscopy, 278, 279–280f, 280 sintering temperature, 280 thermal properties, 279 Bioceramics applications, 5 ceramics, definition, 4 classification, 5, 6t

438Index

Bioceramics (Continued) disadvantages, 5 first generation, 5–6 nacre-based substitutes, 5 orthopedic implants, 5 properties, 4 second generation, 6–7 third generation, 7–8, 8f Biocompatibility assessment, 34 definition, 34 FDA guideline, 34 genotoxicity studies, 35 ISO 10993-22, 35 medical devices characteristics, 36 classification, 36 reproductive toxicity, 36 risk management process, 34 safety evaluation studies, 34 toxicological evaluation, 35 Bioengineered tissue mimic, 47 Biofilm formation, 83 antibacterial nanoparticles, 83–84 antibacterial strategy, 86, 87f biomaterial’s surfaces, 84 bulk fluid, 84 chemical composition, 83–84 development, 83, 84t grafting1 (see Grafting) hybrid process, 87 ion implantation/plasmas, 87, 109–110 mammalian/microbial cell adhesion, 85 microbial biofilms, 85 protein attachment, 86 roughness parameter, 84 surfaces modifications, 86 surface topography, micro-nanoscale bacterial adhesion, 88 bone-implant applications, 90 CFUs, 88–89 Cicada-type titania nanowire arrays, 88 electron microscopy images, 88–89 grafting integrin-binding peptidic ligands, 90 high-aspect-ratio antimicrobial surface, 89–90

mechanical-bactericidal nanostructures, 88 parameters, 87–88 peptidoglycan membrane, 89 physicochemical mechanism, 89–90 PMMA, 88–89 stainless steel, 90 surface microtopography, 89–90 topography modifications, 88–89 wet etching technique, 90 thin organic/inorganic films, 86 antibacterial carbon-based materials, 100 antibacterial coating, 100 antibacterial properties, 90 chitosan, 91 copper, 92, 94t GO, 100 graphene, 100 hydrothermal method, 100–101 metals/metal oxides, 91 metal sulfide NPs, 100 polymeric molecules, 95, 98–99t self-cleaning properties, 100 silver, 91–92, 93t Staphylococcus aureus, 100 sulfur NPs, 100 TNTs, 95–100 zinc, 92–95, 96–97t Bioglass, 275 Bioinert ceramics, 5–6 Bioluminescence imaging (BLI), 365 Biomimetic tissue, 47 Blasting, 318 Bone homeostasis, 354 Bone-implant contact (BIC), 360 Bone morphogenetic proteins (BMPs), 57–58, 413 Bone sialoprotein (BSP), 413 Bovine serum albumin (BSA), 207–208, 346 Brownian movement, 104 Bulk polymerization, 151 C Calcium carbonate (CaCO3), 172–173 Calcium phosphates (CaPs), 48–49, 174–175, 176t bioceramics, 6–7 microneedle arrays, 391

Index439

Calcium polyphosphate (CaPP), 225 Candida albicans, 85–86 Capillary electrochromatography, 142 Carbon nanotubes (CNTs), 185, 186–187t, 212–214 Carboxymethyl cellulose (CMC), 11–12, 167–168 Cartilage engineering, 65 Cell imprinting cell-membrane-MI strategy, 156–157 cell membrane molecules, 156 whole cell imprinting strategy, 157–158, 157f Cements, 238–239 Ceramic nanostructures. See also Bioceramics collagen fibers, 48 nano-bioactive glass, 50–51 nano-hydroxyapatite, 48–50, 49f organic and inorganic compound, 48 structure, 48, 48f CFUs. See Colony-forming units (CFUs) Chitosan (CS), 11, 63, 65, 91, 95, 168–170, 185–188 N-Chloramines, 106 Cicada-type titania nanowire arrays, 88 C18 linoleic acid, 104 CNTs. See Carbon nanotubes (CNTs) Coated microneedles, 379–380, 383 Coatings, 237–238, 322–323 Cobalt-based alloys, 3–4 Collagen fibers, 9–10 Colony-forming units (CFUs), 88–89 Composites, 239–240, 240f and hybrid materials, 171–172 nanostructures bioactive glass nanocomposites, 60–64 cartilage engineering, 65 hydroxyapatite (see Hydroxyapatite nanocomposites) laponite nanoparticles, 65 metal nanocomposites, 64 osteogenesis, 65 osteoinductive growth factor, 65 Copper, 92, 94t Cryo-field emission gun-scanning electron microscopy (Cryo-FEG-SEM), 228–229, 229f

D DDSs. See Drug delivery systems (DDSs) Dental implants, 1, 5–6, 13, 16 Deoxycholic acid (DCA), 142 Dielectrophoresis (DEP) advantages, 215 cell and tissue manipulation, 215 formation, 215 GelMA C2C12 myotubes, 217 CNTs-GelMA hydrogels, 217–218 hydrogels, 215, 216f mechanical properties, 215–217 negative DEP, 214–215 nerve regeneration, 218 PEDOT:PSS, 218 principle, 214–215 N,N′-Diethyl (4-vinylphenyl)amidine (DEVPA), 142 Dissolving microneedles, 380, 384, 388 DNA-based vaccines, 383, 387 Drug delivery systems (DDSs), 31, 143, 165 E Elastic modulus, 207 Electroplating, 382 Electrospinning technique, 52, 57, 60, 205 Emulsion polymerization, 152 Enterococcus faecalis, 50–51 Equal channel angular pressing (ECAP), 320 Etching, 320 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), 209 Ethylene diamine (EDA), 104 Extended release (ER), 166 Extracellular matrix (ECM), 9–10, 152–153, 203 F Flame spray synthesis, 267 Fluorescein isothiocyanate (FTIC), 207–208 5-Fluorouracil, 188 Foreign body giant cells (FBGCs), 338 Foreign body reaction (FBR), 338 Freeze-dried scaffolds, 9–10 Friction stir processing (FSP), 320 Fungal colonization, 85–86

440Index

G Gentamicin sulfate (GS), 265 Glass-ceramics controlled crystallization, clinical applications, 281–285, 283–285f, 284t nanostructures, 285–288, 287–288f Glycoside hydrolase enzymes, 10–11 Grafting, 426 antibacterial agents, 101 antibiotic agents, 108–109 antimicrobial peptides, 106–108 N-chloramines, 106 N-halamines, 105 PET, 106 polymer brushes, 101–102 quaternary ammonium (see Quaternary ammonium (QAs) compounds) Graphene (G), 100, 180–183 sensors, 389–390 Graphene oxide (GO), 100, 183–184, 186–187t GUMMETAL, 59 H N-Halamines, 105 High-performance liquid chromatography, 142 Hollow microneedle, 380 Hot pressure torsion (HPT), 320 Hyaluronic acid scaffold, 56 Hydrogel-forming microneedles, 380–381 Hydrogels, 166–167 Hydrothermal method, 100–101 Hydroxyapatite (HA) nanocomposites application, 60 electrospinning, 60 inorganic minerals, 60 microscopy images, 60, 63f organic collagen matrix, 60 silicate nanoparticles, 60 stiffness, 60 Hydroxycarbonate apatite (HCA) surface layers, 256 I Ibuprofen, 178 Immunoassays, 142

Inorganic nanoparticles biochemical factors, 294 biomaterials nanocomposite colloidal gels, 302–303, 305f nanocomposite hydrogels, 301–302 regenerative medicine, 301 cancer therapy, 293–294 cross-linkers, 303–306 implant coatings antimicrobial activity, 299–300 bone and dental implants, 300 medical implants, 299 MSC, 297–299, 298f nanoscale topographic surfaces, 297–299 nano-topographies, 297–299 PGLA fiber, 297–299 surface chemistry, 297 surface topography, 297 iron deficiency, 293–294 physical-chemical properties, 293–294 polymeric nanoparticles, 293–294 stem cell tracking, 294–297, 297f International Union of Pure and Applied Chemistry (IUPAC), 259 In vivo tests antibacterial action BLI, 365 electrodeposition treatment, 364–365 intramedullary antibacterial effects, 364–365 nano-porous titania surface layer, 364–365 nano-textured surfaces, 365–366 PI, 365 silver-immobilized HA-IP6-Ti, 365, 366f bioactivity bone area, 360 bone-implant contact, 360 detaching test, 361–364, 363–364f mesh structure, 360, 361f micro-computed tomography technique, 360 optical microscope images, 361, 362f pull-out test, 364 synthetic materials, 360 titanium implants, 361, 362f

Index441

Iron-doped apatite nanoparticles, 242 Iron oxide magnetic nanoparticles, 203–204 Iron oxide nanoparticles (IONPs), 59 Iron oxide superparamagnetic nanoparticles, 209 K Keratin, 10, 14t L Laponite nanoparticles, 65 Laser-peening technology, 318 Laser surface texturing, 110 Laser treatment, 57 Long-chain polymers, 103 M Magnetic nanomaterials, 205 Magnetic nanoparticles, 203–204, 204f Magnetite Fe3O4 nanoparticles (MNPs), 175–178 Magnetron sputtering (MS), 109 Matrix metalloproteinase (MMPs), 156 Medical device regulation (MDR) biocompatibility, 34–36 classification, 31 requirements, 31, 32–33t sterility, 36–38 sterilization validation, 33 technical standards, 31 Medicinal products (MPs), 29–31 Melt quenching process, 265 Mesenchymal stem cells (MSCs), 9–10, 297–299, 298f Mesoporous bioactive glasses (MBG), 259 Mesoporous silica nanoparticles (MSN), 179–180, 295, 297f Mesoporous silicate materials, 263 Metal-based antimicrobial (MBAM) macromolecules, 91–92 Metal microneedle, 382 Metal nanocomposites, 64 Metal nanoscale systems antibiotics bacterial infections, 414 biological effects, 414, 415–416t metronidazole, 417 simvastatin, 417

surface nanostructuring methods, 414 tetracycline, 417 vancomycin, 417 BMPs, 413 BSP, 413 endogenous extracellular matrix proteins, 407 growth factors, 407, 408–411t nanotopography, 407 OPN, 412–413 organic compounds, 407 peptides, 407, 408–411t polyphenols and natural molecules, vegetal origin, 418, 419t proteins, 407, 408–411t SIBLING glycoprotein, 413 stability and durability over time and physiological conditions, 423–426, 427t sterilization effects, 418–423, 421t, 422f, 424–425t surface functionalization, 407, 408–411t TGF-β1 signaling, 412 type I collagen, 407–412 Metal nanostructures applications, 58 cellular internalization, 58 characteristics, 58 growth factor-mediated mechanisms, 59 GUMMETAL, 59 IONPs, 59 nanoparticle surface charge, 58 plasma spray method, 58 silver nanoparticles, 58 sol-gel coating, 59 Ti metal implant, 59 TiO2 nanoparticles, 59 ZnO nanoparticles, 59 Metal sulfide NPs, 100 Metronidazole, 417 Microabrasion/polishing, 320 Micro-computed tomography (μCT) technique, 360 Microfibrous scaffold, 52, 52f Microneedle arrays, 184–185 calcium phosphate particles, 391 characteristics, 379 coated, 379–380 cosmetics, 382

442Index

Microneedle arrays (Continued) dissolving, 380 drug delivery, 379 applications, 383 cancer treatment, 385–387 clinical trials, 384 nanoparticles, 385 phase transition microneedles, 385 two-layer microneedles, 385 vaccines, 383–384 “wrappable” microneedle patch, 388 fabrication techniques, 381–382 graphene sensors, 389–390 hollow, 380 hydrogel-forming, 380–381 influenza vaccination, 379 metal, 382 PDMS, 390 phase-changing nanoparticles, 389–390 sensing applications, 388–389 solid, 379 timeline, 2, 379 transdermal development, 389, 390f types, 379, 381f wearable devices, 390 Microporous scaffold, 52, 52f Minimum inhibitory concentrations (MICs), 357 MMN. See Montmorillonite (MMN) MMPs. See Matrix metalloproteinase (MMPs) MNPs. See Magnetite Fe3O4 nanoparticles (MNPs) Molecularly imprinted polymers (MIPs) advantages, 142 applications, 141 bind macrocyclic metal complexes, 142 capillary electrochromatography, 142 DCA, 142 DDSs, 143 design and synthesis AIBN, 151 binding capacity, 147 bulk polymerization, 151 chemical structures, functional monomers, 149, 150f covalent approach, 148 emulsion polymerization, 152 imprinting efficiency, 147

mechanical stability, 151 molecular memory, 146 multistep swelling method, 152 noncovalent approach, 148–149 porogen, 151 precipitation polymerization, 151–152 prepolymerization complex, 149 semicovalent approach, 149 suspension polymerization, 152 TEMED, 151 template-monomers interactions, 147, 148f template-specific cavities, 146–147, 147f thermolysis/photolysis, 151 three-dimensional architecture, 147 functional monomers, 141 high-performance liquid chromatography, 142 immunoassays, 142 label-free molecular sensors, 142–143 plastic antibodies, 143 recognition properties, 141 regenerative medicine artificial organs, 143–144 cell-cell and cell-matrix interactions, 145 embryonic stages, 145 inflammatory stage, 144 natural healing proces, 144 proliferative phase, 144–145 remodeling stage, 145 wound repair process, 144, 144f scaffolds, 143 separation and quantification, 141–142 soil bacteria, 142 supercritical fluid chromatography, 142 technology, 145–146 template, 141 thermal stability, 142 tissue engineering, 143 advantages, 153 applications, 153 cell imprinting (see Cell imprinting) GRGDS peptide imprinted scaffold, 154 intelligent biomaterial/scaffold, 154, 154f macrolevel structures, 152 MMP-9 imprinted scaffold, 156

Index443

peptide imprinted polymers, 156 proteins imprinted PMMA scaffolds, 156 scaffolds preparation, 154, 155f “smart” systems, 154 synthetic biomaterials, 153 Monoclinic calcium pyrophosphate dihydrate (m-CPPD) crystals, 224 Montmorillonite (MMN), 185–188, 189t MPs. See Medicinal products (MPs) MS. See Magnetron sputtering (MS) MSCs. See Mesenchymal stem cells (MSCs) MSN. See Mesoporous silica nanoparticles (MSN) Multistep swelling method, 152 N Nanochannels, 56 Nanocomposite colloidal gels, 302–303, 305f Nanocomposite hydrogels, 301–302 Nanohydrogels, 166–167 Nanoneedles biological functions, 391 cell behavior, 393, 394f cell electrical signaling, 396–397 cell mechanics, 395–396 cell stimulation, 392 classification, 391–392 drug delivery, 392 fluorescent molecules, 392 intracellular environment, 393–395 nano-delivery, 397–398, 398f tumor cells, 398 Nanopatch arrays, 383–384 Nanoscale delivery systems, 57–58 Nanostructured calcium phosphates. See also Calcium phosphates (CaPs) biomineralizations apatitic compositions, 226 biological relevance, 224, 224f biomimetic synthetic analogs, 226 CaPP, 225 chemical formula, 225 endochondral calcification, 225 extracellular mineralizations, 223 intracellular mineralizations, 223 m-CPPD crystals, 224 nanocrystalline apatites, 226 orthophosphate ions, 225

sHA, 225 surface reactivity, 226 t-CPPD crystals, 224 thermodynamic stability, 226–228, 227t bone mineral “apatitic core”, 228 biomaterial processing, 229–230 “biomimetic” analogs, 228 characterization methods, 234–236, 235f cryo-FEG-SEM, 228–229, 229f nanocrystalline apatites, 230 “nanosized” apatite compounds, 231 physicochemical characteristics, 228–229 physicochemical properties, 231–234, 232f precipitation stage, 229–230 SBF, 230–231 sHA, 228 “surface ionic hydrated layer”, 228 bone regeneration, 242–243 chemical resemblance, 223 hard tissue engineering cements, 238–239 coatings, 237–238 composites, 239–240, 240f nanocrystalline apatites, 236 osteoinductive, 236 scaffolds, 236–237 nanomedicine applications, 240–242, 241f New Legislative Framework, 29 Nickel-titanium (Ni–Ti) alloys, 4 Nitrogen plasma immersion method, 109 Nonantibacterial matrix, 91 Nontoxic polysaccharides, 172 O Optical waveguide lightmode spectroscopy (OWLS), 349 Organic/inorganic hybrids, 171 Orthopedic implants, 5, 13, 275 Osseointegration biological functionalization, 16 biological process, 14–15 chemical characteristics, 15–16 definition, 13–14 implant collar, 16 implant failure factors, 15 in vivo tests (see In vivo tests)

444Index

Osseointegration (Continued) metal biomaterials bacterial contamination, 337 bone formation, 339–342, 340–341f, 343t dental implants, 337 infection issue, 342–344 inflammatory response, 338–339 joint replacements, 337 nanoscale features, 15 nanostructured metal implants, clinical trials, 367 structured surfaces, 15 surface chemistry (see Surface chemistry, metal implants) surface topographies, 16 Osteopontin (OPN), 412–413 Oxygen functional groups, 57 P Patterned nanocomposites additive manufacturing, 218–219 ECM, 203 electric field CNTs, 212–214 DEP (see Dielectrophoresis (DEP)) electrical properties, 218 magnetic assembly, 218 magnetic field advantage, 211 anisogels fabrication process, 211, 214f hydrogel matrix, 211 iron oxide superparamagnetic nanoparticles, 209 magnetic nanoparticles, 203–204, 204f PLGA, 211 polymeric matrix, 209 programmed anisotropic distribution, 211 remote control, 209 self-assembly process (see Selfassembly process) slip-casting process, 209, 212f soft hydrogel, 211 multifunctional biomaterials, 203 multiscale anisotropic organizations, 203 soft hydrogels, 203 stimuli-responsive materials, 218–219

PCL. See Poly(ε-caprolactone) (PCL) pDNA. See Plasmid DNA (pDNA) PEG. See Polyethylenglicole (PEG) PEGp. See Polyethyleneglycol molecules (PEGp) PEI. See Polyethylenimine (PEI) PEO. See Plasma electrolytic oxidation (PEO) Peptide imprinted polymers, 156 Peptides, 407, 408–411t PET. See Polyethylene terephthalate (PET) PGA. See Poly(glycolic acid) (PGA) PHAs. See Poly(hydroxyalkanoates) (PHAs) Phase separation, 52 Phase transition microneedles, 385 Photo-induced grafting, 11–12 Photon intensity (PI), 365 PLA. See Polylactic acid (PLA) Plasma electrolytic oxidation (PEO), 110 Plasma-enhanced chemical vapor deposition, 110 Plasma spray method, 58 Plasma treatment, 56–57 Plasmid DNA (pDNA), 57–58 PMMA. See Polymethylmethacrylate (PMMA) Poly(dimethylsiloxane) (PDMS), 390 Poly(glycolic acid) (PGA), 12 Poly(hydroxyalkanoates) (PHAs), 9 Poly(l-lactic acid) (PLLA), 52, 56 Poly(lactic-co-glycolide) acid (PLGA), 12, 57–58, 211, 417 Polyanionic polysaccharides, 11 Polyethyleneglycol molecules (PEGp), 241 Polyethylene terephthalate (PET), 106 Polyethylenglicole (PEG), 11–12 Polyethylenimine (PEI), 64 Poly (ε-caprolactone) (PCL), 16–17, 417 Polyhydroxyurethane films, 105 Polylactic acid (PLA), 12 Polylactic/glycolic acid (PGLA) fiber, 297–299 Polymeric nanostructures. See also Polymers bulk modification, 57–58 collagen fibers, 51 controlling structures aligned nanofibers, 56 core-shell nanofiber structure, 56 functional nanopolymers, 53, 54–55t hyaluronic acid scaffold, 56 nanochannels, 56

Index445

nonwoven fiber matrix, 53 polymethylmethacrylate (PMMA), 53 SEM images, 53, 53f square lattice symmetry patterns, 53–56 fabrication of, 51–52, 52f polymeric nanofibers, 51 surface modification, 56–57 Polymerization process, 148 Polymer matrix, 47, 64, 148, 207, 209 Polymers bioactive glass nanocomposites, 60–64 degradation rates, 9 hydroxyapatite nanocomposites application, 60 electrospinning, 60 inorganic minerals, 60 microscopy images, 60, 63f organic collagen matrix, 60 silicate nanoparticles, 60 stiffness, 60 mechanical properties, 9 metal nanocomposites, 64 natural polymers, 9 bacteria-produced polyester-type polymers, 9 ECM, 9 PHAs, 9 polysaccharides, 10–12 proteins, 9–10 synthetic polymers, 9 acidic degradation, biodegradable polymers, 13, 14t aliphatic polyesters, 12–13 autocatalysis, 13 mechanical properties, 12 PGA, 12 physical properties, 12 PLA, 12 PLGA, 12 polyester degradation, 12 Polymethylmethacrylate (PMMA), 53, 88–89, 156 Poly(3,4-ethylenedioxythiophene):polystyrenesulfonate (PEDOT:PSS), 218 Polysaccharide-based hybrid materials, 10–12 chemical and physical properties, 165

classification cellulose, 167–168 chemical modifications, 167–168 chemical structure, 167–168, 168f CMC, 167–168 CS, 168–170 sodium alginate, 170–171 structure and properties, 167 controlled drug delivery, 165 drug delivery system, 165 drug-release, 165 ER, 166 hydrogels, 166–167 inorganic phases CaCO3, 172–173 calcium phosphate, 174–175, 176t clays, 188, 189t CNTs, 185, 186–187t controlled drug-release system, 188 5-fluorouracil, 188 graphene, 180–183 graphene oxide, 183–184, 186–187t microneedle arrays, 184–185 MMN, 185–188, 189t MNPs, 175–178 noble metals, 180, 181–182t pH-sensitive hybrid aerogel, 184 polymeric carriers, 172 silicon dioxide, 179–180 TiO2, 178 ZnO, 178–179 molecular release, composites and hybrid materials, 171–172 nanohydrogels, 166–167 nanomaterials-based products, 166–167 sustained drug delivery, 166 Polystyrene magnetic beads (PSMBs), 207–208 Polyvinyl alcohol (PVA), 260 Precipitation polymerization, 151–152 Proteins, 9–10, 152–153, 338–342, 344–350, 407, 408–411t Pull-out test, 364 Push-out test, 364, 364f Q Quantum dots (QDs), 295 Quartz crystal microbalance (QCM) technique, 350

446Index

Quaternary ammonium (QAs) compounds abrasion-resistant QAs-based coatings, 104–105 antibacterial mechanism, 102 Brownian movement, 104 C18 linoleic acid, 104 covalent linkage, 102–103 cytotoxicity, 105 disadvantages, 103–104 EDA, 104 etherification, 104 ex situ and in situ strategies, 103 Gram-positive and Gram-negative bacteria, 103 hydrophilic/hydrophobic properties, 102 long-chain polymers, 103 nonleaching surface graftingfunctionalization, 102 polycations, 104 redox-initiated graft copolymerization, 104 solvent-free processes, 105 surface charge density threshold, 102 thermal-curing coating process, 104–105 R Redox-initiated graft copolymerization, 104 RegenOss, 68 ROBAX, 276–277 Runt-related transcription factor 2 (RUNX2) protein, 57–58 S SBF. See Simulated body fluid (SBF) Scanning force microscopy (SFM), 349 SCOPUS, 275–276, 276–277f Self-assembly process acrylamide/bis-acrylamide monomers solution, 205–207 agarose hydrogel, 207, 208f anisotropic distribution, 205–207 APTS, 207 BSA, 207–208 bulk stiffness, 209 direction-dependent thermogenesis, 207 EDC, 209 elastic compression modulus, 207 elastic modulus, 207 electrospinning technique, 205 enhanced magneto-thermal properties, 205

FTIC, 207–208 heat-activated polymerization, 205–207, 206f magnetic coupling, 205–207 magnetic nanomaterials, 205 magnetic nanostructures, 205 magneto-thermal effect, 207 manufacturing process, 209, 211f nanomagnetic system, 205 polymeric matrix, 207 PSMBs, 207–208 three-dimensional (3D) topography, 208 Severe plastic deformation (SPD), 320 SFM. See Scanning force microscopy (SFM) sHA. See Stoichiometric CaP hydroxyapatite (sHA) Shaking flask method, 357 Shot peening, 318 SIBLING glycoprotein, 413 Silica matrix, 295 Silicate-based nanoceramics, 6–7 antibacterial effect, 257 apatite precipitation, 258 biomedical applications, 255 bone matrix formation and calcification, 255 chemical compositions, 258, 258t composites, 267–268 condensation and repolymerization, 256 glass-ceramics, 255 glass/glass-ceramic biomaterials, 256 growth factors and mineralization, 256–257 HCA surface layers, 256 interfacial reactions, 256, 257f macrophages, 257 materials technology, 255 MBG, 259 nanofibrous silicate bioceramics, 259–260 porous silicate bioceramics, 258–259 regenerative medicine biological functions, 260 bone repair, 260–262 drug delivery, 263–265, 264f metallic ions, 260, 261f tissue engineering, 260 wound healing, 262–263, 262f SBF, 259 silicate-based bioactive glasses, 255 silicate-based nanoparticles, 259

Index447

Si-OH, 256 soft tissues, 256 synthesis and process, 265–267, 266f Silicate nanoparticles, 60 Silicon dioxide, 179–180 Silk, 9–10 Silver, 91–92, 93t Silver nanoclusters (Ag NCs), 91–92 Silver nanoparticles, 58, 64, 68 Simulated body fluid (SBF), 57, 230–231, 259 Simvastatin, 417 Slip-casting process, 209, 212f Sodium alginate, 170–171 Soft hydrogels, 203, 211 Sol-gel glass nanofibers, 260 Sol-gel method, 59, 63–64, 265–266, 266f Solid microneedles, 379 Spark plasma sintering, 236–237 SPD. See Severe plastic deformation (SPD) SPIO. See Superparamagnetic iron oxide (SPIO) SPR. See Surface plasmon resonance (SPR) Stainless steel (AISI 316L), 3 Staphylococcus aureus, 100 Stem cell tracking, 294–297, 297f Sterilization methods, 37–38 Stoichiometric CaP hydroxyapatite (sHA), 225 Stoichiometric hydroxyapatite, 228 Streptococcus salivarius, 109 Sulfur NPs, 100 Supercritical fluid chromatography, 142 Superparamagnetic iron oxide (SPIO), 295 Surface chemistry, metal implants antibacterial action, 356, 359t apatite precipitation, 350–352, 351–354f bacterial adhesion, 355, 358t bone homeostasis, 354 cell adhesion and spreading, 354–355 characteristics, 355 in vitro tests, antibacterial biomaterials disk diffusion test, 357 film contact method, 356–357 loop-type laboratory biofilm reactor, 357–360 minimum inhibitory concentrations, 357 shaking flask method, 357 nano-textured titanate layer, 355 protein adsorption

atomic force microscopy, 347–348 charge transfer, 344–345 dynamic mechanism, 348–349 ellipsometry, 349 hydrogen bonding, 344–345 hydrophobic surfaces, 345–346 incubator environment, 348 isoelectric point, 346 optical waveguide lightmode spectroscopy, 349 quartz crystal microbalance technique, 350 rotational diffusion, 348 saturation coverage, 348 scanning force microscopy, 349 soft proteins, 345–347 superhydrophilic surfaces, 346 surface charge, 346 surface features, 344–345, 345f surface modifications, 344 surface plasmon resonance, 349 surface topography, 348 time-of-flight secondary ion mass spectrometry, 350 water molecules, 344–345 zeta potential gap, 347 Surface immobilization, 101–102 Surface nanostructuring methods, 414 Surface plasmon resonance (SPR), 349 Surface topographies biocompatibility, 316 biological response, 324–327 biological system, 316, 317f biometallic materials, 322–323 design, 323–324, 326f finite element simulations, 315 grooves, 323 hierarchical architecture, 323 irregular topography, 323 mechanical properties, 327 orthopedic devices, 315 osseointegration, 316 properties, 315–316 surface chemistry and charge, 327–328 surface modification biomedical application, 318, 319f blasting, 318 coating, 322–323 etching, 320

448Index

Surface topographies (Continued) laser-peening technology, 318 microabrasion/polishing, 320 micro and nanomodification, 316 nanostructuring, 320 patterning/texturing, 318, 321–322 physicochemical properties, 316 shot peening, 318 ultrasound, 320–321 thermomechanical process, 315 tissue regeneration, 316 ultrafine grains, 323 wettability, 328 Surveillance, 41 Suspension polymerization, 152 T Tetracycline, 417 N,N,N′,N′-Tetramethylenediamine (TEMED), 151 Thermal-curing coating process, 104–105 Thermogravimetry (TG), 235 Time-of-flight secondary ion mass spectrometry (ToF-SIMS), 350 Tissue regeneration, 47 clinical performance and benefit, 39 clinical products in vitro and in vivo studies, 65 nanostructured products, 65, 66–67t polymer nanoparticles, 68 RegenOss, 68 silver nanoparticles, 68 MDR (see Medical device regulation (MDR)) nano-manufacturing, 40 “nano-product” commercialization, 40 nanostructured biomaterials, 69–70 nanostructured devices, 40 patterned nanocomposites (see Patterned nanocomposites) postmarket activities surveillance, 41 vigilance, 41–42 products/processes, 40 quality control, 40–41, 44f regulatory landscape “ancillary” medicinal substance, 30 drug delivery product, 31 industrial products, classification, 29

vs. intellectual property, 44–45 “mechanism of action”, 29 medical device, 29 MPs, 29 nanomaterials, definition, 30 nanoparticles, 30 New Legislative Framework, 29 roles and responsibilities competent authorities, 43 manufacturers, 42–43 notified bodies, 43–44 safety, 38–39 surface topographies (see Surface topographies) Tissue repair and regeneration. See also Tissue regeneration biomaterial design, 1–2 biomedical materials, 1–2 ceramics applications, 5 classification, 5, 6t definition, 4 disadvantages, 5 first generation, 5–6 nacre-based substitutes, 5 orthopedic implants, 5 properties, 4 second generation, 6–7 third generation, 7–8, 8f material chemistry and processing technologies, 1 metals see also Metal nanostructures applications, 2 bioactive behavior, 2–3 biological requirements, 2–3 cobalt-based alloys, 3–4 complications, 2–3 elastic deformations, 2 implants failure, causes, 2, 2f mechanical characteristics, 2 multifunctional surfaces, 3 Ni–Ti alloys, 4 plastic deformations, 2 properties, 2 titanium and titanium alloys, 3 nanomaterials, 18 nanoparticles applications, 18–19 cancer management, 19

Index449

characteristics, 19 imaging probes, 19 nonspecific drugs, 20 stem cell engineering, 21 “theranostic” agents, 19, 20f osseointegration (see Osseointegration) polymers (see Polymers) scaffolds, 16–18 Titania nanotubes (TNTs), 95–100 Titanium and titanium alloys, 3, 108 Titanium oxide (TiO2) nanoparticles, 59, 178 ToF-SIMS. See Time-of-flight secondary ion mass spectrometry (ToF-SIMS) Triclinic calcium pyrophosphate dihydrate (t-CPPD) crystals, 224 Two-layer microneedles, 385 U Ultrasound, 320–321 Ultraviolet (UV) illumination, 295

V Vancomycin, 417 Vigilance, 41–42 W Wet etching technique, 90 Wollastonite nanowires, 259–260 Wound healing, 262–263, 262f Wound repair process, 144, 144f X X-ray diffraction, 234 Z Zinc, 92–95, 96–97t Zinc oxide (ZnO) nanoparticles, 59, 178–179 Zirconia implants, 5–6 Zwitterionic polymers, 86