Advanced Technologies and Polymer Materials for Surgical Sutures 0128197501, 9780128197509

Citation preview

Woodhead Publishing Series in Biomaterials

ADVANCED TECHNOLOGIES AND POLYMER MATERIALS FOR SURGICAL SUTURES Edited by

SABU THOMAS Vice Chancellor Mahatma Gandhi University & Director, International and Inter University Centre for Nanoscience and Nanotechnology, Mahatma Gandhi University Kottayam, Kerala, India

PHIL COATES Professor of Polymer Engineering Department of Mechanical and Energy Systems Engineering University of Bradford, Bradford, United Kingdom

BEN WHITESIDE Professor of Precision Manufacturing Department of Mechanical and Energy Systems Engineering University of Bradford, Bradford, United Kingdom

BLESSY JOSEPH Post-Doctoral Research Fellow Business Innovation and Incubation (BIIC) Mahatma Gandhi University Kottayam, Kerala, India

KARTHIK NAIR New Product Introduction Lead Summit Medical Ltd., Bourton on the water, United Kingdom

Woodhead Publishing is an imprint of Elsevier 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, OX5 1GB, United Kingdom Copyright © 2023 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. ISBN: 978-0-12-819750-9 For information on all Woodhead Publishing publications visit our website at https://www.elsevier.com/books-and-journals Publisher: Matthew Deans Acquisitions Editor: Sabrina Webber Editorial Project Manager: Rafael G. Trombaco Production Project Manager: Prasanna Kalyanaraman Cover Designer: Christian Bilbow Typeset by TNQ Technologies

Contributors Samson Afewerki Division of Engineering in Medicine, Department of Medicine, Brigham & Women’s Hospital, Harvard Medical School, Boston, MA, United States; Division of Health Science and Technology, Harvard University e Massachusetts Institute of Technology, MIT, Cambridge, MA, United States R. Anjana Centre for Nanotechnology Research, Vellore Institute of Technology, Vellore, Tamil Nadu, India Hemand Aravind Aromatic and Medicinal Plant Research Station, Kerala Agricultural University, Thrissur, Kerala, India Allan Babu School of Energy Materials, Mahatma Gandhi University, Kottayam, Kerala, India Emilio Bucio Departamento de Química de Radiaciones y Radioquímica, Instituto de Ciencias Nucleares, Universidad Nacional Autónoma de México, Circuito Exterior, Ciudad Universitaria, CDMX, Mexico Haritha R. Das School of Energy Materials, Mahatma Gandhi University, Kottayam, Kerala, India Lorena Duarte-Peña Departamento de Química de Radiaciones y Radioquímica, Instituto de Ciencias Nucleares, Universidad Nacional Autónoma de México, Circuito Exterior, Ciudad Universitaria, CDMX, Mexico Amira J. Fragoso-Medina Departamento de Ciencias Químicas, Facultad de Estudios Superiores Cuautitlán, Universidad Nacional Autónoma de México, Mexico City, Mexico Samarah Vargas Harb Department of Materials Engineering (DEMa), Federal University of São Carlos (UFSCar), São Carlos, São Paulo, Brazil Jemy James University Bretagne Sud, Lorient, France M.S. Jisha School of Bioscience, Mahatma Gandhi University, Kottayam, Kerala, India

ix

x

Contributors

Blessy Joseph Business Innovation and Incubation (BIIC), Mahatma Gandhi University, Kottayam, Kerala, India Nandakumar Kalarikkal International and Inter University Centre for Nanoscience and Nanotechnology, Mahatma Gandhi University, Kottayam, Kerala, India Saravanan Krishnan Creative Carbon Labs Pvt Ltd., Chennai, Tamil Nadu, India Hiran Mayookh Lal School of Energy Materials, Mahatma Gandhi University, Kottayam, Kerala, India Anderson O. Lobo LIMAV - Interdisciplinary Laboratory for Advanced Materials, Department of Materials Engineering, UFPI - Federal University of Piauí, Teresina, Piauí, Brazil Felipe López-Saucedo Departamento de Química de Radiaciones y Radioquímica, Instituto de Ciencias Nucleares, Universidad Nacional Autónoma de México, Circuito Exterior, Ciudad Universitaria, CDMX, Mexico Deepthy Menon Amrita Centre for Nanosciences and Molecular Medicine, Amrita Vishwa Vidyapeetham, Kochi, Kerala, India Teena Merlin School of Bioscience, Mar Athanasios College for Advanced Studies, Tiruvalla, Kerala, India Karthik Nair Summit Medical Ltd., Bourton on the water, United Kingdom Ashwin Kumar Narasimhan Department of Biomedical Engineering, SRM Institute of Science and Technology, Kattankulathur, Chennai, Tamil Nadu, India Neethu Ninan Clinical and Health Sciences, University of South Australia, Adelaide, South Australia, Australia Smrithi Padmakumar Amrita Centre for Nanosciences and Molecular Medicine, Amrita Vishwa Vidyapeetham, Kochi, Kerala, India Thella Shalem Rahul Department of Biomedical Engineering, SRM Institute of Science and Technology, Kattankulathur, Chennai, Tamil Nadu, India Alejandro Ramos-Ballesteros Notre Dame Radiation Laboratory, University of Notre Dame, Indiana, United States

Contributors

xi

Guillermo U. Ruiz-Esparza Division of Engineering in Medicine, Department of Medicine, Brigham & Women’s Hospital, Harvard Medical School, Boston, MA, United States; Division of Health Science and Technology, Harvard University e Massachusetts Institute of Technology, MIT, Cambridge, MA, United States Morvarid Saeinasab Department of Biomedical and Electronics Engineering, School of Engineering, University of Bradford, Bradford, United Kingdom Farshid Sefat Department of Biomedical and Electronics Engineering, School of Engineering, University of Bradford, Bradford, United Kingdom; Interdisciplinary Research Centre in Polymer Science & Technology (Polymer IRC), University of Bradford, Bradford, United Kingdom Rukhsar Shah Department of Biomedical and Electronics Engineering, School of Engineering, University of Bradford, Bradford, United Kingdom Mohammad-Ali Shahbazi Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy University of Helsinki, Helsinki, Finland Thiago Domingues Stocco Faculty of Medical Science, State University of Campinas (UNICAMP), Campinas, São Paulo, Brazil; University of Santo Amaro, Santo Amaro, São Paulo, Brazil Louise Taylor Department of Biomedical and Electronics Engineering, School of Engineering, University of Bradford, Bradford, United Kingdom Sharin Maria Thomas Centre for Nanotechnology Research, Vellore Institute of Technology, Vellore, Tamil Nadu, India Sabu Thomas International and Inter University Centre for Nanoscience and Nanotechnology, Mahatma Gandhi University, Kottayam, Kerala, India Arya Uthaman School of Energy Materials, Mahatma Gandhi University, Kottayam, Kerala, India Smitha Vijayan School of Bioscience, Mar Athanasios College for Advanced Studies, Tiruvalla, Kerala, India

xii

Contributors

Ximu Zhang Chongqing Key Laboratory of Oral Disease and Biomedical Sciences, Chongqing Municipal Key Laboratory of Oral Biomedical Engineering of Higher Education, Stomatological Hospital of Chongqing Medical University, Chongqing, Sichuan Province, China; State Key Laboratory of Oral Diseases & National Clinical Research Center for Oral Diseases & Dept. of Preventive Dentistry, West China Hospital of Stomatology, Sichuan University, Chengdu, Sichuan Province, China Wei Zhang State Key Laboratory of Polymer Materials Engineering, Polymer Research Institute at Sichuan University, Chengdu, Sichuan Province, China

CHAPTER 1

Advances in biopolymer based surgical sutures Blessy Joseph1, Jemy James2, Nandakumar Kalarikkal3 and Sabu Thomas3 1

Business Innovation and Incubation (BIIC), Mahatma Gandhi University, Kottayam, Kerala, India; University Bretagne Sud, Lorient, France; 3International and Inter University Centre for Nanoscience and Nanotechnology, Mahatma Gandhi University, Kottayam, Kerala, India

2

1.1 Introduction Over the years, there has been a dramatic growth of the wound closure market. Traditionally materials like silk, cotton, horsehair, animal tendons and intestines, and wire made of precious metals were in operative procedures. The limitations and risks associated with such wound closure devices demanded the need for efficient and cost-effective techniques for wound healing. Although there have been significant advances in tissue adhesives and other mechanical wound closure devices, sutures have been the preferred choice for surgeons. Sutures can be defined as the materials used to uphold tissues together normally after a trauma or surgery [1]. They can be natural or synthetic materials that can provide adequate mechanical strength during tissue fixation. The art of suturing can be found in the Egyptian mummified resins, in which they have used woolen threads, plant fibers, hair, and tendons. Suturing techniques were documented in 500 BCE (Before Common Era) by Indian surgeon Sushruta in “Sushruta Samhita [2].” Metal wires were first applied in the human body by French physicists Lapayode and Sicre in 1775 to set a broken humerus (upper arm bone)[3]. A fundamental change was witnessed following Second World War, after which polymer sutures and stainless steel became superior. The selection of suture material is dependent on the physical and biological characteristics of the suture as well as the type of tissue to be healed. Sutures are made from synthetic or natural polymers. Synthetic polymers are not readily degradable. They accumulate and can have a long-term detrimental effect on ecosystems. The tunable physical characteristics of biopolymers make them a reliable material for the fabrication of sutures. Biopolymers can be obtained from natural sources or synthesized chemically from Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00008-5

© 2023 Elsevier Ltd. All rights reserved.

1

2

Advanced Technologies and Polymer Materials for Surgical Sutures

biological material or entirely biosynthesized by living organisms[4]. They are easily biodegradable as they are obtained from renewable sources. The term “biodegradation” generally refers to degradation by microorganisms. The polymer is broken down into carbon dioxide and water which forms food for microorganisms[5]. Biopolymers as surgical sutures have gained considerable attention because of their unique properties like biocompatibility and biodegradability. Biopolymers can adopt more precise and defined 3D shapes and structures when compared to synthetic polymers having more simple and random organization[6]. This makes biopolymers attractive for in vivo applications. They are generally classified into three categories based on the nature of repeating units they are composed of (i) polysaccharides, often carbohydrate structures (cellulose, chitin, starch, alginate, etc.); (ii) polypeptides made of amino acids (collagen, actin), and (iii) polynucleotides deoxynucleic acid (DNA) and ribonucleic acid (RNA)) (Fig. 1.1). This chapter intends to provide an overview of the biopolymers used for suture fabrication, their physical and biological properties, and how these properties facilitate wound repair. Sterilization techniques used for sutures have also been discussed in this chapter.

1.2 Polymers as suture materials In the past sutures made of natural materials like dried animal gut, animal hair (e.g., horse hair), silk, tendons, and plant fibers (e.g., linen, cotton) were widely used [7,8]. The technological advancements in polymer science paved way for the development of sutures with diverse materials having excellent mechanical and physical properties. There has been a large-scale expansion and evolution of the research and business in the area of materials for biomedical applications. Still, sutures and staples are the most used material in the biomedical industry. Sutures are to be used in many cases where natural wound closure is difficult and external

Figure 1.1 Classification of biopolymers according to their structure.

Advances in biopolymer based surgical sutures

3

reinforcement is highly essential. Biopolymer based absorbable sutures are much preferred to nonabsorbable sutures. Sutures are generally classified as absorbable and nonabsorbable based on whether they degrade or not after performing the intended function. Nonabsorbable sutures need to be removed by doctors, hence causing additional discomfort to patients. Whereas absorbable sutures degrade within the body usually by hydrolysis or with the aid of proteolytic enzymes[9].

1.3 Biopolymers Environmental problems arise with the continued use of synthetic polymers. Intensive research has been carried out in this direction, possibly replacing synthetic polymers with natural ones. As mentioned above, biopolymers are obtained from biological sources. Hence, the use of biopolymers offers an eco-friendly approach. They are decomposed by microorganisms or natural processes like availability of moisture, sunlight, etc. which is environmentally friendly when compared to petroleum based synthetic polymers releasing toxic byproducts into the surroundings. Biopolymers are employed in diversified fields such as food packaging, drug delivery, tissue engineering, etc. Although they are biocompatible, many of them lack sufficient mechanical properties desired for medical applications. Most often they are crosslinked or modified with materials like glutaraldehyde, citric acid, poly (carboxylic acids), and so forth[10]. Crosslinkers like glutaraldehyde can be cytotoxic hence greener approaches are also being explored. Nanoparticles are also used to enhance the properties of biopolymers. The interaction between biopolymers and nanoparticles results in nanocomposites with improved functionalities like antimicrobial property, tensile strength, thermal stability, or water resistance. Many researchers have investigated the ability of silver nanoparticles (AgNPs ) to improve the antimicrobial properties of biopolymers, wherein cost-effective methodologies could be formulated for developing wound dressings or food packaging films. Cellulose paper coated with silver-gold nanoparticles displayed improved antibacterial activity against E.coli [11]. Another work reported the synthesis of silver-cellulose hybrids which showed excellent antibacterial activity against E.coli and S.aureus whereas pure cellulose (Microcrystalline cellulose) didn’t exhibit any activity against the respective microbial strains[12]. Although several biopolymers find promising applications in the biomedical sector, we will be concentrating on polymers like cellulose, collagen, silk, chitosan, chitin, polyhydroxyalkanoates (PHA), and PLA which particularly fit well for the suture industry.

4

Advanced Technologies and Polymer Materials for Surgical Sutures

Cellulose is the naturally occurring homopolymer consisting of b-1, 4 linked glucan chains. Being inherently biodegradable and low-cost material, cellulose finds immense application in healthcare [13]. Cellulose materials try to self-assemble and form an extended network by both intramolecular and intermolecular hydrogen bonds, which makes them relatively stable. Chitin is a sustainable biopolymer due to its abundance. Structurally, chitin is N-acetyl glucosamine and the main source are crustaceans like crabs, shellfish, etc. The deacylated form of chitin known as chitosan consists of N-acetyl glucosamine and glucosamine moieties. Both chitin and chitosan are versatile enough to be processed to any form like sponges, gels, or scaffolds, thereby finding many applications in tissue engineering and drug delivery[14]. Natural silk fibers are produced by arthropods like silkworms or spider. Mulberry silkworms (Bombyx mori) are most commonly reared to produce silk. They have a core-shell structure consisting of 3 components, a heavy chain fibroin, a light chain fibroin, and a third small glycoprotein, known as the P25 protein. These proteins are coated with hydrophilic sericins. Silk materials are used as sponges, films, or sutures for applications like ligament tissue engineering, hepatic tissue engineering, cartilage tissue engineering, and so on [15e17]. Poly(lactic acid) (PLA) is a biodegradable polyester produced from the monomer, lactic acid (LA) by mechanisms like direct polycondensation (DP) and ring-opening polymerization (ROP). The tunable physicochemical properties and biocompatibility of PLA make it suitable for biomedical applications. Collagen is a major structural protein in animals and forms a vital part of the extracellular matrix. It provides tensile strength to tendons and ligaments and also elasticity to the skin. It has a 3D architecture comprising of a righthanded bundle of three parallel, left-handed poly proline II-type helices [18]. Source of collagen includes bovine skin and tendons, porcine skin, marine organisms like sponges, fish, and jelly fish. It is used for soft tissue repair, dental applications, and as scaffolds for tissue engineering [19,20]. Polyhydroxyalkanoates (PHA) are naturally synthesized polyesters accumulated as energy storage material inside the cellular structure of various microorganisms.

1.4 Biopolymers for sutures 1.4.1 Collagen Collagen nanofibrils (CoNF) have a great potential for being mechanically strong but biodegradable sutures. They play a major role in tissue

Advances in biopolymer based surgical sutures

5

engineering as being the key component of the extracellular matrix. It is the most abundant protein in the human body and imparts structural integrity and strength to the tissues[21]. The use of collagen as a modern biomaterial began in 1881. Joseph Lister and William Macewen (Fig. 1.2) reported the advantages of catgut, a collagen-rich biomaterial prepared from the small intestine of sheep[23]. Untreated catgut sutures are often processed from dead animal tissue, hence causing infections[24]. They are often used in the case of subcutaneous or fatty tissue[25]. Collagen sutures were modified with heparin for sustained release of platelet-derived growth factor-BB (PDGF-BB). Tendon-derived cells seeded on PDGF-BB incorporated collagen sutures showed 50% greater proliferation than untreated collagen sutures[26]. This could be because collagen provides active chemical sites for conjugating growth factors. Collagen has also been used to coat surgical sutures to improve their functionalities. Polyester/polyethylene sutures coated with collagen were evaluated for their response to bone and tendon cells[27]. Collagen coating was found to stimulate proliferation and adhesion of cells in collagen coated sutures when compared to uncoated one. 1.4.2 Polylactic acid (PLA) PLA is one of the most popular biodegradable and bio-based polymers. PLA is used to prepare biodegradable polymer sutures [28]. The biocompatibility of the polymers has been extensively studied, and it has been proven to be one of the best biopolymers for biomedical applications like sutures etc. [29]. PLA is a polymer derived from LA and its structure makes it easily breakable during metabolism and thereby making it easier to be excreted from the body [30]. Degradation occurs through enzymatic or hydrolytic scission of ester bonds. The degradation of PLA depends on its

Figure 1.2 SEM images of PLA suture loaded with PM-Ds: (A) 100
times, (B) 1000
times. (Reproduced with permission from Ref. [22].)

6

Advanced Technologies and Polymer Materials for Surgical Sutures

molecular weight, crystallinity, presence of fillers, etc. Recently, Liu et al. reported the fabrication of PLA sutures loaded with PLA microspheres containing drug[22]. Initially PLA microspheres containing drug gentamicin sulfate was prepared (PM-Ds). Further, this drug loaded microspheres were loaded onto the PLA sutures (PM-Ds/PLA). The mechanical properties were analyzed which showed an increase in the properties of the drug loaded suture when compared to the neat suture. A sustained release of the drug up to 8 days could be achieved. As evident from the scanning electron microscopy images, the microspheres entered the gaps of the suture fibers, and stuck to them firmly which could have resulted in the prolonged release of the drug (Fig. 1.2). In another study, biopolymers like chitosan, alginate, and the blends of these polymers were coated on the surface of PLA sutures. The mechanical studies were carried out. Some of the drugs based on antibiotic sensitivity was chosen and was introduced into the sutures using surface treatment method like dip coating. The drug release studies and antimicrobial activity proved that the drug-coated bio polymeric sutures were effective in wound closing and wound healing [31]. Poor biocompatibility and cellular affinity are major problem encountered with PLA sutures. To improve the surface hydrophilicity, PLA sutures were initially treated with lipase followed by grafting with chitosan [32]. It’s evident from the SEM images that initially the untreated sutures had a smooth surface. Once grafted with chitosan, in some places chitosan united and led to a rougher surface and large friction coefficient. However, hydrophilicity was greatly improved. Blends of PLA and polycaprolactone compatibilized with Ethyl Ester LLysine Triisocyanate (LTI) were melt-spun to produce suture threads of diameter 0.3 mm 1.0 phr of LTI was found to be the most suitable composition for producing sutures, at higher loadings the sutures were too rigid. The suture threads didn’t induce any bacterial growth [33]. 1.4.3 Silk Silk is a protein polymer whose characteristics are slow degradation and good mechanical strength. Silk is preferred for cardiovascular, neurological, and ophthalmic procedures [34]. The ease of handling and improved knot security properties makes silk superior among other sutures. But their use is hindered due to the high inflammatory reactions posed by them [35,36]. Bacterial attachment to silk sutures was compared to commercially available Monocryl Plus suture [37]. From Fig. 1.3 it is evident that the

Advances in biopolymer based surgical sutures

7

Figure 1.3 Scanning electron microscope images of (A) silk suture knot material and (B) Monocryl Plus suture knot material. Microorganisms and cellular detritus are highly visible in silk sutures. (Reproduced with permission from Ref. [37].)

microorganisms were highly colonized around the suture knot of silk suture when compared to that of Monocryl Plus suture. Maintaining sterile conditions in the wound has always been a hurdle after suturing. Medical devices and sutures contribute about 45% of nosocomial infections or hospital-acquired infections [38]. Antibacterial sutures play a pivotal role in combating surgical site infections[39]. Once a biofilm is formed on the surface of a suture, it becomes resistant to traditional antimicrobials. Once bacteria colonize a suture, local methods to treat bacterial infections become inadequate. Hence, several strategies to prevent bacterial adherence have been proposed by researchers including the addition of antibiotics, nanoparticles, biomaterials, etc. Sutures impregnated with antibiotics have been found to prevent the adherence of bacteria and biofilm formation [40]. Tetracycline hydrochloride (TCH), a bacteriostatic drug is found to exhibit activity against a wide range of gram-positive and gram-negative microorganisms[41]. The efficacy of TCH-treated sutures was studied by Viju and Thilagavathi [42]. As was expected, untreated silk sutures promote the growth of E.coli and S.aureus. Synergistic chitosan and TCH drug was exploited to develop antimicrobial silk sutures for preventing microbial infections [43]. Such combinations can provide a prolonged antibacterial effect. AgNPs have been widely used as an antibacterial agent[44,45]. AgNPs exhibits their antimicrobial potential through various mechanisms. The anchoring of AgNPs to

8

Advanced Technologies and Polymer Materials for Surgical Sutures

microbial cells, followed by penetration into the cells, reactive oxygen species and free radical generation, and modulation of microbial signal transduction pathways have been recognized as the most prominent ways of antimicrobial action [46]. AgNPs were coated on silk sutures to impart antibacterial properties [47]. Mechanical strength was retained after the addition of AgNPs; however, a significant reduction in bacterial growth was achieved. Cytotoxicity studies using 3T3 mouse embryonic fibroblast cells showed 82% cell viability for silver treated samples. This showed that the silver treatment did not affect their proliferative capacity. Surface modification of silk fibroin suture, AASF (antheraea assama, popularly known as golden silk; found only in certain parts of Assam) was achieved by grafting polypropylene (PP) onto silk fibroin sutures[48]. Here the sutures were first sterilized using argon and then low-temperature plasma grafting of PP onto sterilized sutures was done to achieve the desired biofunctionalities. Here the modified suture showed more biocompatibility and improved wound healing when compared to the untreated ones. In vivo studies were conducted in three groups. The first group was sutured with AASF, the second with argon plasma-treated AASF (AASFAr) and the third group with PP grafted AASF sutures (PP-AASF). The histopathology studies on the 14th postoperative day show the presence of inflammatory cells in group A characterized by lesser collagen formation (Fig. 1.4). Group B shows a considerably fair amount of collagen formation with slight infiltration in and around hair follicles. Whereas PPAASF sutured group (Group C) shows highly accelerated wound healing activity. Moreover, a greater amount of hair follicles was also present when compared to the other groups.

Figure 1.4 Histologic evaluation of wound healing on 14th postoperative day. Histopathological section of the sample collected from the incised wound of (A) group A, (B) group B, and (C) group C animals shows inflammatory cell inflammation (IN) in and around the hair follicle (HF) as well as subepidermal tissue. Proliferation of fibrous connective tissue (CT) indicates faster healing of group B and group C as compared to group A animals. (Reproduced with permission from Ref. [48].)

Advances in biopolymer based surgical sutures

9

1.4.4 Chitin & chitosan Chitin and chitosan are polymers derived from marine animals and is some of the most available biopolymers other than cellulose. However, some of the challenges make its usage cumbersome [49]. Though chitin is highly biocompatible, nontoxic, and biodegradable, along with its antimicrobial effect, there are still more challenges to overcome to exploit its huge potential for prospective applications [50]. Chitosan is a potent antimicrobial agent and its antimicrobial activity can be attributed to its cationic nature [51]. The positively charged chitosan molecules interact with negatively charged microbial cell membranes leading to the disruption of the microbial membrane [52]. Sutures were fabricated from chitin having good mechanical strength [53]. No allergic reactions or inflammation was seen. The chitin suture was absorbed in about 4 months in rat muscles. The accelerated degradation can be mainly due to the action of lysozyme. Chitin nanofibrils are used as nanofillers for reinforcing polymers to obtain nanocomposites with enhanced stability, especially in the case of bioresorbable sutures [54]. Chitosan stimulates tissue regeneration and prevents scar formation. The mechanical strength of chitosan is very low; hence, it is mainly exploited as suture coatings. Chitosan has been used for coating silk sutures [55]. Silk sutures coated with chitosan also showed excellent antibacterial efficacy [56]. A modified derivative of chitosan known as hydroxyl propyl trimethyl ammonium chloride (HACC) chitosan coated on Vicryl suture showed excellent antibacterial activity and also displayed good biocompatibility [57]. HACC is a water-soluble modified derivative of chitosan that exhibits good antibacterial activity [58,59]. HACC coated sutures effectively prevented biofilm formation when compared to triclosan-coated sutures. Prabha et al. showed that extracted chitosan (EC) from crab shells showed higher inhibition of biofilm formed by mixed species[60]. The antibacterial and antifungal effects of Vicryl absorbable sutures coated with chitosan, uncoated sutures, and commercially available triclosan-coated sutures were studied against S. epidermidis and C. albicans (Fig. 1.5). The uncoated suture (control) and sutures soaked with acetic acid (Vehicle control) did not show any antibacterial or anticandidal activity. The commercial triclosan-coated sutures exhibited only antibacterial activity and did not show any anticandidal activity. EC immobilized sutures exhibited good antimicrobial activity against both strains compared to commercially available chitosan (CC).

10

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 1.5 Antimicrobial activities of the impregnated suture against S. epidermidis and C. albicans: (A) Control; (B) VC; (C) EC (400 mg/mL); (D) CC (400 & 450 mg/mL); (E) Triclosan coated Reproduced with permission from Ref. [60].

1.4.5 Polyhydroxyalkanoate (PHA) PHA is a microbial polyester having excellent biocompatibility and biodegradability. Poly(3-hydroxybutyrate) (PHB) is the most widespread member of the polyhydroxyalkanoate family and is produced under unbalanced growth conditions like depletion of essential nutrients such as nitrogen, phosphorus, or magnesium [61]. Poly (4-hydroxybutyrate) (P4HB) is a typical PHA type used for the fabrication of surgical materials. The most well-known product, and the first approved by the US Food and Drug Administration, is the TephaFLEX suture fabricated from P4HB [62]. In vivo studies of PHA sutures implanted intramuscularly over 1 year showed that animals that received the sutures were in good health condition during the period of study. No adverse reactions were observed, and functional characteristics of the animals were also not affected [63]. Polyhydroxyalkanoate sutures decreased tendency to curl were fabricated by extrusion and orientation of the fibers [64]. The resulting fibers had an elongation to break from about 17% to about 85% and Young’s modulus of less than 350,000 psi. He et al. evaluated the biocompatibility of monofilament made from poly (3-hydroxybutyrate-co-3-hydroxyhexanoate) (PHBHHx) and a multifilament made from poly(3-hydroxybutyrate-co-3hydroxyvalerate) (PHBV) and PLA blend [65]. The PHBHHx fiber and the PHBV/PLA fiber showed remarkable biocompatibility to be used as surgical sutures.

Advances in biopolymer based surgical sutures

11

1.4.6 Cellulose Plants are the major source of cellulose, the most abundant and easily available carbohydrate polymer on earth. Bacterial cellulose, an alternate and highly purified form of cellulose is produced by aerobic bacteria, mainly of the genus Acetobacter [66]. Their unique nanostructure, excellent water retaining capacity, good mechanical strength, and high crystallinity makes them preferred choice in biomedical applications [67]. Bacterial cellulose nanocrystals (BCNC) were used to reinforce chitin (RC) fibers to form BCNC/RC yarns[68]. The fibers were produced by wet spinning technology for application as surgical sutures. In vitro studies showed good biocompatibility and in vivo studies revealed good wound healing with BC coated yarns. However, the knot pull tensile strength of all coated yarns was lower than uncoated ones. Oxidized cellulose is highly biocompatible and has great antibacterial properties against a variety of pathogens. The ability of oxidized cellulose as suture material was studied by Li et al. who explored the effect of tempo oxidation treatment on the physical and mechanical properties of TORC (TEMPO-mediated oxidation of regenerated cellulose) sutures [69]. The carboxyl content in the suture materials was controlled by varied oxidation times. It could be seen that TEMPO oxidation significantly influenced the degradation of sutures as evaluated from the hydrolysis test performed by immersing the sutures in physiological saline for 7, 14, 21, and 28 days (Fig. 1.6). The carboxyl groups introduced in the sutures due to TEMPO

Figure 1.6 The in the vitro degradation rate of TORC and different TORC at different oxidation times (15, 30, 45, 60, and 90 min after PBS impregnation (mean  S.D., n ¼ 10)). (Reproduced with permission from [69].)

12

Advanced Technologies and Polymer Materials for Surgical Sutures

oxidation leads to increased molecular chain spacing and reduction in molecular interatomic force whereby water easily penetrates the fiber resulting in breakage.

1.5 Sterilization of sutures Sterilization technology plays a prominent role in the biomedical field because bacterial colonization remains a big problem with medical implants or devices. Every surgical procedure is associated with a certain risk of contamination and hence the sutures must be well sterilized to prevent bacterial adherence. Despite the widespread use of sutures in the 19th century, suture associated infections were a major concern. Lord Joseph Lister made remarkable contributions to the history of sterilization [2]. He pointed out that whatever be the cause of wound infection, carbolic acid could prevent or halt its further progress [70].In 1869, he developed aseptic silk sutures treated with carbolic acid followed by sutures from sheep intestine known as catgut sutures(catgut sutures treated with 5% chromic acid). Claudius introduced the concept of using potassium iodide for suture sterilization in 1902, and the infection rate was further reduced [71]. A process for sterilizing catgut sutures and ligatures using heat was invented in 1958 [72]. Here catgut sutures and ligatures sealed in a container in the presence of aqueous isopropanol solution were sterilized by heat. Ethicon Inc. started using electron beam accelerators for suture sterilization in 1957 [73]. According to the European Norm 556 sterility is defined as the state of being free from viable microorganisms (1
10e6)[74]. Generally, sterilization techniques can be classified as physical and chemical. Physical sterilization involves sterilization using heat and radiations whereas chemical sterilization involves the use of chemicals like ethylene oxide, hydrogen peroxide, formaldehyde, b-propiolactone, etc. Certain sterilization procedures result in stiffening of sutures and hence the selection of appropriate sterilization techniques is critical. The tensile strength of Virgin silk suture treated by thermal methods of sterilization was found to decrease [75]. Sterilization of collagen sutures with b-propiolactone showed no significant loss of strength of the finished sutures, hence can be used as an alternative to heat sterilization[76].

1.6 Conclusion and future perspectives The growing environmental concerns have led to increased research in the field of biopolymers. Natural polymers or biopolymers are derived from

Advances in biopolymer based surgical sutures

13

living organisms. These have the added advantage of being biodegradable, biocompatible, renewable, and reduced antigenicity. Sutures are materials of immense importance in the biomedical field. Along with ease of handling and biocompatibility, the mechanical properties of suture materials are a major factor that affects the overall suture quality. Despite the modern technological advancements in the materials in the methodology perspectives, biopolymeric sutures have an important role in wound healing. The progressive techniques like onsite evaluation of the wound healing, easy to use sutures or wound closure methods, smart sutures, and other wound closure devices and products are to be looked up in the future. No sutures can be called ideal as such. The vital concern is surgical site infection after surgery. Although antibacterial sutures delivering antibiotics have been developed, maintaining all desirable biological and morphological features in a single suture is still a matter of research. The growing misuse of antibiotics requires more alternatives to combat surgical site infections. The merging of the biopolymeric sutures and nanotechnology will give a boost to the surgical industry where properties like better antimicrobial activity, faster wound healing, etc. could be achieved. The inventions and innovations in suture fabrication have a huge potential to be applied for the betterment of the patients and making their lives better.

References [1] C. Dennis, S. Sethu, S. Nayak, L. Mohan, Y.Y. Morsi, G. Manivasagam, Suture materials - current and emerging trends, J. Biomed. Mater. Res. A 104 (2016) 1544e1559, https://doi.org/10.1002/jbm.a.35683. [2] T.M. Muffly, A.P. Tizzano, M.D. Walters, The history and evolution of sutures in pelvic surgery, J. R. Soc. Med. 104 (2011) 107e112, https://doi.org/10.1258/ jrsm.2010.100243. [3] J.M. Seitz, M. Durisin, J. Goldman, J.W. Drelich, Recent advances in biodegradable metals for medical sutures: a critical review, Adv. Healthc. Mater. 4 (2015) 1915e1936, https://doi.org/10.1002/adhm.201500189. [4] A.M. Smith, S. Moxon, G.A. Morris, Biopolymers as wound healing materials, in: Wound Heal. Biomater, 2016, https://doi.org/10.1016/B978-1-78242-456-7.00013-1. [5] N.R. Nair, V.C. Sekhar, K.M. Nampoothiri, A. Pandey, Biodegradation of biopolymers, Curr. Dev. Biotechnol. Bioeng. (2016), https://doi.org/10.1016/B978-0444-63662-1.00032-4. [6] S. Mohan, O.S. Oluwafemi, N. Kalarikkal, S. Thomas, S.P. Songca, Biopolymers e application in nanoscience and nanotechnology, in: Recent Adv. Biopolym, 2016, https://doi.org/10.5772/62225. [7] P.P. Dattilo Jr., P.P. Dattilo Jr., M.W. King, M.W. King, N.L. Cassill, N.L. Cassill, J.C. Leung, J.C. Leung, Medical textiles: application of an absorbable barbed bidirectional surgical suture, J. Text. Appar. Technol. Manang. 2 (2002).

14

Advanced Technologies and Polymer Materials for Surgical Sutures

[8] A. Goel, Surgical sutures - a review, Delhi J. Ophthalmol. (2016), https://doi.org/ 10.7869/djo.161. [9] B. Joseph, A. George, S. Gopi, N. Kalarikkal, S. Thomas, Polymer sutures for simultaneous wound healing and drug delivery e a review, Int. J. Pharm. 524 (2017), https://doi.org/10.1016/j.ijpharm.2017.03.041. [10] N. Reddy, R. Reddy, Q. Jiang, Crosslinking biopolymers for biomedical applications, Trends Biotechnol. (2015), https://doi.org/10.1016/j.tibtech.2015.03.008. [11] T.T. Tsai, T.H. Huang, C.J. Chang, N. Yi-Ju Ho, Y.T. Tseng, C.F. Chen, Antibacterial cellulose paper made with silver-coated gold nanoparticles, Sci. Rep. 7 (2017) 1e10, https://doi.org/10.1038/s41598-017-03357-w. [12] Y.Y. Dong, L.H. Fu, S. Liu, M.G. Ma, B. Wang, Silver-reinforced cellulose hybrids with enhanced antibacterial activity: synthesis, characterization, and mechanism, RSC Adv. (2015), https://doi.org/10.1039/c5ra19758a. [13] N.K. Blessy Joseph, H.J. Maria, S. Thomas, Nanocellulose: health care applications, in: M. Mishra (Ed.), Encycl. Polym. Appl, CRC Press, 2018, pp. 1829e1852. [14] A. Anitha, S. Sowmya, P.T.S. Kumar, S. Deepthi, K.P. Chennazhi, H. Ehrlich, M. Tsurkan, R. Jayakumar, Chitin and chitosan in selected biomedical applications, Prog. Polym. Sci. (2014), https://doi.org/10.1016/j.progpolymsci.2014.02.008. [15] Y. Wang, D.J. Blasioli, H.J. Kim, H.S. Kim, D.L. Kaplan, Cartilage tissue engineering with silk scaffolds and human articular chondrocytes, Biomaterials (2006), https:// doi.org/10.1016/j.biomaterials.2006.03.050. [16] J. Chen, G.H. Altman, V. Karageorgiou, R. Horan, A. Collette, V. Volloch, T. Colabro, D.L. Kaplan, Human bone marrow stromal cell and ligament fibroblast responses on RGD-modified silk fibers, J. Biomed. Mater. Res. 67A (2003) 559e570, https://doi.org/10.1002/jbm.a.10120. [17] L. Meinel, S. Hofmann, V. Karageorgiou, L. Zichner, R. Langer, D. Kaplan, G. Vunjak-Novakovic, Engineering cartilage-like tissue using human mesenchymal stem cells and silk protein scaffolds, Biotechnol. Bioeng. 88 (2004) 379e391, https:// doi.org/10.1002/bit.20252. [18] M.D. Shoulders, R.T. Raines, Collagen structure and stability, Annu. Rev. Biochem. (2009), https://doi.org/10.1146/annurev.biochem.77.032207.120833. [19] R. Parenteau-Bareil, R. Gauvin, F. Berthod, Collagen-based biomaterials for tissue engineering applications, Materials (2010), https://doi.org/10.3390/ma3031863. [20] R. Khan, M.H. Khan, Use of collagen as a biomaterial: an update, J. Indian Soc. Periodontol. (2013), https://doi.org/10.4103/0972-124X.118333. [21] S.A. Sell, M.J. McClure, K. Garg, P.S. Wolfe, G.L. Bowlin, Electrospinning of collagen/biopolymers for regenerative medicine and cardiovascular tissue engineering, Adv. Drug Deliv. Rev. 61 (2009) 1007e1019, https://doi.org/10.1016/ j.addr.2009.07.012. [22] S. Liu, G. Wu, X. Zhang, J. Yu, M. Liu, Y. Zhang, P. Wang, X. Yin, J. Zhang, F. Li, M. Zhang, Preparation and properties of poly (lactic acid) (PLA) suture loaded with PLA microspheres enclosed drugs (PM-Ds), J. Text. Inst. (2019), https://doi.org/ 10.1080/00405000.2019.1610999. [23] S. Chattopadhyay, R.T. Raines, Collagen-based biomaterials for wound healing, October 141 (2008) 520e529, https://doi.org/10.1016/j.surg.2006.10.010.Use. [24] R.G.C. Gerard, V. Yu, Sutures and anchoring devices, in: McGlamry’s Compr. Textb. Foot Ankle Surg, 2001, pp. 139e174. [25] E.J. Holder, The story of catgut, Postgrad. Med. (1949), https://doi.org/10.1136/ pgmj.25.287.427. [26] M. Younesi, B.O. Donmez, A. Islam, O. Akkus, Heparinized collagen sutures for sustained delivery of PDGF-BB: delivery profile and effects on tendon-derived cells invitro, Acta Biomater. (2016), https://doi.org/10.1016/j.actbio.2016.05.036.

Advances in biopolymer based surgical sutures

15

[27] A.D. Mazzocca, M.B. McCarthy, C. Arciero, A. Jhaveri, E. Obopilwe, L. Rincon, J. Wyman, G.A. Gronowicz, R.A. Arciero, Tendon and bone responses to a collagencoated suture material, J. Shoulder Elbow Surg. (2007), https://doi.org/10.1016/ j.jse.2007.02.113. [28] K. Hamad, M. Kaseem, H.W. Yang, F. Deri, Y.G. Ko, Properties and medical applications of polylactic acid: a review, Express Polym. Lett. 9 (2015). [29] L. Avérous, Synthesis, properties, environmental and biomedical applications of polylactic acid, in: Handb. Biopolym. Biodegrad. Plast. Prop. Process. Appl, 2012, https:// doi.org/10.1016/B978-1-4557-2834-3.00009-4. [30] R. Langer, A. Basu, A.J. Domb, Special issue: polylactide (PLA) based biopolymers, Adv. Drug Deliv. Rev. (2016), https://doi.org/10.1016/j.addr.2016.11.002. [31] O.L. Shanmugasundaram, R.V.M. Gowda, D. Saravanan, Drug release and antimicrobial studies on polylactic acid suture, Int. J. Biotechnol. Mol. Biol. Res. 2 (2011) 80e89. http://www.academicjournals.org/journal/IJBMBR/article-abstract/ 7BDBFE141300. [32] S. Liu, M. Liu, G. Wu, X. Zhang, J. Yu, Y. Zhang, P. Wang, X. Yin, Enhanced surface hydrophilicity of polylactic acid sutures treated by lipase and chitosan, Textil. Res. J. (2018), https://doi.org/10.1177/0040517518811936. [33] A. Visco, C. Scolaro, A. Giamporcaro, S. De Caro, E. Tranquillo, M. Catauro, Threads made with blended biopolymers: mechanical, physical and biological features, Polymers (2019), https://doi.org/10.3390/polym11050901. [34] G.H. Altman, F. Diaz, C. Jakuba, T. Calabro, R.L. Horan, J. Chen, H. Lu, J. Richmond, D.L. Kaplan, Silk-based biomaterials, Biomaterials 24 (2003) 401e416, https://doi.org/10.1016/S0142-9612(02)00353-8. [35] R.H.H. Tan, R.J.W. Bell, B.A. Dowling, A.J. Dart, Suture materials: composition and applications in veterinary wound repair, Aust. Vet. J. (2003), https://doi.org/10.1111/ j.1751-0813.2003.tb11075.x. [36] F. Javed, M. Al-Askar, K. Almas, G.E. Romanos, K. Al-Hezaimi, Tissue reactions to various suture materials used in oral surgical interventions, ISRN Dent. (2012), https:// doi.org/10.5402/2012/762095. [37] S. Sala-Pérez, M. López-Ramírez, M. Quinteros-Borgarello, E. Valmaseda-Castellón, C. Gay-Escoda, Antibacterial suture vs silk for the surgical removal of impacted lower third molars. A randomized clinical study, Med. Oral Patol. Oral Cir. Bucal. 21 (2016) e95ee102, https://doi.org/10.4317/medoral.20721. [38] C.L. He, Z.M. Huang, X.J. Han, Fabrication of drug-loaded electrospun aligned fibrous threads for suture applications, J. Biomed. Mater. Res. A (2009), https:// doi.org/10.1002/jbm.a.32004. [39] E. Laas, C. Poilroux, C. Bézu, C. Coutant, S. Uzan, R. Rouzier, E. Chéreau, Antibacterial-coated suture in reducing surgical site infection in breast surgery: a prospective study, Int. J. Breast Cancer (2012), https://doi.org/10.1155/2012/819578. [40] D. Leaper, P. Wilson, O. Assadian, C. Edmiston, M. Kiernan, A. Miller, G. BondSmith, J. Yap, The role of antimicrobial sutures in preventing surgical site infection, Ann. R. Coll. Surg. Engl. 99 (2017) 439e443, https://doi.org/10.1308/ rcsann.2017.0071. [41] I. Chopra, M. Roberts, Tetracycline antibiotics: mode of action, applications, molecular biology, and epidemiology of bacterial resistance, Microbiol. Mol. Biol. Rev. 65 (2001) 232e260, https://doi.org/10.1128/MMBR.65.2.232e260.2001. [42] S. Viju, G. Thilagavathi, Characterization of tetracycline hydrochloride drug incorporated silk sutures, J. Text. Inst. (2013), https://doi.org/10.1080/ 00405000.2012.720758. [43] S. Viju, G. Thilagavathi, Fabrication and characterization of silk braided sutures, Fibers Polym. 13 (2012) 782e789, https://doi.org/10.1007/s12221-012-0782-8.

16

Advanced Technologies and Polymer Materials for Surgical Sutures

[44] R. Najafi-taher, B. Ghaemi, S. Kharazi, S. Rasoulikoohi, A. Amani, Promising antibacterial effects of silver nanoparticle-loaded tea tree oil nanoemulsion: a synergistic combination against resistance threat, AAPS PharmSciTech (2017), https://doi.org/ 10.1208/s12249-017-0922-y. [45] A.J. Kora, R.B. Sashidhar, Biogenic silver nanoparticles synthesized with rhamnogalacturonan gum: antibacterial activity, cytotoxicity and its mode of action, Arab. J. Chem. (2018), https://doi.org/10.1016/j.arabjc.2014.10.036. [46] T.C. Dakal, A. Kumar, R.S. Majumdar, V. Yadav, Mechanistic basis of antimicrobial actions of silver nanoparticles, Front. Microbiol. (2016), https://doi.org/10.3389/ fmicb.2016.01831. [47] S. De Simone, A.L. Gallo, F. Paladini, A. Sannino, M. Pollini, Development of silver nano-coatings on silk sutures as a novel approach against surgical infections, J. Mater. Sci. Mater. Med. 25 (2014) 2205e2214, https://doi.org/10.1007/s10856-014-5262-9. [48] D. Gogoi, A.J. Choudhury, J. Chutia, A.R. Pal, M. Khan, M. Choudhury, P. Pathak, G. Das, D.S. Patil, Development of advanced antimicrobial and sterilized plasma polypropylene grafted MUGA (antheraea assama) silk as suture biomaterial, Biopolymers 101 (2014) 355e365, https://doi.org/10.1002/bip.22369. [49] D. Elieh-Ali-Komi, M.R. Hamblin, Chitin and chitosan: production and application of versatile biomedical nanomaterials, Int. J. Adv. Res. 4 (2016) 411e427. [50] C.K.S. Pillai, W. Paul, C.P. Sharma, Chitin and chitosan polymers: chemistry, solubility and fiber formation, Prog. Polym. Sci. (2009), https://doi.org/10.1016/ j.progpolymsci.2009.04.001. [51] S. Ahmed, S. Ikram, Chitosan based scaffolds and their applications in wound healing, Achiev. Life Sci. 10 (2016) 27e37, https://doi.org/10.1016/j.als.2016.04.001. [52] T. Dai, M. Tanaka, Y.Y. Huang, M.R. Hamblin, Chitosan preparations for wounds and burns: antimicrobial and wound-healing effects, Expert Rev. Anti Infect. Ther. (2011), https://doi.org/10.1586/eri.11.59. [53] M. Nakajima, K. Atsumi, K. Kifune, K. Miura, H. Kanamaru, Chitin is an effective material for sutures, Jpn. J. Surg. (1986), https://doi.org/10.1007/BF02470609. [54] I.P. Dobrovol’skaya, I.A. Kasatkin, V.E. Yudin, E.M. Ivan’kova, V.Y. Elokhovskii, Supramolecular structure of chitin nanofibrils, Polym. Sci. A 57 (2015) 52e57, https:// doi.org/10.1134/S0965545X15010022. [55] D. Sudha, B. Dhurai, T. Ponthangam, Development of herbal drug loaded antimicrobial silk suture, Indian J. Fibre Text. Res. 42 (2017) 286e290. [56] S. Viju, G. Thilagavathi, Effect of chitosan coating on the characteristics of silk-braided sutures, J. Ind. Textil. (2013), https://doi.org/10.1177/1528083711435713. [57] Y. Yang, S.B. Yang, Y.G. Wang, S.H. Zhang, Z.F. Yu, T.T. Tang, Bacterial inhibition potential of quaternised chitosan-coated VICRYL absorbable suture: an in vitro and in vivo study, J. Orthop. Transl. 8 (2017) 49e61, https://doi.org/10.1016/ j.jot.2016.10.001. [58] L. Marcotte, J. Barbeau, M. Lafleur, Permeability and thermodynamics study of quaternary ammonium surfactants - phosphocholine vesicle system, J. Colloid Interface Sci. (2005), https://doi.org/10.1016/j.jcis.2005.05.060. [59] M. Crismaru, L.A.T.W. Asri, T.J.A. Loontjens, B.P. Krom, J. De Vries, H.C. Van Der Mei, H.J. Busscher, Survival of adhering staphylococci during exposure to a quaternary ammonium compound evaluated by using atomic force microscopy imaging, Antimicrob. Agents Chemother. (2011), https://doi.org/10.1128/AAC.05062-11. [60] S. Prabha, J. Sowndarya, P.J.V.S. Ram, D. Rubini, B. Hari, W. Aruni, P. Nithyanand, Chitosan-coated surgical sutures prevent adherence and biofilms of mixed microbial communities, Curr. Microbiol. 78 (2021) 502e512. [61] J. Chee, S. Yoga, N. Lau, S. Ling, R.M.M. Abed, Bacterially produced polyhydroxyalkanoate ( PHA ): converting renewable resources into bioplastics, in: Current

Advances in biopolymer based surgical sutures

[62] [63] [64] [65] [66]

[67] [68]

[69]

[70] [71] [72] [73] [74] [75] [76]

17

Research, Technology and Education Topics in Applied Microbiology and Microbial Biotechnology, 2010, pp. 1395e1404. Brigham, Sinskey, Applications of polyhydroxyalkanoates in the medical industry, Int. J. Biotechnol. Wellness Ind. (2012), https://doi.org/10.6000/1927-3037.2012.01.01.03. E.I. Shishatskaya, T.G. Volova, A.P. Puzyr, O.A. Mogilnaya, S.N. Efremov, Tissue response to the implantation of biodegradable polyhydroxyalkanoate sutures, J. Mater. Sci. Mater. Med. (2004), https://doi.org/10.1023/B:JMSM.0000030215.49991.0d. S. Rizk, Non-curling Polyhydroxyalkanoate Sutures, US8084125, 2011. Y. He, Z. Hu, M. Ren, C. Ding, P. Chen, Q. Gu, Q. Wu, Evaluation of PHBHHx and PHBV/PLA fibers used as medical sutures, J. Mater. Sci. Mater. Med. 25 (2014) 561e571. M. Moniri, A. Boroumand Moghaddam, S. Azizi, R. Abdul Rahim, A. Bin Ariff, W. Zuhainis Saad, M. Navaderi, R. Mohamad, Production and status of bacterial cellulose in biomedical engineering, Nanomaterials 7 (2017) 257, https://doi.org/ 10.3390/nano7090257. F. Esa, S.M. Tasirin, N.A. Rahman, Overview of bacterial cellulose production and application, Agric. Agric. Sci. Procedia 2 (2014) 113e119, https://doi.org/10.1016/ j.aaspro.2014.11.017. H. Wu, G.R. Williams, J. Wu, J. Wu, S. Niu, H. Li, H. Wang, L. Zhu, Regenerated chitin fibers reinforced with bacterial cellulose nanocrystals as suture biomaterials, Carbohydr. Polym. 180 (2018) 304e313, https://doi.org/10.1016/ j.carbpol.2017.10.022. H. Li, F. Cheng, C. Chávez-Madero, J. Choi, X. Wei, X. Yi, T. Zheng, J. He, Manufacturing and physical characterization of absorbable oxidized regenerated cellulose braided surgical sutures, Int. J. Biol. Macromol. (2019), https://doi.org/10.1016/ j.ijbiomac.2019.05.030. M. Worboys, Joseph Lister and the performance of antiseptic surgery, Notes Record Roy. Soc. Lond. (2013), https://doi.org/10.1098/rsnr.2013.0028. A. Davey, C.S. Ince, Fundamentals of Operating Department Practice, 2015, https:// doi.org/10.1017/CBO9781316529874. B. Alfred, Sterilization of Surgical Catgut Sutures and Ligatures, 2832664, 1958. R. Singh, D. Singh, A. Singh, Radiation sterilization of tissue allografts: a review, World J. Radiol. (2016), https://doi.org/10.4329/wjr.v8.i4.355. M.J. Abreu, M.E. Silva, L. Schacher, D. Adolphe, Recycling of textiles used in the operating theatre, Recycl. Textil. (2006) 183e202, https://doi.org/10.1533/ 9781845691424.4.183. G.N. Shuttleworth, L.F. Vaughn, H.B. Hoh, Material properties of ophthalmic sutures after sterilization and disinfection, J. Cataract Refract. Surg. 25 (1999) 1270e1274, https://doi.org/10.1016/S0886-3350(99)00156-X. E.L. Ball, A.C. Dornbush, G.M. Sieger, F.E. Stirn, J.C. Vitucci, J.F. Weidenheimer, E.L.L. Laboratories, P. River, A.C. Dornbush, G.M. Sieger, F.E. Stirn, J.C. Vitucci, J.F. Weiden, Sterilization of Regenerated Collagen Sutures with f3-Propiolactone, 1960, pp. 269e272.

CHAPTER 2

Functionalization of sutures Felipe López-Saucedo1, Alejandro Ramos-Ballesteros2 and Emilio Bucio1 1

Departamento de Química de Radiaciones y Radioquímica, Instituto de Ciencias Nucleares, Universidad Nacional Autónoma de México, Circuito Exterior, Ciudad Universitaria, CDMX, Mexico; 2 Notre Dame Radiation Laboratory, University of Notre Dame, Indiana, United States

2.1 Introduction It is well known that postoperative infections, and complications derived from them, are among the main causes that delay and lengthen the healing process. Postoperative infections are recurrent and responsible for a prolonged hospital stay, extra intake of antibiotics, and, if the infection progress, additional surgeries and treatments that may cause death. This issue is a worldwide concern to solve, and several efforts have been made to optimize the postoperative phase since it is as important as the surgery itself. Although it is known that the causes of infection are diverse, suturing is particularly relevant because the zone exposed to the intervention is susceptible to pathogenic attacks hosted on the suture surface [1]. Biohazard usually occurs at the time of insertion by contact with opportunistic skin microorganisms, but also for the migration of microorganisms from preexisting foci of infection in the patient. A suture thread is a biomedical device (natural or synthetic) allocated to connect blood vessels or to approximate tissues to accelerate healing. Sutures are widely used in surgery above other methods such as staples, tapes, or laser cautery [2] due to easy sterilization, multipurpose, flexibility, handling, strong knotting, strain at rupture, elastic modulus, hypoallergenic, and ability to avoid the formation of biofilms around suture [3]. In previous sections were also summarized the desirable characteristics that materials for suture threads should comply like: ❖ High biocompatibility and nontoxicity ❖ Easy manipulation for the surgeon (folding, knotting, etc.) ❖ Easy sterilization without compromising material integrity ❖ Hypoallergenic ❖ Absorbable upon completion of its function (preferably) ❖ Inhibit bacterial growth Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00006-1

© 2023 Elsevier Ltd. All rights reserved.

19

20

Advanced Technologies and Polymer Materials for Surgical Sutures

It is precisely the latter feature (antibacterial activity), where a high volume of research has been focused lately on some researches where it is achieved. One of the possible solutions to decrease bacterial growth (and therefore infections) is using suture threads with an agent to inhibit bacterial growth in a localized and long-lasting manner [4,5]. The inclusion of superficial agents adds new properties to the existing ones, enhancing efficiency and safety, and fortunately, the functionalization achieves excellent results without compromising the integrity of the polymeric suture threads. Currently, there is no ideal material able to cover all the required properties, therefore, the surgeon must choose among the assortment of suture materials according to the type of wound, length, organ, exposition, and patient condition [6]. In addition to this, surgical interventions expose the skin tissue to damage, and, consequently, marks and/or scars can be permanent, so the aesthetic variable must also be considered. The objectives of suture modification are based on finding and standardizing experimental methodologies, as well as comparing characteristics such as biocompatibility, susceptibility to biofilm proliferation and toxicity of materials before and after their processing [7]. Suture functionalization includes various strategies going from impregnation and coatings to those methods where the surface is modified using high-energy radiation, such as plasma treatment or gamma radiation.

2.2 Suture materials: from hairs to antibacterial biopolymers Since ancient times, humanity has used cotton ties, hair, and other natural fibers to approximate tissues. Egyptian civilization at the dawn of 3000 B.C., employed different types of cords for mummification, but it is not discarded the use of threads for medical purposes [8]. In Arabia, around 900 B.C., surgical procedures were perfectioned and animal-origin absorbable sutures equivalent to modern catgut were used. Meanwhile in India, around 600 and 500 B.C., the Sushruta Samhita (Sanskrit text on medicine and surgery) already recommended the use of different suture materials including cotton, leather, and even horsehair; as well as other suture techniques that were compiled and described in one of the first medicine manuscripts ever [9]. In Europe, Galen of Pergamum (129e201 A.D.), who lived in the region of present-day Turkey, wrote several books on the use of sutures in surgical procedures, which turned into the annals of the

Functionalization of sutures

21

occidental medicine in the Classical Age. By the Middle Age, silk sutures began to be used, this is a natural fiber formed by a nonabsorbable polymer. Already in contemporary times, in the early 20th century, Dr. William Halstead (1852e1922), who was a pioneer of modern surgery in the United States, recommended the use of silk sutures and even tested with silver sutures threads in hernia surgeries [10]. Starting the first half of the 20th century, during “the boom” of the exploration and exploitation of oil derivatives, many polymer materials began to be developed and were proved in a variety of products, including the first synthetic sutures. At this time in history, it was forward in the design of the first synthetic suture polymers; for example, the first methods for obtaining polyamides and polyesters were established, moreover, increasing demand for materials such as polypropylene (PP) allowed the production of strong suture monofilaments [11]. Undoubtedly, necessity is the mother of invention, and with humanity’s progress, the optimization of sutures has become evident. There is still a long way to go, but thanks to the new techniques of surface modification, it seems that the path is traced. With functionalization, not only the inherent properties of the original material are preserved, like hardness, flexibility, thermal and chemical resistance, but also new features are added including hydrophilicity, charge, surface area, and antimicrobial properties.

2.3 Suture types Surface modification and specific functionalization of sutures depend on composition and application. If the suture is natural or synthetic, the functional groups at the surface will be susceptible to different reactions; therefore, different methods are required. Suture threads are classified according to several criteria, but the most common are biodegradability, origin, and macroscopic structure (Table 2.1).

Table 2.1 Suture classification. Suture thread

Degradability Origin Structure

Absorbable Natural Metallic Monofilament

Nonabsorbable Synthetic No metallic (organics) Multifilament

22

Advanced Technologies and Polymer Materials for Surgical Sutures

The monofilament sutures, as their name suggests, consist of a unique fiber. This configuration is adequate to prevent biofilm’s growth since it has a uniform surface; while the multifilament suture is made up of several twisted or intertwined filaments, making them prone to lodge microorganisms in the larger superficial area. However, monofilaments have less resistance to suturing compared to multifilaments that have a higher friction coefficient. Although in favor of multifilament sutures, they display greater tensile strength, firmer knot tying, and easier handling. Also, the multifilament-type suture generally has better flexibility and is easier to tie compared to the monofilament suture [2,9]. For pragmatic purposes of this document, the functionalization of absorbable and nonabsorbable sutures is discussed in detail in the next sections.

2.4 Biocompatibility studies for functionalized sutures In vitro studies for functionalized sutures are necessary to assess toxicity and possible adverse reactions. Therefore, the progression for new material development after synthesis and characterization implies cytotoxicity studies to verify if these materials are compatible with humans [12]. These tests, in conjunction with complementary (physical-chemical-mechanical properties), help to discern viability for specific functionalization. In the case of suture thread modifications also is monitored the possible changes that occurred compared with the pristine suture threads. Tests to determine the biocompatibility of materials are usually carried out in human and mammalian cell lines; for example, one of the most used is the embryonic tissue cell line from mouse fibroblasts (Mus musculus) BALB 3T3. This cell line is widely used as primary control because of its extreme sensitivity, and inherent capacity to be easily inhibited by contact with surrounding toxic substances [13]. There are pre-established qualitative and quantitative kits that help to identify toxic or biocompatible materials or drugs. One of these kits to determine cell proliferation is the WST-1 assay. This is a colorimetric test that spectrophotometrically quantifies tetrazolium salts that are degraded to formazan by the biological functions of the cell. The chemical name of WST-1 is 4-[3-(4-iodophenyl)-2-(4-nitrophenyl)-2H-5tetrazolium]-1,3-benzene disulfonate, a molecule that degrades through a complex mitochondrial succinate-tetrazolium-reductase system, which is only active in living cell mitochondria [14].

Functionalization of sutures

23

Another staining-based indicator used for cell viability bioassays is the reagent resazurin, traded under the name Alamar Blue from ThermoFicher USA, which is blue in color and is irreversibly reduced to the resofurin compound, which is red. This reagent has been tried in a range of biological systems and different cell types such as bacteria, fungi, protozoa, and cultured mammalian cells [15]. Although quantification of cytotoxicity may be calculated by measuring the “maximum noncytotoxic concentration” (limit at which there is not an alteration in the cell morphology and metabolism), the parameter known as 50% cytotoxic inhibitory concentration (CI50%) is the most used. CI50% is defined as the concentration of lethality for half of the cells, compared with controls and a blank [16]. Regarding suture materials grafted with vinyl monomers such as N-vinylimidazole (NVIm), hydroxyethylmethacrylate (HEMA), and N-isopropylacrylamide (NIPAAm) [17], glycidyl methacrylate (GMA), or acrylic acid (AAc) [18], it has been verified with cytocompatibility studies that functionalization does not promote the formation and proliferation of biofilms (i.e., bacteria colonies). For this reason and biocompatibility properties, derived hydrophilic polymers are used in the manufacture of several biomedical devices [19].

2.5 Functionalization The functionalization of surface in suture threads is generally performed through two approaches: coating and graft functionalization. Each with special characteristics according to the intended use of the final product, besides, some matrices are not prone to be functionalized (or coated) due to the charge, polarity, hydrophilicity, or chemical resistance of the components. 2.5.1 Coating in fibers The definition of coating is the covering of any surface. It is characterized because there are no significant changes in the inner composition of the modified material, and the coating usually does not cause a drastic change in the mechanical properties, being the main objective of coatings only to achieve a difference in the reactivity of the surface and protection (to prevent corrosion). In other words, a coating is conducted for a pragmatic improvement in the surface reactivity without compromising other parts of the material [20].

24

Advanced Technologies and Polymer Materials for Surgical Sutures

Concerning fiber-structured materials, the techniques have been adapted to the development of tailored sutures. Coatings in suture threads are directed to increase the lifetime, minimize the tissue damage, and stimulate the cell recovery, but mainly to provide antibacterial features in drug-delivery systems. The different techniques used for the coating of fibers allow acquiring specific features, some alternatives of dip-coating and electrodeposition are mentioned below. It is important to mention that even though coatings increase resistance and add new features to the material, does not present synergy or mutually enhanced properties that come from the chemical combination that functionalization provides. 2.5.1.1 Dip-coating Dip-coating applied in suture threads is like dip processes in films or pellets. This is the simplest fundament to achieve a uniform coating containing a bioactive molecule. The procedure consists of preparing a dissolution with the bioactive principle in an adequate liquid medium, and subsequently, carried to dryness or it is submitted to curing. It is expected that the compound of interest remained attached to the surface of the material without decomposing [21]. In other words, the molecule of interest is dissolved, and then a third component is added, which plays a double role of adhesive agent (to the surface of the material) and agglutinant (with the bioactive molecule). The reaction conditions such as pH or ionic strength may be easily controlled to increase the electrostatic interactions of the coating with the surface of the material and promote layer growth. Between the layers, at the interface, there are partial positive and negative charges whose interactions are responsible to keep the suture in one piece during and after the surgical procedure, but at the same time, keeping enough reactivity for a successful controlled release. Examples of dipcoating in absorbable sutures illustrate better this principle (Fig. 2.1). In the first example, vicryl suture threads were coated with a blend of collagen and gentamicin; the protein and the antibiotic were dissolved in acetic acid. The system works as follows: the gentamicin is embedded in the collagen coating, while the collagen is well adhered to the surface of the suture, leading this way to an antibacterial absorbable suture [22]. In another example, the coating of a suture was performed using a biomacromolecule, the messenger RNA, which was loaded directly onto poly(lactide-co-glycolic acid), (PLGA) threads. The mRNA is loaded to stimulate the regeneration of epidermal tissue; the composite mRNAPLGA activates the expression of growth factors in the human

Functionalization of sutures

25

Figure 2.1 Coating in sutures allows the immobilization of bioactive compounds.

keratinocyte cell line (HaCat). This coating was carried out by dissolving the mRNA in a stock solution of PLGA/ethyl acetate. Afterwards, the mRNA was properly treated dropwise with the plasmid transfection agents Viromer RED or Lipofectamine 2000. The whole procedure is achieved at room temperature [23]. 2.5.1.2 Electrodeposition It is an electrochemical method employed to produce metallic coatings in situ. In this process, colloids or nanoparticles are formed in solution and attached to the fiber, shaping a layer. This method is perfect for the modification of metal threads. Recently a copper wire was coated with Ni using electrodeposition, the Ni layer allowed the loading of TiO2 nanoparticles in a homogeneous distribution along the wire surface; this coating would be useful for antibacterial sutures [24]. The versatility of this method enables the modification even in fragile fibers, as the poly(lactic acid), coated with an alloy of Tie6Ale4V through cathodic electrodeposition. The electrodeposition was performed into a glass cell, using graphite anodes. The coated poly(lactic acid) resisted chemical corrosion after 1 week immersed in distilled water, while in Ringer’s solution, the coating almost disappeared [25]. Prevention of spoilage in a uniform layer helps to keep good performance and functionality; therefore, an efficient covering may be achieved through electrostatic layer-by-layer deposition. This method allows the option to choose the number of cycles to acquire functional surfaces. Thus, this type of materials are composed of alternated positive/negative layers and repeats the process until achieving the desired thickness. The

26

Advanced Technologies and Polymer Materials for Surgical Sutures

advantages provided by this method are the control of coating morphology and uniform thickness [26,27]. 2.5.2 Grafted sutures A graft copolymer may be obtained from a matrix polymer and a covalently bonded monomer in chains (a method called “grafting from”) or directly by the union of entire chains of polymers onto another polymer (“grafting to”). In both types of functionalization, constituents of the copolymer can partially retain their identity, and/or possibly the final product may acquire different qualities than the polymers by separate. According to the polymer structure, the functionalization of sutures using grafting sometimes requires activation with appropriate functional groups to induce reaction with the monomers on a polymer matrix. Grafting can be made through polymerization with acidic or alkaline initiators, but the integrity of the sutures can be compromised, and therefore, their functionality. Thus, free-radical copolymerization is a more common method for suture functionalization. Chemical initiators such as peroxides, peracids, azo derivatives, or inorganic compounds are traditionally used in copolymerization via free radicals [28], although the alternative with high-energy radiation presents interesting advantages over chemical methods, specifically for sutures. Initiators such as plasma, accelerated electrons, or gamma radiation can be used. The mechanism is based on the formation of radicals in the backbone polymer, just as the chemicalinitiator method alternative, but avoiding contamination and thermal treatment, which could lead to degrading the suture. Indeed, the use of ionizing radiation (at an adequate dosage) as a means of sterilization of suture materials, has proved high effectivity and efficiency without counterproductive effects [29]. Expanding on the advantages of free-radical grafting in copolymer synthesis, it allows some flexibility concerning the control of reaction variables, in both homogeneous and heterogeneous systems, since reactions can be improved by controlling factors such as temperature and reaction time [30]. Regarding the free-radical polymerization mechanism for grafting (Fig. 2.2), this proceeds in a similar way to the generic chain reaction, with steps of initiation, propagation, and termination. In general, hydrophobic grafted copolymers with vinyl monomers and hydrophilic groups, such as NIPAAm, HEMA, and NVIm; yields materials with enhanced hydrophilic character and potential for a wide variety of

Functionalization of sutures

27

Figure 2.2 Grafting via free-radical on inactive surfaces using peroxide groups as initiator.

chemical functionalization on the surface including. Derivatizations, modifications, immobilizations, or drug loading-releasing systems (Fig. 2.3). For example, grafting by preirradiation oxidative method of PP-gNIPAAm films [31], PP-g-HEMA [32]; and simultaneous irradiation to obtain PP-g-NVIm [33], are affordable options for the functionalization of “chemically inert” matrix such as PP. Investigations regarding PP film modifications became the source of inspiration to modify PP sutures and other types of suture polymers. Different factors affect the degree of copolymerization; this is particularly true for “grafting from” reactions employing either high-energy radiation or chemical initiators. It is quite pragmatic to control the degree of copolymerization by slight adjusts to the reaction conditions. The variables necessary to consider in a copolymerization reaction are the concentration of monomer, solvent, and temperature [34]. 2.5.2.1 Monomer Regarding the grafting of vinyl monomers, their reactivity may be diametrically opposite since some monomers easily polymerize, and others require higher amounts of energy to activate the vinyl groups. Regarding

Figure 2.3 Grafting on PP with highly biocompatible vinyl monomers like NIPAAm, NVIm, and HEMA.

28

Advanced Technologies and Polymer Materials for Surgical Sutures

the monomer to be grafted, it must be stable under an air atmosphere and possess double bonds sufficiently reactive to form covalent bonds with the matrix. Alkenes with electron-withdrawing groups are suitable for reacting, e.g., styrene, acrylate, or methacrylate derivatives, being the methacrylate analogs more adequate to control the chain length via a radical mechanism because intermediaries are more energetically stable. Also, the radical polymerization of N-vinyl monomers as N-vinylcarbazoles, N-vinylindoles, N-vinylpyrrolidones, N-vinylcaprolactams, N-vinylimidazoles, or azoles in general, may be easily controlled, achieving not only a narrow distribution in chain length but also a defined architecture [35]. Then remote groups or heteroatoms contained in the molecule influence the polymerization and its reaction rate. Besides conventional radical initiators, there are two methods to obtain grafting employing high-energy radiation; this is through direct irradiation and preirradiation oxidative [36]. Preirradiation oxidative is a method with better control of reaction variables that requires a matrix able to host or stabilize peroxide and hydroperoxide species on its surface. This method is ideal when the monomer is prone to homopolymerization. Otherwise, when the monomer is low reactive, direct irradiation is maybe the best option to achieve the corresponding graft. 2.5.2.2 Solvent Proper selection of solvent for the modification of sutures is perhaps as crucial as the choice of reactants and materials. The solvent must cover some technical aspects follow mentioned. The first role of the solvent is to dissolve the monomer while the suture is not affected. The solvent only should help to swell the suture increasing its reactivity on the surface, but the solvent should not deform the suture or dissolve it because the mechanical properties of the products will be affected. The second characteristic is related to mechanism because some solvents are radical scavengers and, opposite, others are radical generators [37]. Also, under ionic mechanisms, the polarity and ionic strength of solvents are involved in the polymerization rate, inhibiting or favoring the chain reaction [38]. The third point to consider for a solvent is the capability to be innocuous and to avoid side reactions that would cause undesirable subproducts, waste, or toxic residues. The last characteristic that must satisfy an ideal solvent is the easiness of removal, and if it is possible, recycling. Moreover, the choice of a solvent must fulfill the normative for the synthesis of medical devices, i.e., solvents destinated for biomedical devices and sanitary materials are

Functionalization of sutures

29

restricted to those with no toxicity, biodegradable, biocompatible, and non-halogenated [39]. 2.5.2.3 Temperature of reaction An external stimulus is necessary to start free-radical reactions through initiators or peroxy-/hydroperoxy-species formed by ionizing radiation. Usually an increase in temperature or exposure to light provide enough energy for homolytic bond cleavage and starts the polymerization process. Higher kinetic energy increases molecular diffusion and the number of effective collisions, leading to effective chain growth. The temperature necessary to break the RO-OR peroxide bonds and RO-OH hydroperoxides to initiate the chain polymerization reaction is relatively low and should not exceed the glass transition temperature of the pristine suture. In some cases, is possible to reach initiation energy at room temperature and even lower [40]. Nevertheless, the formation of several active sites needs heat and long reaction times to obtain copolymerization with a high degree of grafting. In some cases, when the grafting is difficult to obtain, due to low reactivity or interference from the solvent (inhibitor), increasing the temperature may force reaction to take place, as long as the polymeric matrix and reactants resist heating [41]. 2.5.3 Stimuli-responsive polymers on sutures There are different definitions for stimuli-responsive materials, beyond epistemology, is a fact that all materials in this category share special properties, such as when they are exposed to an external stimulus, said materials experiment with physical or chemical changes that produce a measurable effect. Even in the literature, there is more than one word to define polymers with a response to external stimuli, it is common to find words like “smart”, “stimuli-responsive”, or “intelligent” to denote the same kind [42]. In any case, the absence of systematization or homogenization in definitions does not impede the exploration and take of benefits that these materials bring us. Stimuli-responsive polymers can be classified according to the type of material, the stimulus to which they respond, and the response to the stimulus. The first refers to the structural aspects, for example, in polymers, there are crosslinked structures, interpenetrating networks, hydrogels, “comb” copolymers, shape-memory polymers, etc. The second classification is based on the stimulus to which the materials respond such as temperature, pH, light, electric field, magnetic field, or molecular recognition.

30

Advanced Technologies and Polymer Materials for Surgical Sutures

Finally, the third classification is according to the response that the material shows versus certain stimuli including swelling/collapse, charge, color change, physical-state change (solid, liquid, or gas), photoluminescence, or conductivity. Regarding sutures, it is possible to modify their surface with smart polymers. The polymer chains on the modified sutures are activated when getting in contact with a fluid medium (blood plasma or body fluids), then change in the configuration as swelling or collapse take place. The degree of swelling may be controlled by external factors such as pH and/or temperature. The swelling responsiveness is used in drug-delivery systems to control the release rate of antibiotics, antiinflammatories, or analgesics [43]. Interesting investigations have been reported in the field, for example, the poly(NIPAAm), is a thermo-responsive polymer that in an aqueous solution has a low critical solution temperature (LCST) of around 32
C and may reach around 36
C when is modified. LCST is a property found in some thermo-responsive materials that switch their behavior from hydrophilic to hydrophobic. The LCST is a parameter used in the design of drug delivery systems. The reason is drug loading may be carried out at low temperatures, while the release is favored at corporal temperature or slightly under it. For this reason, poly(NIPAAm) is usually investigated for biomedical purposes, either as a homopolymer [44] or as a copolymer [45]. Also, pH-responsive polymers such as poly(NVIm) are commonly used as biomedical materials [46] because they combine antimicrobial properties with their capability of metal retention and chelation [47].

2.6 Functionalization of nonabsorbable sutures Nonabsorbable sutures are designed to remain stable for long periods, in both internal and external suturing. Mechanical properties such as resistance and tensile strength of nonabsorbable sutures are adequate for many types of procedures (short or long period) depending on the evolution of the wound. Nonabsorbable sutures are utilized to close dermal tissue, which removal normally takes 10e14 days, although the time may change depending on the location and wound environment. The use of nonabsorbable sutures is also destined to remain indefinitely or permanently in tissues where wound healing is difficult, or even when the tissue lacks the strength to remain attached, and is latent the risk of the wound being exposed again [9].

Functionalization of sutures

31

Expanding on the use of nonabsorbable sutures, these are used in organs and tissues including vital organs for the body’s functioning, for example, the heart and blood vessels, whose rhythmic movement requires the suture to remain more than 3 weeks [3]. Another reason for preferring nonabsorbable sutures is because fluid secretion of certain organs, such as the bladder, causes that absorbable sutures to degrade before the wound heals properly. Sometimes, absorbable sutures induce irritation or inflammation in the intervened zone, provoking even biofilm growth, and consequently an infection. In counterpart, nonabsorbable sutures produce a minimal inflammatory reaction in the tissues, usually do not adhere to the tissues since the body does not recognize them, and maintain the knot correctly until they are removed [48]. Due to the superior mechanical properties of nonabsorbable suture materials, that is, greater tensile strength and durability, great efforts have been made to functionalize them. Since many nonabsorbable materials do not show important chemical reactivity, traditional activation methods for their functionalization are limited (Fig. 2.4), thus its functionalization using a high energy source, for example, gamma radiation or plasma techniques, is viable [49]. 2.6.1 Polypropylene sutures The use of gamma rays as a means for free-radical formation to functionalize with acrylic monomers various apparently “inert” materials, such as PP, is widely documented [33,50]. In the first decade of this century, the grafting of sutures employing gamma radiation was already extensively known. For example, acrylonitrile (AN) grafted on PP through

Figure 2.4 Degree of functionalization and bioactivity of suture threads depends on the type and reactivity of chemical groups in the surface.

32

Advanced Technologies and Polymer Materials for Surgical Sutures

preirradiation oxidative technique gives PP-g-AN sutures where the conservation of the mechanical properties of the original material is pointed out, with a small increase of 5% in its toughness [51]. These grafting materials can subsequently be hydrolyzed to obtain an AAc derivative, i.e., a PP-g-AAc suture that shows stability in both neutral and basic pH media. The hydrolysis functionalizes the surface of the suture with carboxylic groups capable to load/transfer/release tetracycline hydrochloride under certain environmental conditions. The overall process resulted in a suture with effective antimicrobial properties against in vitro cultures of Escherichia coli (E. coli), Klebsiella pneumonea (K. pneumonea), and Staphylococcus aureus (S. aureus) [52,53]. Our Laboratory has taken up the line of research on suture threads and has developed graft materials mediated by gamma radiation with AAc and GMA on PP sutures, for the irreversible immobilization of vancomycin through the formation of a covalent bond with the epoxy group of the PPg-GMA graft. In addition, drug release studies were performed with the PP-g-AAc suture, with positive results in adhesion and inhibition tests in vitro against S. aureus [18]. These examples consolidate the fundamental idea of chemical functionalization or loading in suture threads with potential antimicrobial activity. 2.6.1.1 Functionalization with azoles Azole-derived heterocycle structures are known to be part of a wide variety of drugs such as antivirals, antiinflammatories, analgesics, antidepressants, and anticancer, but mainly for being used as antimicrobials [54e56]. If we focus our attention only on antimicrobial activity, azoles are an excellent option given their inherent biocidal properties [57]. As an example, we have the quaternary salts of butyl-, hexyl- and octylimidazolium, which have in vitro activity against E. coli, S. aureus, B. subtilis, P. fluorescens and S. cerevisiae [58]. Furthermore, according to recent works, imidazole oligomers have been shown to eliminate >99.7% of S. aureus and E. coli within the first 30 s of contact, simply with the formation of imidazolium bromides [59]. Although these polymeric materials are not yet commercially used in drugs, work is underway to graft these oligomers into biomaterials for the manufacture of sutures with passive antibacterial properties, the advantage of which is that they are not invasive and only act in the contact area [60]. A very convenient method for imidazolium group formation is through the quaternization of amines with methyl iodide (MeI). When using

Functionalization of sutures

33

iodinated compounds, the activation energy of the reaction is lower compared with analogous halogenated derivatives (bromides and chlorides) [61,62]. In addition, it has the advantage that the N-alkylation reaction can be carried out in nonpolar solvents and at room temperature. The product of the reaction is an imidazolium iodide derivative that potentiates the antimicrobial activity of the material surface. In previous studies, NVIm had already been grafted into other medical grade materials such as PVC to later functionalize them with MeI and endow the surface with bacteriostatic activity [60], so the same methodology was easily transferred to the functionalization of NVIm-grafted sutures [63]. The advantage is that quaternization with alkyl iodides can be followed visually with the darkening of the suture and quantification can be determined gravimetrically, where antimicrobial activity has been found to be related to the degree of grafting (Fig. 2.3). With respect to the antimicrobial properties of modified suture threads after MeI treatment, in previously grafted PP sutures, it was found that the suture inhibition zone [(PP-g-HEMA)-g-NVIm]/MeI increased progressively depending on the MeI concentration. Properly, the zone of inhibition created by [(PP-g-HEMA)-g-NVIm]/MeI and [(PP-gNIPAAm)-g-NVIm]/MeI in the S. aureus strain was greater than in the E. coli strains, where inhibition was moderate [63]. These results indicate that imidazolium is selective to attacking bacteria, which is crucial evidence when looking for specific antimicrobial capabilities in modified materials. 2.6.1.2 Functionalization with Ag A different approach to modifying nonabsorbable sutures is through the incorporation of Ag into the surface. Studies of the effect of colloidal silver on bacterial inhibition in S. aureus colonies have been quantified positively [64]; also, it is possible to find several papers towards the therapeutic value and possible side effects of silver [65]. Both Ag(I) and Ag(0) show antimicrobial activity. Specially, there are interesting studies when silver nitrate, AgNO3, is used in treatments to combat diseases of viral origin [66]. The most used source to obtain Ag(I) is AgNO3, which possesses an antimicrobial effect and desirable characteristics such as high solubility and relative biocompatibility. Specific studies conducted on S. aureus and E. coli reveal that Ag(I) ions kill bacteria thanks to a mechanism that damages the membrane and interferes with metabolic activity at the intracellular level [67]. Silver does not necessarily have to be in the solution. It is possible to incorporate nanoparticles that destroy the target cell. There is still no

34

Advanced Technologies and Polymer Materials for Surgical Sutures

consensus on the mechanisms associated with the activity of nanoparticles (cell damage) and their correlation with particle size [68], and still research is to be made. In the reduction methods, in general, the ionic salts of Ag(I) can be photo-reduced in a homogeneous aqueous medium using intermediate Ag(I) complexes with organic binders [69]. In situ chemical reduction of Ag(I) with NaBH4 on porous poly(NVIm) beads have also been recently used, resulting in Ag(0) nanoparticles in an acidic medium, with the ability to control and eliminate S. aureus and E. coli [70]. While in Ag photoreduction, unlike chemical methods, the formation of waste that requires further purification is avoided and large amounts of energy are not required. Photoreduction is carried out in a controlled manner in reactions with polymers such as polyethylene diamine [71], allowing the formation of Ag(0) nanoparticles, in an eco-friendly way [72], easy, and sustainable [39]. An important mechanism through, which polymer chains reduce Ag, possibly proceeds through the complexation of the Ag salt with the heteroatoms of the carbonyl groups [73], hydroxyl, amide, or imidazole [74], as the case may be. Such intermediates are formed between the grafted chains or on the surface of the material. In a second step; either by the action of sunlight, thermal heating, the action of a chemical reducer, or a combination; the reduced Ag particles are formed, remaining adhered “on” or “among” the polymer chains whose noncovalent interactions between the metal and polymer are strong enough to remain as a unit in the suture filament. In recent years, a string of articles on Ag suture functionalization has shown different angles of the same principle about loading or fixing metal on the biomaterial surface. Adhesion is facilitated if the suture surface meets one or more of the following properties: high surface area (high porosity), high density of reactive functional groups on the surface, and swelling capacity. Some more representative examples are mentioned in the following paragraphs. Modification of PP sutures with vinyl monomers was discussed in previous sections, but we will delve into some investigations. In a study carried out by our working group, graft copolymers of PP-g-HEMA-gNVIm, PP-g-NIPAAm-g-NVIm, and PP-g-NVIm were compared; all of them loaded with Ag(0) by photoreduction with natural light [75]. The loading with Ag was made using aqueous solutions of different concentrations of AgNO3 (10-10,000 mg L1) and under environmental conditions around (25
C), which meant an advance in the methodology, since

Functionalization of sutures

35

heating and the use of reducing agents such as sodium citrate or thiosulfite were avoided. For the reduction to be effective, the suture must be able to swell in an aqueous medium, which guarantees maximum absorption of Agþ ions between the chains. Swelling occurs within the first 4 h, practically reaching the limit after 24 h of reaction. The experiments showed that more concentrated solutions and with higher percentages of swelling in water loaded more Ag. In addition, by comparison, the pristine PP sutures, as well as the grafted PP monofilaments were subjected to treatment with AgNO3 and were prepared for the in vitro bioactivity tests against S. aureus and E. coli strains. The results showed that the pristine PP positive controls did not inhibit the bacteria growth since the PP surfaces have practically no chemical reactivity and were not capable of loading Ag; while the grafted suture threads treated with AgNO3 showed an outstanding ability to inhibit microbial growth after 24 h. Inhibition patterns also provide us with interesting information. [(PP-gHEMA)-g-NVIm/Ag was found to produce inhibition halos from concentrations as low as 10 mg L1, and showed an increase in the zone of inhibition consistently to 1000 mg L1 to then remain constant up to 10,000 mg L1. On the other hand, the grafted copolymer [(PP-gNIPAAm)-g-NVIm/Ag inhibited bacterial growth for 100 mg L1. Another important fact is that Ag particles are more effective in S. aureus (Gram-positive) cultures compared to E. coli (Gram-negative) cultures. This was due to the fact that the S. aureus strains were more susceptible to low concentrations of Ag for the two grafts analyzed when an increase in the inhibition halo was observed consistently up to 10,000 mg L1 (8 mm); while in cultures with E. coli, only the maximum inhibition zone of 5 mm was obtained [75]. The results indicate that the activity of the Ag particles is subject to both, the graft percentages, the type of polymer graft, and the amount of metal adhered to the surface. The BALB/3T3 cell lines were tested with the same sutures used to determine the inhibitory capacity, that is, with [(PP-g-HEMA)-g-NVIm (36/22%)]/Ag and [(PP-g-NIPAAm)-gNVIm (20/22%)]/Ag. Thanks to the WST-1 assays, cytocompatibility was determined to be acceptable after 24 h of incubation, in addition, greater cell survival was observed with the material [(PP-g-HEMA)-g-NVIm (36/ 22%)]/Ag, which implies that PHEMA is a highly biocompatible polymer. 2.6.2 Modified silk sutures Silk suture is a special case, this type is a braided multifilament coated with beeswax. It is a natural polymeric suture made from long-lasting fiber proteins, for this reason, it can be considered nonabsorbable, although after

36

Advanced Technologies and Polymer Materials for Surgical Sutures

1 year it loses tensile strength and suffers degradation, according to reports from in vivo studies [76]. The reason to modify silk sutures lies in their great flexibility and maneuverability as well as tying strength. In fact, silk fibers have been modified at a structural level to accelerate their biodegradation as in the work by Jo et al. where 4-hexylresorcinol was incorporated into silk fibers. 4-Hexylresorcinol fulfilled a double function, first to promote proteolysis and second as an antiseptic to inhibit bacterial growth [77]. In silk fibers, as with other sutures, there is a risk of bacterial overgrowth. This is the reason why, in 2020, Baygar carried out the modification of silk sutures through the application of a coating with propolis [78], taking advantage of the fact that silk contains eCO- and eNH- groups, propolis can efficiently adhere to the surface. Once the coating was done, the loading was carried out with Ag nanoparticles of biogenic origin. The results indicated the synthesis of a functional and broad-spectrum antibacterial capacity suture. In summary, of the different methods for functionalization of nonabsorbable sutures, the Ag inclusion technique on the surface is presented as the alternative with the greatest projection thanks to its biocompatibility, flexibility and inhibition capacity under different chemical environments, in addition to high yields in metal loading and ease of obtaining.

2.7 Functionalization of absorbable sutures Absorbable sutures are composed of materials assimilable by organic tissues and are digested by proteolytic enzymes or hydrolyzed by tissue fluids after a certain time of exposure, examples of sutures with these characteristics are Catgut Simple, Chromic Catgut, Vicryl, PDS, Dexon, Maxon, Resorba PGA, etc. Unlike non-absorbable sutures, equipping suture threads with antimicrobial properties is an idea already grounded in commercial absorbable threads. For example, polyglycolic acid sutures (merchandised under the name Vicryl from Ethicon) have been coated with Ag nanoparticles [79]. Using proven and available antibiotics in commercial sutures seems the most intuitive and feasible way to modify sutures. Modifications can be designed according to a certain release profile, but in general, it is desirable that the diffusion of antibiotics reach maximum effectiveness at least 2 days after the application. There are already reports about modifications with triclosan (a powerful antibacterial and fungicide) for the functionalization of absorbable sutures with positive results, so it is plausible to think that in the

Functionalization of sutures

37

future more of these antibiotic/suture combinations will become more common [80,81]. 2.7.1 Functionalization with silver In Gallo 2016 work, Ag clusters were deposited on Coated VICRYL (polyglactin 910) suture by AgNO3 photoreduction [82]. These multifilament-like sutures were tested in bacterial cultures of S. aureus and E. coli, inhibiting both even after 21 days. The Ag particles act through suture degradation-controlled release mechanisms, where half of the Ag load is released within the first 7 days. It is worth mentioning that the sutures were tested before and after the Ag was added and that the useful life of the suture (preprogrammed) does not seem to be significantly affected. In fact, in the degradation kinetics experiment, physiological conditions were simulated with phosphate buffer solutions at pH 7.4 for 4 weeks, it seems that the degradation of the suture helps to the diffusion of the Ag particles and, therefore, inhibits any bacterial activity. Regarding the functionality, stability, and duration of the suture threads, microscopic and elemental analyzes show that there is uniformity in Ag deposition along the surface. The MTT cytotoxicity results support the conclusion that these types of sutures are suitable for use in the closure of certain types of wounds. Two years later in 2018 Gallo and his collaborators used the same type of absorbable suture as Polyglactin 910 to improve Ag adherence but previously coated it with silk sericin [83]. The sutures were activated in a basic medium with sodium hydroxide, in this way, the adhesion of the silk sericin is promoted, to load the Ag by dipping it in a AgNO3 solution in a later step. The sutures were tested in the presence of strains of S. aureus and E. coli, showing inhibition in microbial proliferation for up to 21 days. 2.7.2 Chitin sutures An entirely organic alternative to absorbable sutures is to use sutures made from chitin fibers. Chitin is a glucose derivative present in arthropod exoskeletons and some crustaceans and mollusks. Its linear structure allows fibrous structures to be built, with adequate resistance for absorbable sutures. Prolonged exposure to tissues does not produce adverse reactions or at least is minimal, allowing an appropriate environment for healing. Compared to other sutures such as catgut and Dexon, chitin retains its mechanical properties (tensile strength, Young modulus, and flexibility) for longer days [84].

38

Advanced Technologies and Polymer Materials for Surgical Sutures

The development of more and better chitin sutures has been enhanced thanks to research focused on improving the physical properties of the material. It is possible to reinforce the fibers by doping with cellulose nanocrystals of bacterial origin [85]. The nano cellulose is added to the chitin powder together with a solution of urea and NaOH. In a second process, the solution is subjected to a “wet-spinning” process to interweave the fibers that constitute the suture threads [86]. The advantage of these sutures is that they are fully biodegradable, biocompatible, and promote cell regeneration, this helps to accelerate wound healing, demonstrating that an antibiotic is not necessary to reduce the risk of infection. 2.7.3 Caprolactam sutures gentamicin/silver loaded As we have already noticed, absorbable sutures can be loaded with a commercial antibiotic or Ag, but also with a combination of both types of antimicrobials, such as caprolactam sutures that were built in one step from the polymer (beads) and mixing them with dispersed Ag and a solution of gentamicin. The mixture was treated by electrospinning using a “spinneret” and applying a voltage of 12 kV. Once the fibers of different gauges (3e12 mm width) were obtained, they were wound into multifilament of different diameters. The sutures are sterilized and subjected to biocompatibility tests with HaCaT cell lines with acceptable results to be used as medical healing devices. As for the antibacterial activity, this was performed with colonies of Pseudomonas aeruginosa, and it was observed that the combination of drugs actually helps, also in the wound healing tests, there is no healing interference with respect to the control sutures. 2.7.4 Drug-loading on absorbable sutures In addition to antibiotics, absorbable sutures can also be loaded with antiinflammatory drugs. The reason for adding this type of drug in suture materials is justified based on the prevention of possible damage caused by some tissue reaction, since, as mentioned, combined factors such as the nature of the material and exposure to tissue can cause a response by the immune system in the presence of a foreign body. There are reported works where vicryl absorbable sutures are modified, which basically consists of a copolymer composite of PLGA with diclofenac [87]. In this work, diclofenac was loaded directly into a mixture of dichloromethane and N-N-dimethylformamide (DMF), the mixture was impregnated through

Functionalization of sutures

39

electrospray. The release of diclofenac was monitored through UV-Vis and the results indicate that more than 50% of the drug is released within the first 48 h and continues to release until day 10. In the same way as with other sutures, by modifying them with nonaggressive reaction methods and conditions, the mechanical properties of the sutures are not affected and are functional.

2.8 Conclusions Modifying the structures and surfaces of the suture threads is a reliable and practical way to develop biomedical materials with enhanced or tailor-made properties. Whether in absorbable or nonabsorbable sutures, it is possible to make superficial modifications by various methods, which helps in the healing process and, consequently, the patient recovery. The methods for superficial modifications can be categorized into two main groups, chemical (such as grafts, derivatizations, and complexations) and physical (such as coatings, loading, and doping). Surface modification of sutures is a novel field in constant change and improvement. The search continues so that the operation of the original materials does not appreciably change and improve the healing. Furthermore, and in accordance with what is explained in this chapter, it can be concluded that the functionalization of sutures with antimicrobial properties is affordable. The main objective is to get safer devices for the postoperative healing process.

Acknowledgments This work was supported by Dirección General de Asuntos del Personal Académico, Universidad Nacional Autónoma de México under Grant IN202320.

References [1] S. Kathju, et al., Chronic surgical site infection due to suture-associated polymicrobial biofilm, Surg. Infect. 10 (5) (2009) 457e461, https://doi.org/10.1089/sur.2008.062. [2] C. Dennis, et al., Suture materials d current and emerging trends, J. Biomed. Mater. Res. 104 (6) (2016) 1544e1559, https://doi.org/10.1002/jbm.a.35683. [3] R.E. Abhari, et al., Synthetic sutures: clinical evaluation and future developments, J. Biomater. Appl. 32 (3) (2017) 410e421, https://doi.org/10.1177/ 0885328217720641. [4] D. Leaper, et al., Antimicrobial sutures and prevention of surgical site infection: assessment of the safety of the antiseptic triclosan, Int. Wound J. 8 (6) (2011) 556e566, https://doi.org/10.1111/j.1742-481X.2011.00841.x. [5] D. Leaper, et al., The role of antimicrobial sutures in preventing surgical site infection, Ann. R. Coll. Surg. Engl. 99 (6) (2017) 439e443, https://doi.org/10.1308/ rcsann.2017.0071.

40

Advanced Technologies and Polymer Materials for Surgical Sutures

[6] M. Byrne, A. Aly, The surgical suture, Aesthetic Surg. J. 39 (Suppl. ment_2) (2019) S67eS72, https://doi.org/10.1093/asj/sjz036. [7] B. Joseph, et al., Polymer sutures for simultaneous wound healing and drug delivery e a review, Int. J. Pharm. (2017), https://doi.org/10.1016/j.ijpharm.2017.03.041. [8] C.C. Snyder, On the history of the suture, Plast. Reconstr. Surg. 58 (4) (1976) 401e406, https://doi.org/10.1097/00006534-197610000-00001. [9] A. Goel, Surgical sutures - a review, Delhi J. Ophthalmol. (26) (2016) 159e162, https://doi.org/10.7869/djo.161. [10] D.J. Barillo, D.E. Marx, Silver in medicine: a brief history BC 335 to present, Burns 40 (S1) (2014) S3eS8, https://doi.org/10.1016/j.burns.2014.09.009. [11] H.A. Maddah, Polypropylene as a promising plastic : a review, Am. J. Polym. Sci. 6 (1) (2016) 1e11, https://doi.org/10.5923/j.ajps.20160601.01. [12] A. Bruinink, R. Luginbuehl, Evaluation of biocompatibility using in vitro methods: interpretation and limitations, in: C. Kasper, F. Witte, R. Pörtner (Eds.), Tissue Engineering III: Cell - Surface Interactions for Tissue Culture, Springer, Berlin, Heidelberg, 2011, pp. 117e152, https://doi.org/10.1007/10_2011_111. [13] D. Poburski, R. Thierbach, Improvement of the BALB/c-3T3 cell transformation assay: a tool for investigating cancer mechanisms and therapies, Sci. Rep. 6 (1) (2016) 32966, https://doi.org/10.1038/srep32966. [14] L.-M. Yin, et al., Long term and standard incubations of WST-1 reagent reflect the same inhibitory trend of cell viability in rat airway smooth muscle cells, Int. J. Med. Sci. 10 (1) (2013) 68e72, https://doi.org/10.7150/ijms.5256. [15] S.N. Rampersad, Multiple applications of alamar blue as an indicator of metabolic function and cellular health in cell viability bioassays, Sensors 12 (9) (2012) 12347e12360, https://doi.org/10.3390/s120912347. [16] A.L. Inselman, et al., Reevaluation of the embryonic stem cell test, Int. J. Regul. Sci. 1 (1) (2013) 32e49. [17] F. López-Saucedo, C. Alvarez-Lorenzo, et al., Radiation-grafting of vinyl monomers separately onto polypropylene monofilament sutures, Radiat. Phys. Chem. 132 (2017) 1e7, https://doi.org/10.1016/j.radphyschem.2016.11.006. [18] M. García-Vargas, et al., Acrylic polymer-grafted polypropylene sutures for covalent immobilization or reversible adsorption of vancomycin, Int. J. Pharm. 461 (1e2) (2014) 286e295, https://doi.org/10.1016/j.ijpharm.2013.11.060. [19] J.-P. Montheard, M. Chatzopoulos, D. Chappard, 2-Hydroxyethyl methacrylate (HEMA): chemical properties and applications in biomedical fields, J. Macromol. Sci., Part C 32 (1) (1992) 1e34, https://doi.org/10.1080/15321799208018377. [20] A.S.H. Makhlouf, Current and advanced coating technologies for industrial applications, in: A.S.H. Makhlouf, I.B.T.-N. Tiginyanu, F. U.-T. (Eds.), Nanocoatings and Ultra-thin Films, first ed., Woodhead Publishing, 2011, pp. 3e23, https://doi.org/ 10.1533/9780857094902.1.3. [21] X. Tang, X. Yan, Dip-coating for fibrous materials: mechanism, methods and applications, J. Sol. Gel Sci. Technol. 81 (2) (2017) 378e404, https://doi.org/10.1007/ s10971-016-4197-7. [22] Y. Yang, et al., Bacterial inhibition potential of quaternised chitosan-coated VICRYL absorbable suture: an in vitro and in vivo study, J. orthop. Transl. 8 (2017) 49e61, https://doi.org/10.1016/j.jot.2016.10.001. [23] A. Link, et al., Development of a novel polymer-based mRNA coating for surgical suture to enhance wound healing, Coatings 9 (6) (2019) 374e390, https://doi.org/ 10.3390/coatings9060374. [24] S. Saad, et al., Effect of TiO2-nanoparticles on Ni electrodeposition on copper wire, Metals 10 (3) (2020) 406e419, https://doi.org/10.3390/met10030406.

Functionalization of sutures

41

[25] D.C. Parada Quinayá, et al., Obtención, electrodeposición y caracterización de un recubrimiento polímerico bioabsorbable a partir de ácido L - láctico para aplicaciones biomédicas, Ingeniare Rev. Chil. Ing. 17 (3) (2009) 365e374, https://doi.org/ 10.4067/S0718-33052009000300010. [26] T.M. Benedetti, et al., Electrostatic layer-by-layer deposition and electrochemical characterization of thin films composed of MnO2 nanoparticles in a room-temperature ionic liquid, Langmuir 24 (7) (2008) 3602e3610, https://doi.org/10.1021/la702347x. [27] J. Lyklema, L. Deschênes, The first step in layer-by-layer deposition: electrostatics and/ or non-electrostatics? Adv. Colloid Interface Sci. 168 (1e2) (2011) 135e148, https:// doi.org/10.1016/j.cis.2011.03.008. [28] H.-G. Elias, Free radical polymerization, in: H.-G. Elias (Ed.), Macromolecules, Springer US, Boston, MA, 1984, pp. 681e736, https://doi.org/10.1007/978-1-48992809-2_6. [29] E.A. Kovtun, et al., Stability of polycaproamide suture thread with coatings in radiation sterilization, Fibre Chem. 31 (6) (1999) 455e459, https://doi.org/10.1007/ BF02358259. [30] N.-J. Huang, D.C. Sundberg, Fundamental studies of grafting reactions in free radical copolymerization. I. A detailed kinetic model for solution polymerization, J. Polym. Sci. Polym. Chem. 33 (15) (1995) 2533e2549, https://doi.org/10.1002/ pola.1995.080331502. [31] M. Kumari, B. Gupta, S. Ikram, Characterization of N-isopropyl acrylamide/acrylic acid grafted polypropylene nonwoven fabric developed by radiation-induced graft polymerization, Radiat. Phys. Chem. 81 (11) (2012) 1729e1735, https://doi.org/ 10.1016/j.radphyschem.2012.05.018. [32] A. Ramírez-Jiménez, et al., Radiation-grafting of 2-hydroxyethylmethacrylate and oligo (ethylene glycol) methyl ether methacrylate onto polypropylene films by one step method, Radiat. Phys. Chem. 81 (1) (2012) 27e32, https://doi.org/10.1016/ j.radphyschem.2011.09.002. [33] E. Zavala-Lagunes, et al., Synthesis and characterization of stimuli-responsive polypropylene containing N-vinylcaprolactam and N-vinylimidazole obtained by ionizing radiation, Mater. Sci. Eng. C 67 (2016) 353e361, https://doi.org/10.1016/ j.msec.2016.05.044. [34] H.I. Meléndez-Ortiz, G. Burillo, E. Bucio, Aplicaciones de radiación gamma y UV para el injerto molecular y el diseño de matrices para aplicaciones terapéuticas, in: H. de Sousa, M.E.M. Braga, A. Sosnik (Eds.), Biomateriales aplicados al diseño de sistemas terapéuticos avanzados. Coimbra, Imprensa da Universidade de Coimbra, 2015, pp. 453e484, https://doi.org/10.14195/978-989-26-0881-5_11. [35] K. Nakabayashi, H. Mori, Recent progress in controlled radical polymerization of Nvinyl monomers, Eur. Polym. J. 49 (10) (2013) 2808e2838, https://doi.org/10.1016/ j.eurpolymj.2013.07.006. [36] E. Bucio, G. Burillo, Radiation-induced grafting of sensitive polymers, J. Radioanal. Nucl. Chem. 280 (2) (2009) 239e243, https://doi.org/10.1007/s10967-009-0505-9. [37] J.D. Mota-Morales, et al., Free-radical polymerizations of and in deep eutectic solvents: green synthesis of functional materials, Prog. Polym. Sci. 78 (1) (2018) 139e153, https://doi.org/10.1016/j.progpolymsci.2017.09.005. [38] B.A. Miller-Chou, J.L. Koenig, A review of polymer dissolution, Prog. Polym. Sci. 28 (8) (2003) 1223e1270, https://doi.org/10.1016/S0079-6700(03)00045-5. [39] P.T. Anastas, M.M. Kirchhoff, Origins, current status, and future challenges of green chemistry, Acc. Chem. Res. 35 (9) (2002) 686e694, https://doi.org/10.1021/ ar010065m.

42

Advanced Technologies and Polymer Materials for Surgical Sutures

[40] A. Székely, M. Klussmann, Molecular radical chain initiators for ambient- to lowtemperature applications, Chem. Asian J. 14 (1) (2018) 105e115, https://doi.org/ 10.1002/asia.201801636. [41] S. Yang, et al., Effect of reaction temperature on grafting of g-aminopropyl triethoxysilane (APTES) onto kaolinite, Appl. Clay Sci. 62 (63) (2012) 8e14, https:// doi.org/10.1016/j.clay.2012.04.006. [42] S. Guragain, et al., Multi-stimuli-responsive polymeric materials, Chem. Eur J. 21 (38) (2015) 13164e13174, https://doi.org/10.1002/chem.201501101. [43] M. Wei, et al., Stimuli-responsive polymers and their applications, Polym. Chem. 8 (1) (2017) 127e143, https://doi.org/10.1039/C6PY01585A. [44] X. Li, et al., Thermoresponsive PNIPAAM bottlebrush polymers with tailored sidechain length and end-group structure, Soft Matter 10 (12) (2014) 2008e2015, https://doi.org/10.1039/c3sm52614c. [45] J.E. López-Barriguete, T. Isoshima, E. Bucio, Development and characterization of thermal responsive hydrogel films for biomedical sensor application, Mater. Res. Express 5 (4) (2018) 45703, https://doi.org/10.1088/2053-1591/aabb0c. [46] E.B. Anderson, T.E. Long, Imidazole- and imidazolium-containing polymers for biology and material science applications, Polymer 51 (12) (2010) 2447e2454, https:// doi.org/10.1016/j.polymer.2010.02.006. [47] M. Takafuji, et al., Preparation of poly(1-vinylimidazole)-grafted magnetic nanoparticles and their application for removal of metal ions, Chem. Mater. 16 (10) (2004) 1977e1983, https://doi.org/10.1021/cm030334y. [48] H. Sun, L. Xie, An introduction of structure, synthesis and safety concerning polypropylene application, in: L.P. Silva, E.F. Barbosa (Eds.), Polypropylene, Nova Sciences Publishers, Inc., New York, 2013, pp. 1e10. [49] S. Saxena, et al., Development of a new polypropylene-based suture: plasma grafting, surface treatment, characterization, and biocompatibility studies, Macromol. Biosci. 11 (3) (2011) 373e382, https://doi.org/10.1002/mabi.201000298. [50] H.I. Meléndez-Ortiz, et al., Binary graft modification of polypropylene for antiinflammatory drug-device combo products, J. Pharmaceut. Sci. 103 (4) (2014) 1269e1277, https://doi.org/10.1002/jps.23903. [51] R. Jain, et al., Preparation of antimicrobial sutures by preirradiation grafting of acrylonitrile onto polypropylene monofilament. II. Mechanical, physical, and thermal characteristics, J. Appl. Polym. Sci. 93 (3) (2004) 1224e1229, https://doi.org/ 10.1002/app.20543. [52] B. Gupta, et al., Preparation of antimicrobial sutures by preirradiation grafting of acrylonitrile onto polypropylene monofilament. III. Hydrolysis of the grafted suture, J. Appl. Polym. Sci. 94 (6) (2004) 2509e2516, https://doi.org/10.1002/app.21211. [53] B. Gupta, R. Jain, H. Singh, Preparation of antimicrobial sutures by preirradiation grafting onto polypropylene monofilament, Polym. Adv. Technol. 19 (12) (2008) 1698e1703, https://doi.org/10.1002/pat.1146. [54] J.A. Maertens, History of the development of azole derivatives, Clin. Microbiol. Infect. 10 (2004) 1e10, https://doi.org/10.1111/j.1470-9465.2004.00841.x. [55] V. Gupta, V. Kant, A review on biological activity of imidazole and thiazole moieties and their derivatives, Sci. Int. 1 (7) (2013) 253e260, https://doi.org/10.17311/ sciintl.2013.253.260. [56] A.S.W. Oak, J.W. Baddley, B.E. Elewski, Systemic antifungals, in: P.S. Yamauchi (Ed.), Biologic and Systemic Agents in Dermatology, Springer International Publishing, Cham, 2018, pp. 425e450, https://doi.org/10.1007/978-3-319-66884-0_40. [57] P. Borowiecki, et al., Synthesis and antimicrobial activity of imidazolium and triazolium chiral ionic liquids, Eur. J. Org Chem. (4) (2013) 712e720, https://doi.org/ 10.1002/ejoc.201201245.

Functionalization of sutures

43

[58] K.M. Docherty, C.F. Kulpa Jr., Toxicity and antimicrobial activity of imidazolium and pyridinium ionic liquids, Green Chem. 7 (4) (2005) 185e189, https://doi.org/ 10.1039/b419172b. [59] S.N. Riduan, et al., Ultrafast killing and self-gelling antimicrobial imidazolium oligomers, Small 12 (14) (2016) 1928e1934, https://doi.org/10.1002/smll.201600006. [60] H.I. Meléndez-Ortiz, et al., Modification of medical grade PVC with Nvinylimidazole to obtain bactericidal surface, Radiat. Phys. Chem. 119 (2016) 37e43, https://doi.org/10.1016/j.radphyschem.2015.09.014. [61] G. Glockler, Carbonehalogen bond energies and bond distances, J. Phys. Chem. 63 (6) (1959) 828e832, https://doi.org/10.1021/j150576a013. [62] R.J. Ouellette, J.D. Rawn, Introduction to organic reaction mechanisms, in: R.J. Ouellette, J.D. Rawn (Eds.), Organic Chemistry, second ed., Elsevier, 2018, pp. 51e86, https://doi.org/10.1016/B978-0-12-812838-1.50003-7. [63] F. López-Saucedo, G.G. Flores-Rojas, et al., Achieving antimicrobial activity through poly(N-methylvinylimidazolium) iodide brushes on binary-grafted polypropylene suture threads, MRS Commun. 7 (4) (2017) 938e946, https://doi.org/10.1557/ mrc.2017.121. [64] R. Goggin, et al., Colloidal silver: a novel treatment for Staphylococcus aureus biofilms? Int. Forum Allergy Rhinol. 4 (3) (2014) 171e175, https://doi.org/10.1002/alr.21259. [65] S. Chernousova, M. Epple, Silver as antibacterial agent: ion, nanoparticle, and metal, Angew. Chem. Int. Ed. 52 (6) (2013) 1636e1653, https://doi.org/10.1002/ anie.201205923. [66] S. Ebrahimi, et al., Efficacy of 10% silver nitrate solution in the treatment of common warts: a placebo-controlled, randomized, clinical trial, Int. J. Dermatol. 46 (2) (2007) 215e217, https://doi.org/10.1111/j.1365-4632.2007.02955.x. [67] W.K. Jung, et al., Antibacterial activity and mechanism of action of the silver ion in Staphylococcus aureus and Escherichia coli, Appl. Environ. Microbiol. 74 (7) (2008) 2171e2178, https://doi.org/10.1128/AEM.02001-07. [68] G. Franci, et al., Silver nanoparticles as potential antibacterial agents, Molecules 20 (5) (2015) 8856e8874, https://doi.org/10.3390/molecules20058856. [69] H. Hada, et al., Photoreduction of silver ion in aqueous and alcoholic solutions, J. Phys. Chem. 80 (25) (1976) 2728e2731, https://doi.org/10.1021/j100566a003. [70] M.A. Mudassir, et al., Development of silver-nanoparticle-decorated emulsiontemplated hierarchically porous poly(1-vinylimidazole) beads for water treatment, ACS Appl. Mater. Interfaces 9 (28) (2017) 24190e24197, https://doi.org/10.1021/ acsami.7b05311. [71] S. Tan, et al., Synthesis of positively charged silver nanoparticles via photoreduction of AgNO3 in branched polyethyleneimine/HEPES solutions, Langmuir 23 (19) (2007) 9836e9843, https://doi.org/10.1021/la701236v. [72] V. Manikandan, et al., Green synthesis of silver oxide nanoparticles and its antibacterial activity against dental pathogens, 3 Biotech 7 (1) (2017) 72e80, https://doi.org/ 10.1007/s13205-017-0670-4. [73] M.N. Siddiqui, et al., Synthesis and characterization of poly(2-hydroxyethyl methacrylate)/silver hydrogel nanocomposites prepared via in situ radical polymerization, Thermochim. Acta 643 (2016) 53e64, https://doi.org/10.1016/j.tca.2016.09.017. [74] M. McCann, et al., Synthesis, structure and biological activity of silver(I) complexes of substituted imidazoles, Polyhedron 56 (2013) 180e188, https://doi.org/10.1016/ j.poly.2013.03.057. [75] F. López-Saucedo, et al., Antimicrobial silver-loaded polypropylene sutures modified by radiation-grafting, Eur. Polym. J. 100 (2018) 290e297, https://doi.org/10.1016/ j.eurpolymj.2018.02.005.

44

Advanced Technologies and Polymer Materials for Surgical Sutures

[76] G. Thilagavathi, S. Viju, Silk as a suture material, in: A. Basu (Ed.), Advances in Silk Science and Technology, Elsevier, Coimbatore, India, 2015, pp. 219e232, https:// doi.org/10.1016/B978-1-78242-311-9.00011-2. [77] Y.-Y. Jo, et al., Accelerated biodegradation of silk sutures through matrix metalloproteinase activation by incorporating 4-hexylresorcinol, Sci. Rep. 7 (2017) 42441, https://doi.org/10.1038/srep42441. [78] B. Tuba, Characterization of silk sutures coated with propolis and biogenic silver nanoparticles (AgNPs); an eco-friendly solution with wound healing potential against surgical site infections (SSIs), Turk. J. Med. Sci. (2020) 258e266, https://doi.org/ 10.3906/sag-1906-48. [79] C.H. Ho, et al., Long-term active antimicrobial coatings for surgical sutures based on silver nanoparticles and hyperbranched polylysine, J. Biomater. Sci. Polym. Ed. 24 (13) (2013) 1589e1600, https://doi.org/10.1080/09205063.2013.782803. [80] S. Rothenburger, et al., In vitro antimicrobial evaluation of coated VICRYL* plus antibacterial suture (coated polyglactin 910 with triclosan) using zone of inhibition assays, Surg. Infect. 3 (s1) (2002) s79es87, https://doi.org/10.1089/sur.2002.3.s1-79. [81] S. Hoshino, et al., A study of the efficacy of antibacterial sutures for surgical site infection: a retrospective controlled trial, Int. Surg. 98 (2) (2013) 129e132, https:// doi.org/10.9738/CC179. [82] A.L. Gallo, et al., Efficacy of silver coated surgical sutures on bacterial contamination, cellular response and wound healing, Mater. Sci. Eng. C 69 (2016) 884e893, https:// doi.org/10.1016/j.msec.2016.07.074. [83] A.L. Gallo, M. Pollini, F. Paladini, A combined approach for the development of novel sutures with antibacterial and regenerative properties: the role of silver and silk sericin functionalization, J. Mater. Sci. Mater. Med. 29 (8) (2018) 133, https://doi.org/ 10.1007/s10856-018-6142-5. [84] M. Nakajima, et al., Chitin is an effective material for sutures, Jpn. J. Surg. 16 (6) (1986) 418e424, https://doi.org/10.1007/BF02470609. [85] H. Wu, et al., Regenerated chitin fibers reinforced with bacterial cellulose nanocrystals as suture biomaterials, Carbohydr. Polym. 180 (2018) 304e313, https://doi.org/ 10.1016/j.carbpol.2017.10.022. [86] H. Wu, et al., A novel multifunctional biomedical material based on polyacrylonitrile: preparation and characterization, Mater. Sci. Eng. C 62 (2016) 702e709, https:// doi.org/10.1016/j.msec.2016.02.026. [87] B.K. Huh, et al., Surgical suture braided with a diclofenac-loaded strand of poly(lacticco -glycolic acid) for local, sustained pain mitigation, Mater. Sci. Eng. C 79 (2017) 209e215, https://doi.org/10.1016/j.msec.2017.05.024.

CHAPTER 3

Improving the therapeutic value of sutures Lorena Duarte-Peña1, Amira J. Fragoso-Medina2, Emilio Bucio1 and Felipe López-Saucedo1 1 Departamento de Química de Radiaciones y Radioquímica, Instituto de Ciencias Nucleares, Universidad Nacional Autónoma de México, Circuito Exterior, Ciudad Universitaria, CDMX, Mexico; 2 Departamento de Ciencias Químicas, Facultad de Estudios Superiores Cuautitlán, Universidad Nacional Autónoma de México, Mexico City, Mexico

3.1 Content This chapter discusses the role of sutures, their improvement throughout time, and their therapeutic value, including general concepts and examples. First, the history of sutures and some of their characteristics are mentioned, followed by the classification and features of current threads. Also, the modification of sutures, as a tendency to improve their bioactive properties, comprising evaluation criteria such as structural modifications (i.e., fiber dimensions, topography, and microstructure), chemical modification, studies of drug delivery, and new materials with stimuli responsiveness (pHand thermo-responsive polymers) is discussed. Finally, a brief conclusion summarizes the importance of suture improvement.

3.2 General concepts Since immemorial times, the wish to maintain physical health has been an inherent characteristic of humanity. Certainly, the human lifestyle or modus vivendi has changed through history according to needs, feeding, and resources. In this context of health, sutures have played an important role as a medical device since ancient civilizations, so there is an extensive list of materials, which includes plants, minerals, and animal parts, as well as synthetic products, which have served for this purpose. The new advances in technology, combined with the use of resources, have been determinant in wound care and helped to increase the life expectancy. The main function of a suture is the approximation of tissues and blood vessels to facilitate closure of open wounds caused by surgeries or traumas, which is necessary to avoid bleeding and minimizing risks of infections Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00003-6

© 2023 Elsevier Ltd. All rights reserved.

45

46

Advanced Technologies and Polymer Materials for Surgical Sutures

helping to successful patient recovery. The most common disadvantages of the current suture materials are their tendency to be attacked by microorganisms, which increases the morbidity and sometimes causes the rejection of the material by the treated tissue. Therefore, recent studies have been focused on the development of bioactive sutures able to act as cell growth supports to boost tissue regeneration and with antimicrobial properties. 3.2.1 History of sutures Health care has been attended by different characters throughout history, from shamans, witches, healers, or quacks in the far past, to medical professionals in the modern era. First medical devices and treatments were obtained from the empirical knowledge of the community and depended on their environmental resources. For example, some ancient Egyptian documents (redacted 3500 years ago) describe healing processes; the papyrus of Edwin Smith contain war wound treatments and anatomical descriptions [1]; and the Ebers papyrus is a pharmacopeia indicating the treatment of diseases and injuries with using of over 700 substances from natural or minerals origin [2]. Besides, the advancement of the medicine was influenced by the coexistence of cultures in each continent, being the most relevant the Arabica and European cultures that were strongly influenced by Greek culture [3,4]. Even nowadays, the Indian medical system is influenced by the “Sushruta-Samhita” for the treatment of diseases and the balance of the human body. This book series describes the use of medical materials (minerals, animal materials, plants, and blends preparations) and dates from 500 BCE [5]. In ancient India, different materials from vegetable sources were used, for example linen and hemp for external sutures, and absorbable materials for internal sutures [5]. Thereby, the development of biomedical materials was adapted to cover the demands of populations of each age, for example, the suture materials has been used for treatment of wounds since the Middle Age, and although the lack of sanitation, the use of sutures became a usual practice, being silk the most employed material for wound closure [6]. Already in the dawn of the 20th century, the casualties of the First World War triggered the development of medicine, including advances in suturing procedures [7]. During the war, a fast and precise technique of suturing to treat the wounds produced in combat was necessary because it was indispensable to extend the service of the soldiers or at least to guarantee their survival. As a result, the basic principles of emergency wound care and henceforth were settled

Improving the therapeutic value of sutures

47

and the interest to obtain materials with biodegradable characteristics increased, because absorbable synthetic suture materials need less medical care. During these years, synthetic suture materials started crowding out natural ones since they were considered to offer advantages in some regards as production and versatility. Among pioneer countries in synthetic polymer materials, Germany and the United States of America stood out, where were developed some polymers still indispensable and widely useful in the current days, such as polyamides, polyesters, and polypropylene. These developments immortalized the names of scientists, as Wallace Carothers (1896e1937), an inventor chemist leader in some of the polymer research at Dupont for around one decade [8]. Currently, the production of synthetic sutures has been widely extended, as well as the development of absorbable sutures, which are necessary for specific procedures since allow reducing the manipulation of the wound [9]. Therefore, the research on absorbable sutures represents a milestone in the field of medicine. Gradually, the absorbable materials were introduced in the market, the list includes in 1931, the first synthetic polyvinyl alcohol fibers; in 1946, the polyamide suture; in 1960, the polyester suture; and in 1970, a polyglycolic acid braided suture thread [10]. Since the end of the twentieth century, the seeking for new synthetic suturing materials has been a priority. Now are available materials derived from glycolic acid, hydrocarbons, or fluorinated molecules. Also, new molecules were developed to use them as absorbable suture materials, such as 2-octyl-cyanoacrylate, a molecule used as a topical adhesive for superficial skin cut in pediatric, approved by the FDA in 1998 [11]. Besides, the technology around sutures has been improving, as sterilization has perfected, achieving better products with superior performance, for example, the LAT anesthetic gel (lidocaine 1.5%, adrenaline 0.1%, and tetracaine 1%), which is applied topically around the suture zone of uncomplicated wounds [12]. Following is shown a map of the evolution of the material suture for a long time (Fig. 3.1), which allows observing the advance of the suture’s technology and the continuous need to improve them. 3.2.2 Characteristics and classification of sutures The use of sutures is limited according to their composition and function. In certain types of surgical interventions, the physical properties and the biochemical reactivity of the thread are contraindicated being desirable those sutures that minimize risks and damage to the patient’s tissue. Some of positive characteristics are collected in Table 3.1.

48

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 3.1 Evolution of the material suture. (Credit: Original Figure.)

Table 3.1 Characteristics of ideal sutures.

1. 2. 3. 4. 5. 6. 7. 8. 9. 10. 11.

Ensure the continuity of the skin Initiate and promote the wound healing process Predictable behavior Offer a balance of the advantages and disadvantages, considers, of its application Minimally reactive in tissue and not predisposed to bacterial growth Manipulability, due to their scale of use or manufacturing Accessible raw materials Variable or null degradation time of a suture (depending on the material of origin or manufacture) Uniform tensile strength provided by the joint Minimize the damage by friction Easy sterilization

It is possible to classify suture threads according to different criteria, such as its origin, its execution, its structure, or its effect on the organism (permanence or trauma produced), some examples are shown in Fig. 3.2. The existence of different classification criteria results in a wide range of characteristics that can be observed based on: 1. The suture origin, in this classification there are two groups. The first group is conformed of sutures that have a natural origin (animal, vegetal, or mineral) such as silk, cotton, or steel sutures. The second group corresponds to the synthetic origin, these materials may be obtained from total synthesis (as most of polymers) or obtained from the modification of natural materials.

Improving the therapeutic value of sutures

49

Figure 3.2 Sutures according the classification. (Credit: Original Figure.)

2. Another classification of the sutures is based on the execution, that is to say, how they are applied. In this context, sutures can be manual if their application requires manipulation by hand for the action of suturing, such as sutures using laparoscopy [13]; or mechanical sutures, when the suture is performed with aided by a stapling mechanism, such as surgical staple applicators, which are an instrument to perform a mechanical suture between a blood vessel and a synthetic prosthesis [14]. 3. A third classification is based on the structure of the suture thread, which could be monofilament structure, which implies that it is a medical instrument made up of a filament of some specific material as CATGUTTM. Or a multifilament structure that is characterized by making up of more than one filament, within this subclassification there are different types among which are the type sutures braided (MERSILENETM), twisted (SURGICAL suture in STAINLESS STEEL SS), and covered (VYCRILTM) [15]. 4. The last classification is based on the effect on the organism (permanence or trauma produced). According to which, the sutures can be absorbable when are used in surgical procedures, so the suture material is discarded naturally by its degradation and not need medical intervention, as MONOCRYLTM or absorbable surgical sutures based on a lactide containing block terpolymer [16]; or not absorbable if their

50

Advanced Technologies and Polymer Materials for Surgical Sutures

materials resist the action of living mammalian tissues as such NUROLONTM and require manual removal by medical professionals. Although there are different types of sutures, all of them share desirable characteristics as minimal or none toxicity, a low infection rate, a rapid elimination from the body (once its function is fulfilled), a quick execution, a minimum requirement for anesthesia to the well-being of the patient, reduction or elimination of risk for the person who executes the suture, among others. The proper suture depends on the wound; when a medical professional applies a suture, they must choose the suture type bases on acquired and developed criteria with their experience and knowledge, besides they must consider other aspects such as the existence of the material or availability, the kind and localization of the injury, and the normativity. It becomes necessary to highlight the preparation of the suture material because it causes different consequences in the organism according to their ability to injure tissues. It should be noted that suture area trauma are produced when the suture thread is no attached to the needle and must be threaded at the time of suturing [17,18], while deatraumatic sutures are ready to use, they have the thread incorporated in the needle or the thread has a less thickness such as the atraumatic needle [19], in this regard, the development of better surgical sutures has been achieved when atraumatic indices were improved [20]. Some examples of commercial sutures validate the potential of these during the recovery of patients; as stainless steel sutures, which were developed and designed for atrial reconstruction [21]; or polymeric sutures for odontology that were made from natural absorbable biocompatible polymers (polyoxyalkanoates, collagen, chitin, alginate, etc.) to obtain of resistant, elastic, biocompatible, and absorbable thread [22]. 3.2.3 Characteristics of commercial sutures With this variety of materials, the criteria for the choice of a suture can be complex, in other words, when increasing the characteristics considered of a suture, the choice of material becomes more complicated. However, once it is diagnosticated the needing of a suture in an internal or external wound, a type of thread must be selected as soon as possible. The options are various and correct healing will depend on this choice. For this reason, a deep knowledge of sutures may help to make a better decision. The following examples can be mentioned where the contraindications that involve the use of specific suture material are emphasized. The information displayed

Improving the therapeutic value of sutures

51

requires analysis to form an opinion, and data is provided from EthiconTM [10]. Following are showed some commercial sutures, their uses, advantages, and contraindications. First, materials without severe contraindications, in Table 3.2 is showed the comparison between PRONOVATM and MERSILENETM, two commercial products that are synthetic, nonabsorbable, and without known contraindications, which are used for soft tissue, cardiovascular, ophthalmic, and neurological surgery. Being the difference between them that the MERSILENETM is braided while PRONOVATM is monofilament. Another material that can be used in similar conditions (treatment of soft tissue, cardiovascular, ophthalmic, and neurological) but sometimes causes tissue reactions in patients with allergies is the PERMA-HANDTM. A suture classified as braided, of natural origin, and nonabsorbable, which is produced from silk fibroin and whose main advantage is its high tensile strength. Following are mentioned suture materials with contraindications. The first case corresponds to materials that do not present enough mechanical yield for some applications. CATGUT CHROMICTM contraindications include tissue reactions in patients with allergies and it is not indicated where extended retention of tensile strength is required. On the other hand, the application of MONOCRYLTM reports about 60%e70% tensile strength retention at 7 days of postimplantation. Both examples correspond to absorbable monofilaments, but the CHROMIC CATGUTTM is from natural origin and MONOCRYLTM is from synthetic origin. Another criterion is directly related to the contraindications of the product, as in the case of the material PDS IITM that retains 70% of tensile strength at 2 weeks postimplantation. Or the VICRYLTM that is applied in conjunctival sutures and may cause localized irritation. These materials are an example of the same type of suture classified as monofilament, synthetic, and absorbable. Besides, some sutures, main of nonabsorbable materials, cause difficulties Table 3.2 PRONOVATM and MERSILENETM comparison. Trade name

Composition

Advantages

MERSILENE

Poly (ethylene terephthalate)

PRONOVA

Blend of poly (vinylidene fluoride) and poly (vinylidene fluoride-cohexafluoropropylene)

Minimal acute inflammatory reaction. Do not adhere to tissue. Resistant against biocontamination. Efficient as a pull-out suture.

52

Advanced Technologies and Polymer Materials for Surgical Sutures

when applied, as the PROLENETM and the SURGICAL STAINLESSSTEEL a metallic monofilament or twisted multifilament. In the case of PROLENETM, knotting is more difficult than others as nylon (this was common suture material). While, application of SURGICAL STAINLESS-STEEL can cause a possible cutting, pulling, or fragmentation at the surrounding skin. The composition, use, and advantages of these sutures are shown in Table 3.3. Finally, materials not indicated in specific cases, for example, NUROLONTM may be used where permanent retention of tensile strength is required, being PANACRYLTM contraindicated in ophthalmic, cardiovascular, or neurological tissue. These sutures are braided threads and Table 3.3 Some suture materials with contraindications. Trade name/ composition

Use

Advantages

CATGUT CHROMIC/natural proteins, mostly collagen MONOCRYL/ copolymer glycolide and ε-caprolactone PDS II/polyester poly (p-dioxanone)

Ophthalmic and soft tissue.a

Totally absorbed by enzymes.

Subcuticular and soft tissue.b

Absorption predictable. Virtually inert.

Cardiovascular, orthopedic, gynecologic, ophthalmic, plastic, digestive, and colonic.

VICRYL (polyglactin 910)/copolymer of lactide and glycolide

Ophthalmic.

PROLENE/ polypropylene

Cardiovascular, plastic, and orthopedic.

SURGICAL STAINLESS-STEEL/ 316L stainless steel

Implants, prostheses, abdominal wall, sternum, skin, orthopedic, and neurosurgery.

Soft and pliable. Low microorganism growth. Tensile strength for up to 6 weeks. Wound support provided for at least 14 days. Lower tissue reaction. Do not adhere to tissue. Biologically inert. Efficient as a pull-out suture. Nontoxic. Low tissue reactivity. Good flexibility and tensile strength. Fine wire size. Good tying and holding of knot.

a

Not indicated in cardiovascular and neurological surgery. Exemption in neural, cardiovascular, ophthalmic, and microsurgery.

b

Improving the therapeutic value of sutures

53

Table 3.4 Materials not indicated in specific cases. Trade name/composition

Use

Advantages

NUROLON/nylon 6 or nylon 6,6

Soft tissue, cardiovascular, ophthalmic, and neurological.

PANACRYL/copolymer lactide and glycolide, coating with copolymer caprolactone and glycolide

Soft tissue, orthopedic, tendon, ligament, and reattachment to bone.

Good feel and handle. High tensile strength. Minimal tissue reaction. Extended wound support. Retains 80% tensile strength at 90 days of postimplantation.

synthetics, but NUROLONTM is nonabsorbable, while PANACRYLTM is absorbable; Table 3.4 shows the use and advantages of these materials. In addition to the exposed criteria, an increasing number of factors can be included that will lead to making a better decision of the suture procedure to choose, such as the type of sutures (according to the analysis of all necessary material, such as the classification of the needles needed and the practical aspects of its use), anesthetics (their effectiveness, administration, and need) other types of sutures (such as approach points and adhesives), the increase and improvement of suture techniques, taking into account the importance of the asepsis (defined as the set of procedures that aim to prevent the penetration of germs in the site that does not contain them). The point is to highlight that the development of this topic is still booming, so continuous knowledge of these is necessary. This manuscript makes the disclaimer about the functionality of a type of suture if there is not an expert diagnostic. Also, it is necessary to verify the current list of approved suture materials in each country, doing special emphasis on the current availability of materials. It is attempted only to provide information on the characteristics of the most used sutures.

3.3 Suture modification: bioactive devices as the future of the suture technology Until now, sutures have been used as passive agents during the wound healing process. Although currently sutures are made of inert and biocompatible materials and used in aseptic environments, the fact of implanting an external material in the body generates an immune response,

54

Advanced Technologies and Polymer Materials for Surgical Sutures

which alters the cellular function and affects the healing process, increasing inflammation and pain [23]. Furthermore, these devices tend to be easily colonized by bacteria, increasing the risk of infection. Surgical site infections (SSI) are one of the most frequent nosocomial infections, around 2%e4% of all surgeries present SSI [24], the risk of them occurring varies from 0.6% in outpatient and lightly invasive surgeries to 20% in highly invasive surgeries and preexisting risk factors such as diabetes, obesity, age, and poor nutrition [25,26]. Due to this, recent researches have sought to increase the therapeutic value of sutures by developing bioactive sutures, which allow overcoming current difficulties, improving the interaction of the suture with tissue, and incorporating properties that help to accelerate the healing process, decrease the risk of infection, or relieve inflammation and pain [27]. Improving the features of sutures can be realized in structural or chemical modifications. The type of modification will depend on the features of tissue involved, depth of the wound, and expected result. 3.3.1 Structural modification The successful implant acceptance depends largely on the structural characteristics of implanted material since it constitutes the support for cell growth and organization. An ideal bioactive suture should simulate the biological environment of the tissue that will be regenerated [28]. In humans, as in all mammals, cells are embedded in the extracellular matrix (ECM) that provides support for cell communication, organization, and proliferation. ECM is formed by a network of reticulum proteins, salts, chemokines, and growth factors. These proteins are generally fibrous (types of collagen) and on tissue, they are found forming filaments of 25e500 nm in diameter and lengths in the order of micrometers, which make up threedimensional structures unique to each tissue, which are characterized by their porosity, tortuosity, pore size, and interconnectivity [29]. 3.3.1.1 Fiber dimensions Recent studies seek the production of sutures whose fiber structure simulates the structure of the ECM to increase the affinity between the suture and the tissue since the fiber structure affects the body’s response to the implanted suture. A better interaction could allow the suture to act as temporary structural support for the regeneration of damaged tissue. A first approximation to structures with the dimensions of the ECM is polymeric fibers with diameters less than 1000 nm, also called the submicron scale [23]. These materials show a surface area to volume ratio similar to the ECM filaments,

Improving the therapeutic value of sutures

55

so they could be applicable in biomedical instruments, especially for the manufacture of biomimetic sutures that help to reconstruct tissues with high percentages of ECM, such as tendons or ligaments [30]. Currently, the submicron dimensions fibers are produced by the electrospinning method. These fibers show homogeneity and high mechanical resistance. In the electrospinning technique, a polymeric solution is driven through a dispensing needle that is in front of a cylindrical metallic collector for fibers, then an electric field is established between the needle and collector by the application of a potential difference. When the electric field allows overcoming the surface tension of the polymeric solution, the drop at the end of the needle is distorted, which generates a conical geometry, known as the Taylor cone, and expels an electrostatically charged nanometric flow towards the collector that generates of the fibers [31]. He et al. used this method to manufacture submicron fibers of poly(L-lactic acid) (PLA) loaded with tetracycline hydrochloride to manufacture bioactive sutures, which had a greater affinity for tissues while providing antimicrobial protection [32]. Also, recently supercritical CO2 was used to impregnate ketoprofen in PLA suture threads; the loaded sutures released ketoprofen prolongedly for up to 3 months [33]. 3.3.1.2 Topography and microstructure of the suture The topography (roughness) and the microstructure of the suture are other important aspects to achieving biomimetic sutures that stimulate tissue regeneration. Roughness is a measure of the surface alterations or relief of the material, which directly influences cell adhesion and proliferation [34]. Three roughness scales can be distinguished: macro (millimeters to 100 mm), micro (100e100 nm), and nano (less than 100 nm); the cellular response produced by each will depend on the type of cell, for example, large cells (osteoblasts, neurons) have a greater affinity for macro-roughness, while smaller cells such as endothelial and epithelial cells respond better to nano-roughened materials [35,36]. On the other hand, the microstructure comprises porosity and pore interconnectivity. The biomaterial must strike a balance between high mechanical resistance and a microstructure that allows cell permeation and the formation of tissue, which replaces the material as it degrades [37]. The appropriate microstructure depends on the tissue with it will interact, for example, each cell type has a greater affinity for certain pore size, human skin fibroblasts have shown accelerated growth in PLLA and poly (L-lacticco-glycolic acid) (PLLA/PLGA) scaffolds with pores of approximately

56

Advanced Technologies and Polymer Materials for Surgical Sutures

160 mm [38], while endothelial cells have grown favorably in silicon nitride supports with a pore size of less than 80 mm [39]. Currently, this area is seeking the development of very porous biomaterials with a size pore suitable to the cell or tissue to allow proper adherence. A small pore size would limit cell migration, while a too-large size pore hinders the strong interactions between cell and material decreasing the mechanical stability of the system. Besides, high interconnectivity allows a better distribution of nutrients and metabolites to regulate the local pH and to improve the cell growth in the interior of the material [34]. 3.3.2 Chemical modification The molecular conformation in the surface of sutures determines its physicochemical characteristics such as hydrophilicity, biocompatibility, biodegradability, biological activity, and chemical resistance. Good chemical composition is essential to achieve bioactive sutures that allow a constant interaction with the biological medium, exercising a function that stimulates the healing of a wound or relieves pain. Different active agents can be incorporated into the main suture structure, either superficially or as a coating. The suture design must balance good biological functionality with the mechanical integrity necessary to guarantee the stability of surgical stitches until the wound heals. Below are presented examples of modified sutures endowed with antimicrobial properties through drug delivery systems. 3.3.2.1 Antimicrobial sutures SSI are generally caused by bacterial colonization of the suture surface, these are the most frequent postoperative complications estimated to occur in 2%e4% of procedures, increasing morbidity, treatment costs, and mortality [26]. The bacterial attack occurs with the formation of biofilms, which are usually treated with antibiotics by systemic administration, which leads to an inefficient supply of the drug and generates side effects as the appearance of antibiotic-resistant strains [40]. Sutures with antimicrobial properties aim to decrease these complications, without affecting the physical and chemical characteristics require for the function of the material. Two ways of adding this antimicrobial characteristic to sutures have been studied, the incorporation of an antimicrobial agent within the structure of the material and incorporation in the surface of the material, which can be by coating with polymer, nanoparticles (NPs), proteins, or antimicrobial agents. Fig. 3.3 shows a schematic representation of these modifications.

Improving the therapeutic value of sutures

57

Figure 3.3 Schematic representation of the ways to incorporate an antimicrobial agent. (Credit: Original Figure.)

3.3.2.2 Incorporation of antimicrobial agents in sutures Different natural and synthetic agents, which have antimicrobial activity, can be incorporated into the structure of commonly used polymers for sutures manufacture [41]. At the molecular level, antibiotics, and antibacterial molecules such as quaternary amines have been incorporated in the polymeric structure [42]. For example, López-Saucedo et al. modified polypropylene sutures with a binary graft of N-isopropyl acrylamide and Nvinyl imidazole, using ionizing radiation. Grafted sutures were later functionalized with methyl iodide to form quaternary amines that provide antimicrobial activity to suture, which was verified against Gram-negative and Gram-positive bacteria [43]. The main disadvantage of incorporating antibiotics is that their activity can be affected during the manufacturing processes, so a chemical modification of the suture is preferred that allows the antibiotic to interact later as a drug loading and release system. Another alternative is the use of these antimicrobial materials as raw materials for suture fabrication when physical characteristics of the material meet the application requirements. Many polymers with antibiotic characteristics allow the formation of biocompatible and mechanically resistant sutures. For example, chitin, a natural polysaccharide of N-acetyl glucosamine, and its derivatives have been shown to promote healing and decrease the risk of infection when they are used in wound treatments [44].

58

Advanced Technologies and Polymer Materials for Surgical Sutures

Shao et al. developed a bioabsorbable multifilament diacetyl chitin suture, which showed viability for wound healing in vivo tests on rats and mechanical properties comparable to currently used polymeric sutures. These sutures presented an increase in the regeneration rate of injured tissue, mechanical integrity up to 63% for 14 days (no rupture), and complete degradation 42 days after implantation [45]. Zwitterionic polymers too have antimicrobial features and can be used to manufacture sutures, these are neutral polymers that contain anionic and cationic groups in their structures and have been studied for their high biocompatibility and their antimicrobial and antifouling properties. In 2017, Chen et al. reported the synthesis of a copolymer of 3-dimethyl (methacyloyloxyethyl) ammonium propane sulfonate (DMAPS) and acrylic acid (AAc), (DMAPS-co-AAc) to form a zwitterionic biomaterial that showed antimicrobial properties with a minimal inhibitory concentration for Escherichia coli of 850 mg/mL DMAPS and shape memory [46]. 3.3.2.3 Surface incorporation of antimicrobial agents in sutures The incorporation of an antibacterial agent on suture surface to endow them with antibacterial capacity can be by coating with inorganic or organic materials. Among the coatings that have been studied, the coatings with bioactive glasses (BGs), silver nanoparticles (Ag NPs), and proteins show promising results to avoid infections and increase the rate of tissue regeneration. BGs are biocompatible ceramics that have a strong interaction with living tissues, which stimulates damaged tissue regeneration, while also provides antibacterial protection [47]. BGs are mostly composed of a mixture of silicates, borates, and phosphates that degrades in biological conditions without repercussions. BGs tend to be fragile and of low mechanical resistance, due to which are usually used as coatings for different implants [48]. Bretcanu et al. coated PLGA sutures with a bioactive inorganic glass of calcium and sodium phosphosilicates (45% SiO2, 24.5% Na2O, 24.5% CaO, and 6% P2O5), known as Bioglass 45S5, by the slurry dipping method. The coated sutures showed homogeneous surfaces with little decrease in mechanical resistance and biocompatibility increase [49]. Besides, it was found that the strong interaction between BA coated sutures and tissues is related to the formation of hydroxyapatite crystals on the suture surface, by contact with biological fluids [50]. AgNPs are widely used as coatings in different medical tools [51,52]. AgNPs are clusters of silver atoms, with sizes between 1 and 100 nm, which

Improving the therapeutic value of sutures

59

are commonly obtained by chemical or light-induced reduction of silver nitrate, and have high antibacterial activity, antiinflammatory effects, and stimulate tissue regeneration [53]. Although AgNPs show low toxicity in humans, their concentration and size should be controlled to avoid harmful effects [54]. The bactericidal action mechanism of AgNPs is not fully elucidated, but the mechanisms of contact death and ion-mediated death are the most accepted [55]. The first mechanism is based on their physical and chemical properties, AgNPs can interact with the plasmatic membrane, anchoring themselves and producing physical alterations that damage the cellular equilibrium. Furthermore, when AgNPs enter the cell, interact with cellular structures such as ribosomes, inhibiting the expression of some proteins, or with biomolecules such as deoxyribonucleic acid (DNA), producing alterations. In any case, this interaction leads to cell death [53,55]. The second mechanism involves the union of the silver cations (Agþ) released from the AgNPs with the sulfhydryl groups present in the different proteins, generating alterations in their functionality. Furthermore, Agþ can inhibit cell growth by forming nucleic complexes with DNA and increasing the oxidative stress of the cell [56,57]. Due to the expression of multiple mechanisms of action, which can become synergistic, the development of bacterial resistance to the action of AgNPs is complicated and rarely occurs. Different studies have been carried out to incorporate AgNPs to polymeric sutures to endow them with antimicrobial activity, for example, AgNPs biologically synthesized (using Streptomyces sp.) were deposited in nonabsorbable silk sutures [58]; AgNPs sutures exhibited activity against Candida albicans, Escherichia coli, and Staphylococcus aureus in vitro studies. Furthermore, a high affinity was observed between AgNPs and the suture and it was determined that under the conditions of the experiment, AgNPs do not affect cell viability [59]. Also, hybrid systems have been attempted, Ciraldo et al. covered polyacid glycolic sutures with two layers, using the dip-coating method, the first was a polymeric chitosan layer and the second a silver-doped bioactive mesoporous glass (Ag-MBC). The coated suture showed antimicrobial activity against Gram-positive and Gram-negative bacteria, which was attributed to the synergistic action of Ag with the superficial hydroxyapatite crystals formed on contact with the simulated body fluid solution (SBF) [60]. Coatings of proteins modified with antimicrobial peptides have been studied to improve the antimicrobial properties and biocompatibility of polymer sutures, as an alternative to decreasing SSI, without the use of

60

Advanced Technologies and Polymer Materials for Surgical Sutures

antibiotic, which prevents the formation of drug-resistant strains. For example, coated a commercial silk suture with spider silk modified with 6mer-HNP1 antimicrobial domains, HNP1 is a broad-spectrum antimicrobial peptide from the alpha defensin family, which is widely found in mammals and promotes the development of the immune response [61]; through dip-coating method. Coated sutures showed low cytotoxicity on human fetal lung fibroblasts and good hemocompatibility in hemolysis tests, without affecting the mechanical properties of the original suture. Besides, they achieved significant inhibition of proliferation and biofilm formation against methicillin-resistant Staphylococcus aureus and Escherichia coli with in vitro tests [62]. 3.3.3 Drug delivery sutures Drug delivery systems permit greater effectiveness to the drug, because they can provide a release with temporal control, local control, or a combination of both. Controlled release over time allows the administration of a constant dose of the drug without exceeding the therapeutic maximums for prolonged periods, it is especially useful with drugs that are metabolized rapidly or in treatments that require constant concentrations over time. On the other hand, localized release allows the release in specific sites, which reduces the side effects associated with systemic administration, also maintaining a high concentration of drug at the target site [63]. Sutures are presented as a good alternative for drug delivery since they can be used as a localized delivery system due to the way of application, which can be easily combined with a temporal control system. The type of system and the method of delivery will depend on the suture raw material, the application site and the drug to be released, which may be an antibiotic, an antiinflammatory, or an analgesic, among others, each one of which gives particulars properties to the suture [64,65]. Polymers commonly used for the manufacture of sutures have difficulties to store hydrophilic drugs due to their hydrophobic nature and have slow rates of degradation, which affects the timely delivery of the drug, therefore, different modified systems are sought to overcome these disadvantages. Sutures may be coated with drugs by direct or indirect methods. Dip methods are direct methods, in which the sutures are immersed in a drug solution for a determinate time and later are dried, loaded drug by this process is generally released by diffusion, providing a localized release. In indirect methods, the drug is loaded into a different matrix before coating

Improving the therapeutic value of sutures

61

the suture as coatings by electrospinning, or the suture can be superficially modified to have a higher affinity for the active agent, such as coating by graft; these coatings provide more control over drug release [66]. Several drugs, including antibiotics and antiinflammatories, have been deposited on the surface of sutures; the best example is triclosan-coated PLGA sutures, which showed characteristics so favorable than were approved in 2002 by the USA Food and Drug Administration as the first antimicrobial suture for use in humans [41]. Triclosan is a broad-spectrum bactericide and fungicide, used in many grooming and medical products, due to its effectiveness as antibacterial and low toxicity. This is a polychlorinated compound (Fig. 3.4) slightly soluble in water and with a good affinity for polymers with functional groups in its structure [67], its action mechanism is given by permeation in the bacterial cell membrane, once in the cytoplasm, it affects different functions, including RNA synthesis and lipid synthesis [68]. Zurita et al., coated the polyglycolic acid sutures with a copolymer (10:60:30) of lactide, ε-caprolactone, and trimethylene carbonate loaded with triclosan, by slurry dip method. The coating copolymer showed miscibility with triclosan, loading around 7000 mg g1 when placed in 3 mL of 1% (w/v) solution, the triclosan release profiles of the coated sutures showed first-order behavior with 100% release in a lipophilic nutrient medium at 37 C after 24 h, in this case, the release mechanism was diffusion, since the total release of the drug took place before the degradation of the matrix occurred [69]. Other drugs have been studied as suture coatings by dip method, for example, chlorhexidine that is a broad-spectrum antiseptic, which can easily bind to the cell membrane, destabilizing it and leading to bacteria death due

Figure 3.4 Dip-coated sutures to drug delivery. (Credit: Original Figure.)

62

Advanced Technologies and Polymer Materials for Surgical Sutures

to it is positively charged, was impregnated in sutures. The drug was diluted in palmitic acid as a carrier, achieving a significant antibacterial effect in vitro for 96 h on Staphylococcus aureus strains, besides little cytotoxic effect was observed at chlorhexidine concentrations less than 11 mg/cm [70]. Fatty acid as a carrier provides a lubricating effect, in addition to allowing a drug homogeneous concentration and reducing the release rate since its solubility in aqueous media is low [71]. Levofloxacin, another broadspectrum antibiotic, from the quinolone group (Fig. 3.4), was used dissolved in poly (ε-caprolactone) to coat braided silk sutures; the coating process was performed by immersion of the silk fibers into a levofloxacin solution before braiding for suture formation. The coated sutures showed significant bacterial inhibition against Staphylococcus aureus and Escherichia coli, with a continuous drug release for 5 days [72]. The release profile was adjusted to the Higuchi model, indicating that the predominant release mechanism in the system is diffusion. These sutures also comply with the mechanical specifications of the U.S. Pharmacopoeia and have acceptable cytotoxicity according to ISO 10993-5. As mentioned above, electrospinning is a technique that allows the manufacture of polymeric fibers of submicron sizes that can form sheets or threads [73]. Electrospinning coatings involve wrapping the suture in an electrospun sheet. In this way, the coating does not damage the original structure of suture, so which can keep its mechanical properties. Lee et al. covered the commercial suture VICRYLTM with an electrospun sheet of PLGA loaded with ibuprofen, which is a nonsteroidal antiinflammatory used to relieving pain in the treatment of wounds. The ibuprofen-loaded sheet of around 29 mm thickness was braided around the suture and cured at 47 C, a temperature higher than the glass transition temperature of PLGA, but lower than the decomposition temperature of ibuprofen, to achieve a better grip of the coating; the release profiles showed drug eluding for 4 days, coated sutures had pain relief similar to oral treatment but using lower doses, in tests with rats by the animal pain-induced model. And presented good biocompatibility that is associated with the polymer used as the base of the coating [74]. The administration of antiinflammatories and analgesics in a localized way allows a faster and less traumatic recovery [75]. Grafting modification involves the formation of branches on the main polymer chain with a different composition polymer, through which it is intended to give a specific characteristic to the material such as increasing its mechanical resistance, providing hydrophilicity or stimuli response, or incorporating functional groups that allow specific molecule anchoring [76].

Improving the therapeutic value of sutures

63

Grafting formation may be carried out by chemical methods or induced by ionizing radiation. In the case of sutures, these modifications allow them to alter their characteristics to provide them with the ability to load and release drugs. In this context, polypropylene sutures were modified with 5% acrylonitrile graft by the oxidative preirradiation method using ionizing radiation; acrylonitrile was hydrolyzed to form carboxyl groups that allowed the immobilization of tetracycline, a broad-spectrum antibiotic, Fig. 3.5 shows schematic representation. Release profiles displayed continuous drug elution for 4 days, and in vitro tests proved that sutures inhibited Escherichia coli, Klebsiella pneumonia, and Staphylococcus aureus. Also, in vivo assays in albino mice proved that modified sutures can avoid infection [77]. Grafting polymer also can be incorporated to suture as raw material to bind an active agent, for example, sutures were manufactured by electrospinning method with PLGA, poly(ethylene glycol) and poly(ethyleneimine), to form a copolymer, this grafting was used to load heparin, an anticoagulant commonly used that in sutures can act as a microvascular antithrombotic. This PLGA copolymer has negative charge, which allowed the electrostatic binding to heparin and a longer release. Local delivery of heparin avoids the formation of thrombus and the occlusions during the anastomosis of blood vessels after surgery, which is a common complication in these procedures. Tests in vitro of heparin release were controlled for 20 days with a total release of 68%, while control sutures released 60% in 4 h. These results were consistent with longer clotting times using the graft suture [78].

Figure 3.5 Grafting modification of polypropylene sutures with acrylonitrile to load tetracycline. (Credit: Original Figure.)

64

Advanced Technologies and Polymer Materials for Surgical Sutures

The graft of stimuli sensitive polymers has become a good alternative to drug delivery systems, because of these materials present a conformational or electronic change when are exposed to external stimuli, such as temperature, pH, ionic strength, magnetic field, or light, which can trigger the drug release process in a controlled way. So, for example, polypropylene sutures were modified through grafting of AAc by ionizing radiation (gamma rays), to load and release vancomycin by reversible ionic interactions, taking advantage of the pH sensitivity of polyacid acrylic. The release profiles showed continuous elution for 2 h and in vitro tests displayed Staphylococcus aureus inhibition, which is promising for obtaining tools that decrease the risk of infection [79]. Biodegradable sutures are suitable for delivery systems controlled by matrix erosion, where drug release is determined by system degradation, which is a result of the breakdown of a weak interaction or a bond by hydrolysis or enzymatic action [80]. Fig. 3.6 shows an erosion release

Figure 3.6 Schematic representation of drug release by erosion of suture. (Credit: Original Figure.)

Improving the therapeutic value of sutures

65

scheme, where the drug begins to release as the suture degrades, continuing to release even after the mechanical stability of the suture is affected. In these systems, the rate of release is directly related to the erosion of matrix. To develop erosion-controlled drug delivery systems, the biodegradable materials must be coupled to the application, to be biocompatible, with low cytotoxicity, highly resistance, and with affinity toward drug release; their main disadvantage is that adding the drug to the matrix decreases its mechanical properties. Silva et al. manufactured an erosion-controlled release suture with a combined system of chitosan and N-acetyl-D-glucosamine (GlcNAc), using the wet spinning method. GlcNAc is a drug that stimulates tissue formation and relieves pain. In vitro release tests showed a gradual release with a maximum of 60% after 30 days, being consistent with tests of the degradation of the suture, which occurred around 30 days in buffer solution [81]. Although drug release was controlled, the GlcNAc suture showed a decrease in its mechanical resistance compared to the chitosan suture, possibly due to the GlcNAc interference in the intermolecular interactions of chitosan. 3.3.4 Stimuli responsive systems These materials have the particularity of response at external stimuli, which could be chemical or physics, with a change in their conformation as a result of a variation in the interaction with the environment. This feature finds applications in different areas of science and technology, including the development of biomedical devices, among them sutures [82]. Stimuliresponsive materials undergo physical or chemical changes in their properties due to small variations in the medium. These changes are characterized by being drastic at the structural level (external or internal), as well as changes in their photochemical, electromagnetic, or bioactive properties [83]. Before continuing with the description of stimuli-responsive materials, it is convenient to define “the swelling” since this property is the fundamental factor in the improvement of a surface. Swelling is a phenomenon manifested as the increase in the volume of a solid or semisolid by the adsorption of a liquid or a gas. When a polymer is in contact with a fluid phase, the fluid diffuses through the solid structure, solvating it from the outside to the inner. Then, the small molecules of the solvent are distributed around macromolecules causing the swelling and weakening intermolecular forces, as a result, the material becomes softer and more ductile.

66

Advanced Technologies and Polymer Materials for Surgical Sutures

Swelling is considered as a partial dissolution process, being the solubility limited by the inherent characteristics of the polymer-solvent system. The swelling capacity depends on the system, the hydrogels, for example, are polymers with high swelling capability, which hydrophilicity brings them soft and elastic features in water solution, besides increasing their surface and weight, while in the dehydrated state, they keep a structure typical of solids and loss the characteristics mentioned. For these characteristics, hydrogels are widely used for sanitary, care, and cosmetic products [84,85]. About the swelling and behavior of stimuli-responsive materials, the macromolecules that make up the stimuli-responsive polymer are solubilized and pass from a monophasic to a biphasic state near the transition point, giving to the formation of reversible sol-gel hydrogels. At this transition point, the smart grid networks undergo a chain reorganization, so the network goes from a collapsed to a reversibly swelled dynamic state. Stimuli-responsive surfaces change their hydrophilic behavior, commonly depending on the sensitive interface provided by stimuli such as pH or temperature [86,87]. 3.3.4.1 pH-responsive polymers These types of materials can be defined as polyelectrolytes that contain acid or basic groups in their structure, able to accept or release protons in response to a change in pH. The ratios of charges and concentration of ions Hþ or OH in the medium yields electrostatic repulsion/attraction forces cause an increase or decrease in the volume of the polymer. This transition between the swelled and collapsed state is influenced by factors in the medium such as pH, ionic strength, and type of counterion. Usually, “critical pH” is the name assigned to the point where this transition takes place. The transition from the swelled to the collapsed state is explained by changes in the osmotic pressure of counterions that neutralize the network charges [88]. Fig. 3.7 shows a representation of a pH-responsive material with acid groups that above their critical pH are ionized as anions. Polymers with pH-responsive contain weak acid groups (such as carboxylic acids) or basic groups (such as amines). For example, AAc has a dissociation constant (pKa) equal to 4.25 [89], then above this pH, the carboxylic groups are ionized. This leads to electrostatic repulsion between the chains that can interact with water to cause a high degree of swelling at low pH. The shift in critical pH will depend on the type of chains containing donor- or electro-attractor groups to achieve the transition from contraction to swelling, for example, the length of the carbon chains in an

Improving the therapeutic value of sutures

67

Figure 3.7 Schematic representation of pH responsive material. (Credit: Original Figure.)

aliphatic chain of ammonium salts is capable of modifying critical pH of organic acid groups, where the hydrophobic interactions of their aliphatic chains must be broke to allow solvation with water molecules [90]. 3.3.4.2 Thermo-responsive polymers Thermo-responsive polymers display a critical temperature (CT) in dissolution, which corresponds to a reversible behavior from hydrophilic to hydrophobic. These polymers are characterized by a coil-to-globule transition, an illustration of this phenomenon is shown in Fig. 3.8 In the thermo-responsive polymer chains, the amphiphilic balance is required, so that, the hydrophilic part is responsible for interacting with the solvent by intermolecular interactions as hydrogen bond, dipole-dipole, or ion-dipole allowing the open coil structure, while the hydrophobic allows the auto association of the chains for the favoring of globular phase [91]. The

Figure 3.8 Schematic representation of coil-to-globule transition in thermoresponsive materials, dependent on critical temperature (CT). (Credit: Original Figure.)

68

Advanced Technologies and Polymer Materials for Surgical Sutures

addition of surfactants significantly affects these systems because their amphiphilic nature cause alterations in the thermo-response, even nullifying the transition [85]. Two types of sensitivity may be displayed, polymers with Lower Critical Solution Temperature (LCST) and polymers with Upper Critical Solution Temperature (UCST). Polymers with LCST are hydrophilic at temperatures under the CT, the hydrophilic behavior decreases significantly after this temperature. That is, the transition of linear polymer from an open coil to a globular form occurs. The transition corresponds to the region in which the enthalpy contribution of the intermolecular water-polymer bonds is lower than the entropy of the system. The LCST only depends on the concentration (very dilute or very concentrated) and molecular weight of the polymers; under standard conditions, it is attributed to the hydrophobic/hydrophilic balance, when a copolymer with a specific composition is reformulated, this balance can change as well as the LCST. The binding with monomers, substituents, or hydrophilic molecules in general increases the LCST, while the binding of hydrophobic species decreases it. Materials such as poly(N-isopropylacrylamide) (PNIPAAm), poly(diethylamine) (PDEAM), poly(methyl vinyl ether) (PMVE), and poly(N-vinylcaprolactam) (PNVCL) are widely used in biomedical applications since their LCST is close to body temperature [92]. Polymers with UCST are poorly swelling or insoluble in water at temperatures below the UCST, but heating above this temperature, the solubility and affinity are enhanced. This is because the polymerepolymer and solventesolvent interactions are weakened. Then, the polymer chains display the transition from a globular structure to an open coil. These materials seldom find applications as biomedical devices since their UCST is normally at temperatures around 50 C, but attempts are made to reach the application range [93]. Poly(methacrylic acid) (PMAAc) and poly(acrylic acid) (PAAc) are examples of polymers with this behavior [94]. 3.3.4.3 Stimuli responsive sutures In the case of sutures, the relevant are the changes concerning their surface because the zone of surgery is prone to infection due to a weakened immune system, where the contact between the sutured tissue and the external environment is exposed to several attacks by possible exogenous surrounding pathogens [95]. The said biocontamination may be treated either controlling, stopping growth, inhibiting, or eradicating them. One way to achieve an aseptic surface is by a stimuli-responsive polymer with

Improving the therapeutic value of sutures

69

adequate characteristics. These materials are useful to endow sutures with pH- or thermo-responsive features, which can help with phenomena such as the load and release of agent actives that may be harnessed to administrate molecules of interest, one example more detailed of this application is shown in the section of drug delivery systems [41]. The stimuli-responsive materials of polymer origin have the greatest potential for applications in sutures since they can swell or collapse their chains in the presence of a solvent. In this way, the behavior improved surfaces with stimuli-responsive polymers can be manipulated at convenience in a specific fluid medium. Furthermore, some stimulus-responsive polymers also have shapememory characteristics, so they can go from a deformed state to their permanent or original state when subjected to an external stimulus such as heat, light, electrical or magnetic force. This very striking property allows the manufacture of self-adjusting sutures, which would facilitate knotting processes in small spaces, reducing the risk of additional tissue damage, due to which, self-adjusting sutures have high applicability in cardiovascular, ophthalmological, and obstetric surgeries [96]. Jing et al. developed selfadjusting sutures with a mixture (1:3) of thermoplastic polyurethane and poly(ε-caprolactone), two polymers that have thermally induced shapememory. These sutures showed a 90% shape-recovery at temperatures around 35 C, with a deformation temperature of 61 C [97]. A behavior attributed to the balance between amorphous and crystalline segments in the polymeric structures. Also, the material displayed good biocompatibility. For the use of this type of sutures, they are initially heated over their deformation temperature, later stretched them to their temporary shape, and are placed on the wounds causing minimal stress. Finally, when the stimulus is removed, the suture recovers its original shape adjusting the tissues involved.

3.4 Conclusion The development of bioactive sutures, which play a dynamic role in wound healing processes, represents the future of this type of medical device. Sutures are in direct contact with injured tissues, which makes them an ideal alternative to influence cellular regeneration. Advances in materials science have allowed the manufacture of sutures with a similar structure to the biological environment that can act as supports for cell growth while keeping the wound sealed. Besides, due to the possibility of using chemically modified materials for sutures manufacture, new developers involve

70

Advanced Technologies and Polymer Materials for Surgical Sutures

sutures with antibacterial properties, sutures that can load and release drug, and selfadjusting sutures. These are an alternative to decreasing the SSI index and the risks of operative complications, as well as to relieve pain and increase the speed of tissue regeneration.

Acknowledgments This work was supported by Dirección General de Asuntos del Personal Académico, Universidad Nacional Autónoma de México under Grant IN202320.

References [1] A. Vargas, M. López, C. Lillo, M.J. Vargas, El papiro de Edwin Smith y su trascendencia médica y odontológica, Rev. Med. Chile 140 (2012) 1357e1362, https:// doi.org/10.4067/S0034-98872012001000020. [2] A. Hallmann-Mikołajczak, Ebers Papyrus. The book of medical knowledge of the 16th century B.C. Egyptians, Arch. Hist. Filoz Med. 67 (2004) 5e14. [3] J.E. González, Z.V. Camejo, Esclepios (Esculapio) y su familia en la mitología y la medicina Occidental, Salus 19 (2015) 26e32. [4] J.E. González, Z. Camejo, Mitología y medicina I: dioses griegos primigenios de la medicina Occidental, Salus 18 (2014) 33e40. [5] B. Deepa, K. Pushpalatha, Contributions of sushruta to anatomy, Anat. J. Africa 3 (2014) 362e365. [6] J.L. Cameron, William stewart halsted. Our surgical heritage, Ann. Surg. 225 (1997) 445e458, https://doi.org/10.1097/00000658-199705000-00002. [7] S. Stucchi-Portocarrero, La primera guerra mundial y su impacto en la psiquiatría, Rev. Neuro Psiquiatr. 77 (2014) 139e143. [8] H. Hopf, Serendipity: accidental discoveries in science, Angew. Chem. 102 (1990) 1210e1211, https://doi.org/10.1002/ange.19901021038. [9] D.M. Adelman, K.G. Cornwell, Fundamentals of extracellular matrix biomaterial assimilation, Plast. Reconstr. Surg. Glob. Open 8 (2020) e2635, https://doi.org/ 10.1097/GOX.0000000000002635. [10] Johnson, E. Johnson, Ethicon Wound Closure Manual [WWW Document], 2005. Available from: https://www.sutureonline.com/wound-closure-manual (Accessed 14 May 2020). [11] A. Vargas, C. Foncea, A. Campolo, Adhesivos de cianoacrilato en cirugía oral y maxilofacial, Rev. clínica periodoncia, Implantol. y Rehabil. oral 10 (2017) 107e110, https://doi.org/10.4067/S0719-01072017000200107. [12] T. Valls Durán, E. Díaz Sanisidro, L. Nadal González, Uso del gel LAT para suturar heridas en niños, Pediatría Atención Primaria 11 (2009) 575e595. [13] L. Cuevas-López, C.M. Echeverri-Gómez, Diseño y evaluación de un tejido sintético para la simulación de nudos y suturas mediante laparoscopia, Rev. Colomb. Cirugía 33 (2018) 362e370, https://doi.org/10.30944/20117582.83. [14] R.J. Segura Iglesias, Instrumento de grapado quirúrgico vascular para realizar una sutura mecánica entre un vaso sanguíneo y una prótesis sintética, Angiologia 67 (2015) 266e272, https://doi.org/10.1016/j.angio.2014.12.002. [15] D. Gazivoda, D. Pelemis, G. Vujaskovic, A clinical study on the influence of suturing material on oral wound healing, Vojnosanit. Pregl. 72 (2015) 765e769, https:// doi.org/10.2298/VSP140401064G.

Improving the therapeutic value of sutures

71

[16] W. Channuan, J. Siripitayananon, R. Molloy, G.R. Mitchell, Defining the physical structure and properties in novel monofilaments with potential for use as absorbable surgical sutures based on a lactide containing block terpolymer, Polymer 49 (2008) 4433e4445, https://doi.org/10.1016/j.polymer.2008.07.058. [17] F. Dávila, D. Tsin, G. González, M.R. Dávila, J. Lemus, U. Dávila, Use of percutaneous needles in the feasibility of single-port laparoscopic cholecystectomy, Cirugía Española 92 (2014) 261e268, https://doi.org/10.1016/j.cireng.2013.01.011. [18] S. Matsuzaki, M. Jitsumori, T. Hara, S. Matsuzaki, S. Nakagawa, T. Miyake, T. Takiuchi, A. Kakigano, E. Kobayashi, T. Tomimatsu, T. Kimura, Systematic review on the needle and suture types for uterine compression sutures: a literature review, BMC Surg. 19 (2019) 196e201, https://doi.org/10.1186/s12893-019-0660-z. [19] A. van Oosterwyk, Atraumatic sutures can be made locally, Trop. Doct. 34 (2004) 95e96, https://doi.org/10.1177/004947550403400213. [20] V.P. Sergeev, E.P. Plygan’, E.A. Ivashchenko, N.I. Vagin, N.N. Baglei, Development of methods of manufacturing modified sutures corresponding to current requirements in medicine, A review. Fibre Chem. 34 (2002) 433e439, https://doi.org/10.1023/ A:1022964309414. [21] C.A. Erazo-Cortés, C. Gutiérrez-Gómez, M. Mucharraz-Diaz, Sutura metálica con aguja doble armada para reconstrucción auricular: experiencia en su fabricación y uso, Cirugía Plástica Ibero-Latinoam. 39 (2013) 293e297, https://doi.org/10.4321/S037678922013000300012. [22] J. Canales, C. Espinoza-Montes, M. Alarcón-Palacios, Material de suturas en periodoncia e implantes, Rev. Estomatol. Hered. 23 (2014) 148e153, https://doi.org/ 10.20453/reh.v23i3.26. [23] R.E. Abhari, J.A. Martins, H.L. Morris, P.-A. Mouthuy, A. Carr, Synthetic sutures: clinical evaluation and future developments, J. Biomater. Appl. 32 (2017) 410e421, https://doi.org/10.1177/0885328217720641. [24] S.I. Berríos-Torres, C.A. Umscheid, D.W. Bratzler, B. Leas, E.C. Stone, R.R. Kelz, C.E. Reinke, S. Morgan, J.S. Solomkin, J.E. Mazuski, E.P. Dellinger, K.M.F. Itani, E.F. Berbari, J. Segreti, J. Parvizi, J. Blanchard, G. Allen, J.A.J.W. Kluytmans, R. Donlan, W.P. Schecter, Centers for disease control and prevention guideline for the prevention of surgical site infection, JAMA Surg. 152 (2017) 784e791, https:// doi.org/10.1001/jamasurg.2017.0904. [25] S. Kathju, L. Nistico, L. Hall-Stoodley, J.C. Post, G.D. Ehrlich, P. Stoodley, Chronic surgical site infection due to suture-associated polymicrobial biofilm, Surg. Infect. 10 (2009) 457e461, https://doi.org/10.1089/sur.2008.062. [26] C.D. Owens, K. Stoessel, Surgical site infections: epidemiology , microbiology and prevention, J. Hosp. Infect. 70 (2008) 3e10, https://doi.org/10.1016/S0195-6701(08) 60017-1. [27] F. Alshomer, A. Madhavan, O. Pathan, W. Song, Bioactive sutures: a review of advances in surgical suture functionalisation, Curr. Med. Chem. 24 (2017) 215e223, https://doi.org/10.2174/0929867324666161118141724. [28] C.K.S. Pillai, C.P. Sharma, Review paper: absorbable polymeric surgical sutures: chemistry, production, properties, biodegradability, and performance, J. Biomater. Appl. 25 (2010) 291e366, https://doi.org/10.1177/0885328210384890. [29] C.T. Laurencin, L.S. Nair, Nanotechnology and tissue engineering: the scaffold, in: S. Challa, R. Kumar (Eds.), Nanotechnologies for the Life Sciences, Wiley-VCH Verlag, Weinheim, 2008, pp. 1e65. [30] S. Lyu, C. Huang, H. Yang, X. Zhang, Electrospun fibers as a scaffolding platform for bone tissue repair, J. Orthop. Res. 31 (2013) 1382e1389, https://doi.org/10.1002/ jor.22367.

72

Advanced Technologies and Polymer Materials for Surgical Sutures

[31] B. Ramesh, K.M. Cherian, A.O.J. Fakoya, Fabrication and electrospinning of 3D biodegradable poly-l-Lactic acid (PLLA) nanofibers for clinical application, Methods Mol. Biol. 2125 (2020) 119e128, https://doi.org/10.1007/7651_2019_213. [32] C. He, Z. Huang, X. Han, Fabrication of drug-loaded electrospun aligned fibrous threads for suture applications, J. Biomed. Mater. Res. 89A (2009) 80e95, https:// doi.org/10.1002/jbm.a.32004. [33] M. Champeau, I.T. Coutinho, J.-M. Thomassin, T. Tassaing, C. Jérôme, Tuning the release profile of ketoprofen from poly(l-lactic acid) suture using supercritical CO2 impregnation process, J. Drug Deliv. Sci. Technol. 55 (2020) 101468, https://doi.org/ 10.1016/j.jddst.2019.101468. [34] H.-I. Chang, Y. Wang, Cell responses to surface and architecture of tissue engineering scaffolds, in: D. Eberli (Ed.), Regenerative Medicine and Tissue Engineering e Cells and Biomaterials, InTech, London, 2011, pp. 569e588, https://doi.org/10.5772/ 21983. [35] T.W. Chung, D.Z. Liu, S.Y. Wang, S.S. Wang, Enhancement of the growth of human endothelial cells by surface roughness at nanometer scale, Biomaterials 24 (2003) 4655e4661, https://doi.org/10.1016/S0142-9612(03)00361-1. [36] F. Han, J. Wang, L. Ding, Y. Hu, W. Li, Z. Yuan, Q. Guo, C. Zhu, L. Yu, H. Wang, Z. Zhao, L. Jia, J. Li, Y. Yu, W. Zhang, G. Chu, S. Chen, B. Li, Tissue engineering and regenerative medicine: achievements, future, and sustainability in Asia, Front. Bioeng. Biotechnol. 8 (2020) 1e35, https://doi.org/10.3389/fbioe.2020.00083. [37] D. Rana, K. Ram, M. Leena, J. Constanza, J. Campos, P. Ibarra, M. Ramalingam, Surface functionalization of nanobiomaterials for application in stem cell culture, tissue engineering, and regenerative medicine, Biotechnol. Prog. 32 (2016) 554e567. [38] J. Yang, G. Shi, J. Bei, S. Wang, Y. Cao, Q. Shang, G. Yang, W. Wang, Fabrication and surface modification of macroporous poly(L-lactic acid) and poly(L-lactic-coglycolic acid) (70/30) cell scaffolds for human skin fibroblast cell culture, J. Biomed. Mater. Res. 62 (2002) 438e446, https://doi.org/10.1002/jbm.10318. [39] A.K. Salem, R. Stevens, R.G. Pearson, M.C. Davies, S.J.B. Tendler, C.J. Roberts, P.M. Williams, K.M. Shakesheff, Interactions of 3T3 fibroblasts and endothelial cells with defined pore features, J. Biomed. Mater. Res. 61 (2002) 212e217, https:// doi.org/10.1002/jbm.10195. [40] M.J. Henry-Stanley, D.J. Hess, A.M.T. Barnes, G.M. Dunny, C.L. Wells, Bacterial contamination of surgical suture resembles a biofilm, Surg. Infect. 11 (2010) 433e439, https://doi.org/10.1089/sur.2010.006. [41] C. Dennis, S. Sethu, S. Nayak, L. Mohan, Y. Morsi, G. Manivasagam, Suture materials d current and emerging trends, J. Biomed. Mater. Res. 104 (2016) 1544e1559, https://doi.org/10.1002/jbm.a.35683. [42] B. Hu, X. Chen, Y. Zuo, Z. Liu, X. Xing, Dual action bactericides: quaternary ammonium/N-halamine-functionalized cellulose fiber, J. Appl. Polym. Sci. 131 (2014) 1e7, https://doi.org/10.1002/app.40070. [43] F. López-Saucedo, G.G. Flores-Rojas, E. Bucio, C. Alvarez-Lorenzo, A. Concheiro, O. González-Antonio, Achieving antimicrobial activity through poly(Nmethylvinylimidazolium) iodide brushes on binary-grafted polypropylene suture threads, MRS Commun. 7 (2017) 938e946, https://doi.org/10.1557/mrc.2017.121. [44] C.K.S. Pillai, W. Paul, C.P. Sharma, Chitin and chitosan polymers: chemistry, solubility and fiber formation, Prog. Polym. Sci. 34 (2009) 641e678, https://doi.org/ 10.1016/j.progpolymsci.2009.04.001. [45] K. Shao, B. Han, J. Gao, Z. Jiang, W. Liu, W. Liu, Y. Liang, Fabrication and feasibility study of an absorbable diacetyl chitin surgical suture for wound healing, J. Biomed. Mater. Res. B Appl. Biomater. 104 (2016) 116e125, https://doi.org/10.1002/ jbm.b.33307.

Improving the therapeutic value of sutures

73

[46] S. Chen, Z. Mei, H. Ren, H. Zhuo, Integrated antimicrobial and shape memory zwitterionic copolymers, Mater. Lett. 189 (2017) 74e77, https://doi.org/10.1016/ j.matlet.2016.11.019. [47] S. Hu, J. Chang, M. Liu, C. Ning, Study on antibacterial effect of 45S5 bioglass, J. Mater. Sci. Mater. Med. 20 (2009) 281e286, https://doi.org/10.1007/s10856-0083564-5. [48] F. Baino, E. Verne, Glass-based coatings on biomedical implants: a state-of-the-art review, Biomed. Glas. 3 (2017) 1e17, https://doi.org/10.1515/bglass-2017-0001. [49] O. Bretcanu, E. Verné, L. Borello, A.R. Boccaccini, Bioactivity of degradable polymer sutures coated with bioactive glass, J. Mater. Sci. Mater. Med. 15 (2004) 893e899, https://doi.org/10.1023/B:JMSM.0000036277.99450.a2. [50] A.R. Boccaccini, A.G. Stamboulis, A. Rashid, J.A. Roether, Composite surgical sutures with bioactive glass coating, J. Biomed. Mater. Res. B Appl. Biomater. 67 (2003) 618e626, https://doi.org/10.1002/jbm.b.10047. [51] A.-C. Burdușel, O. Gherasim, A.M. Grumezescu, L. Mogoanta, A. Ficai, E. Andronescu, Biomedical applications of silver nanoparticles: an up-to-date overview, Nanomaterials 8 (2018) 681e705, https://doi.org/10.3390/nano8090681. [52] F. Furno, K.S. Morley, B. Wong, B.L. Sharp, P.L. Arnold, S.M. Howdle, R. Bayston, P.D. Brown, P.D. Winship, H.J. Reid, Silver nanoparticles and polymeric medical devices: a new approach to prevention of infection? J. Antimicrob. Chemother. 54 (2004) 1019e1024, https://doi.org/10.1093/jac/dkh478. [53] K. Chaloupka, Y. Malam, A.M. Seifalian, Nanosilver as a new generation of nanoproduct in biomedical applications, Trends Biotechnol. 28 (2010) 580e588, https:// doi.org/10.1016/j.tibtech.2010.07.006. [54] R. Vazquez-Muñoz, B. Borrego, K. Juárez-Moreno, M. García-García, J.D. Mota Morales, N. Bogdanchikova, A. Huerta-Saquero, Toxicity of silver nanoparticles in biological systems: does the complexity of biological systems matter? Toxicol. Lett. 276 (2017) 11e20, https://doi.org/10.1016/j.toxlet.2017.05.007. [55] Y. Qing, L. Cheng, R. Li, G. Liu, Y. Zhang, X. Tang, J. Wang, H. Liu, Y. Qin, Potential antibacterial mechanism of silver nanoparticles and the optimization of orthopedic implants by advanced modification technologies, Int. J. Nanomed. 13 (2018) 3311e3327, https://doi.org/10.2147/IJN.S165125. [56] L.Z. Flores-López, H. Espinoza-Gómez, R. Somanathan, Silver nanoparticles: electron transfer, reactive oxygen species, oxidative stress, beneficial and toxicological effects, Mini review. J. Appl. Toxicol. 39 (2019) 16e26, https://doi.org/10.1002/jat.3654. [57] B. Lee, M.J. Lee, S.J. Yun, K. Kim, I.-H. Choi, S. Park, Silver nanoparticles induce reactive oxygen species-mediated cell cycle delay and synergistic cytotoxicity with 3bromopyruvate in Candida albicans, but not in Saccharomyces cerevisiae, Int. J. Nanomed. 14 (2019) 4801e4816, https://doi.org/10.2147/IJN.S205736. [58] A.M. Othman, M.A. Elsayed, N.G. Al-Balakocy, M.M. Hassan, A.M. Elshafei, Biosynthesis and characterization of silver nanoparticles induced by fungal proteins and its application in different biological activities, J. Genet. Eng. Biotechnol. 17 (2019) 8, https://doi.org/10.1186/s43141-019-0008-1. [59] T. Baygar, N. Sarac, A. Ugur, I.R. Karaca, Antimicrobial characteristics and biocompatibility of the surgical sutures coated with biosynthesized silver nanoparticles, Bioorg. Chem. 86 (2019) 254e258, https://doi.org/10.1016/j.bioorg.2018.12.034. [60] F.E. Ciraldo, K. Schnepf, W.H. Goldmann, A.R. Boccaccini, Development and characterization of bioactive glass containing composite coatings with ion releasing function for antibiotic-free antibacterial surgical sutures, Materials 12 (2019) 423, https://doi.org/10.3390/ma12030423. [61] X. Zhang, A. Jiang, B. Qi, H. Yu, Y. Xiong, G. Zhou, M. Qin, J. Dou, J. Wang, Secretion expression of human neutrophil peptide 1 (HNP1) in Pichia pastoris and its

74

[62]

[63] [64] [65]

[66] [67] [68] [69] [70]

[71]

[72] [73] [74]

[75] [76]

Advanced Technologies and Polymer Materials for Surgical Sutures

functional analysis against antibiotic-resistant Helicobacter pylori, Appl. Microbiol. Biotechnol. 102 (2018) 4817e4827, https://doi.org/10.1007/s00253-018-8982-9. A.R. Franco, E.M. Fernandes, M.T. Rodrigues, F.J. Rodrigues, M.E. Gomes, I.B. Leonor, D.L. Kaplan, R.L. Reis, Antimicrobial coating of spider silk to prevent bacterial attachment on silk surgical sutures, Acta Biomater. 99 (2019) 236e246, https://doi.org/10.1016/j.actbio.2019.09.004. G. Tiwari, R. Tiwari, S. Bannerjee, L. Bhati, S. Pandey, P. Pandey, B. Sriwastawa, Drug delivery systems: an updated review, Int. J. Pharm. Investig. 2 (2012) 2e11, https://doi.org/10.4103/2230-973x.96920. A. Anand, P. Das, S.K. Nandi, B. Kundu, Development of antibiotic loaded mesoporous bioactive glass and its drug release kinetics, Ceram. Int. 46 (2020) 5477e5483, https://doi.org/10.1016/j.ceramint.2019.10.264. C. Yu, X. Tan, Z. Xu, G. Zhu, W. Teng, Q. Zhao, Z. Liang, Z. Wu, D. Xiong, Smart drug carrier based on polyurethane material for enhanced and controlled DOX release triggered by redox stimulus, React. Funct. Polym. 148 (2020) 1e12, https://doi.org/ 10.1016/j.reactfunctpolym.2020.104507. A. Arora, G. Aggarwal, J. Chander, P. Maman, M. Nagpal, Drug eluting sutures: a recent update, J. Appl. Pharmaceut. Sci. 9 (2019) 111e123, https://doi.org/10.7324/ JAPS.2019.90716. M.A. Alfhili, M.-H. Lee, Triclosan: an update on biochemical and molecular mechanisms, Oxid. Med. Cell. Longev. 2019 (2019), https://doi.org/10.1155/2019/ 1607304. A.B. Dann, A. Hontela, Triclosan: environmental exposure, toxicity and mechanisms of action, J. Appl. Toxicol. 31 (2011) 285e311, https://doi.org/10.1002/jat.1660. R. Zurita, J. Puiggalí, A. Rodríguez-Galán, Triclosan release from coated polyglycolide threads, Macromol. Biosci. 6 (2006) 58e69, https://doi.org/10.1002/ mabi.200500147. A. Obermeier, J. Schneider, S. Wehner, F.D. Matl, M. Schieker, R. von EisenhartRothe, A. Stemberger, R. Burgkart, Novel high efficient coatings for anti-microbial surgical sutures using chlorhexidine in fatty acid slow-release carrier systems, PLoS One 9 (2014) e101426, https://doi.org/10.1371/journal.pone.0101426. A. Obermeier, F.D. Matl, J. Schwabe, A. Zimmermann, K.D. Kühn, S. Lakemeier, R. von Eisenhart-Rothe, A. Stemberger, R. Burgkart, Novel fatty acid gentamicin salts as slow-release drug carrier systems for anti-infective protection of vascular biomaterials, J. Mater. Sci. Mater. Med. 23 (2012) 1675e1683, https://doi.org/10.1007/ s10856-012-4631-5. X. Chen, D. Hou, L. Wang, Q. Zhang, J. Zou, G. Sun, Antibacterial surgical silk sutures using a high-performance slow-release carrier coating system, ACS Appl. Mater. Interfaces 7 (2015) 22394e22403, https://doi.org/10.1021/acsami.5b06239. B. Joseph, A. George, S. Gopi, N. Kalarikkal, S. Thomas, Polymer sutures for simultaneous wound healing and drug delivery e a review, Int. J. Pharm. 524 (2017) 454e466, https://doi.org/10.1016/j.ijpharm.2017.03.041. J.E. Lee, S. Park, M. Park, M.H. Kim, C.G. Park, S.H. Lee, S.Y. Choi, B.H. Kim, H.J. Park, J.-H. Park, C.Y. Heo, Y.B. Choy, Surgical suture assembled with polymeric drug-delivery sheet for sustained, local pain relief, Acta Biomater. 9 (2013) 8318e8327, https://doi.org/10.1016/j.actbio.2013.06.003. M.L. Hakim, N. Nahar, M. Saha, M.S. Islam, H.M. Reza, S.M. Sharker, Local drug delivery from surgical thread for area-specific anesthesia, Biomed. Phys. Eng. Express 6 (2020) 15028, https://doi.org/10.1088/2057-1976/ab6a1e. A. Setia, Applications of graft copolymerization: a revolutionary approach, in: V.K. Thakur (Ed.), Biopolymer Grafting: Applications, Elsevier, 2017, pp. 1e44.

Improving the therapeutic value of sutures

75

[77] B. Gupta, R. Jain, H. Singh, Preparation of antimicrobial sutures by preirradiation grafting onto polypropylene monofilament, Polym. Adv. Technol. 19 (2008) 1698e1703, https://doi.org/10.1002/pat.1146. [78] S. Bae, M.J. Dibalsi, N. Meilinger, C. Zhang, E. Beal, G. Korneva, R.O. Brown, K.G. Kornev, J.S. Lee, Heparin-eluting electrospun nanofiber yarns for antithrombotic vascular sutures, ACS Appl. Mater. Interfaces 10 (2018) 8426e8435, https://doi.org/ 10.1021/acsami.7b14888. [79] M. García-Vargas, C. González-Chomón, B. Magariños, A. Concheiro, C. AlvarezLorenzo, E. Bucio, Acrylic polymer-grafted polypropylene sutures for covalent immobilization or reversible adsorption of vancomycin, Int. J. Pharm. 461 (2014) 286e295, https://doi.org/10.1016/j.ijpharm.2013.11.060. [80] M.L. Bruschi, Main mechanisms to control the drug release, in: M.L. Bruschi (Ed.), Strategies to Modify the Drug Release from Pharmaceutical Systems, Woodhead Publishing, Cambridge, 2015, pp. 37e62, https://doi.org/10.1016/B978-0-08100092-2.00004-7. [81] M.C.D. Silva, R.D.C.A. Leal, H.N.D. Silva, M.V.L. Fook, Biodegradable suture threads as controlled drug delivery systems, Mater. Res. Innovat. 24 (2020) 161e165, https://doi.org/10.1080/14328917.2019.1619982. [82] D. Das, N. Mulchandani, A. Kumar, V. Katiyar, Fabrication of stimuli-responsive polymers and their composites: candidates for resorbable sutures, in: V. Katiyar, A. Kumar, N. Mulchandani (Eds.), Advances in Sustainable Polymers, Springer, Singapore, 2020, pp. 121e144, https://doi.org/10.1007/978-981-15-1251-3_6. [83] A. Kumar, A. Srivastava, I.Y. Galaev, B. Mattiasson, Smart polymers: physical forms and bioengineering applications, Prog. Polym. Sci. 32 (2007) 1205e1237, https:// doi.org/10.1016/j.progpolymsci.2007.05.003. [84] J. Maitra, V.K. Shukla, Cross-linking in hydrogels e a review, Am. J. Polym. Sci. 4 (2014) 25e31, https://doi.org/10.5923/j.ajps.20140402.01. [85] Y. Qiu, K. Park, Environment-sensitive hydrogels for drug delivery, Adv. Drug Deliv. Rev. 53 (2001) 321e339, https://doi.org/10.1016/S0169-409X(01)00203-4. [86] M.R. Aguilar, C. Elvira, A. Gallardo, B. Vázquez, J.S. Román, Smart polymers and their applications as biomedicals, in: N. Ashammakhi, R. Reis, E. Chiellini (Eds.), Topics in Tissue Engineering, Biomaterials and Tissue Engineering Group, 2007, pp. 225e248. https://www.oulu.fi/spareparts/ebook_topics_in_t_e_vol3/abstracts/ aguilar_01.pdf. [87] M.R. Aguilar, J. San Román, Introduction to smart polymers and their applications, in: M.R. Aguilar, J.B.T. San Román (Eds.), Smart Polymers and Their Applications, Woodhead Publishing, Cambridge, 2014, pp. 1e11, https://doi.org/10.1533/ 9780857097026.1. [88] S. Dai, P. Ravi, K.C. Tam, pH-Responsive polymers: synthesis, properties and applications, Soft Matter 4 (2008) 435e449, https://doi.org/10.1039/B714741D. [89] Y. Yin, Responsive Photonic Nanostructures: Smart Nanoscale Optical Materials, Smart Materials Series, Royal Society of Chemistry, Cambridge, 2013, https://doi.org/ 10.1039/9781849737760. [90] T. Yoshida, T.C. Lai, G.S. Kwon, K. Sako, pH- and ion-sensitive polymers for drug delivery, Expet Opin. Drug Deliv. 10 (2013) 1497e1513, https://doi.org/10.1517/ 17425247.2013.821978. [91] G. Kamath, S.A. Deshmukh, G.A. Baker, D.C. Mancini, S.K.R.S. Sankaranarayanan, Thermodynamic considerations for solubility and conformational transitions of polyN-isopropyl-acrylamide, Phys. Chem. Chem. Phys. 15 (2013) 12667e12673, https:// doi.org/10.1039/c3cp44076a.

76

Advanced Technologies and Polymer Materials for Surgical Sutures

[92] Z. Zhang, Switchable and Responsive Surfaces and Materials for Biomedical Applications, Woodhead Publishing, Cambridge, 2015, https://doi.org/10.1016/C2013-016356-8. [93] P. Zarrintaj, M. Jouyandeh, M.R. Ganjali, B.S. Hadavand, M. Mozafari, S.S. Sheiko, M. Vatankhah-Varnoosfaderani, T.J. Gutiérrez, M.R. Saeb, Thermo-sensitive polymers in medicine: a review, Eur. Polym. J. 117 (2019) 402e423, https://doi.org/ 10.1016/j.eurpolymj.2019.05.024. [94] J. Seuring, S. Agarwal, Polymers with upper critical solution temperature in aqueous solution: unexpected properties from known building blocks, ACS Macro Lett. 2 (2013) 597e600, https://doi.org/10.1021/mz400227y. [95] C. Von Eiff, W. Kohnen, K. Becker, B. Jansen, Modern strategies in the prevention of implant-associated infections, Int. J. Artif. Organs 28 (2005) 1146e1156, https:// doi.org/10.1177/039139880502801112. [96] W. Zhao, L. Liu, F. Zhang, J. Leng, Y. Liu, Shape memory polymers and their composites in biomedical applications, Mater. Sci. Eng. C 97 (2019) 864e883, https:// doi.org/10.1016/j.msec.2018.12.054. [97] X. Jing, H.-Y. Mi, H.-X. Huang, L.-S. Turng, Shape memory thermoplastic polyurethane (TPU)/poly(ε-caprolactone) (PCL) blends as self-knotting sutures, J. Mech. Behav. Biomed. Mater. 64 (2016) 94e103, https://doi.org/10.1016/ j.jmbbm.2016.07.023.

CHAPTER 4

Evaluating the mechanical properties of sutures Sharin Maria Thomas1, R. Anjana1, Blessy Joseph3, Nandakumar Kalarikkal2 and Sabu Thomas2 1

Centre for Nanotechnology Research, Vellore Institute of Technology, Vellore, Tamil Nadu, India; International and Inter University Centre for Nanoscience and Nanotechnology, Mahatma Gandhi University, Kottayam, Kerala, India; 3Business Innovation and Incubation (BIIC), Mahatma Gandhi University, Kottayam, Kerala, India

2

4.1 Introduction Sutures are thread or strings made up of various materials for the primary closure of the tissues due to an injury or surgery and to foster the healing process [1]. Sutures approximate the tissues together. Sutures must be strong enough to hold the tissues together and flexible enough to tie into knots. The suturing technique generally involves a needle which is attached with a thread (suture). Sutures have gained wide attention in wound closures compared to other biomaterials such as staples/ligating clips, tissue adhesives, and medical tapes. An ideal suture should exhibit optimal mechanical properties such that it should possess high tensile strength and it should loosen up simultaneously as the tissue regenerates [2]. Sutures should be biocompatible, easily conformable, and should be able to accommodate the wound edema [3]. A secure knot with minimal infection and less scarring is highly beneficial. A knot is formed by intertwining the two ends of a suture more than once. The first knotting loop formed is called the approximation loop that is primarily responsible for holding the open ends of tissue together and fixes the wound edges. Additional loops are made further as support for the approximation loop [4]. The mechanical properties of the sutures vary by their composition. Sutures are chosen depending on the particular properties of the tissue. Sutures can be categorized into different types depending on their origin, structure, and spontaneity of degradation. Based on the origin of material, sutures can be classified into natural and synthetic. Natural sutures are made from naturally occurring sources (animal/plant tissues), whereas synthetic are those made in industries. One of the main advantages of synthetic suture is its predictability in the degradation rate (in the case of absorbable suture) inside a biological domain and Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00002-4

© 2023 Elsevier Ltd. All rights reserved.

77

78

Advanced Technologies and Polymer Materials for Surgical Sutures

its reproducibility. This helps in avoiding undesirable tissue reactions after the suture has lost most of the mechanical properties. Various parameters have to be considered in order to optimize the desired sutures according to its area and depth of wound and region of application. United States Pharmacopeia and European Pharmacopeia are the two main widely used official compendiums that have set the standards to classify the surgical sutures. These standards define the range on average diameter, average knot pull tensile strength, the sterility of the suture, mode of needle attachment, extractable color, and also various analyses of soluble harmful compounds if any like chromium in the case of absorbable sutures. In the case of primary wound management, the surgical sutures represent the gold standard [5]. However, sutures are still utilized mainly for the closure of the wounds with the main attention on the structural stability of the sutures, imperative for ameliorating the rate of healing and to tamponade the haemorrhage. The mechanical characteristics of suture materials affect the suture function and thereafter wound healing process. Robust mechanical parameters at an optimal mode could facilitate tissue proliferation. However, due to the increasing occurrence of surgical wound infections, there is a need to integrate drug-delivery with sutures without compromising on the vital mechanical aspects. Conventionally, sutures are manufactured similar to the polyester thread in the garment industry. The raw materials are polymerized, extruded through a nozzle into fibers, and then stretched. It is then again processed accordingly for the multifilament type. Apart from this, electrospun membranes can be cut into smaller widths and twisted to form sutures. Electrospun surgical sutures are an effective approach for this inclusion as the small diameters of nanofibers sustaining good tensile strength and simultaneously facilitating cell adhesion. The electrospun sutures can be functionalized for drug elution and antimicrobial effects. The large surface area of the electrospun membrane means greater active areas. Padmakumar et al. developed core-shell suture yarns using electrospinning. Poly (L-lactic acid) (PLLA) formed the core and poly (lactic-co-glycolic acid) (PLGA) acted as the sheath. To enhance the mechanical properties the spun yarns were heat stretched. Thus the heat stretched fibers exhibited a tensile strength of 242.8  12.3 MPa that was 14 times higher than unheated ones (16.7  1.3 MPa) [6]. This chapter explains the different mechanical properties that are to be considered while manufacturing a suture, the important characterization techniques, and it further discusses the effects of various antibacterial coatings on the mechanical properties of sutures.

Evaluating the mechanical properties of sutures

79

4.2 Mechanical properties The fundamental mechanical properties of sutures are vital for efficient functioning and wound patching. The mechanical properties of the sutures are of great importance in short-term restoring the lost strength of tissue. Counting on the patient’s innate tendency to restore wounds is not reliable in severe conditions of diabetics [7] and other inflammatory diseases. A bioactive photosynthetic suture seeded with therapeutic recombinant growth factors and altered microalgae, an active oxygen carrier facilitating the healing process have been investigated by Centeno-Cerdas et al. This novel work depicted improved tissue regeneration along with stable mechanical properties [8]. Electrospun PLGA loaded with local anesthetic was reported by Weldon et al. for perioperative anesthetics that reduces the demand for postoperative opioids [9]. The mechanical properties remained unaffected over time; however, above an optimal amount of drug load can hasten suture degradation before the stipulated time. There will be postoperative complications due to the weakening of the wound site. So improving the mechanical properties of suture is important. Pasternak et al. coated sutures with cross-linked fibrinogen film and bound doxycycline into the film which gave 17% higher breaking strength and 20% enhanced energy uptake at failure [10]. Chen et al. have developed antibacterial sutures by twisting strips of gentamicin/pluronic F127-silver/PCL (poly c-capro-lactone) nanofiber from electrospun membrane. The ultimate tensile strength of the antibacterial suture was lower than the clear PCL suture [11]. Crucial decline in the mechanical properties were observed in the case of berberine incorporated electrospun polyurethane fibers. This fall off is because of the excess loading of berberine that will lead to larger micropores, thus loosening the microstructure of the fiber. This can reduce the mechanical property of the fiber. However, once optimized the fiber depicted in vitro shape memory effect and antiinflammatory responses [12]. From the above said works, it can be seen that by improving the performance of sutures by functionalizing will enhance or decrease the mechanical properties. So, it is important to choose the materials for improvising suture efficiency without compromising the suture mechanics. The mechanical properties like tensile strength, knot strength, elasticity, etc. determines the properties of the suture. The custom tensile testing set up for sutures is given in Fig. 4.1. In an ideal case, the mechanical property of the suture should match that of wound healing rate. To avoid unexpected burst opening of the wound,

80

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 4.1 Solid model image of the custom tensile mounts. (Reproduced with permission from Ref. [2].)

knot strength should be high. The suture should be able to conform to the present stage of wound repair especially during the initial stage of edema. 4.2.1 Tensile strength Tensile strength is the maximum tensile load a body can withstand before failure divided by its cross-sectional area. Tensile strength determines the threshold strength up to which a suture material withstands breakage when knotted [13]. According to USP standards (US Pharmacopeia, Est 1937) [14], the diameter of sutures is denoted by the number of zeroes. For instance, 4-0 means 0000 and 6-0 means 000000. Smaller diameter has more zeroes in the suture size (6-0 is smaller in diameter than 4-0) and lesser will be the tensile strength. Elongation of suture materials can lead to higher tensile load and will henceforth result in a decrease in diameter of the suture that will be detrimental for tissue healing. Braided sutures have a lower elongation rate compared to monofilament sutures. A high load will lead to breakage or gaps due to slippage of knots thus transpiring to mechanical instability [15]. Unknotted polydioxane and knotted polyglyconate

Evaluating the mechanical properties of sutures

81

Figure 4.2 Modulus lines are shown for the nonabsorbable sutures. Note the two separate lines used for the initial and secondary moduli of the biphasic sutures. (Reproduced with permission from Ref. [2].)

permitted the most elongation rate prior to achieving the breaking strength. This indicated greater accommodation of postoperative edemas. Considering the elastic modulus of suture materials, certain nonabsorbable sutures depict a biphasic stage where the materials show early stages of elongation upon exertion of load. The biphasic curves resulted in earlier elongation rate as shown in Fig. 4.2 for Nylon and polybutester. Nonabsorbable silk (8701 MPa) and absorbable polyglactin 910 (9320 MPa) depicted the highest initial elastic modulus. Multifilaments which are made up of braided or twisted with many filaments show better tensile strength than monofilament sutures [16]. 4.2.2 Knot strength Knot strength is the force required to create a slip or break in the knot. This force is directly proportional to the coefficient of friction of the suture. A suture material with high coefficient of friction will have good knot security but this friction will come up with negative effects to the tissue by being abrasive and contributing unwanted resistance [17]. The precision of the knot is important. Knot strength should be optimum such that if not properly tied, it adds to the bulk, leading to nonuniformity. The knot strength varies among different multifilament sutures due to the technical difference in the braiding and twisting process. Knotting greatly affects the mechanical properties of the sutures. It is highly dependent on the plasticity

82

Advanced Technologies and Polymer Materials for Surgical Sutures

of the suture material and the coefficient of the static friction. Of the nonabsorbable suture materials, nylon depicted the highest unknotted failure load and polypropylene, being most widely used for in-vivo, had the highest knotted failure load. For absorbable sutures, the Glycomer 631 displayed highest failure load in unknotted and knotted polyglyconate depicted highest failure load [2]. Silk sutures exhibit good knot security and easiness in knotting which is generally acknowledged by the clinical experiences [18,19]. 4.2.3 Breaking strength The maximum threshold of the tensile strength above which the suture fails by tearing or ripping of the suture or the skin is called the breaking strength of the suture [20]. This point is either achieved gradually with time when the tissue regains strength or due to the excessive tension applied on the suture while applying knots. However, the suture should be flexible as well. If the mobility of the sutured area is more, higher will be the tensile strength required for the suture which would result in demand for higher value of breaking strength, both of which being very important parameters for a successful suture. Wound edema demands for higher mobility resulting in increased breaking strength due to greater tensile strength that is a prerequisite for a successful suture. Breaking strength is highly dependable on the suture technique followed and the technique which gives the highest breaking strength for a particular suture will be considered as the best technique. Rashid et al. investigated the breaking strength of different polypropylene sutures [21]. They have used the Instron model as the testing machine for carrying out the study. Common basic suturing techniques are continuous running, simple interrupted, locking, purse, far-and-near, and horizontal mattress. 4.2.4 Knot-pull tensile strength Knot-pull tensile strength is the maximum tensile stress that can be applied on the ears of a knot. Tests for knot pull strength have particular significance in defining the quality of suture. Knotting generally reduces the tensile strength of the sutures. A secure knot can be defined as a knot which fails by breaking rather than untying. There are various types of knots used depending on the surgeon and or type of wound like square knot, surgeon’s (friction) knot, simple knot etc. A few of them are as shown in Fig. 4.3.

Evaluating the mechanical properties of sutures

83

Figure 4.3 Digital images showing the configuration of (A) square knot, (B) surgeon’s square knot and (C) surgeon’s granny knot. (Reproduced with permission from Ref. [22].)

Few layers are added to the knot to secure it properly and prevent any slippage. Throws in a knot are made by crossing the ends to form a loop and then wrapping one end of the suture around the other. Number of throws has a crucial role in knot security and tensile load failure. A square knot with five throws is found to have maximum tensile strength when tested on individual sutures made of polyester, polydioxanone, polypropylene, and polyglactin 910 [23]. The tensile strength of the knot in polyamide (PA) material reported a high stress concentration in and around the knot [24]. The knot strength depends on the size of the knot and the suture material. The tensile strength is generally lower for sutures of smaller size. It can be increased by reinforcing the supporting knot to prevent slippage of the knot when excess load is applied. A knot pull tensile test described by USP is measured similar to any other tensile testing (UTMUniversal Testing Machine) where the knot is centered in between the grips. The test is carried out in relation to the type of knot given by the surgeon versus the shape of the clamps used to stretch the sutures [25]. Effective knot-pull tensile strength is of great importance considering the risk of wound splitting, but it should not be greater than the tensile strength of the tissue [26]. 4.2.5 Wound breaking strength A critical effect of the wound repair process is restoration of the mechanical properties of tissue strength. The wound breaking strength is defined as the limit of tensile strength at which the wound edges separate in the case of a healing wound [27]. The analysis of wound breaking strength contributes to the evaluation of the aggregate healing process. If enough wound strength is not achieved during the healing process, the net effect may be a wound failure [28]. There are basically two methods for measuring the tensile strength of the wound: depending on the increased intraluminal

84

Advanced Technologies and Polymer Materials for Surgical Sutures

pressure for the disruption of wounds of hollow viscera (internal organs) and the lateral forces required to disrupt a wound. Several instruments were designed and fabricated by using these methods to improve the measurability [29e35].The tensile strength of a healed wound can only reach up to 80% approx. that of an unwounded tissue. Musculoaponeurotic suture model was experimented in a group of mice to investigate the scar breaking strength where polypropylene suture scars showed a higher breaking strength than silk or polyglactin sutures [36]. 4.2.6 Elasticity Elasticity is the innate property in the suture material to get back to its original shape after being stretched. Elasticity helps accommodate wound edemas and offers distinctive wound edge appositions upon recoil. The modulus of elasticity is the measure of resistance upon elongation before the breaking point. It is depicted as the slope of the stressestrain curve. Monofilament polydioxanone depicted sustained high elasticity even when incubated in phosphate buffer solution for 28 days and showed highest yield to failure strain in contrast to glycolide/lactide copolymer and polyglactin 910 [37]. Low elasticity can lead to tissue tear due to excess load. Highly elastic nylon threads are favorable for peripheral sutures in tendons with reduced load while maintaining the mechanical stability [38]. 4.2.7 Plasticity Plasticity of the suture is the ability of the material to undergo deformation with the applied force and has less tendency to return to its original initial dimension. High plasticity hinders the approximation of wound edges. Above the optimal conditions, the suture cannot recoil and it is beneficial for the proper circulation in tissues postwound healing. But if the sutures turn out to be plastic prior wound healing, the function of the suture acts up. Prolene is a highly plastic suture being utilized mainly because of its smooth surface, it won’t cut the tissue [39]. Certain degrees of plasticity can be tolerated when used in intradermal stitches that take into account the smoothness over plasticity. Otherwise, plasticity is not a much preferred characteristic in sutures that are supposed to be pliable and flexible. 4.2.8 Memory Memory of the suture is interlinked with elasticity and plasticity. It lies in an equally critical distance from both. It is simply the ability of the suture

Evaluating the mechanical properties of sutures

85

material to return to its original position after being deformed. Memory of the material can be a measure of the stiffness of the material. Sutures with a high packaging memory do not guarantee a secure knot as they tend to return to their winding or kink form. Surgeons find it difficult to sew using high memory sutures. One easiest way to determine the memory is to suspend the suture and analyze the time it takes to recoil to its original state. Except for monofilament sutures like Monocryl, Biosyn, and Gore-Tex, all other monofilaments depict good memory, hence less favorable for sutures. The memory of monofilaments could be reduced by using a conjugate spinning method. Park et al., using this method, let fine strands of polymeric fiber into another polymer matrix by developing sea or island shaped biocomponents onto the monofilaments. These modified monofilaments rendered commendable pliability and exceptional knot security [40]. 4.2.9 Pliability Pliability is the ease of using a suture in securing a knot. It can also be regarded in terms of flexibility for easier handling of sutures. Even though multifilament sutures depict high mechanical properties, they are highly pliable compared to monofilament sutures [41]. Pliability is one of the sought after structural properties in sutures for deep wound closures [42]. The pliable sutures facilitate formation of self-tightening and secure knots for tissue closure with significant pull strength. Considerable pliability is demonstrated by the multifilament sutures, especially the braided ones when weighed against monofilament sutures [43]. 4.2.10 Capillarity Capillarity in sutures is the tendency of the surrounding fluids to get absorbed and then transferred by the suture material. By this effect, it is possible for the unwanted bacteria and microorganisms to pass through it and reach the entire wounded area, hence making it infectious. Capillary effect is more prominent in multifilament than monofilament sutures. To prevent capillarity due to potential contamination and suture fistulas [44], several attempts like suture coating, jacketing, polishing, etc. An open suture is also similar to the capillary which encourages wound infection. The studies based on the effect of capillary for the bacterial intrusion and transport was done as early as in the 1970s. Higher rate of infection was observed in multifilament sutures due to greater capillarity when compared

86

Advanced Technologies and Polymer Materials for Surgical Sutures

with the monofilament sutures [44]. Independent of the material type, braided multifilament silk and nylon sutures attracted thrice more bacteria than monofilament sutures of the same [45] Geiger et al. reported that the multifilament sutures of PDSII, Biosyn, Prolene, Monocryl, and Maxon facilitated E. coli growth than the monofilament ones [46]. 4.2.11 Abrasion Abrasive wear or Abrasion is an inconsistency in the suture material leading to decreased breaking strength of the materials. Abrasive defects can tear or damage the healing tissues. Ultrahigh molecular weight polyethylene core filaments provided greater abrasion resistance against PET and its smooth morphology has a lubricating effect in abrasion [47], hence highly applicable in open orthopedic surgeries. The suture anchor axis and the rotational orientation of the anchor’s eyelet with respect to the plane are found to worsen the effects of abrasion. Circular or semicircular configurations with large diameter and sharp unfriendly edges of the anchor eyelets cause abrasive effects such as tissue tear. Some of the least suited SEM images of anchor eyelets that cause abrasion are as shown in Fig. 4.4. Elimination of edgy designs and employing more polished surfaces for anchor eyelets can enhance the failure strength during cyclic loading. This further offers better abrasion resistance with multi orientation freedom [48].

Figure 4.4 SEM images of various eyelet designs namely (a) Fastin (b) Ultrafix (c) Corkscrew (d) Revo (e) GII (f) PeBA with unfriendly edges. (Reproduced with permission from Ref. [48].)

Evaluating the mechanical properties of sutures

87

4.3 Characterization techniques 4.3.1 Universal testing machine (UTM) The fundamental mechanical properties of the suture can be analyzed by a universal testing machine. Various mechanical parameters such as the failure load, elongation, and knot breakage can be characterized. The analysis is conducted by placing the sutures on the two opposite hooks and is tested by the application of a certain standard force by maintaining a mandible separation speed between the two hooks. Fig. 4.5 displays an Instron testing machine with suture position [49]. Failure load is the maximum load the suture can withstand. Elongation is the displacement suture experiences prior to the breaking point. Knot breakage is the point at which the suture breaks upon the application of load. The stressestrain curves help determine the peak tensile strength and elongation. The obtained data is analyzed with the help of multivariate analysis of variance and a Tukey studentized range test. 4.3.2 Abrasive testing The abrasive resistance of the sutures through an anchor eyelet can be studied. A custom-made device was used to analyze the suture abrasion

Figure 4.5 Suture position within the Instron clamps. (Reproduced with permission from Ref. [49].)

88

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 4.6 Abrasion testing apparatus. (A) Electronically controlled motor. (B) Wheel. (C) Plunger. (D) Polyurethane block containing anchor. (E) Weight. Note that the construct is being tested at a 45 sutureeanchor angle. (Reproduced with permission from Ref. [50].)

primarily happened during arthroscopic or keyhole surgery, where the suture abrade against the bone through an anchor eyelet. The testing machine contained an electronic motor (A), wheel (B), plunger (C), Polyurethane (PU) block (D) and a standard weight (E) as shown in Fig. 4.6. The plunger attached to the wheel rotates due to the motor converting the rotational motion to a linear motion. The anchor is attached to the PU blocks at a variable inclination angle such that it is centered near the suture and the suture is allowed to pass through the eyelet in a to and fro motion to cause the abrasion. The revolutions and the speed of occurrence can be set. The varying parameters were the suture type, anchor angle, and wet or dry conditions. The abrasive resistance due to cyclic loading can be investigated. The results obtained were compared by ANOVA and Bonferroni corrections. The abrasive effects could be minimized by proper orientations angles between the anchor and sutures and by using suitable suture types [40].

4.4 Effect of antibacterial coating on mechanical properties Sutures can be a source for the bacterial infection. It is one of the major causes for the increasing number of deaths and high medical expenses [51,52]. Microorganisms incubate in the suture and then spread to form a biofilm. The formation of the biofilm leads to high risk of Surgical Site Infections (SSIs) and once this happens, the elimination of these microorganisms is difficult. The common bacteria like Escherichia coli, Staphylococcus aureus, Enterococci that find sutures as a convenient place to form biofilm

Evaluating the mechanical properties of sutures

89

which can make long-term infections and may lead to prolonged treatments [53]. One way to prevent the bacterial adhesion is by improving the surface sutures [54,55]. Materials like triclosan, chlorhexidine, and tetracycline can be used to achieve the antibacterial property. Silicone and wax of low density are used to smoothen the surface. The tensile strength of certain sutures (PGA, silk, polyester, and chromic catgut) remained the same with this coating [56]. Triclosan was a common material used for coatings for achieving the antibacterial property [57]. It is proven to be efficient in reducing SSI. But using triclosan as a coating in sutures have raised many concerns on environmental pollution, food allergy, developing bacterial resistance, and cross-resistance and a wide range of health risks [58,59]. Therefore, the need for developing effective and harmless coating for antibacterial sutures is crucial. Nanosilver particles can be used to treat this issue which does not compromise and even improve the tensile strength and knot strength as well as reduce degradation [52]. It is important to note that this advancement was due to the selection of solvent (e.g.,: PEG) for the proper dispersion of nanosilver particles. Braided nonabsorbable PA suture was coated with a mixture of citric acid, sodium hypophosphite, and chitosan which showed remarkable antibacterial effect against E. coli and P. aeruginosa without compromising its tensile and knot pull strength [60]. Franco et al. have put forward an antimicrobial coating by using the recombinant DNA technology. Spider silk chimeric proteins were designed and functionalized with antimicrobial peptides and were dip-coated onto the silk sutures. They studied inhibitory effects against MRSA (methicillinresistant Staphylococcus aureus) and E. coli and reported a significant reduction in the microbial activity simultaneously sustaining the desirable tensile and knot strength.

4.5 Conclusion Surgical sutures are the most primitive and the longest used implements for the wound closure. With the advent of synthetic materials, there came more consistency in the type of precursors used and it turned out to be more reproducible because of the well-defined processing conditions. The instigated synthetic sutures were inert and biocompatible in nature. A lot of attention was focused on the absorbable type particularly the monofilaments with multifunctional attributes such as antibacterial or antimicrobial effects in order to reduce or nullify the postoperative infections. The mechanical properties are the foundational traits governing the closure of

90

Advanced Technologies and Polymer Materials for Surgical Sutures

wounds. These include tensile strength, modulus of elasticity, elongation at break, security and strength of the knot, compliance and the pliability of the sutures. Sutures should sustain their strength until the wound restoration is satisfactory. The tensile breaking strength being independent of the chemical structure or size can be normalized for the comparison of sutures based on dimensions. The modulus of elasticity derived from the tensile test can provide insights into the bending stiffness. Other handling properties such as memory and pliability also fall under mechanical properties. Sutures with low memory are desirable or else they easily return to their initial form from the knot form. The testing of some of these properties can be done by UTM and other custom designed testing setups. The standards for selection of suture materials are done in coherence with the United States Pharmacopoeia (USP) or European Pharmacopoeia (EP). Although the suture material meets the mechanical requirements for mending the tissues, there are certain other challenges like SSIs and tissue trauma. This issue is addressed by introducing antibacterial coatings on the surface of sutures. Another coating used in sutures is to make the surface of sutures smooth for easiness while suturing the tissue. Coatings with similar chemical composition as that of suture are favorable. Coating like Poloxamer 188 and calcium stearate which are absorbable and nonabsorbable sutures are treated with materials like wax, silicon and fluorocarbons to make the surface smooth. However, there are many challenges that limit the operational quality of the surgical sutures. Each suture varies depending on their utility, good knot run down, knot security, high tensile strength, good pliability, and suture materials compatible with the body. And in teams of coatings, it can impart bending stiffness to the sutures. So, there is a need for a universal standard that determines the suture properties, proper classification and comparison of sutures. There is great potential for biodegradable sutures with optimal parameters that opens a wide area of extensive study in the field suture material research.

References [1] R.L. Moy, B. Waldman, D.W. Hein, A review of sutures and suturing techniques, J. Dermatol. Surg. Oncol. 18 (9) (1992) 785e795, https://doi.org/10.1111/j.15244725.1992.tb03036.x. [2] S.E. Naleway, W. Lear, J.J. Kruzic, C.B. Maughan, Mechanical properties of suture materials in general and cutaneous surgery, J. Biomed. Mater. Res. B Appl. Biomater. 103 (4) (2015) 735e742, https://doi.org/10.1002/jbm.b.33171.

Evaluating the mechanical properties of sutures

91

[3] C.C. Chu, Mechanical properties of suture materials: an important characterization, Ann. Surg. 193 (3) (1981) 365e371, https://doi.org/10.1097/00000658-19810300000021. [4] R.C. Webster, E.G. McCollough, P.R. Giandello, R.C. Smith, Skin wound approximation with new absorbable suture material, Arch. Otolaryngol. 111 (8) (1985) 517e519, https://doi.org/10.1001/archotol.1985.00800100065008. [5] C.E. Azmat, M. Council, Wound Closure Techniques, in: StatPearls [Internet], StatPearls Publishing, Treasure Island (FL), 2022. PMID: 29262163. [6] S. Padmakumar, et al., Electrospun polymeric coreesheath yarns as drug eluting surgical sutures, ACS Appl. Mater. Interfaces 8 (11) (2016) 6925e6934. [7] F.M. Davis, A. Kimball, A. Boniakowski, K. Gallagher, Dysfunctional wound healing in diabetic foot ulcers: new crossroads, Curr. Diabetes Rep. 18 (1) (2018) 1e8. [8] C. Centeno-Cerdas, et al., Development of photosynthetic sutures for the local delivery of oxygen and recombinant growth factors in wounds, Acta Biomater. 81 (2018) 184e194. [9] C.B. Weldon, et al., Electrospun drug-eluting sutures for local anesthesia, J. Contr. Release 161 (3) (2012) 903e909. [10] B. Pasternak, et al., Doxycycline-coated sutures improve mechanical strength of intestinal anastomoses, Int. J. Colorectal Dis. 23 (3) (2008) 271e276. [11] S. Chen, et al., Twisting electrospun nanofiber fine strips into functional sutures for sustained co-delivery of gentamicin and silver, Nanomed. Nanotechnol. Biol. Med. 13 (4) (2017) 1435e1445. [12] W. Zhou, et al., Berberine-incorporated shape memory fiber applied as a novel surgical suture, Front. Pharmacol. 10 (2020) 1506. [13] S.V. Khiste, V. Ranganath, A.S. Nichani, Evaluation of tensile strength of surgical synthetic absorbable suture materials: an in vitro study, J. Periodontal Implant Sci. 43 (3) (2013) 130. [14] E. Açan, O. Hapa, F.A. Barber, Mechanical properties of suture materials, in: U. Akgun, M. Karahan, P.S. Randelli, J. Espregueira-Mendes (Eds.), Knots in Orthopedic Surgery, Springer, Berlin, Heidelberg, 2018, pp. 21e31. [15] J.-C. Kim, Y.-K. Lee, B.-S. Lim, S.-H. Rhee, H.-C. Yang, Comparison of tensile and knot security properties of surgical sutures, J. Mater. Sci. Mater. Med. 18 (12) (2007) 2363e2369. [16] C. McGarrigle, et al., Extruded monofilament and multifilament thermoplastic stitching yarns, Fibers 5 (4) (2017), https://doi.org/10.3390/fib5040045. Art. no. 4. [17] J.B. Trimbos, Security of various knots commonly used in surgical practice, Obstet. Gynecol. 64 (2) (1984) 274e280. [18] N. Tomita, S. Tamai, K. Ikeuchi, Y. Ikada, Effects of cross-sectional stress-relaxation on handling characteristics of suture materials, Bio Med. Mater. Eng. 4 (1) (1994) 47e59, https://doi.org/10.3233/BME-1994-4105. [19] N. Tomita, S. Tamai, T. Morihara, K. Ikeuchi, Y. Ikada, Handling characteristics of braided suture materials for tight tying, J. Appl. Biomater. 4 (1) (1993) 61e65, https:// doi.org/10.1002/jab.770040108. [20] G.V. Poole Jr., J.W. Meredith, N.D. Kon, M.B. Martin, E.H. Kawamoto, R.T. Myers, Suture technique and wound-bursting strength, Am. Surg. 50 (10) (1984) 569e572. [21] R. Rashid, M. Sartori, L.E. White, M.T. Villa, S.S. Yoo, M. Alam, Breaking strength of barbed polypropylene sutures: rater-blinded, controlled comparison with nonbarbed sutures of various calibers, Arch. Dermatol. 143 (7) (2007) 869e872, https://doi.org/ 10.1001/archderm.143.7.869. [22] F. Debbabi, S.B. Abdessalem, S. Limem, New test methods to evaluate the performance of dermatological braided sutures from both the doctor and the patient sides, J. Text. Inst. 102 (6) (2011) 548e557.

92

Advanced Technologies and Polymer Materials for Surgical Sutures

[23] T.M. Muffly, N. Kow, I. Iqbal, M.D. Barber, Minimum number of throws needed for knot security, J. Surg. Educ. 68 (2) (2011) 130e133. [24] D. Carr, A. Heward, R. Laing, B. Niven, Measuring the strength of knotted suture materials, J. Text. Inst. 100 (Mar. 2009) 51e56, https://doi.org/10.1080/ 00405000701608177. [25] H. Planck, O. Weber, C. Elser, M. Renardy, K. Mayr, M. Milwich, Test methods for surgical sutures, in: H. Planck, M. Dauner, M. Renardy (Eds.), Medical Textiles for Implantation, Springer, Berlin, Heidelberg, 1990, pp. 231e242. [26] A.J. Dart, C.M. Dart, 7.38 suture material: conventional and stimuli responsive, Compr. Biomater. (2011) 573e587, https://doi.org/10.1016/B978-0-08-0552941.00245-2. [27] S. Najibi, R. Banglmeier, J.M. Matta, M. Tannast, Material properties of common suture materials in orthopaedic surgery, Iowa Orthop. J. 30 (2010) 84. [28] R.L. Gamelli, L.-K. He, Incisional wound healing, in: L.A. DiPietro, A.L. Burns (Eds.), Wound Healing, Humana Press, Totowa, NJ, 2003, pp. 37e54. [29] S.C. Harvey, The velocity of the growth of fibroblasts in the healing wound, Arch. Surg. 18 (4) (1929) 1227e1240. [30] T.H. Lanman, T.H. Ingalls, Vitamin C deficiency and wound healing: an experimental and clinical study, Ann. Surg. 105 (4) (1937) 616. [31] E.L. Howes, S.C. Harvey, The strength of the healing wound in relation to the holding strength of the catgut suture, N. Engl. J. Med. 200 (25) (1929) 1285e1291. [32] J.B. Hartzell, W.E. Stone, The relationship of the concentration of ascorbic acid of the blood to the tensile strength of wounds in animals, Surg. Gynecol. Obstet. 75 (1942) 1e7. [33] C.M. Jones, M.K. Bartlett, A.E. Ryan, G.D. Drummey, The effect of sulfanilamide powder on the healing of sterile and infected wounds: with special reference to tensile strength and ascorbic acid content in the scar, N. Engl. J. Med. 229 (17) (1943) 642e646. [34] A.C. Berger, et al., The angiogenesis inhibitor, endostatin, does not affect murine cutaneous wound healing, J. Surg. Res. 91 (1) (2000) 26e31. [35] D. Greenhalgh, R.L. Gamelli, R.S. Foster Jr., A. Chester, Inhibition of wound healing by Corynebacterium parvum, J. Surg. Res. 41 (2) (1986) 209e214. [36] D. Miro, M.V. Julià, A. Sitges-Serra, Wound breaking strength and healing after suturing noninjured tissues, J. Am. Coll. Surg. 180 (6) (1995) 659e665. [37] C.M. Kearney, C.T. Buckley, F. Jenner, P. Moissonnier, P.A.J. Brama, Elasticity and breaking strength of synthetic suture materials incubated in various equine physiological and pathological solutions, Equine Vet. J. 46 (4) (2014) 494e498. [38] K. Nozaki, R. Mori, K. Ryoke, Y. Uchio, Comparison of elastic versus rigid suture material for peripheral sutures in tendon repair, Clin. Biomech. 27 (5) (2012) 506e510. [39] B.S. Bloom, D.J. Goldberg, Suture material in cosmetic cutaneous surgery, J. Cosmet. Laser Ther. 9 (1) (2007) 41e45. [40] J.N. Im, J.K. Kim, H.-K. Kim, K.Y. Lee, W.H. Park, Characteristics of novel monofilament sutures prepared by conjugate spinning, J. Biomed. Mater. Res. Part B Appl. Biomater. Off. J. Soc. Biomater. Jpn. Soc. Biomater. Aust. Soc. Biomater. Korean Soc. Biomater. 83 (2) (2007) 499e504. [41] C. Dennis, S. Sethu, S. Nayak, L. Mohan, Y. Morsi, G. Manivasagam, Suture materials e current and emerging trends, J. Biomed. Mater. Res. A 104 (6) (2016) 1544e1559, https://doi.org/10.1002/jbm.a.35683. [42] M. Indhumathi, S. Kumar, Application of antibacterial suture materials in oral and maxillofacial surgery, Drug Invent. Today 12 (1) (2019). [43] D. Kirsch, S. Marczyk, Multifilament barbed suture, 2013.

Evaluating the mechanical properties of sutures

93

[44] B. Osterberg, B. Blomstedt, Effect of suture materials on bacterial survival in infected wounds. An experimental study, Acta Chir. Scand. 145 (7) (1979) 431e434. [45] T.E. Bucknall, H. Ellis, Abdominal wound closureea comparison of monofilament nylon and polyglycolic acid, Surgery 89 (6) (1981) 672e677. [46] D. Geiger, et al., Capillary activity of surgical sutures and suture-dependent bacterial transport: a qualitative study, Surg. Infect. 6 (4) (2005) 377e383. [47] E. Savage, C.J. Hurren, S. Slader, L.A. Khan, A. Sutti, R.S. Page, Bending and abrasion fatigue of common suture materials used in arthroscopic and open orthopedic surgery, J. Orthop. Res. 31 (1) (2013) 132e138. [48] D.D. Bardana, R.T. Burks, J.R. West, P.E. Greis, The effect of suture anchor design and orientation on suture abrasion: an in vitro study, Arthrosc. J. Arthrosc. Relat. Surg. 19 (3) (2003) 274e281. [49] T.D. Schiller, E.A. Stone, B.S. Gupta, In vitro loss of tensile strength and elasticity of five absorbable suture materials in sterile and infected canine urine, Vet. Surg. 22 (3) (1993) 208e212. [50] I.K. Lo, S.S. Burkhart, K. Athanasiou, Abrasion resistance of two types of nonabsorbable braided suture, Arthrosc. J. Arthrosc. Relat. Surg. 20 (4) (2004) 407e413. [51] A.J. Mangram, T.C. Horan, M.L. Pearson, L.C. Silver, W.R. Jarvis, Guideline for prevention of surgical site infection, 1999, Am. J. Infect. Control 27 (2) (1999) 97e134, https://doi.org/10.1016/S0196-6553(99)70088-X. [52] B. James, R. Ramakrishnan, A.S. Aprem, Development of environmentally safe biodegradable, antibacterial surgical sutures using nanosilver particles, J. Polym. Environ. 29 (2021) 2282e2288, https://doi.org/10.1007/s10924-021-02048-y. [53] B. Joseph, A. George, S. Gopi, N. Kalarikkal, S. Thomas, Polymer sutures for simultaneous wound healing and drug delivery e a review, Int. J. Pharm. 524 (1) (2017) 454e466, https://doi.org/10.1016/j.ijpharm.2017.03.041. [54] D. Leaper, et al., Healthcare associated infection: novel strategies and antimicrobial implants to prevent surgical site infection, Ann. R. Coll. Surg. Engl. 92 (6) (2010) 453e458, https://doi.org/10.1308/003588410X12699663905276. [55] T. Hranjec, B.R. Swenson, R.G. Sawyer, Surgical site infection prevention: how we do it, Surg. Infect. 11 (3) (2010) 289e294, https://doi.org/10.1089/sur.2010.021. [56] F. Debbabi, S. Gargoubi, M.A. Hadj Ayed, S.B. Abdessalem, Development and characterization of antibacterial braided polyamide suture coated with chitosan-citric acid biopolymer, J. Biomater. Appl. 32 (3) (Sep. 2017) 384e398, https://doi.org/ 10.1177/0885328217721868. [57] I. Ahmed, et al., The use of triclosan-coated sutures to prevent surgical site infections: a systematic review and meta-analysis of the literature, BMJ Open 9 (9) (2019) e029727. [58] A. Apisarnthanarak, N. Singh, A.N. Bandong, G. Madriaga, Triclosan-coated sutures reduce the risk of surgical site infections: a systematic review and meta-analysis, Infect. Control Hosp. Epidemiol. 36 (2) (2015) 169e179. [59] M.-F. Yueh, et al., The commonly used antimicrobial additive triclosan is a liver tumor promoter, Proc. Natl. Acad. Sci. Unit. States Am. 111 (48) (2014) 17200e17205. [60] A. Munan, A.J. Khan, M. Khan, G. Abbas, S. Khan, M.I. Nazir, Fakhar-e-Alam, A. Iqbal, A novel method for preparation of antibacterial and atraumatical surgical sutures, Drug Discov. 15 (35) (2021) 132e139.

CHAPTER 5

Polymers for surgical sutures Samson Afewerki1, 2, *, Samarah Vargas Harb3, *, Thiago Domingues Stocco4, 5, *, Guillermo U. Ruiz-Esparza1, 2 and Anderson O. Lobo6 1

Division of Engineering in Medicine, Department of Medicine, Brigham & Women’s Hospital, Harvard Medical School, Boston, MA, United States; 2Division of Health Science and Technology, Harvard University e Massachusetts Institute of Technology, MIT, Cambridge, MA, United States; 3Department of Materials Engineering (DEMa), Federal University of São Carlos (UFSCar), São Carlos, São Paulo, Brazil; 4Faculty of Medical Science, State University of Campinas (UNICAMP), Campinas, São Paulo, Brazil; 5University of Santo Amaro, Santo Amaro, São Paulo, Brazil; 6LIMAV - Interdisciplinary Laboratory for Advanced Materials, Department of Materials Engineering, UFPI - Federal University of Piauí, Teresina, Piauí, Brazil

5.1 Introduction One of the biggest challenges within the medical practice is the innovation and improvements in technologies for the closure of wounds or sutures [1]. The general technologies comprise physically perforating materials, for example, staples or sutures. These approaches have several limitations and challenges such as the risk of infections [2], cause continues pain, not always effective and in some cases can result in leakage at the site of closure. To overcome these challenges and limitations polymer-based sutures can be employed to hold body tissues together or ligate blood vessels, after a surgery or accidental injury [3]. Depending on the damaged site, specific features are required to withstand the natural conditions of the body, but the utmost property for a suture material is its tensile strength, which can be tailored by the composition and thickness of the yarn. Aside the strength, other important properties to be considered are absorbability, sterility, high knot security, lack of allergic reaction, and ease of handling [4]. In addition to these characteristics of biomaterials, in general, other criteria used for suture selection are based on the properties of the tissues involved, such as the specific healing rate; wound condition and general health of the patient, potential postoperative complications, personal preference and experience of the surgeon, and economic reasons [5,6]. Among the extensive portfolio of materials currently available, synthetic and natural polymers have been the most frequently targeted. *

These authors contributed equally to this work.

Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00004-8

© 2023 Elsevier Ltd. All rights reserved.

95

96

Advanced Technologies and Polymer Materials for Surgical Sutures

Over the past millennia, a huge number of suture materials have been developed by scientists and used by physicians, dentists, and veterinarians [7]. Original sutures were made from biological materials, such as silk and gut (made by twisting together strands of purified collagen), but most modern sutures are synthetic, including absorbable (poly(glycolic acid) (PGA), poly(lactide-co-glycolic acid) (PLGA), poly(lactic acid) (PLA), polycaprolactone (PCL), poly(4-hydroxybutyrate) (P4HB) and polydioxanone (PDO or PDS)) as well as nonabsorbable polymers (polytetrafluoroethylene (PTFE), nylon, poly(ethylene terephthalate) (PET), polypropylene (PP), polybutester (copolymer composed of polyglycol terephthalate and polybutylene terephthalate) and poly(vinylidene fluoride) (PVDF)) (Fig. 5.1) [8]. Additionally, stainless steel has also been used as sutures due to its high tensile strength [9], and applied in abdominal wound closure, intestinal anastomosis, hernia repair, sternal closure, and for certain orthopedic procedures [9,10]. Additionally, very recently, Afewerki et al. disclosed the engineering of multifunctional surgical bactericidal nanofibers with tunable mechanical and biological properties comprising the integrated strategy of combining electrospinning-, plasma treatment, and direct surface modification stratagem [11]. The devised nanofibers were employed in abdominal hernia repair, which showed good biointegration, blood vessel formation, and tissue growth. Furthermore, the nanofibers with their antibacterial

Figure 5.1 Examples of some of the most employed absorbable and nonabsorbable polymers for surgical sutures and their structures.

Polymers for surgical sutures

97

properties could be a good candidate for the treatment of abdominal hernia repair and prevent any future infections [11]. Sutures are generally categorized by their (i) absorption (nonabsorbable or absorbable), (ii) yarn construction (monofilament or multifilament/ twisted/braided), (iii) origin (synthetic or natural), (iv) presence of dye (undyed or dyed to enhance visibility in tissue), (v) presence of coating, to improve biocompatibility or to provide antibacterial property [4], and (vi) thickness, normally called size [7]. Other characteristics that are often considered are the capillarity of the suture, tissue reactivity and the rate of wound healing in the area [12]. Some commercial availably surgical sutures and their properties are summarized and systematic classified based on natural and absorbable, natural and nonabsorbable, synthetic and absorbable, and synthetic and nonabsorbable characteristics in Table 5.1. The yarn can be constructed by a single filament (called monofilament) or by multiple filaments that can be twisted or braided into bundles (Fig. 5.2). The size of the suture is defined by the United States Pharmacopeia (U.S.P.) ranging from #10 (diameter of 1.2 mm) to #12e0 (diameter of 0.001 mm) [7] (Tables 5.2 and 5.3). The diameter for a given U.S.P. size differs depending on the suture origin (natural or synthetic) and absorbability (nonabsorbable or absorbable). Thicker sutures present higher knot-pull tensile strength and are normally used for orthopedics, while thinner sutures are commonly used for ophthalmic. Absorbable materials naturally degrade in the body over time and the byproducts are eliminated by urine. The degradation rate depends on the material and can take days or even months. Many synthetic suture polymers are primarily degraded by hydrolysis of their ester bonds [15]. However, natural polymers, such as collagen and silk fibroin, are degraded by catalyzed proteolysis, that is the breakdown of proteins through the hydrolysis of peptide bonds catalyzed by cellular enzymes called proteases [15]. Typically, the biodegradability of polymer sutures is investigated using in vitro assays in which the absorbable suture is immersed in a medium capable of simulating body fluid characteristics, such as phosphate-buffered saline and HANKs’ balanced salt solution composed of 8.0 g/L sodium chloride (NaCl), 0.4 g/L potassium chloride (KCl), 0.14 g/L calcium chloride (CaCl), 0.1 g/L magnesium chloride (MgCl2), 0.06 g/L magnesium sulfate (MgSO4), 0.06 g/L potassium phosphate monobasic (KH2PO4), 0.06 g/L disodium phosphate (Na2HPO4), 0.35 g/L sodium bicarbonate (NaHCO3), 1.0 g/L glucose (C6H12O6). After a preestablished time interval, the biodegradation behavior can be analyzed by observing the

Table 5.1 Commercially available sutures.

Natural and absorbable

Natural and non-absorbable Synthetic and absorbable

Material

Brand/Company

Yarn construction

Plain gut Plain gut Chromic gut

CP medical Atramat, Internacional Farmacéutica CP medical

Twisted Twisted Twisted

Chromic gut

Atramat, Internacional Farmacéutica

Twisted

Linen

LIN, Peters surgical

Twisted

PLGA (polyglactin 910) PLGA

Braided or monofilament Braided

PLGA with triclosan PLGA

Coated Vicryl, Ethicon ( Johnson & Johnson) Coated Vicryl rapide, Ethicon ( Johnson & Johnson) Coated Vicryl plus, Ethicon ( Johnson & Johnson) Polysorb, Covidien, Medtronic

Braided

PLGA PGA PGA

Velosorb, Covidien, Medtronic Visorb synthetic, CP medical Bondek plus, Teleflex

Braided Braided Braided

PGA PGA-PCL PGA-PCL with triclosan PDS PDS PDS with triclosan P4HB

OPTIME, Peters surgical Atramat, Internacional Farmacéutica MONOCRYL PlusAntibacterial, Ethicon ( Johnson & Johnson) Monodek, Teleflex PDSII, Ethicon ( Johnson & Johnson) PDS plus antibacterial, Ethicon ( Johnson & Johnson) Monomax, B. Braun

Coating

Antibacterial property

‒ ‒ Treated with a chromic salt solution Treated with a chromic salt solution ‒

No No No

No

Braided Monofilament Monofilament

PLGA (polyglactin 370) and calcium stearate PLGA (polyglactin 370) and calcium stearate PLGA (polyglactin 370) and calcium stearate PLGA and calcium stearoyl lactylate PLGA and calcium stearate PCL and calcium stearate Polycaprolactone copolyglycolic acid ‒ ‒ ‒

Monofilament Monofilament Monofilament

‒ ‒ ‒

No No Yes

Monofilament

‒

No

Braided

No No

No Yes No No No No No No Yes

Synthetic and non-absorbable

PTFE PVDF Nylon6,6 Nylon6,6 Nylon6,6

Cytoplast, Osteogenics biomedical PREMIO, Peters surgical Teleflex Trelon, B. Braun Supramid, B. Braun

Nylon6,6 Silk Silk

CARDIONYL, Peters surgical Teleflex Mersilk, Ethicon ( Johnson & Johnson) CP medical Perma Sharp, Hu-Friedy manufacturing company DemeBondTM, DemeTech Corporation Polydek, Teleflex Ethibond Excel, Ethicon ( Johnson & Johnson) NovafilTM, Medtronic Medtronic ACIER, Peters surgical

PP PP PET PET PET Polybutester Stainless steel Stainless steel

Monofilament Monofilament Monofilament Braided Pseudomonofilament Monofilament Braided Braided

‒ ‒ Silicone Polyamide 6

Yes No No No No

‒ Wax Beeswax

No No No

Monofilament Monofilament

‒ ‒

No No

Braided

‒

No

Braided Braided

PTFE ‒

No No

Monofilament Monofilament Monofilament

‒ ‒ ‒

No No No

100

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 5.2 Structure of monofilament and multifilament sutures. SEM micrographs of (A) monofilament PDO suture, (B) braided PGA suture, and (C) twisted PDO suture. (Reproduced with permission Copyright: © 2018 Ercan et al. and Copyright: © 2020 Rashid et al., licensed under A creative Commons Attribution License (CC BY) [13,14].)

changes in the surface morphology of the sutures by scanning electron microscopy and by determining the degradation rate from measuring the suture weight loss during the process [5,16]. Absorbable sutures are normally applied in internal body tissues, with the exceptions of stressful internal environments, such as heart or bladder, where nonabsorbable sutures are normally preferred. Nonabsorbable materials are also commonly used for skin wound closure, where the sutures can be removed after a few weeks [17]. Polymers used for surgical suture are often recognized as foreign materials within the body, trigging a host of immune response and leading to inflammation [4]. In this context, to minimize potential risks to patients, it is essential that the biocompatibility of sutures have to be evaluated. Although it is not possible to generalize which biocompatibility tests should be performed, since the test depends on the material, type of device and, mainly, the application, the International Organization for Standardization (ISO) provides a series of guidelines that can assist in selection of the most appropriate assay. The ISSO 10993, which has the general title of “Biological evaluation of medical devices,” consists of a set of standardized tests

Polymers for surgical sutures

101

Table 5.2 List of absorbable suture size as defined by the United States Pharmacopeia (U.S.P.). Collagen suture

U.S.P. size

12e0 11e0 10e0 9e0 8e0 7e0 6e0 5e0 4e0 3e0 2e0 0 1 2 3 4 5

Diameter range (mm)

0.040e0.049 0.050e0.069 0.070e0.099 0.10e0.149 0.15e0.199 0.20e0.249 0.25e0.339 0.35e0.399 0.40e0.499 0.50e0.599 0.60e0.699 0.70e0.799 0.80e0.899

Knot-pull tensile strength (N)

‒ 0.44 0.69 1.76 3.73 7.55 12.2 19.6 27.2 37.3 44.2 57.8 68.6

Synthetic suture

Diameter range (mm)

Knot-pull tensile strength (N)

0.001e0.009 0.010e0.019 0.020e0.029 0.030e0.039 0.040e0.049 0.050e0.069 0.070e0.099 0.10e0.149 0.15e0.199 0.20e0.249 0.25e0.299 0.30e0.399 0.40e0.499 0.50e0.599 0.60e0.699 0.60e0.699 0.70e0.799

‒ ‒ 0.24 0.49 0.69 1.37 2.45 6.67 9.32 17.4 26.3 38.2 49.8 62.3 71.5 71.5 ‒

to assess biocompatibility that comprise, for example, in vitro assays as tests for cytotoxicity (ISO 10,993e5) [18], and in vivo assays as well (ISO 10993e10) [19]. Usually, multifilament and absorbable suture materials are more reactive than monofilament and nonabsorbable sutures [12]. Additionally, the yarn itself can be a vehicle for bacterial contamination, and therefore, increasing the chances of a surgical site infection [20]. Here again, multifilament sutures are more likely to contribute to the wicking of bacteria and fluids into the wound, due to the capillary action [21]. Although the multifilament sutures present higher tissue reactivity and capillarity, they display better handling characteristics [12]. Furthermore, bioactive materials that can enhance suture function and capability have been at the forefront of suture technology. Alshomer and

102

Advanced Technologies and Polymer Materials for Surgical Sutures

Table 5.3 List of nonabsorbable suture size as defined by the United States Pharmacopeia (U.S.P.). Knot-pull tensile strength (N) U.S.P. size

Diameter range (mm)

Class Ia

Class IIb

Class IIIc

12e0 11e0 10e0 9e0 8e0 7e0 6e0 5e0 4e0 3e0 2e0 0 1 2 3 and 4 5 6 7 8 9 10

0.001e0.009 0.010e0.019 0.020e0.029 0.030e0.039 0.040e0.049 0.050e0.069 0.070e0.099 0.10e0.149 0.15e0.199 0.20e0.249 0.25e0.299 0.30e0.399 0.40e0.499 0.50e0.599 0.60e0.699 0.70e0.799 0.80e0.899 0.90e0.999 1.00e1.099 1.100e1.199 1.200e1.299

0.01 0.06 0.194 0.424 0.59 1.08 1.96 3.92 5.88 9.41 14.1 21.2 26.7 34.5 47.8 60.4 71.4 88.6 ‒ ‒ ‒

‒ 0.05 0.14 0.28 0.39 0.59 1.08 2.26 4.51 6.47 10.0 14.2 17.8 24.9 36.1 ‒ ‒ ‒ ‒ ‒ ‒

0.02 0.20 0.59 0.68 1.08 1.57 2.65 5.30 8.04 13.3 17.6 33.3 46.7 57.8 89.3 112 133 156 178 201 224

Class I: suture composed of silk or synthetic fibers of monofilament, twisted, or braided construction where the coating, if any, does not significantly affect thickness. Class II: suture composed of cotton or linen fibers, or coated natural or synthetic fibers where the coating affects thickness. c Class III: suture composed of metal wire. a

b

coauthors have defined bioactive sutures as “biomaterials that are engineered to have controlled tissue interaction to optimize wound/defect healing, in addition to their essential function in tissue approximation” [4]. Beyond their traditional function, bioactive sutures play a major role as a vessel to host and delivery drugs (e.g., antimicrobial-, anti-inflammatory-, and anesthetics drugs), growth factors (e.g., vascular endothelial growth factor (VEGF), recombinant human growth/differentiation factor-5 (rhGDF-5), and recombinant human platelet-derived growth factor-BB (rhPDGF-BB)), active nanoparticles (silver and bioglass (BG)), peptides (RGD (arginine-glycineaspartic acid) and polylysine), proteins (intracellular adhesion molecule 1

Polymers for surgical sutures

103

(ICAM-1), fibronectin and fibrinogen), and cells (mesenchymal stem cells, osteoblasts, tenocytes, and embryonic stem cells) to traumatic sites [4]. Besides, the need for surgical procedures has increased over the time due to worldwide aging population, which has boosted the search for innovative and high performance materials in suture technology [22]. The market for surgical sutures is dominated by the global leading enterprises like Johnson & Johnson, Medtronic, Covidien, Teleflex, CP Medical, Peters Surgical, DemeTech, Samyang Biopharmaceuticals, B. Braun, Internacional Farmacéutica, among others. Johnson & Johnson has been a pioneer in wound healing, ever since the company created the world’s first mass-produced sterile sutures made of either gut or silk in 1887, and kept innovating with the release of antibacterial sutures containing triclosan in 2003 [23]. The purest form of triclosan (IRGACARE MP) is a broadspectrum antimicrobial agent that prevents bacteria from congregating on the suture, reducing the risk of developing a surgical site infection by almost a third [24].

5.2 Types of polymeric surgical sutures and their applications 5.2.1 Natural polymers 5.2.1.1 Gut Gut, also known as catgut, is made of twisted collagen fibers usually harvested from beef tendon or from the intestine of sheep, cattle or goats [25]. There are two types of gut used for suture: (i) plain gut, composed of collagen slender strands woven together and further precision grounded to form a suture with uniform diameter, and (ii) chromic gut, when treated with a chromic salt solution promote the crosslinking of the collagen fibers. The chromic gut presents reduced tissue reaction, enhanced tensile strength and higher resistance to body enzymes, thus slowing down the absorption process [26]. The two types of sutures naturally degrade in the body catalyzed by proteolysis, with complete absorption after 90 days for chromic suture and 70 days for plain gut suture. Common uses of these suture materials include general closure, ophthalmic, orthopedics, obstetrics/gynecology, gastro-intestinal tract surgery, urology, and bowel anastomosis. When selecting a suture, its tensile strength, knot strength, handling property, and degradation rate should be taken into account, aside the tissue characteristics such as reactivity, wound healing rate and mechanical

104

Advanced Technologies and Polymer Materials for Surgical Sutures

properties [22]. Every specific application has its essential requirements for sutures and will highly depend on the properties of the suture, this kind of information and possible applications of the sutures can be found in the manufacturer instructions. 5.2.1.2 Silk On the contrary of gut, silk is regarded as a nonabsorbable material according to the U.S.P. definition, because complete biodegradation requires approximately 2 years [15,27]. Silk suture is primarily composed of silk fibroin, that is a natural protein produced by the domestic silkworm Bombyx mori, and it is bioinert and relatively inexpensive [28]. Compared to collagen and PLA, silk fibroins have better mechanical properties like strength and toughness [27]. It is commonly coated with wax for easy pull out and applied for skin closure, gastrointestinal, cardiovascular surgery, plastic surgery, ophthalmic, and neurological procedures [29]. Silk suture has high capillarity, and should be avoided in contaminated wounds [12]. In this context, Jo and coauthors have reported the modification of silk sutures with 4-hexylresorcinol (4HR), which is a well-known antiseptic agent, to incorporate antimicrobial property and to achieve biodegradability [15]. In this study, silk sutures containing 12 wt.% of 4HR were compared to untreated silk (Woorhi Medical) and PLGA sutures (coated Vicryl, Ethicon (Johnson & Johnson)). The incorporation of the 4HR increased the expression of matrix metalloproteinase (MMP) in RAW264.7 cells, such as MMP-2, -3, and -9, which can digest a wide spectrum of proteins including silk fibroin. As a result, only 59.5% of the 4HR-silk suture remained after 11 weeks, which was similar to the results obtained for the PLGA degradation (56.4% remained), on the other hand, very different from the residual amount of bare silk suture (91.5% remained). In addition to displaying biodegradation rate similar to PLGA suture, the 4HR-treated silk also exhibited antimicrobial activity against six pathogens (Staphylococcus aureus (S. aureus), Streptococcus sanguinis (S. sanguinis), Actinomyces naeslundii (A. naeslundii), Streptococcus gordnonii (S. gordnonii), Escherichia coli (E. Coli), and Actinomyces odontolyticus (A. odontolyticus)) as evidenced by inhibition zone assay [15]. Moreover, silk sutures have been also modified with peptides [30], BG [31] and silver ions (Agþ) [31,32]. Kardestuncer et al. demonstrated the ability of silk-RGD to stimulate human tenocyte adhesion, proliferation, and differentiation, with the aim of achieving faster and stronger interaction at the tendon to bone interface after tendon reconstruction surgery [30]. Besides, Blaker and coworkers have investigated the

Polymers for surgical sutures

105

use of Agþ containing BG (AgBG) as a coating for silk suture (Mersilk, Ethicon (Johnson & Johnson)) [31]. In vitro assay through immersion in simulated body fluid (SBF) solution confirmed the bioactivity of the AgBG-silk suture, with the formation of bonelike hydroxyapatite after only 3 days of immersion. Another attempt to modify silk suture with Ag þ has been proposed by De Simone et al., who developed an effective and lowcost antibacterial silver coating by implementing an innovative photochemical deposition process [32]. The sutures were dipped in the silver solution and then exposed to UV radiation, which produced silver clusters on the surface of the suture. The silk suture containing Agþ presented significantly inhibited microbial colonization, with reduction of 81% on S. aureus and 78% on E. coli, and only slightly affected fibroblasts viability (82% cell viability compared to 91% of untreated suture) [32]. 5.2.2 Synthetic and absorbable polymers The most explored absorbable polymers for sutures applications are PGA, PCL, and PLA, and their blends. They present less associated tissue inflammation than the silk and plain or chromic catgut [12]. Their mechanical strength and rate of hydrolytic degradation can be controlled by the blend composition and by altering their physical properties, such as their molecular weights, degree of crystallinity and glass transition temperature (Tg) [33]. 5.2.2.1 PGA-PCL blend The combination of PGA with PCL at 75:25 ratio, named poliglecaprone 25 or PGC25, is extensively applied for human and veterinary use, in general as soft tissue approximation and/or ligation [34]. It presents an excellent handling property (flexible and easy to tie), smooth tissue passage, lower incidence of infection and trauma due to smooth surface, higher strength compared to catgut, and total absorbability after 110 days by hydrolysis process. De Lima and coworkers compared the PGC25 (Monocryl, Ethicon (Johnson & Johnson)) with a nonabsorbable suture composed of nylon (Mononylon ETHILON, Ethicon (Johnson & Johnson)) as intradermal suture for skin closure in women undergoing their first cesarean section, which was removed after about 7e10 days [35]. The cesarean is the most frequent surgery in women, and its esthetic outcome is a constant concern [36]. This clinical trial was performed with 60 women undergoing their first cesarean section, and 6 months after the operation the authors took photographs of the scars and evaluated the hypertrophy, color, and

106

Advanced Technologies and Polymer Materials for Surgical Sutures

width [35]. The scars from patients treated with PGA-PCL were significantly less hypertrophic, thinner, and had more acceptable color, demonstrating that the intradermal suture with PGC25 for skin closure after cesarean incision provides better esthetic outcome (Fig. 5.3) [35]. 5.2.2.2 PGA-PLA blend Another important type of suture material is obtained by the combination of PGA with PLA. For example, the coated Vicryl suture (Ethicon (Johnson & Johnson)) is composed of a copolymer made from 90% PGA and 10% PLA, known as polyglactin 910, and coated with polyglactin 370 (copolymer of 30% PGA and 70% PLA) and calcium stearate [37,38]. The coated Vicryl is usually braided, but a monofilament version is also available for use in ophthalmic practice [39e41]. An equivalent material is produced by Medtronic, named Polysorb and composed of 10% PGA and 90% PLA, with a coating of glycolide and ε-caprolactone [42,43]. Both Vicryl and Polysorb decomposes in 56e70 days within the body, and they are

Figure 5.3 (A) Photo of the scar. (B) Value of the hypertrophy and scar staining: GI (Nylon) and GII (PLGA). (C) Values of the scar width for GI (Nylon) and GII (PLG 25). G1: Group 1; GII: Group 2; SD: Standard Deviation; E (1e4): Evaluators. (Reproduced with permission Copyright: © 2020 Lima et al., licensed under A creative Commons Attribution License (CC BY) [35].)

Polymers for surgical sutures

107

designated for soft tissue approximation and ligation applications [43e45]. There is also a faster absorption version of these types of sutures, called coated Vicryl rapide (total absorption in 42 days) and Velosorb (40e50 days). These are indicated for use in soft tissue approximation where only short term wound support is required, such as ophthalmic surgery and skin closure, particularly in pediatric surgery, episiotomies, circumcision, and closure of oral mucosa [46]. Vicryl suture has been extensively modified with different types of coatings to improve the biocompatibility [31,47e49]. Cummings and coworkers have coated Vicryl suture with rhPDGF-BB, using a dip-coating process in rhPDGF-BB solution for 30 min followed by air-drying, to repair tendon injuries, which showed a noticeable increase in tendon tensile strength [47]. A burst release of rhPDGF-BB from the sutures was observed after the first hour of incubation, followed by a continuous and gradual release of growth factor through 48 h. In a similar work, Dines et al. coated Vicryl suture with rhGDF-5/gelatin and demonstrated its beneficial effect on rat tendon fibroblasts [48]. The author also used a dip-coating process to coat the suture, and here approximately 95% of the rhGDF-5 release occurred within 24 h, followed by complete release by 48 h. Another attempt to improve bioactivity was performed by Boccaccini et al., using 45S5 Bioglass coating deposited by a slurry-dipping technique [49]. A stable slurry was prepared by dissolving 47 wt.% Bioglass particles in water, and used to coat the Vicryl suture by immersion during 3 min. Following immersion, the samples were dried at room temperature in a humid atmosphere to avoid microcrack formation on the coating. The adhesion strength of the coating or the release of Bioglass particles was not quantitatively determined; however, the high bioactive character of the composite suture was confirmed by the formation of hydroxyapatite crystals after 7 days of immersion in SBF solution. On the other hand, as vide supra mentioned, Blaker et al. silverdoped bioactive glass powder (AgBG) to coat Vicryl, and confirmed the formation of bonelike hydroxyapatite on the coated suture after only 3 days of immersion in SBF solution [31]. Furthermore, Johnson & Johnson has also launched a version of Vicryl with antibacterial property (coated Vicryl Plus) [50], nevertheless recent findings on triclosan toxicity [51] and triclosan-resistance bacteria have raised potential concerns over the use of this strategy [52]. As an alternative, silver nanoparticles (AgNPs) have been used to coat Vicryl suture through layer-by-layer deposition [53]. The silver nanoparticle solutions were prepared by photo-induced reduction under UV lamp of silver nitrate in

108

Advanced Technologies and Polymer Materials for Surgical Sutures

dilute solution of polymethacrylic acid. The deposition of the nanoparticles layer was intercalated with a poly-diallyl dimethylammonium chloride (PDADMAC) solution to immobilize the AgNPs onto the suture. After 20 deposition of each layer, the coated sutures were allowed to dry overnight. In vitro antibacterial assay against E. coli showed a significant growth inhibition from the silver coated suture. Moreover, immunohistochemistry in the intestinal anastomosis model and burst pressure measurement in healed anastomosis confirmed less inflammatory, cell infiltration and better mechanical properties. Another strategy to impart antibacterial property comprises the use of cefotaxime sodium (CFX-Na), a third generation antibiotic with broad spectrum, but this alternative was tested for PLA suture [54]. Pure PLA fibers fabricated by single-phase electrospinning were compared with PLA/CFX-Na nanofibers obtained by blend electrospinning and with PLA/CFX-Na coreesheath nanofibers fabricated by coaxial electrospinning. For fiber preparation, the PLA was dissolved in 2,2,2-trifluoroethanol (TFE) and the CFX-Na was dissolved in water. For the fabrication of coreesheath nanofibers the solutions were injected separately, while for blend nanofiber suture the CFX-Na solution was mixed with the PLA solution. An in vitro study indicated that CFX-Na release from both the composite nanofibers consisted of a low initial burst release followed by a sustained and slow release over a prolonged period of time. The coreesheath suture exhibited a relatively constant rate of drug release over a much longer duration, due to the presence of drug trapped deep inside the core layer. An inhibition zone experiment showed that both PLA-CFX-Na sutures had favorable antibacterial properties against E. coli and S. aureus when compared to pure commercial PLA suture. Additionally, Obermeier and coworkers have investigated the use of chlorhexidine (the golden standard in oral antiseptics) as an alternative to triclosan [55]. The authors coated a commercial braided suture made of PGA (Gunze) with chlorhexidine, using two types of fatty acids (palmitic or lauric acid) to optimize the drug release. The coating solutions were prepared by dissolving the fatty acids and chlorhexidine in ethanol, and posteriorly the sutures were immersed in these solutions and placed on a thermo-shaker for 2 min at 35  C and 150 rpm and finally dried for at least 2 h. The coated sutures showed an initial fast elution of chlorhexidine and a subsequent continuous slow drug release over 96 h, with 70% release of chlorhexidine carried by lauric acid and 46% release for palmitic acid, both at drug content of 33 mg/cm, showing that the palmitic acid lead to a slower drug release over time (Fig. 5.4A and B). Both samples presented

Polymers for surgical sutures

109

Figure 5.4 Chlorhexidine release from PGA suture using (A) lauric acid and (B) palmitic acid as drug carrier. (C) Diclofenac release from PCL suture using hydrotalcite as carrier. (D) Bupivacaine release from PLGA suture. (E) VEGF release from PDS suture. (A) and (B) Reproduced with permission Copyright: © 2014 Obermeier et al. licensed under A creative Commons Attribution License (CC BY) [55]. (C) Reproduced with permission Copyright: © 2014 Catanzano et al. licensed under Elsevier by Ref. [56]. (D) Reproduced with permission Copyright: © 2012 Weldon et al. licensed under Elsevier by Ref. [57]. (E) Reproduced with permission Copyright: © 2014 Bigalke et al. under Acta Materialia Inc [58].)

high antimicrobial efficacy against S. aureus for up to 5 days, and acceptable cytotoxicity levels [55]. A different drug carrier was reported by Catanzano et al., who used synthetic hydrotalcite (magnesium/aluminum (Mg/Al) hydroxycarbonate) for the sustained delivery of the antiinflammatory drug diclofenac in PCL sutures [56]. Although a significant reduction in tensile strength (breaking stress of 190 MPa, compared to 400 MPa for pure PCL), the PCL-hydrotalcite-diclofenac suture presented controlled release over 55 days (Figure 5.4C), and reduction of inflammatory responses. Another example of a controlled release system was presented by Weldon and coauthors, who fabricated PLGA sutures with the local anesthetic bupivacaine using the electrospinning technique [57]. It was noted that the sutures released their entire drug payload over the course of 12 days, mitigating the need for postoperative opioid analgesics (Fig. 5.4D). 5.2.2.3 P4HB Furthermore, P4HB is a homopolymer of 4HB and presents a chemical structure similar to PGA and PCL, differing only by the number of methylene groups (1, 3, or 5) in the polymer backbone (Fig. 5.1) [59].

110

Advanced Technologies and Polymer Materials for Surgical Sutures

Similar to PGA and PCL, P4HB degrades by hydrolysis of the ester bonds, producing 4HB that is quickly metabolized and eliminated via the Krebs cycle. However, unlike PGA and PCL, P4HB belongs to a diverse class of biopolyesters called polyhydroxyalkanoates (PHAs) that are produced naturally by microorganisms [60]. The Monomax suture (B. Braun) was the first commercial P4HB product to be launched, and it is indicated for closure of the abdominal wall [59]. P4HB sutures are exceptionally strong, retaining approximately 50% of its initial tensile strength after 12 weeks, and substantially degrades in 1 year [61]. Williams et al. have shown that P4HB monofilament are 35% stronger than PDS suture (PDSII, Ethicon (Johnson & Johnson)) and 16% stronger than PP suture (Prolene, Ethicon (Johnson & Johnson)) [59]. In addition, P4HB suture has the highest pliability of any commercially available monofilament absorbable suture, and present excellent knot strength and security [62]. 5.2.2.4 PDS Additionally, PDS is a homopolymer of p-dioxanone, introduced in 1984 by Ethicon (Johnson & Johnson) [61]. PDS is a colorless, crystalline, and biodegradable polyester. It absorbs slowly over a period of 6e7 months, and it is best suited for use in general orthopedic surgery, pediatric cardiovascular surgery, ophthalmic, general, subcuticular, and fascia closure applications. Bigalke et al. coated PDS suture (PDSII, Ethicon (Johnson & Johnson)) with VEGF/PLA blend [58]. For the coating deposition, PLA and VEGF were separately dissolved in chloroform and then mixed together to achieve PLA/VEGF coatings containing 0.1 and 1 mg of VEGF. An in vitro release study showed for the PLA/VEGF-coated suture material with higher VEGF load a 18% release within 5 days, while the lower VEGF loaded-suture presented 9% release within the same period (Fig. 5.4E). The PLA/ VEGF(1 mg)-coated suture lead to improved cell viability in vitro and enhanced angiogenesis and vascularization in vivo [58]. 5.2.3 Synthetic and nonabsorbable polymers 5.2.3.1 Nylon Nylon is a generic designation for a class of polyamides, composed by repeating units linked by amide bonds, similar to the peptide bonds in proteins. Nylon 6,6 (nylon six to six, nylon 6/6 or nylon 66) and nylon six are the two most common for textile and plastic industries [63]. Nylon 6,6 is synthesized by polycondensation of hexamethylenediamine and adipic acid, forming the chemical structure presented in Fig. 5.1. Polyamides

Polymers for surgical sutures

111

sutures are usually composed by Nylon 6,6, producing nonabsorbable, smooth, tough and elastic sutures, that generate minimum tissue reactivity. On the other hand, their knot security is poor, and multiple throws are required to properly close a wound [12]. They are commonly used in both human and veterinary medicine for general and skin closure, cardiovascular, ophthalmic, and neurological procedures. Li and coauthors have modified commercially available nylon sutures (Supramid, B. Braun), consisting of a core of polyamide 6,6 and a sheath of polyamide 6, focusing on improved delivery of growth factors for tendon repair application [64]. The authors swollen the fibers into a methanol/calcium chloride (CaCl2) solution and then freeze-dried to generate micrometer-sized pores in the sheaths, that efficiently loaded rhPDGF (disulfide-linked dimers consisting of two 12.0e13.5 kDa polypeptide chains) using fibrin as a carrier material (Fig. 5.5). PDGF has been successfully used to assist tendon healing due to its ability to promote chemotaxis and mitogenesis of mesenchymal cells, enhancing the collagen organization and vascularity [64]. The PDGF-nylon sutures presented sustained release of the growth factor without compromising their mechanical properties, and supported the proliferation of human mesenchymal stem cells (hMSCs).

Figure 5.5 (A) SEM images of the pristine (a, c) and modified (b, d) nylon sutures. (B) In vitro release of PDGF from the modified nylon suture. (C) Live/dead staining of hMSCs after culture for 72 h on the (a, b) pristine, (c, d) modified, and (e, f) PDGFloaded suture. (Reproduced with permission Copyright: © 2016 Li et al. [64].)

112

Advanced Technologies and Polymer Materials for Surgical Sutures

5.2.3.2 PP Moreover, PP sutures have been used in all surgical branches, especially cardiovascular surgery [65] orthopedics [66], traumatology (tendons) [67], ophthalmology [68] and plastic surgery [69]. They present smooth texture, elasticity, nonporous surface, low tissue reactivity, and no capillary effect, suited for stitches in infected wounds. In comparison to nylon, PP suture has better knot security and pulls smoothly through tissues [12]. 5.2.3.3 PET Additionally, PET sutures, commonly called polyester suture, are commonly used for cardiovascular surgeries [70], general closure [71], ophthalmic [72] and neurological procedures [73]. PET suture has low tissue reactivity, good handling characteristics, high tensile strength, and knot security. It can be uncoated or coated with PTFE or polybutylene. The coating allows for tissue passage with less friction and minimizes capillarity [12]. Yao and coworkers have investigated the use of PET sutures (Ethibond Excel, Ethicon (Johnson & Johnson)) coated with poly-L-lysin, intercellular cell adhesion molecule 1 (ICAM-1) and bone marrowe derived stem cells (BMSCs) for tendon repair [74]. In vivo assays showed a statistically greater load to failure level in repaired tendon of rats with cell seeded sutures compared to controls. 5.2.3.4 Polybutester Polybutester is a newer type of polyester composed of a copolymer of polyglycol terephthalate and polybutylene terephthalate. The commercial suture based on polybutester is known as Novafil (Medtronic) and presents hydrophobicity, elasticity, flexibility, fray resistance, and excellent knot security [75]. Compared to nylon, Novafil is less stiff, has a lower memory, and has greater elasticity, as a result, Novafil is capable to accommodate wound edema, reducing suture marks and cut-throughs [12]. In general, it is employed for soft tissue approximation and/or ligation [76], including use in cardiovascular, skin closure and ophthalmic surgery. Pasternak et al. have coated the Novafil suture with doxycycline, which is a substance associated with the inhibition of MMPs [77]. MMPs normally present high activity around sutures inserted into tendon, resulting in tissue breakdown. The coating was applied using plasma treatment, followed by incubation with fibrinogen and doxycycline. The Achilles tendon of rats treated with polybutester-doxycycline suture showed improved suture holding capacity and force of failure [77].

Polymers for surgical sutures

113

5.2.3.5 PVDF and PTFE Furthermore, polymers containing fluorine atoms, such as PVDF and PTFE, have also been used to produce sutures [78]. They are physiological inert, soft, smooth, with excellent knotting properties, easy to handle, and do not present capillary effect [79]. PVDF is used in all surgical branches, including cardiovascular surgery, orthopedics [80], traumatology (tendons) [81], plastic [82] and ophthalmologic surgery [83]. PTFE is commonly used for dental bone grafting and implant procedures where a soft monofilament suture is desirable. PTFE has the advantage of preventing bacterial wicking into surgical sites [84].

5.3 Tissue adhesive polymers as suture candidate Another class of next-generation materials for various suture applications overcoming some of the limitations with traditional staples and sutures are surgical glues or adhesives [85]. These polymers offers great advantages such as being easy to use, able to prevent leakage of fluids, facile application, no requirements for removal, avoiding needlestick injury, and minimal tissue damages [86]. Within this framework, some traditional tissue glues used in clinics are for instance, fibrin sealant (e.g., Tissel) [87], cyanoacrylate based glues (e.g., Histoacryl and Dermabond) [88] and protein based glues (e.g., BioGlue) [89]. However, limitation such as lack of controllable practicability, challenges in their use for minimally invasive procedures, lack on demand activation and controllable adhesion properties have been observed for these clinically approved surgical glues [90]. Therefore, scientist have put dedication and effort to advance these needs and overcome some of the challenges highlighted above [91]. In this context, Annabi et al. disclosed the employment of highly elastic human protein based sealant comprising of the light sensitive methacryloyl modified tropoelastin (MeTro) [92]. The material demonstrated successful in vivo lung sealing in rat models with low toxicity and controllable degradation. Moreover, the crosslinking could be controlled by the light activation, thus simplify its translational and practical application. Here, Lang et al. also disclosed a hydrophobic light-activated adhesive for minimally invasive repair of vessels and heart defects [93]. The poly(glycerol sebacate urethane) based polymer was converted to a patch after crosslinking, which demonstrated strong wet adhesion (after 5 s of UV light exposure) with about 275% stronger adhesion than fibrin sealant. In vivo experiments on a beating heart of a pig (onto the interventricular septum) and carotid artery defect demonstrated the superior

114

Advanced Technologies and Polymer Materials for Surgical Sutures

performance of the adhesive material under highly dynamic and wet environment. Recently we have seen the increase interest in bioinspired tissue adhesives, in particular, the mussel-inspired wet adhesion [94,95]. In this context, Mehdizadeh et al. devised citrate based strong wet bioadhesive based on the mussel mimicry strategy [96]. The adhesive polymer was designed by the combination of citric acid, polyethylene glycol (PEG) and dopamine. Through a 2-component injection procedure, the polymer and the oxidizing solution (sodium periodate), the crosslinking of the polymers could be promoted, which provided successful sutureless wound closure. In vitro adhesion tests comparing the devised adhesive material with the clinical employed fibrin glue demonstrated 2.5e8.0-fold stronger wet tissue adhesion strength. Moreover Liu et al. presented a moldable bioadhesive made of nanosilicate laponite [97] and dopamine modified-PEG [98]. Due to the ability of the composite to undergo auto-oxidation rendered from the dopamine moiety, the material initially underwent reversible crosslinked network and eventually a more compact gel through covalent bonding. These unique abilities presented a material that could fit as moldable sealant to any shape, besides the nanocomposite hydrogel could be injected through a syringe, simplifying its application [98]. In 2018, a study comparing the cyanoacrylate glue performance with conventional sutures in the closure of inguinal hernia skin incisions on randomized control trials was reported [99]. The authors concluded that the tissue adhesive was superior, while both the procedures presented similar safety. Another study compared Histoacryl with suture in the repair of knee meniscal tears, where the biomechanical evaluation demonstrated a better performance from the adhesive material [100]. However, despite that a large number of reports have been disclosed presenting a huge number of new adhesive biomaterials and their proof of concept applications, very few have ended up as clinical products [101]. Hence, overcoming the translational barriers of tissue adhesives is one of the greatest challenges within this research field. Here, Taboada and coauthors, very recently presented a beautiful review on how to overcome the translational barriers of tissue adhesives [102]. The authors highlighted some important aspects to consider when designing translational adhesive materials such as clearly understanding the target tissue surface and environment, the long-term performance, and consideration of the regulatory and development pathways at early stage [102].

Polymers for surgical sutures

115

5.4 Challenges with current technologies Surgical sutures play a crucial role in the success of surgical treatment, and the increase in the number of surgical procedures performed worldwide has led to a consequent increase in the demand for better suture materials [103]. The search for a perfect, ideal suture material has been ongoing for decades and, in all likelihood, will continue in the future, since the current technologies employed in rejoining injured tissue after surgery such as surgical sutures and staples encounter several challenges and limitations [1]. In addition to conceivably inducing damages in the surrounding tissue of the surgery site and some cosmetic challenges, there are potential risk for infections [13] and leakage that could be devasting and cause substantial problems [104]. Recently, Ananda et al. reported a comparative study between the use of skin suture, staples, and adhesive glue for surgical skin closure [105]. The authors concluded that staples were the fastest option, while the adhesive glue provided the best outcomes with regard to less postoperative pain, improved cosmetics, and more cost-effective. Although recent advances have increased the effectiveness of sutures, most of the progress can be attributed to technological advances focused on the field of materials science, especially polymeric sutures [5]. Indeed, polymers, natural or synthetic, absorbable or nonabsorbable, have significant potential and, over the years, through the improvement of materials, such as changes in composition, surface alteration and polymer blend, several sutures have been created with excellent physical and mechanical properties [106]. Interestingly, it is undeniable that the sutures currently available for clinical use has been evolving significantly regarding the manufacturing techniques and applicability; nevertheless, little has been advanced to increase the therapeutic properties of the suture itself [4]. 5.4.1 Bioactive sutures Current efforts are focused on the development of suture materials that have improved mechanical properties, but with additional features, emphasizing biologically active sutures. Bioactive materials that can enhance suture function and capability have been at the forefront of suture technology nowadays [4]. Therefore, the strategy of sutures developed as a yarn of biocompatible material only to mechanically bring the tissues together is over, but instead a multifunctionality approach is of great interest [107]. Moreover, antibacterial sutures have been a historic milestone in the development of novel sutures with additional features [108]. After several

116

Advanced Technologies and Polymer Materials for Surgical Sutures

years of research and development, the first antibacterial suture available for clinical use, VICRYL Plus by Johnson & Johnson (coated polyglactin 910 with triclosan) was approved in 2002 by the Food and Drug Administration (FDA), helping to reduce the risk of infections at the surgical site [4]. However, even though antimicrobial effectiveness has been extensively researched including other methods such as the incorporation of agents into the suture (e.g., chlorhexidine and octenidine) and AgNPs treated sutures, the applications of bioactive sutures are not limited to antimicrobial activity [53,109‒112]. The drug release from the suture (drug delivery suture) can be used to deliver a high drug concentration at the wound area from a wide variety of drugs with potential anesthetic, antiinflammatory, and antineoplastic activity [4,57,113,114]. However, there are several challenges that need to be overcome with drug eluting sutures in order to make them translational and sustainable, such as controlled and sustained drug release, thus avoiding burst release and toxicity to the tissues [114,115]. Prior work on the delivery of bioactive growth factors, the delivery has mainly been focused on coating the surface of a suture with the bioactive compound. Nevertheless, disadvantages with this strategy includes, the limited number of bioactive agents that can be loaded into the suture, which is restricted to a thin coating layer [116]. Further limitation is the quick release (burst release) of a large portion of the agents within the first few hours after implantation [64]. In order to obtain a sustained and controlled release, a recent strategy has been the use of carriers, such as inorganic clays (magnesium and aluminum hydroxycarbonates) [56,117], fatty acid [55], and fibrin [64]. Moreover, the tissue engineering and regenerative medicine strategies have contributed to further transforming the vision of the suture from solely a yarn of biocompatible material, to a biologically active medical device with the aim of not only being a material to mechanically close the sutured wound and prevent infections [110]. For this purpose, cells have been incorporated within the suture material [118]. The main objective of cell seeded biological sutures is to increase the number of healthy cells at the injured site to accelerate the tissue regeneration and repair [119]. Although several types of cells have been evaluated (e.g., osteoblasts and tenocytes) [4], the stem cells have been highlighted for this application and have shown great potential, both pluripotent embryonic cells [120], adiposederived stem cells [121], and mesenchymal stem cells [122]. However, some challenges regarding the use of this technology with current methods

Polymers for surgical sutures

117

remain unresolved. Among the main limitations are the low rate of cell retention at the site, the challenges in targeting cells to a specific region, the time required to expand a cell population and problems associated with genetic mutations of the cells during culture [110]. 5.4.2 Smart sutures Another considerably promising advancement within the area of sutures is the development of smart sutures [123]. This new class of sutures is based on responsive polymers capable of significantly altering their properties under small physical or chemical stimuli [124]. Through shape memory selfknotting and tightening sutures, this type of suture can allow suturing of difficult tissues and wounds, where access is strictly limited. In this context, Lendlein and Langer were the first to introduce the concept of shape memory polymer in sutures applications [125]. The authors developed a smart degradable polyurethane suture that underwent spatial transformation according to temperature. The material had a temporary shape below a critical temperature and acquired a permanent shape at a higher temperature. Thus, after the suture was applied to the site and the temperature increased, the suture material decreased in size, creating a knot with adequate tension in the surrounding tissue [125]. Other intelligent sutures were developed following these principles, in which the suture can be loosely connected in wounds with its temporary shape, and with appropriate stimuli (such as heat, light, solution and applied magnetics or electric field) the suture recovers its original state and forming the knot automatically (selftightening knots) [106,126]. In addition to conformational changes, smart sutures can be used for controlled release of drugs and bioactive agents. Both exogenous stimuli (magnetic fields, ultrasound, electric and light fields) and endogenous stimuli (such as pH, temperature and mechanical load) can be used to control the release of drugs from bioactive sutures for each patient [110]. However, although this study demonstrates that smart sutures have the potential to change when stimulated by physiologically relevant environments, the generation of clinically useful and biologically safe materials still remains a challenge [127]. 5.4.3 Biomimetic sutures Since an ideal bioactive suture should stimulate a regenerative response in the tissue, it is crucial to consider the native tissue environment in which the suture will be applied. The cells within the tissue detect and

118

Advanced Technologies and Polymer Materials for Surgical Sutures

respond to the nanoarchitecture in the native extracellular matrix (ECM) and, therefore, biomimetic materials that resemble the original ECM architecture is also another strategy inherited from tissue engineering concept, which has been applied in the development of novel sutures [110,128]. Here, a trend in tissue engineering is the application of nanotechnology to produce biomimetic scaffolds with dimensions at the nanoscale that is the same scale as the native ECM. Scaffolds composed of nanofibers have high porosity, high surface/volume ratio, promote better cell adhesion and proliferation and facilitate the transport of nutrients and oxygen during regeneration. In addition, the fibers are at the same scale as the size of the ECM components, allowing to simulate the original environment and allowing the cells to behave similarly to native tissue cells [129,130]. In this context, the electrospinning has proven to be a powerful tool for the manufacturing of polymeric nanofibers for tissue engineering application, since it is a simple, low-cost, versatile method capable of forming nanostructured scaffolds [131e134]. Despite these advantages of the electrospun nanofibers, little has been related to their applications as tissue sutures. This is because the nanofibers obtained through electrospinning are generally in nonwoven form, which leads to poor mechanical performance [135]. However, electrospinning does have several unique advantages in providing mechanical functionality and for the production of composite nanofibers, unlike any other processing technology [133,136]. 5.4.4 Translation of basic discoveries in clinical applications Another challenge with current technologies of sutures is their clinical evaluation. There is a need to conduct detailed preclinical studies and evaluate the long-term safety and efficacy in human trials on these emerging sutures, in particularly with novel materials. Moreover, a recent review revealed that many of the sutures currently in use, even though they have been available for decades, have never been clinically evaluated [110]. Furthermore, the regulatory question is also an important challenge to overcome regarding new suture material technologies [137]. Most suture research and development efforts have been focusing on modifying sutures made of materials already approved by the FDA, for example by modifying the surface or by different combinations of these materials. This strategy is probably employed in order to smoothly simplify and promote the fast approval of the new suture materials. Hence, the development of biomaterial sutures completely different from those commercially available and

Polymers for surgical sutures

119

approved by the FDA has been very rare. One reason as mentioned above could be that completely new suture materials that do not have a substantially equivalent predecessor approved by the FDA can be considered as a Class III device, and in that case, a Premarket Approval would be required. The purpose of a PMA is to provide adequate safety and efficacy information for a new material, which requires extensive preclinical and clinical testing, increasing the regulatory burden and cost, and resulting in longer development times [138]. In the case of bioactive sutures containing pharmacological substances, they are automatically classified as a high risk Class III medical device independent of the primary suture material, which also adding extra complexity to the regulatory approval process [139].

5.5 Future perspective and remarks One of the aims in the field of polymeric surgical sutures is to advance materials design providing polymers with the desired properties (e.g., biocompatible, biodegradable, high performer and stable), thus simplifying its use and at the same time promote its mission. This would provide surgeon with easy to use and conduct technologies which are safer and less stress to the tissue, thus avoiding current invasive technologies (e.g., sutures, staples or clips) [140]. In the context of simplicity, an in situ deposition approach of nanofibers via solution blow spinning has been presented [141]. The ease of the technology was demonstrated in the direct deposition in various surgical piglet models such as lung resection, intestinal anastomosis, liver injury, and hernia. These experiments showed successful blocking of the bleeding and leakage in the injury and the quick formation of the protective fiber layer (less than 1 min). Moreover, the development of novel materials has also led to the invention of polymers with the ability to promote healing of tissues and regeneration [142]. Further, remarks on future ideal surgical polymer, they should display multifunctionality and also overcome current challenges and limitation such as potential microbial infection [143], fluid leakage, poor cosmetic, poor healing, and ideally “one fits all,” thus suitable for a wide range of surgical applications and in minimally invasive procedure. All these characteristics should ideally be incorporated without trade-off any of the vital properties (biocompatible, stable, sufficient mechanical property etc.). We believe that the future holds great promise in the advancement and invention of the perfect polymeric

120

Advanced Technologies and Polymer Materials for Surgical Sutures

surgical sutures; nevertheless, to turn the vision into reality, a better understanding of the tissue, its microenvironment and behavior is vital. A couple of examples have already been presented in this chapter, where smart materials (ability to respond to physiological and external stimuli and change their properties) are believed to play a crucial role in this quest providing improved treatments [126]. Despite, that we have seen the increase inventions the last decades, of new polymeric materials as surgical sutures, little is known about their long-term performance and stability. Therefore, we should not rush into translating novel discoveries from basic research into clinics without understanding its long-term safety, biocompatible and stability, in particularly with novel chemistries and structures. Nevertheless, the advancement can also be elevated where new polymeric smart sutures with the ability to monitor the healing and detect potential defects in the injured environment and surrounding will be designed, thus preventing any future infections or failure.

5.6 Conclusion Natural and synthetic polymers have been used as surgical suture to hold body tissues together or ligate blood vessels, after a surgery or accidental injury. Among the large portfolio of biomaterials, synthetic polymers such as PGA, PLA, PCL, P4HB, and PDO are currently the most employed as absorbable suture, and synthetic polymers including nylon, PP, PET, polybutester, PVDF, and PTFE as nonabsorbable suture. The development of bioactive sutures enabled to enhance the biocompatibility with tissues and also to display additional functions such as antimicrobial, antiinflammatory, and anesthetics properties. Another improvement was achieved by the use of smart polymers capable of significantly altering their properties under small physical or chemical stimuli. Recently, sutures have been replaced by polymeric based tissue glues, which are easier to use and provides minimal tissue damage. Although bioactive, smart, and adhesive polymers are promising for the field, there are still some challenges for clinical application.

Acknowledgments Dr. Afewerki gratefully acknowledges the financial support from the Sweden-America Foundation (The Family Mix Entrepreneur Foundation) and Olle Engkvist Byggmästare Foundation.

Polymers for surgical sutures

121

References [1] N. Annabi, A. Tamayol, S.R. Shin, A.M. Ghaemmaghami, N.A. Peppas, A. Khademhosseini, Surgical materials: current challenges and nano-enabled solutions, Nano Today 9 (5) (2014) 574e589. [2] S. Afewerki, N. Bassous, S. Harb, C. Palo-Nieto, G.U. Ruiz-Esparza, F.R. Marciano, T.J. Webster, A.S.A. Furtado, A.O. Lobo, Advances in dual functional antimicrobial and osteoinductive biomaterials for orthopaedic applications, Nanomedicine 24 (2020) 102143. [3] A. Lauto, D. Mawad, L.J.R. Foster, Adhesive biomaterials for tissue reconstruction, J. Chem. Technol. Biotechnol. 83 (4) (2008) 464e472. [4] F. Alshomer, A. Madhavan, O. Pathan, W. Song, Bioactive sutures: a review of advances in surgical suture functionalisation, Curr. Med. Chem. 24 (2) (2017) 215e223. [5] C.K.S. Pillai, C.P. Sharma, Absorbable polymeric surgical sutures: chemistry, production, properties, biodegradability, and performance, J. Biomater. Appl. 25 (4) (2010) 291e366. [6] A.J. Dart, C.M. Dart, Suture material: conventional and stimuli responsive, Compr. Biomater. (2011) 573e587. [7] J. Meyle, Suture materials and suture techniques, Periodontal Pract. Today 3 (4) (2006) 253e268. [8] C.C. Chu, Types and properties of surgical sutures, in: M.W. King, B.S. Gupta, R. Guidoin (Eds.), Biotextiles as Medical Implants, Woodhead Publishing, 2013, pp. 231e273. [9] R. Edlich, D.B. Drake, G.T. Rodeheaver, K.L. Winters, J.A. Greene, W.B. Long III, L. Britt, S.P. Winters, C.C. Scott, K.Y. Lin, SynetureÔ stainless steel suture. A collective review of its performance in surgical wound closure, J. Long-term Eff. Med. Implants 16 (1) (2006) 101e110. [10] S. Najibi, R. Banglmeier, J. Matta, M. Tannast, Material properties of common suture materials in orthopaedic surgery, Iowa Orthop. J. 30 (2010) 84e88. [11] S. Afewerki, et al., Engineering multifunctional bactericidal nanofibers for abdominal hernia repair, Commun. Biol. 4 (2021) 233. [12] A.J. Tsugawa, F.J.M. Verstraete, Suture materials and biomaterials, in: F.J.M. Verstraete, M.J. Lommer (Eds.), Oral and Maxillofacial Surgery in Dogs and Cats, W.B. Saunders, Oxford, 2012, pp. 69e78. [13] U.K. Ercan, F. I_bis, C. Dikyol, N. Horzum, O. Karaman, Ç. Yıldırım, E. Çukur, E.A. Demirci, Prevention of bacterial colonization on non-thermal atmospheric plasma treated surgical sutures for control and prevention of surgical site infections, PLoS One 13 (9) (2018) e0202703. [14] M. Rashid, J. Dudhia, S.G. Dakin, S. Snelling, A. Lach, R. De Godoy, P.A. Mouthuy, R. Smith, M. Morrey, A.J. Carr, Histological evaluation of cellular response to a multifilament electrospun suture for tendon repair, PLoS One 15 (6) (2020) e0234982. [15] Y.-Y. Jo, H. Kweon, D.-W. Kim, M.-K. Kim, S.-G. Kim, J.-Y. Kim, W.-S. Chae, S.-P. Hong, Y.-H. Park, S.Y. Lee, Accelerated biodegradation of silk sutures through matrix metalloproteinase activation by incorporating 4-hexylresorcinol, Sci. Rep. 7 (2017) 42441. [16] S. Liu, G. Wu, X. Chen, X. Zhang, J. Yu, M. Liu, Y. Zhang, P. Wang, Degradation behavior in vitro of carbon nanotubes (CNTs)/poly (lactic acid)(PLA) composite suture, Polymers 11 (6) (2019) 1015. [17] L. Al-Mubarak, M. Al-Haddab, Cutaneous wound closure materials: an overview and update, J. Cutan. Aesthetic Surg. 6 (4) (2013) 178e188.

122

Advanced Technologies and Polymer Materials for Surgical Sutures

[18] International Organization for Standardization (ISO), Biological Evaluation of Medical Devices d Part 5: Tests for in Vitro Cytotoxicity, 2009, p. 34. [19] International Organization for Standardization (ISO), Biological Evaluation of Medical Devices d Part 10: Tests for Irritation and Skin Sensitization, 2010, p. 67. [20] B.D. Masini, D.J. Stinner, S.M. Waterman, J.C. Wenke, Bacterial adherence to suture materials, J. Surg. Educ. 68 (2) (2011) 101e104. [21] T. Grigg, F. Liewehr, W. Patton, T. Buxton, J. McPherson, Effect of the wicking behavior of multifilament sutures, J. Endod. 30 (9) (2004) 649e652. [22] J.A. Greenberg, R.M. Clark, Advances in suture material for obstetric and gynecologic surgery, Rev. Obstet. Gynecol. 2 (3) (2009) 146e158. [23] N.P. Brenwald, A.P. Fraise, Triclosan resistance in methicillin-resistant Staphylococcus aureus (MRSA), J. Hosp. Infect. 55 (2) (2003) 141e144. [24] H.M. Proskin, Advanced oral antibacterial/anti-inflammatory technology: a comprehensive review of the clinical benefits of a triclosan/copolymer/fluoride dentifrice, J. Clin. Dent. 16 (2005) S1eS20. [25] S. Rajendran, S. Anand, A. Rigby, Textiles for Healthcare and Medical Applications, Handbook of Technical Textiles, Elsevier, 2016, pp. 135e168. [26] C.C. Chu, Materials for absorbable and nonabsorbable surgical sutures, in: M.W. King, B.S. Gupta, R. Guidoin (Eds.), Biotextiles as Medical Implants, Woodhead Publishing, 2013, pp. 275e334. [27] Y. Cao, B. Wang, Biodegradation of silk biomaterials, Int. J. Mol. Sci. 10 (4) (2009) 1514e1524. [28] L.-D. Koh, Y. Cheng, C.-P. Teng, Y.-W. Khin, X.-J. Loh, S.-Y. Tee, M. Low, E. Ye, H.-D. Yu, Y.-W. Zhang, M.-Y. Han, Structures, mechanical properties and applications of silk fibroin materials, Prog. Polym. Sci. 46 (2015) 86e110. [29] A.S. Hockenberger, E. Karaca, Investigating fracture mechanisms of some nonabsorbable sutures in vivo, in: S.C. Anand, J.F. Kennedy, M. Miraftab, S. Rajendran (Eds.), Medical and Healthcare Textiles, Woodhead Publishing, 2010, pp. 430e436. [30] T. Kardestuncer, M. McCarthy, V. Karageorgiou, D. Kaplan, G. Gronowicz, RGDtethered silk substrate stimulates the differentiation of human tendon cells, Clin. Orthop. Relat. Res. 448 (2006) 234e239. [31] J. Blaker, S. Nazhat, A. Boccaccini, Development and characterisation of silver-doped bioactive glass-coated sutures for tissue engineering and wound healing applications, Biomaterials 25 (7‒8) (2004) 1319e1329. [32] S. De Simone, A. Gallo, F. Paladini, A. Sannino, M. Pollini, Development of silver nano-coatings on silk sutures as a novel approach against surgical infections, J. Mater. Sci. Mater. Med. 25 (9) (2014) 2205e2214. [33] J. Loo, C. Ooi, F. Boey, Degradation of poly (lactide-co-glycolide)(PLGA) and poly (L-lactide)(PLLA) by electron beam radiation, Biomaterials 26 (12) (2005) 1359e1367. [34] X. Ming, M. Nichols, S. Rothenburger, In vivo antibacterial efficacy of MONOCRYL plus antibacterial suture (Poliglecaprone 25 with triclosan), Surg. Infect. 8 (2) (2007) 209e214. [35] R.J.d. Lima, T.B. Schnaider, A.M.C. Francisco, D. FrancescatoVeiga, Absorbable suture. Best aesthetic outcome in cesarian scar, Acta Cir. Bras. 33 (11) (2018) 1027e1036. [36] J.G. Hofmeyr, N. Novikova, M. Mathai, A. Shah, Techniques for cesarean section, Am. J. Obstet. Gynecol. 201 (5) (2009) 431e444. [37] J. Conn, R. Oyasu, M. Welsh, J.M. Beal, Vicryl (polyglactin 910) synthetic absorbable sutures, Am. J. Surg. 128 (1) (1974) 19e23.

Polymers for surgical sutures

123

[38] H.R. Ford, P. Jones, B. Gaines, K. Reblock, D.L. Simpkins, Intraoperative handling and wound healing: controlled clinical trial comparing coated VICRYL® Plus antibacterial suture (coated polyglactin 910 suture with triclosan) with Coated VICRYL® suture (coated polyglactin 910 suture), Surg. Infect. 6 (3) (2005) 313e321. [39] J. Dhom, D.A. Bloes, A. Peschel, U.K. Hofmann, Bacterial adhesion to suture material in a contaminated wound model: comparison of monofilament, braided, and barbed sutures, J. Orthop. Res. 35 (4) (2017) 925e933. [40] B.K. Huh, B.H. Kim, S.-N. Kim, C.G. Park, S.H. Lee, K.R. Kim, C.Y. Heo, Y.B. Choy, Surgical suture braided with a diclofenac-loaded strand of poly (lactic-coglycolic acid) for local, sustained pain mitigation, Mater. Sci. Eng. C 79 (2017) 209e215. [41] C. Munton, C. Phillips, B. Martin, R. Bartholomew, I. Capperauld, Vicryl (Polyglactin 910): a new synthetic absorbable suture in ophthalmic surgery. A preliminary study, Br. J. Ophthalmol. 58 (11) (1974) 941e947. [42] L. Cong, C. Li, B. Wei, L. Zhan, W. Wang, Y. Xu, V-LocÔ 180 suture in total laparoscopic hysterectomy: a retrospective study comparing Polysorb to barbed suture used for vaginal cuff closure, Eur. J. Obstet. Gynecol. Reprod. Biol. 207 (2016) 18e22. [43] E. Debus, D. Geiger, M. Sailer, J. Ederer, A. Thiede, Physical, biological and handling characteristics of surgical suture material: a comparison of four different multifilament absorbable sutures, Eur. Surg. Res. 29 (1) (1997) 52e61. [44] V. Zhukovskii, Problems and prospects for development and production of surgical suture materials, Fibre Chem. 40 (3) (2008) 208e216. [45] M. Castillo, Suture and Needles Used in Surgery, Mental Conditioning to Perform Common Operations in General Surgery Training, Springer, 2020, pp. 357e367. [46] M.M. Al-Qattan, Vicryl Rapide® versus Vicryl® suture in skin closure of the hand in children: a randomized prospective study, J. Hand Surg. 30 (1) (2005) 90e91. [47] S.H. Cummings, D.A. Grande, C.K. Hee, H.K. Kestler, C.M. Roden, N.V. Shah, P. Razzano, D.M. Dines, N.O. Chahine, J.S. Dines, Effect of recombinant human platelet-derived growth factor-BB-coated sutures on Achilles tendon healing in a rat model: a histological and biomechanical study, J. Tissue Eng. 3 (1) (2012), 2041731412453577. [48] J.S. Dines, M.B. Cross, D. Dines, C. Pantazopoulos, H.J. Kim, P. Razzano, D. Grande, In vitro analysis of an rhGDF-5 suture coating process and the effects of rhGDF-5 on rat tendon fibroblasts, Growth Factors 29 (1) (2011) 1e7. [49] A.R. Boccaccini, A.G. Stamboulis, A. Rashid, J.A. Roether, Composite surgical sutures with bioactive glass coating, J. Biomed. Mater. Res. B Appl. Biomater. 67 (1) (2003) 618e626. [50] T.A. Barbolt, Chemistry and safety of triclosan, and its use as an antimicrobial coating on Coated VICRYL* Plus Antibacterial Suture (coated polyglactin 910 suture with triclosan), Surg. Infect. 3 (S1) (2002) s45es53. [51] G.S. Dhillon, S. Kaur, R. Pulicharla, S.K. Brar, M. Cledón, M. Verma, R.Y. Surampalli, Triclosan: current status, occurrence, environmental risks and bioaccumulation potential, Int. J. Environ. Res. Publ. Health 12 (5) (2015) 5657e5684. [52] S.P. Yazdankhah, A.A. Scheie, E.A. Høiby, B.-T. Lunestad, E. Heir, T.Ø. Fotland, K. Naterstad, H. Kruse, Triclosan and antimicrobial resistance in bacteria: an overview, Microb. Drug Res. 12 (2) (2006) 83e90. [53] S. Zhang, X. Liu, H. Wang, J. Peng, K.K. Wong, Silver nanoparticle-coated suture effectively reduces inflammation and improves mechanical strength at intestinal anastomosis in mice, J. Pediatr. Surg. 49 (4) (2014) 606e613. [54] W. Hu, Z.-M. Huang, X.-Y. Liu, Development of braided drug-loaded nanofiber sutures, Nanotechnology 21 (31) (2010) 315104.

124

Advanced Technologies and Polymer Materials for Surgical Sutures

[55] A. Obermeier, J. Schneider, S. Wehner, F.D. Matl, M. Schieker, R. von EisenhartRothe, A. Stemberger, R. Burgkart, Novel high efficient coatings for anti-microbial surgical sutures using chlorhexidine in fatty acid slow-release carrier systems, PLoS One 9 (7) (2014) e101426. [56] O. Catanzano, S. Acierno, P. Russo, M. Cervasio, M.D.B. De Caro, A. Bolognese, G. Sammartino, L. Califano, G. Marenzi, A. Calignano, Melt-spun bioactive sutures containing nanohybrids for local delivery of anti-inflammatory drugs, Mater. Sci. Eng. C 43 (2014) 300e309. [57] C.B. Weldon, J.H. Tsui, S.A. Shankarappa, V.T. Nguyen, M. Ma, D.G. Anderson, D.S. Kohane, Electrospun drug-eluting sutures for local anesthesia, J. Contr. Release 161 (3) (2012) 903e909. [58] C. Bigalke, F. Luderer, K. Wulf, T. Storm, M. Löbler, D. Arbeiter, B.M. Rau, H. Nizze, B. Vollmar, K.-P. Schmitz, VEGF-releasing suture material for enhancement of vascularization: development, in vitro and in vivo study, Acta Biomater. 10 (12) (2014) 5081e5089. [59] S.F. Williams, S. Rizk, D.P. Martin, Poly-4-hydroxybutyrate (P4HB): a new generation of resorbable medical devices for tissue repair and regeneration, Biomed. Technol. 58 (5) (2013) 439e452. [60] P. Suriyamongkol, R. Weselake, S. Narine, M. Moloney, S. Shah, Biotechnological approaches for the production of polyhydroxyalkanoates in microorganisms and plantsda review, Biotechnol. Adv. 25 (2) (2007) 148e175. [61] E.K. Odermatt, L. Funk, R. Bargon, D.P. Martin, S. Rizk, S.F. Williams, MonoMax suture: a new long-term absorbable monofilament suture made from poly-4hydroxybutyrate, Int. J. Polym. Sci. 2012 (2012). [62] S.F. Williams, D.P. Martin, A.C. Moses, The history of GalaFLEX P4HB scaffold, Aesthetic Surg. J. 36 (Suppl. l_2) (2016) S33eS42. [63] R. Shamey, K. Sinha, A review of degradation of nylon 6. 6 as a result of exposure to environmental conditions, Rev. Prog. Coloration Relat. Top. 33 (1) (2003) 93e107. [64] J. Li, S.W. Linderman, C. Zhu, H. Liu, S. Thomopoulos, Y. Xia, Surgical sutures with porous sheaths for the sustained release of growth factors, Adv. Mater. 28 (23) (2016) 4620e4624. [65] D.V. Feliciano, Cardiac, great vessel, and pulmonary injuries, in: Rich’s Vascular Trauma, 3e Ed., Elsevier, 2016, pp. 71e99. [66] R. Vaishya, A.K. Agarwal, M. Tiwari, A. Vaish, V. Vijay, Y. Nigam, Medical textiles in orthopedics: an overview, J. Clin. Orthop. Trauma 9 (2018) S26eS33. [67] M.C. Jordan, S. Boelch, H. Jansen, R.H. Meffert, S. Hoelscher-Doht, Does plastic suture deformation induce gapping after tendon repair? A biomechanical comparison of different suture materials, J. Biomech. 49 (13) (2016) 2607e2612. [68] T. John, S. Tighe, O. Hashem, H. Sheha, New use of 8-0 polypropylene suture for four-point scleral fixation of secondary intraocular lenses, J. Cataract Refract. Surg. 44 (12) (2018) 1421e1425. [69] S. Lee, N. Isse, Barbed polypropylene sutures for midface elevation, Arch. Facial Plast. Surg. 7 (2005) 55e61. [70] D. Bezuidenhout, P. Zilla, Vascular surgery: biomaterials, in: K.H.J. Buschow, R.W. Cahn, M.C. Flemings, B. Ilschner, E.J. Kramer, S. Mahajan, P. Veyssière (Eds.), Encyclopedia of Materials: Science and Technology, Elsevier, Oxford, 2001, pp. 9513e9518. [71] A.L. Tajirian, D.J. Goldberg, A review of sutures and other skin closure materials, J. Cosmet. Laser Ther. 12 (6) (2010) 296e302. [72] A. Metzger, Polyethylene terephthalate and the pillarÔ palatal implant: its historical usage and durability in medical applications, Biomed. Eng. 11 (9) (1976) 301e306.

Polymers for surgical sutures

125

[73] O. Kawakami, S. Miyamoto, T. Hatano, K. Yamada, N. Hashimoto, Y. Tabata, Accelerated embolization healing of aneurysms by polyethylene terephthalate coils seeded with autologous fibroblasts, Neurosurgery 56 (5) (2005) 1075e1081. [74] J. Yao, C.Y.-L. Woon, A. Behn, T. Korotkova, D.-Y. Park, V. Gajendran, R.L. Smith, The effect of suture coated with mesenchymal stem cells and bioactive substrate on tendon repair strength in a rat model, J. Hand Surg. 37 (8) (2012) 1639e1645. [75] S. Weitzul, R.S. Taylor, Chapter 16 - Suturing technique and other closure materials, in: J.K. Robinson, R.D. Sengelmann, C.W. Hanke, D.M. Siegel, A.C. Bhatia, T.E. Rohrer (Eds.), Surgery of the Skin, Mosby, Edinburgh, 2005, pp. 225e244. [76] S.J. Langley-Hobbs, Chapter 10 - Sutures and general surgical implants, in: S.J. Langley-Hobbs, J.L. Demetriou, J.F. Ladlow (Eds.), Feline Soft Tissue and General Surgery, W.B. Saunders, 2014, pp. 105e116. [77] B. Pasternak, A. Missios, A. Askendal, P. Tengvall, P. Aspenberg, Doxycycline-coated sutures improve the suture-holding capacity of the rat Achilles tendon, Acta Orthop. 78 (5) (2007) 680e686. [78] U. Klinge, B. Klosterhalfen, A. Öttinger, K. Junge, V. Schumpelick, PVDF as a new polymer for the construction of surgical meshes, Biomaterials 23 (16) (2002) 3487e3493. [79] K. Baylón, P. Rodríguez-Camarillo, A. Elías-Zúñiga, J.A. Díaz-Elizondo, R. Gilkerson, K. Lozano, Past, present and future of surgical meshes: a review, Membranes 7 (3) (2017) 47. [80] R. Sellei, E. Bauer, M. Hofman, P. Kobbe, P. Lichte, R. Garrison, H. Pape, K. Horst, Reconstruction of a quadriceps tendon tear using Polyvinylidene fluoride sutures and patellar screw fixation: a biomechanical study, Knee 22 (6) (2015) 535e541. [81] P. Brossard, G. Le Roux, B. Vasse, Acute quadriceps tendon rupture repaired by suture anchors: outcomes at 7 years’ follow-up in 25 cases, Orthop. Traumatol. Surg. Res. 103 (4) (2017) 597e601. [82] C.R. Jecan, C.I. Filip, L. Raducu, D. Hernic, Advanced polymeric biomaterials-an integrated solution for plastic, reconstructive and aesthetic surgery, Key Eng. Mater. (2015) 197e204. [83] E. Stodolak, T. Gumula, R. Leszczynski, J. Wieczorek, S. Blazewicz, A composite material used as a membrane for ophthalmology applications, Compos. Sci. Technol. 70 (13) (2010) 1915e1919. [84] M. Alkadi, F. Alsalleeh, Ex vivo microbial leakage analysis of polytetrafluoroethylene tape and cotton pellet as endodontic access cavity spacers, J. Conserv. Dent. 22 (4) (2019) 381e386. [85] R. Jain, S. Wairkar, Recent developments and clinical applications of surgical glues: an overview, Int. J. Biol. Macromol. 137 (2019) 95e106. [86] J.C. Dumville, P. Coulthard, H.V. Worthington, P. Riley, N. Patel, J. Darcey, M. Esposito, M. van der Elst, O.J. van Waes, Tissue adhesives for closure of surgical incisions, Cochrane Database Syst. Rev. 11 (2014) 1e135. [87] R.H. Fortelny, A.H. Petter-Puchner, K.S. Glaser, H. Redl, Use of fibrin sealant (Tisseel/Tissucol) in hernia repair: a systematic review, Surg. Endosc. 26 (7) (2012) 1803e1812. [88] E.M. Petrie, Cyanoacrylate adhesives in surgical applications, Rev. Adhes. Adhes. 2 (3) (2014) 253e310. [89] C.M. Bhamidipati, J.S. Coselli, S.A. LeMaire, BioGlue in 2011: what is its role in cardiac surgery? J. Extra Corpor. Technol. 44 (1) (2012) P6eP12. [90] Y. Lee, C. Xu, M. Sebastin, A. Lee, N. Holwell, C. Xu, D. Miranda Nieves, L. Mu, R.S. Langer, C. Lin, J.M. Karp, Bioinspired nanoparticulate medical glues for minimally invasive tissue repair, Adv. Healthc. Mater. 4 (16) (2015) 2587e2596.

126

Advanced Technologies and Polymer Materials for Surgical Sutures

[91] A.P. Duarte, J.F. Coelho, J.C. Bordado, M.T. Cidade, M.H. Gil, Surgical adhesives: systematic review of the main types and development forecast, Prog. Polym. Sci. 37 (8) (2012) 1031e1050. [92] N. Annabi, Y.-N. Zhang, A. Assmann, E.S. Sani, G. Cheng, A.D. Lassaletta, A. Vegh, B. Dehghani, G.U. Ruiz-Esparza, X. Wang, S. Gangadharan, A.S. Weiss, A. Khademhosseini, Engineering a highly elastic human proteinebased sealant for surgical applications, Sci. Transl. Med. 9 (410) (2017) eaai7466. [93] N. Lang, M.J. Pereira, Y. Lee, I. Friehs, N.V. Vasilyev, E.N. Feins, K. Ablasser, E.D. O’Cearbhaill, C. Xu, A. Fabozzo, A blood-resistant surgical glue for minimally invasive repair of vessels and heart defects, Sci. Transl. Med. 6 (218) (2014), 218ra6. [94] B.K. Ahn, Perspectives on mussel-inspired wet adhesion, J. Am. Chem. Soc. 139 (30) (2017) 10166e10171. [95] P. Kord Forooshani, B.P. Lee, Recent approaches in designing bioadhesive materials inspired by mussel adhesive protein, J. Polym. Sci., Part A: Polym. Chem. 55 (1) (2017) 9e33. [96] M. Mehdizadeh, H. Weng, D. Gyawali, L. Tang, J. Yang, Injectable citrate-based mussel-inspired tissue bioadhesives with high wet strength for sutureless wound closure, Biomaterials 33 (32) (2012) 7972e7983. [97] S. Afewerki, L.S.S.M. Magalhães, A.D.R. Silva, T.D. Stocco, E.C. Silva Filho, F.R. Marciano, A.O. Lobo, Bioprinting a synthetic smectic clay for orthopedic applications, Adv. Healthc. Mater. 8 (13) (2019) 1900158. [98] Y. Liu, H. Meng, Z. Qian, N. Fan, W. Choi, F. Zhao, B.P. Lee, A moldable nanocomposite hydrogel composed of a Mussel-inspired polymer and a nanosilicate as a fit-to-shape tissue sealant, Angew. Chem. Int. Ed. 56 (15) (2017) 4224e4228. [99] A. Bellad, K.A. Pratap, Effectiveness of tissue adhesive versus conventional sutures in the closure of inguinal hernia skin incisions: a prospective randomized controlled trial, Int. Surg. J. 5 (5) (2018) 1797e1801. [100] I. Ayan, M. Colak, U. Comelekoglu, A. Milcan, O. Ogenler, V. Oztuna, F. Kuyurtar, Histoacryl glue in meniscal repairs (a biomechanical study), Int. Orthop. 31 (2) (2007) 241e246. [101] N. Annabi, K. Yue, A. Tamayol, A. Khademhosseini, Elastic sealants for surgical applications, Eur. J. Pharm. Biopharm. 95 (Pt A) (2015) 27e39. [102] G.M. Taboada, K. Yang, M.J. Pereira, S.S. Liu, Y. Hu, J.M. Karp, N. Artzi, Y. Lee, Overcoming the translational barriers of tissue adhesives, Nat. Rev. Mater. (2020) 1e20. [103] C. Serrano, L. García-Fernández, J.P. Fernández-Blázquez, M. Barbeck, S. Ghanaati, R. Unger, J. Kirkpatrick, E. Arzt, L. Funk, P. Turón, Nanostructured medical sutures with antibacterial properties, Biomaterials 52 (2015) 291e300. [104] D.C. Moore, M.H. Sellers, K.R. Archer, H.S. Schwartz, G.E. Holt, Staples equal sutures for skin closure after soft tissue tumor resection, Clin. Orthop. Relat. Res. 471 (3) (2013) 899e904. [105] B. Ananda, J. Vikram, B. Ramesh, H.M. Khan, A comparative study between conventional skin sutures, staples adhesive skin glue for surgical skin closure, Int. Surg. J. 6 (3) (2019) 775e782. [106] C. Dennis, S. Sethu, S. Nayak, L. Mohan, Y. Morsi, G. Manivasagam, Suture materialsdcurrent and emerging trends, J. Biomed. Mater. Res. A 104 (6) (2016) 1544e1559. [107] J.-M. Seitz, M. Durisin, J. Goldman, J.W. Drelich, Recent advances in biodegradable metals for medical sutures: a critical review, Adv. Healthc. Mater. 4 (13) (2015) 1915e1936. [108] S. Hoshino, Y. Yoshida, S. Tanimura, Y. Yamauchi, T. Noritomi, Y. Yamashita, A study of the efficacy of antibacterial sutures for surgical site infection: a retrospective controlled trial, Int. Surg. 98 (2) (2013) 129e132.

Polymers for surgical sutures

127

[109] F. Matl, J. Zlotnyk, A. Obermeier, W. Friess, S. Vogt, H. Büchner, H. Schnabelrauch, A. Stemberger, K.-D. Kühn, New anti-infective coatings of surgical sutures based on a combination of antiseptics and fatty acids, J. Biomater. Sci. Polym. Ed. 20 (10) (2009) 1439e1449. [110] R.E. Abhari, J.A. Martins, H.L. Morris, P.-A. Mouthuy, A. Carr, Synthetic sutures: clinical evaluation and future developments, J. Biomater. Appl. 32 (3) (2017) 410e421. [111] S.T. Dubas, S. Wacharanad, P. Potiyaraj, Tunning of the antimicrobial activity of surgical sutures coated with silver nanoparticles, Colloids Surf. 380 (1‒3) (2011) 25e28. [112] T. Baygar, N. Sarac, A. Ugur, I.R. Karaca, Antimicrobial characteristics and biocompatibility of the surgical sutures coated with biosynthesized silver nanoparticles, Bioorg. Chem. 86 (2019) 254e258. [113] J.E. Lee, S. Park, M. Park, M.H. Kim, C.G. Park, S.H. Lee, S.Y. Choi, B.H. Kim, H.J. Park, J.-H. Park, Surgical suture assembled with polymeric drug-delivery sheet for sustained, local pain relief, Acta Biomater. 9 (9) (2013) 8318e8327. [114] A. Arora, G. Aggarwal, J. Chander, P. Maman, M. Nagpal, Drug eluting sutures: a recent update, J. Appl. Pharmaceut. Sci. 9 (07) (2019) 111e123. [115] G.N. Salles, M.L. Calió, S. Afewerki, C. Pacheco-Soares, M. Porcionatto, C. Hölscher, A.O. Lobo, Prolonged drug-releasing fibers attenuate Alzheimer’s disease-like pathogenesis, ACS Appl. Mater. Interfaces (2018) 36693e36702. [116] T. Fuchs, C. Surke, R. Stange, S. Quandte, B. Wildemann, M. Raschke, G. Schmidmaier, Local delivery of growth factors using coated suture material, Sci. World J. 2012 (2012). [117] U. Costantino, V. Ambrogi, M. Nocchetti, L. Perioli, Hydrotalcite-like compounds: versatile layered hosts of molecular anions with biological activity, Microporous Mesoporous Mater. 107 (1‒2) (2008) 149e160. [118] P. Laurén, P. Somersalo, I. Pitkänen, Y.-R. Lou, A. Urtti, J. Partanen, J. Seppälä, M. Madetoja, T. Laaksonen, A. Mäkitie, Nanofibrillar cellulose-alginate hydrogel coated surgical sutures as cell-carrier systems, PLoS One 12 (8) (2017) e0183487. [119] J.P. Guyette, M. Fakharzadeh, E.J. Burford, Z.W. Tao, G.D. Pins, M.W. Rolle, G.R. Gaudette, A novel suture-based method for efficient transplantation of stem cells, J. Biomed. Mater. Res. A 101 (3) (2013) 809e818. [120] J. Yao, T. Korotkova, R.L. Smith, Viability and proliferation of pluripotential cells delivered to tendon repair sites using bioactive suturesdan in vitro study, J. Hand Surg. 36 (2) (2011) 252e258. [121] A.K. Reckhenrich, B.M. Kirsch, E.A. Wahl, T.L. Schenck, F. Rezaeian, Y. Harder, P. Foehr, H.-G. Machens, J.T. Egaña, Surgical sutures filled with adipose-derived stem cells promote wound healing, PLoS One 9 (3) (2014) e91169. [122] J.G. Casado, R. Blazquez, I. Jorge, V. Alvarez, G. Gomez-Mauricio, M. OrtegaMuñoz, J. Vazquez, F.M. Sanchez-Margallo, Mesenchymal stem cell-coated sutures enhance collagen depositions in sutured tissues, Wound Repair Regen. 22 (2) (2014) 256e264. [123] X. Jing, H.-Y. Mi, H.-X. Huang, L.-S. Turng, Shape memory thermoplastic polyurethane (TPU)/poly (ε-caprolactone)(PCL) blends as self-knotting sutures, J. Mech. Behav. Biomed. Mater. 64 (2016) 94e103. [124] P.S. Kowalski, C. Bhattacharya, S. Afewerki, R. Langer, Smart biomaterials: recent advances and future directions, ACS Biomater. Sci. Eng. 4 (11) (2018) 3809e3817. [125] A. Lendlein, R. Langer, Biodegradable, elastic shape-memory polymers for potential biomedical applications, Science 296 (5573) (2002) 1673e1676.

128

Advanced Technologies and Polymer Materials for Surgical Sutures

[126] R. Duarah, Y.P. Singh, P. Gupta, B.B. Mandal, N. Karak, Smart self-tightening surgical suture from a tough bio-based hyperbranched polyurethane/reduced carbon dot nanocomposite, Biomed. Mater. 13 (4) (2018) 045004. [127] A. Biswas, A.P. Singh, D. Rana, V.K. Aswal, P. Maiti, Biodegradable toughened nanohybrid shape memory polymer for smart biomedical applications, Nanoscale 10 (21) (2018) 9917e9934. [128] J.D. Kingsley, S. Ranjan, N. Dasgupta, P. Saha, Nanotechnology for tissue engineering: need, techniques and applications, J. Pharm. Res. 7 (2) (2013) 200e204. [129] S. Kumbar, R. James, S. Nukavarapu, C. Laurencin, Electrospun nanofiber scaffolds: engineering soft tissues, Biomed. Mater. 3 (3) (2008) 034002. [130] T.D. Stocco, N.J. Bassous, S. Zhao, A.E. Granato, T.J. Webster, A.O. Lobo, Nanofibrous scaffolds for biomedical applications, Nanoscale 10 (26) (2018) 12228e12255. [131] S. Khorshidi, A. Solouk, H. Mirzadeh, S. Mazinani, J.M. Lagaron, S. Sharifi, S. Ramakrishna, A review of key challenges of electrospun scaffolds for tissue-engineering applications, J. Tissue Eng. Regener. Med. 10 (9) (2016) 715e738. [132] M.M.M. De-Paula, S. Afewerki, B.C. Viana, T.J. Webster, A.O. Lobo, F.R. Marciano, Dual effective core-shell electrospun scaffolds: promoting osteoblast maturation and reducing bacteria activity, Mater. Sci. Eng. C 103 (2019) 109778. [133] A.O. Lobo, S. Afewerki, M.M.M. de Paula, P. Ghannadian, F.R. Marciano, Y.S. Zhang, T.J. Webster, A. Khademhosseini, Electrospun nanofiber blend with improved mechanical and biological performance, Int. J. Nanomed. 13 (2018) 7891. [134] M.M.M. De Paula, N.J. Bassous, S. Afewerki, S.V. Harb, P. Ghannadian, F.R. Marciano, B.C. Viana, C.R. Tim, T.J. Webster, A.O. Lobo, Understanding the impact of crosslinked PCL/PEG/GelMA electrospun nanofibers on bactericidal activity, PLoS One 13 (12) (2018) e0209386. [135] C.-L. He, Z.-M. Huang, X.-J. Han, Fabrication of drug-loaded electrospun aligned fibrous threads for suture applications, J. Biomed. Mater. Res. A 89A (1) (2009) 80e95. [136] S. Liao, B. Li, Z. Ma, H. Wei, C. Chan, S. Ramakrishna, Biomimetic electrospun nanofibers for tissue regeneration, Biomed. Mater. 1 (3) (2006) R45eR53. [137] R. Sridhar, D. Pliszka, H.-K. Luo, C. Lim, H. Keith, S. Ramakrishna, Medical devices regulatory aspects: a special focus on polymeric material based devices, Curr. Pharmaceut. Des. 21 (42) (2015) 6246e6259. [138] G.J. Mattamal, US FDA perspective on the regulations of medical-grade polymers: cyanoacrylate polymer medical device tissue adhesives, Expet Rev. Med. Dev. 5 (1) (2008) 41e49. [139] D.S. Couto, L. Perez-Breva, P. Saraiva, C.L. Cooney, Lessons from innovation in drug-device combination products, Adv. Drug Deliv. Rev. 64 (1) (2012) 69e77. [140] V. Bhagat, M.L. Becker, Degradable adhesives for surgery and tissue engineering, Biomacromolecules 18 (10) (2017) 3009e3039. [141] A.M. Behrens, B.J. Casey, M.J. Sikorski, K.L. Wu, W. Tutak, A.D. Sandler, P. Kofinas, In situ deposition of PLGA nanofibers via solution blow spinning, ACS Macro Lett. 3 (3) (2014) 249e254. [142] M. Mir, M.N. Ali, A. Barakullah, A. Gulzar, M. Arshad, S. Fatima, M. Asad, Synthetic polymeric biomaterials for wound healing: a review, Prog. Biomater. 7 (1) (2018) 1e21. [143] S. Afewerki, X. Wang, G.U. Ruiz-Esparza, C.-W. Tai, X. Kong, S. Zhou, K. Welch, P. Huang, R. Bengtsson, C. Xu, M. Strømme, Combined Catalysis for Engineering Bioinspired, Lignin-Based, Long-Lasting, Adhesive, Self-Mending, Antimicrobial Hydrogels, ACS Nano. 14 (2020) 17004e17017.

CHAPTER 6

Smart sutures Louise Taylor1, Morvarid Saeinasab1, Mohammad-Ali Shahbazi2, Ximu Zhang3, 4, Wei Zhang5, Karthik Nair6 and Farshid Sefat1, 7 1

Department of Biomedical and Electronics Engineering, School of Engineering, University of Bradford, Bradford, United Kingdom; 2Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy University of Helsinki, Helsinki, Finland; 3Chongqing Key Laboratory of Oral Disease and Biomedical Sciences, Chongqing Municipal Key Laboratory of Oral Biomedical Engineering of Higher Education, Stomatological Hospital of Chongqing Medical University, Chongqing, Sichuan Province, China; 4State Key Laboratory of Oral Diseases & National Clinical Research Center for Oral Diseases & Dept. of Preventive Dentistry, West China Hospital of Stomatology, Sichuan University, Chengdu, Sichuan Province, China; 5State Key Laboratory of Polymer Materials Engineering, Polymer Research Institute at Sichuan University, Chengdu, Sichuan Province, China; 6Summit Medical Ltd., Bourton on the water, United Kingdom; 7Interdisciplinary Research Centre in Polymer Science & Technology (Polymer IRC), University of Bradford, Bradford, United Kingdom

6.1 Introduction Smart sutures, otherwise referred to as electronic sutures, refer to the incorporation of a specialized coating or the addition of flexible electronics, including miniature sensors or microfluidic technology [1] into sutures which can be used to retrieve real time data of the wound or targeted tissue. This has become an achievable possibility due to the advancements of microelectronics, in particular the creation of a new class of electronics which includes flexibility, degradability and biocompatibility electronics [2]. The information gathered from the smart sutures provides real-time data that will aid in the wound healing process alongside the use of the data monitoring internal tissues or organs, if used through surgery procedures. The World Health Organization has found that Surgical Site Infections (SSI), where an infection occurring on the any site of the anatomy which occurs within 30 days of the procedure [3], are the most surveyed and frequent type of hospital acquired infections in low and middle income countries and the second most frequent type within Europe and USA. Application of smart sutures varies due to the type of suture used alongside the adaptive quality of the 2D suture threads allows the smart suture to be transformed into a 3D structure through already established textile techniques, further increasing the possibilities of application to become almost endless [1]. Smart sutures have been in development since approximately 2012 [4]; however, there are no current smart sutures in use within practice to date. There is documentative evidence that in vivo Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00012-7

© 2023 Elsevier Ltd. All rights reserved.

129

130

Advanced Technologies and Polymer Materials for Surgical Sutures

Table 6.1 Different categories and parameters for smart sutures. Category

Physical Chemical

Name of characteristic of style of sensor

Temperature Strain pH Glucose Microfluidic Bacteria Resorbable Neutrophils Colorimetric

Development stage

In vivo In vivo In vivo In vitro In vitro Sensors available for future development. Sensors available for future development. Sensors available for future development. Sensors available for future development.

testing has taken place within some types of smart suturing producing positive outcomes from these trials, indicating that smart sutures may become a reality in the clinical environment. There are several types of parameters that are currently being used within these smart sutures trials that are either physical or chemical aspects. The list of each category included in this chapter is illustrated below alongside the stage of development (See Table 6.1).

6.2 Base material of smart suture To date there are several different types of sutures, some of which are an appropriate material to base intended electronic structures on for the use of smart sutures, however this venture into electrical sutures has introduced newer materials that would be more suitable smart sutures while in keeping the original characteristics required for regular suturing. 6.2.1 Paper The use of paper within a suture, for the intension of smart sutures, has been identified by several investigations including Najafabadi et al. (2014) who outlined paper to be a flexible, light weight material that would also be cost effective to fabricate within a large-scale view. Papers additional properties of being porous and permeable to both air and liquid would make an effective adaptive substrate for electronics; however, the inability

Smart sutures

131

to be elastic would damage the electronic aspects and the suture itself when surrounded in an aqueous environment and therefore would have to be further adapted to be an acceptable material. 6.2.2 Polyglycerol sebacate Polyglycerol sebacate, more commonly referred to as PGS, is a synthetic polymer that was first established in 2002 showing good biocompatibility and absorption properties. PGS has been used in a variety of ways within biomedical applications including cardiac, nerve, and adhesives [5]; however, not as a pure suture format. PGS is a biocompatible material that is often used as a base or scaffold with the addition of another more functional element to the intended use of the implant. 6.2.3 Polycaprolactone Polycaprolactone (PCL) is a biodegradable synthetic thermoplastic polymer that is used for many medical applications. Although biocompatible, PCL alongside many other biopolymers do not possess positive biochemical signals to enable interaction with cells, and thus additives are commonly used to create a superior material [6]. Sutures that are PCL based are currently being trialed having the addition of Keratins and Collagens. Ghosh et al. (2017) [6] established that the addition of keratins would increase cell proliferation and adhesion and thus promote natural tissue repair and has been an increasingly researched additive within biomaterials since the 1970s when keratin extraction and purification advances opened the doors to more research and applications of the natural protein [7]. Keratin is not only a great biocompatible organic but also contains intrinsic biological activity which has been seen to promote repair in all types of biological tissues, including nerve regeneration which is an area of medicine that has a low success rate and thus provides a positive possibility within the specialty. Collagens were introduced to PCL fibers by McNeil et al. (2011) [8], in the form of electrospun fibers to create a scaffold that would be implanted into a boney structure to encourage the proliferation of osteoblasts which facilitate bone regeneration. Although this is used specifically with a scaffold platform, the fibers can be adapted into other forms including sutures, alongside other extracellular matrix proteins, to enhance cell adherence and proliferation for targeted tissue repair.

132

Advanced Technologies and Polymer Materials for Surgical Sutures

6.2.4 PCL/PGS blend Najafabadi et al. (2014) identified an opportunity to incorporate PCL with PGS to create a more mechanically strong synthetic polymer, to form the base substrate for biomedical structures, through electrospinning. The electrospun combination of the two polymers was formed into sheets that were then used as the basis of various sensors. Through tensile testing the outcome of the electrospun substrate, the elastic modulus of the substrate mixture was 4.86
0.54 MPA which is comparable to some native tissues such as skin dermis [9]. This then provided the evidence needed to ensure that this combination of polymers would be suitable base for the bioelectronics and would survive the tension and stresses endured during the suturing process. 6.2.5 Cotton Cotton thread has been used as a suture material throughout history with one of the earliest recommendations for the suture material being approximately 500 BCE by Susruta [10]. This nonabsorbable suture is still in use today and has good strength, alongside the ease of knotting; however, when incorporated into a multifilament suture, there is an increased likelihood of inflammation and causes a tissue reaction by being present in some wounds [11]. Cotton could be used as a smart suture base material by using smart nanoparticle coatings to transform the thread into a smart suture as opposed to layering different components onto the cotton suture [12]. Preparation of the cotton suture by Mostafalu et al. (2016) included a plasma cleaning process, which not only removes the added wax coating of the thread, but also adds an eOH bond to the surface, allowing the nanoparticles to be more effective at bonding, using covalent bonds. Nanoparticles that are the most suitable for coating cotton include zinc oxide, titanium, and silver, which are all costeffective allowing mass production possibilities. Once the nanoparticles are added, the cotton threads would then be further coated with a conductive ink, such as silver, silver chloride, CNTs, and polyanilines, to convert the coated thread into a functionable electrode [12]. Once the cotton smart suture has been created sterilization is required that does not alter the structure of the suture, and the most common method of which is by using ultraviolet light. Although in vitro and in vivo testing has been carried out, there have been no long-term biocompatibility studies carried out, which will also ensure that the nanoparticle technology that has been applied to the sutures remain durable, effective, and safe.

Smart sutures

133

6.2.6 Carbon nanotubes Carbon nanotubes (CNT) were first established in Russia and were published in journals for their novel characteristics in 1991 by Lijima13. CNTs can be formed out of any carbon allotrope, such as diamond, grapheme, and graphite, and is a one-dimensional form of the allotrope which has an aspect ratio greater than 100. CNTs present high mechanical strength alongside flexibility and efficient electrical and thermal conductivity, however, have the disadvantage of having a smooth seamless surface which inhibits physical interaction between any added matrix. CNTs also have the absence of any chemical affinities to organic solvents and polymers and due to these properties, CNTs have to be adapted with the application of other materials such as proteins and nucleic acids through bio-nanohybrid systems [13]. 6.2.7 Wicking Wicking is the process of the threads ability to absorb fluids and thus wicking properties of the underlying and overall suture materials should be taken into consideration to ensure that a smart suture possess the suitable wicking properties for the intended process. If a suture is intended to be hydrophobic, such as in a microfluidic function, the suture can be covered with a silicone lubricant which will fill all pores between individual strands blocking as much liquid flow as possible [1]. 6.2.8 Polyurethane Polyurethane is a commonly used suture material that is use as a nonabsorbable monofilament suture in various types of applications including skin closure due to having a high tensile strength alongside good elasticity property [14]. Polyurethane is an elastomer that is used most commonly in a foam state in many everyday items alongside biocompatible products.

6.3 Temperature sensors for smart sutures The uses of temperature properties within suturing, alongside other implantable implants, has already been established with the heat sensitive activation of shape memory polymer sutures [15]. Temperature is an important indicator of the health of an individual with an abnormal temperature being a diagnostic tool to identify many illnesses and conditions including infections. Localized skin temperature abnormalities at a wound site, particularly if elevated from the optimum range of 37 c, can indicate

134

Advanced Technologies and Polymer Materials for Surgical Sutures

abnormal healing progress. Elevated temperature occurs at infected sites due to the immune response triggering an influx of white blood cells to combat the wound infection in days 1e4 of an infection being present [1]. This constant marker of the early onset of infection is an effective parameter to be able to monitor as a suture due to the suture situating within the at-risk wound. There is currently only one tried approach to create a temperature sensitive suture, where a resistive temperature sensor is created and attached onto a suture which reads and sends real time data onto a displaying device, however a second approach of colorimetric designs are discussed with future smart suture designs. The base material for a temperature smart suture can vary, due to a base coating of silicon nanomembrane to then apply the smart material to attach to. Metals are then applied in a set pattern of either serpentine or spiraling which will then become a functioning resistive temperature sensor when electricity is passed though, using an external power source. This method has been popular within current trials and experiments with three different groups using this process with varying suture materials and metals used. Mostafalu et al. (2016) designed a toolkit for thread-based diagnostics including several types of smart suture sensors. The chosen method to create a temperature sensor for this development was a resistive temperature sensor created out of CNT-coated threads with a serpentine pattern onto a woven fabric for testing. CNT-coated threads were functionalized with a carboxylic group and were chosen over the more expensive optimum materials of nickel and platinum however were tested to be effective range between 20 and 40 c which is applicable to most viable biological tissues. This sensor would then need to be attached to an external power source and a display device which can then transmit data remotely through Bluetooth on an Arduino microprocessor. In this trial, the temperature sensor was not investigated in vivo. Kim et al. (2012) have further developed the scope of thermal based smart suture by designing a suture strip which incorporates both a temperature sensor and a thermal actuator. The advantages of being able to control the temperature at the wound site are twofold with not only correcting any abnormal temperatures but also by manipulating the temperature to be optimum for healing at the specific wound site. Maintaining optimum heat can also be used as a treatment method by promoting increased blood flow to the infected area [16]. The designed suture uses strips of polyester which incorporate a single crystal silicone nanomembrane temperature diode on the top side of the strip with a gold microheater on

Smart sutures

135

the reverse side, which is then encased in polydimethylsiloxane (PDMS) as a protective coating. The same process was carried out for an alternative temperature sensor of a platinum resistance thermometer created with a serpentine pattern; however, the gold microheater and thermometer are bifacially placed and are thus wider strips. Electrical elements are then connected to an external power source using contact pads and anisotropic conductive films. Test on the accuracy of the temperature sensors were carried out using a heated hot which identified that both variations of the sutures obtained a resolution of 0.2 c. Tensile stress tests were carried out on both variations using a simple suture stitch and several types of knots to ensure that the components would remain intact during the process. Microheaters were tested using infrared camera showing the change of temperature caused by the heater. Testing two variations of the temperature sensors alongside the suitability of the sensors to withstand tensile stresses of suturing show that these device designs and fabrication methods can be transferable to other suture materials to create further specialization including being bioabsorbable or thinner sutures. In vivo tests were then carried out testing only the new temperature sensor technology using a silk suture substrate to create a thin smart suture with the platinum resistance temperature sensor chosen due to being the cheaper and simpler option to fabricate. During the test, tensile strength was further tested using three varying stitch patterns on a smart suture containing four temperature sensors which read localized temperatures throughout being in situ. In vivo testing has shown that this style of resistance temperature sensor will work while within a wound site and temperature monitoring can occur, however the heater element of the suture was not tested in vivo and thus further investigations will be needed to specifically test the microheating element of the suture.

6.4 pH sensor smart sutures The acidity levels within a wound will differ depending on what stage of healing the wound is in. As a general rule of thumb that acidic environment within a wound indicated effective healing, whereas alkaline environments are indicative of wound that has a severe bacterial infection or a wound that is likely to become a chronic wound [17]. Currently pH is not a regularly measured parameter within wound care management however can be an informative indicator when a wound is struggling to heal naturally where a

136

Advanced Technologies and Polymer Materials for Surgical Sutures

suture that indicates irregular pH being a beneficial tool for further investigation and medication. Alterative applications of pH smart suture would be within the monitoring of the gastrointestinal tract to diagnose diseases such as inflammatory bowel disease, or infections such as Helicobacter pylori. There are two approaches to creating pH smart sutures, with either creating a pH sensor using the suture as a base or from a colorimetric approach with a reactive coating. Wu & Sailor (2009) [18] have used the coating method with the application of a hydrogel film which has been created out of a reaction between chitosan and glycidoxypropyltrimethoxysilane. This hydrogel was then applied to a silicon dioxide coated suture which allows the base suture to be variated. The hydrogel will naturally react with the environments pH levels by changing in volume, when the environment turns acidic to 6.0, swelling occurs and reduces when returned to the normal state of 7.4. This reaction is restricted to a small range of pH change, restricted to the acidic side of the scale, where the normal pH of healthy skin is slightly acidic at 4.7 pH whereas the body’s natural pH internally is between 7.35 and 7.45 [19]. The mechanism of swelling as the response to pH does eliminate the need for power to the sensor however the swelling is not an obvious change with the opportunity to measure the extent of suture swelling while in situ from a clinical application being extremely difficult. Within Mostafalu et al.’s (2016) approach to establishing a toolkit for thread sensors, the pH sensor was created using a potentiometric approach where voltage between two electrodes within the environment, is entered into the Nernst equation to equate the voltage into a pH value. Electrodes were constructed with cotton threads which dipped in silver/silver chloride and polyaniline (PANI) alternatively before being tested in prepared in phosphate-buffered saline (PBS) that had been altered pH ranges between 3 and 8. This calibration process was then followed by using a microfluidic approach of directing a testing fluid though a chicken skin to stimulate the intended environment to ensure that the pH sensor would read and be able to transmit readings wirelessly to a computer. This sensor design was able to detect a stable pH reading within 30 s of being in situ and was tested for a maximum amount of 4 h with effective pH reading. In vivo testing was carried out on a rat model which showed that the sensor was successful in reading pH in both a gastric reading though the stomach and subcutaneously. This suturable pH sensor measures wound fluids accurately as demonstrated; however, the implementation of this sensor would result in the sensor having to be implanted within a wound as

Smart sutures

137

opposed to being used topically or within the incision suture itself. This would then require removing once the use of the sensor had finished and thus increasing opportunities for infections with a second procedure at a wound site. Alternatively, the sensor could be applied topically with the use of microfluidic sutures directing the internal wound fluid and then returning the fluid back to the wound; however, this too could introduce new opportunities for infections, even if the suture and sensor were under a dressing. During this investigation, Mustafalu et al. identified that the sensor would be prone to degrading if continuously used in vivo due to the absorption of proteins within the dye reaction. A possible solution was identified, using antifouling-coating materials; however, this was not trialed with this study, and further illustrates the impracticalities of this style of suture as it currently stands.

6.5 Strain smart sutures Skin is constantly under tensile stress which is apparent when an incision is made as the skin relaxes creating the cut. Strain and stress within the wound healing process is an important factor for the effective healing of wounds and can also be an indicator that inflammation is present [20]. Currently there is no in vivo evidence that suggests inflicting tension to create stress in healing skin wounds is beneficial however it has been theoretically worked out to be 10% by Zöllner et al. (2012) [21] to promote tissue growth. Ultimately the strain smart suture application will remain within the role of inflammation observation until further evidence comes to light with applying tension. From a scientific evidence point of view, strain sutures could also be used within study work for evaluations on tendon repairs or degradation; however, the application of the suture to a tendon could further weaken the tendon and thus is currently restricted to cadaver studies [22]. One method of being able to wirelessly assess the strain on a suture would be to apply spots of iron oxide (Fe3O4) onto the suture material, which was carried out by Najafabadi et al. (2014) using a PGS-PCL mixture. Fe3O4 spots were screen printed onto the suture substrate, which was then tested using a polyimide-insulated copper coil creating strain on the Fe3O4 spots due to the magnetic field. The change in inductance within the Fe3O4 spot movement due to differing magnetic strains applied to the substrate is what is measured to identify the present of strain within the suture. The amount of strain present on this style of sensor

138

Advanced Technologies and Polymer Materials for Surgical Sutures

is measured as a percentage of kHz and has a sensitivity of 0.7% within the first 10% on strain applied and is increased sensitivity of 0.1% at 10%e20% which is in line with other flexible strain sensors. The created strain sensor was not tested in vivo however similarly designed strain sensors using strain-controlled modulation have been used within a wet environment and is therefore expected that this sensor technique should also be applicable to wet environments, due to the magnetic permeability between air and water being very similar. The application of the strain sensor is not evident within this study as only functionality of the device was investigated and not the application methods of incorporating the suture-based sensor within a wound setting. The strain sensor has intentionally been based on a biodegradable suture material with the scope of further development being able to incorporate the sensor within a resorbable smart suture or dressing application. Mostafalu et al. (2016) used the same technique of measuring the strain by observing the electrical resistance; however, the sensor within the toolkit was created out of Polyurethane (PU) threads which were coated with carbon ink and CNTs, both of which are established to be biocompatible. Carboxylic groups were used to functionalize both the carbon ink and CNTs to ensure the secure binding to the threads that were further coated with PDMS as a protective layer. Testing the created strain sensor was achieved by applying constant strain while using strain gauges to register the linear changes of the electrical resistance while being stretched. Due to the inclusion of CNTs, this style of strain sensor was able to withstand higher strains, of up to 100% or GFw3. In vivo evaluation of the strain sensor suture was carried out on rats where three states of wound: closed, semiclosed, and open, of 1 cm incisions were measured using the sensor suture being placed within the wound and secured with a knot on either side. The strain sensor was then connected using crocodile clips to an external power and strain-monitoring equipment. Although this design of strain sensor suture has been tested in vivo, the application was as a straight strain gauge which may not be practical within a wound suture setting and this may only be currently clinically relevant as an external strain monitor for the indication of swelling due to infection within the wound. The supporting electronics would need to be refined and studied further before this style of strain suture could be considered clinical beneficial alongside the use of an external power source making the suture less practicable.

Smart sutures

139

6.6 Glucose smart sutures Monitoring of the glucose levels is common practice, especially within an individual who has diabetes is a daily task, which can be carried out several times a day depending on the type of diabetes, where a fingerpick blood sample is taken to measure blood glucose levels. If the glucose levels are not correctly managed within the blood stream, this can cause several sever medical conditions and thus effective monitoring is the key to maintain and correcting glucose levels. With the intended use of a glucose monitoring suture being implanted and left in situ until deterioration of the suture where it would be replaced, the sensor should be designed to be as long standing as possible to reduce the frequency of procedures needed to replace the device [1]. Glucose presence can be detected by measuring the current that can pass through a fluid between electrodes. Mostafalu et al.’s toolkit (2016) created the electrodes out of carbon and CNT threads for the working electrodes, with carbon threads for the counterelectrode and silver/silver chloride threads as the reference electrode, which can be seen in Fig. 6.1. Once created, the treads were patterned into a woven fabric for testing with glucose solutions with concentration ranges from 2 to 15 mM were added to the working electrode thread use PBS with a normal pH of 7.4 to mimic the body environments as closely as possible. Two pulses with voltages 0.5

Figure 6.1 Glucose sensor. (Reproduced from Ref. [1] with permission from NPG, copyright 2016dOpen Access.)

140

Advanced Technologies and Polymer Materials for Surgical Sutures

and 0 V were applied with a duty cycle of 50%, with the output current measured, where the higher the current indicating the higher concentration of glucose. Despite mimicking the in vivo environments as closely as possible, in vivo studies should still be carried out to ensure that the threads can be implemented in situ alongside establishing a life cycle time frame. Further developments to the design would include an implantable device to power the sensor as well as communicating readings via Bluetooth to a device to read real time data.

6.7 Microfluidic analysis smart sutures Microfluidic system analysis is the controlled direction of fluids through sensors for analysis which would be beneficial within smart sutures due to the ability for the sutures and sensors to be placed in a 3D fashion in an intended tissue or organ to monitor more than one parameter in that area. Microfluidic systems have already been established for several uses include molecular assays on blood samples, identifying and collecting tumor cells that are circulating through the area and or detecting cancer-specific biomarkers; however, these microfluidic systems are currently restricted to small transportation distances and in a planar structure [1].Suturable microfluidic systems that were wireless would be extremely beneficial from a surgical point of view due to the ability to monitor diseased and transplanted organs and tissues alongside, reducing the need for as many regular appointments and investigations which can harm the body overall with increased incidence of exposure to harmful X-rays. Mostafalu et al. (2016) have incorporated microfluid chemical analysis into their toolkit of smart sutures, due to the incorporation of several types of sensors including, temperature, pH, glucose, and strain. The transportation of the fluids was controlled by suture threads that possess either hydrophobic or hydrophilic properties to direct fluids into the intended sensors and keeping others away from the sensor. These transportation sutures do already exist within established clinical practice with the required properties. This toolkit requires an external power source with real time data being delivered via Bluetooth chip to a mobile phone and contained sensors with a suture-based electrode approach, as seen in Fig. 6.2, which can be used for multiple monitoring purposes depending on the measured parameter of electrical aspects of the fluid. The application of this style of analysis by smart sutures may bring disadvantages including exposing a

Smart sutures

141

Figure 6.2 A simple demonstration of an Arduino-based microprocessor for the smart suture designs which requires Bluetooth communications. (Reproduced from Ref. [1] with permission from NPG, copyright 2016dOpen Access.)

wound or targeted tissue to become more prone to infection with a passage of fluids to be brought to the surface for testing; however, if this was implanted within an organ or deeper surface tissue, this would have to be designed into an incased implantable device with a power source which may be more practical depending on the application.

6.8 Resorbable smart sutures Resorbable sutures are in common use since they were introduced in the early 1960s [23] as they reduced the need to interrupt a healing wound and decreases the likelihood of secondary infection. There would be further advantages to developing a smart suture that would also be resorbable using the basis of the suture as an already established resorbable material, such as silk fibroin, PLGA, PCL and collagen, with the addition of electronic material that could also be resorbable, known as physically transient electrics [2]. The challenge that resorbable smart sutures present is to achieve a smart suture that will have provide a time frame that is suitable to the clinical use before degradation [24]. For this reason, silk can be a more suitable substrate to use as the base material of the suture thanks to the ability of adapting the reabsorption rate due to the crystallinity of the silk, however, can be more challenging with the mechanical handling, stability, sterilization, and biopolymer interface [21]. Najafabadi et al. (2014) [2] designed and tested a temperature sensor based on a biodegradable suture substrate of PGS-PCL mixture on a 1:1 ratio using an electrospinning process, creating similar morphology and physical characteristics to paper suture substrates, thus allowing techniques

142

Advanced Technologies and Polymer Materials for Surgical Sutures

of paper-based electronics to be implemented on PGS-PCL sheets. This substrate was then overlayered with silver ink using a screen-printing in a serpentine pattern. Silver was selected as the conductive metal for the resistive temperature sensor due to the natural antibacterial properties alongside the advantage of having no mechanical effect to the PGS-PCL sheets. Tensile testing showed that the smart sensor suture obtained the elastic modulus of 4.87
0.17 MPa which is comparable to the elastic modulus of certain tissues of the body, such as the skin dermis. Electrical conductivity was tested during applied stress and strain cycles, which indicated that initially electrical resistance did not greatly vary until reaching the 20th cycle point but did stay consistent after the third cycle. Natural degradation of electrical components was tested through mimicking the bodily conditions with submerging the suture sheet within sodium hydroxide and PBS solutions separately at 37 C while measuring electrical resistance and weight over 14 days, with the complete degradation of the substrate after 30 days in PBS. Electrical resistance stayed within 10% for the first 10 days; however, degradation rate after day 10 did reduce in PBS. In contrast, the substrate was aggressively degraded through the sodium hydroxide submersion; however, the electrical functionality and general structure of the fabricated sensor were preserved. Further testing in vivo is required; however, the implication of the design would suggest that an external power source would be required to enable the suture remain reabsorbable. Although this is a wholly resorbable smart suture, if an external power source was required, the application of the suture could not be used to close the wound and would be used as a topical sensor for the wound area with the current design of the sensor. Data would be displayed with an as yet undeveloped device or through wireless communication, such as Bluetooth, to another electrical device. The temperature of the wound area in real time will be established for the predicted 30-day period which would cover the timeframe of most postoperative wound infections [25].

6.9 Future smart sutures Identifying areas for future smart sutures would require identifying sensors that can be adapted to be constructed with biocompatible materials, alongside adapting sensors that are being developed for dressing settings as the application of the sensor are almost identical to that of a suture alongside the same situ of the device. Ideally, any suture that is powerless can be seen

Smart sutures

143

as being more advantageous due to no extra external objects next to the wound or under the wound dressing to reduce the likelihood of introducing infection risks to the susceptible break in the skin. 6.9.1 Bacterial detection sensors Bacteria produce a secretion of protein designed to aid in the infiltration of the host immune system [26] which can be used as a tool to identify if bacteria are present within a wound. The proteins are specific to the type of bacteria; however, the most popular wound infection bacteria can be targeted. Currently there are established sensors developed to identify the presence of pyocyanin, which indicated the presence of Pseudomonas aeruginosa, urate, or uric acid, where sudden decrease in levels is a diagnostic marker of Staphylococcus aureus and of Pseudomonas aeruginosa, both have been established by Sharp et al. (2008 & 2010) [27,28]. Both sensors were based with carbon fibers used to monitor the oxidation of the targeted element, although the sensor technique was successful, both designs had disadvantages. Sensors for pyocyanin were very successful; however, the intended implementation was for a wound dressing where a power source and an electrochemical detector would be needed and thus not always practical in every setting. The sensor designed for urea and uric acid unfortunately became fouled with longer use which was not fully correctible within the study time frame. Despite this, the sensor was affective in measuring the presence of the urea and uric acid within the wound, studies of chronic wounds has shown that differing types of chronic wounds can produce different levels of uric acid which can also be present within a healthy wound [29], and therefore, the application of this technique is a complex approach depending on the type of wound. 6.9.2 Neutrophil sensor Neutrophils are a type of white blood cell which comprises part of the immune response to injury and infection. Initially the body responds to injury by sending neutrophils onto the injury site which in turn release enzymes to clean the wound of any necrotic tissue through engulfing. Once cleaned, the concentration of neutrophils should reduce; however, if an infection is within the injury, the body will continue to respond to the presence of the infection with providing neutrophils to the infection site. While neutrophils are present within an injury site, they will prevent the closure of the wounds due to their phagocytic nature [30]. In order to

144

Advanced Technologies and Polymer Materials for Surgical Sutures

monitor the presence of active neutrophils, a method has been established where the enzymes that the neutrophils produce is monitored, specifically Human Neutrophil Elastase (HNE) and Cathepsin G (CatG) which have both been found to be an early warning signs of infection [31]. Edwards et al. (2005) [32] have designed a sensor, made from chromophore linked with a short peptide, in which the enzymes within the wound from the neutrophils will cleavage the peptide. This substrate was then introduced to wounds via wound dressings, including collagen, which is an already established suture material, produced positive results with the increased absorption within infected wound fluid. Presently there is concern for the increased incubation periods of 12þ hours; however, this was tested on wound fluid and not the wound itself. This has prompted developments in 2008 [33] where a cellulose based substrate lowered the incubation time to 2e5 min which has been adapted into a dip stick product. This shows that this technology, although originally designed for wound dressing, has scope to be adapted within a suture structure, such as colorimetric indication with dyes that cleavage the peptide element of the sensor, if it was adapted to a protein. This would then reduce the swelling mechanic that is currently in use with the sensor which would not be appropriate to a suture sensor. 6.9.3 Colorimetric smart sutures Reches et al. (2010) [34] identified five different colorimetrical sutures for the detection of ketones, nitrite, protein (pH monitoring) and glucose for urinal applications and alkaline phosphatase within a plasma application. Alongside this study, Li et al. (2010) [35] designed two cotton-based threads that have a colorimetric response to nitrite ions and uric acid, which can be seen in Fig. 6.3. The basis of all sutures was cotton thread; however, these

Figure 6.3 Color change of threads using nitrite ion and uric acid. (Reproduced from Ref. [34] with permission from ACS, copyright 2010dOpen Access.)

Smart sutures

145

assays were not designed for clinical uses as sutures, but have the potential to be further developed into sutures due to being based on the cotton thread ensuring that the assay dyes do not bleed into the wound. Another colorimetric smart suture design could be adapted from Tamayol et al. (2016) [36] work where a hydrogel fiber which contained mesoporous particles which possessed a pH responsive dye creating a colorimetric reading. The hydrogel which has been adapted is a natural form called Alginate with the pH responsive mesoporous particles being present in the form of beads to reduce the likelihood of the pH dye bleeding into the wound. This study focused on the fiber mechanism rather than the application of the fiber within a suture setting; however, as sutures can be created with many strands such as normal cotton thread, this sensor has potential to have a big impact as a functioning smart suture. This colorimetric pH smart suture would then be able to be quickly visually assessed at dressing changes and negate the need for extra complex structures to the would site as demonstrated in Fig. 6.4.

Figure 6.4 Illustrates the color change mechanism against skin, using pig skin, and sprayed pH solutions. (Reproduced from Ref. [36] with permission from PMC, copyright 2016dOpen Access.)

146

Advanced Technologies and Polymer Materials for Surgical Sutures

6.10 Conclusions Although there are many prongs to the approach of smart sutures, there are few techniques which are close to being accepted into human trials at present. There are such a variety of uses of smart sutures clinically and from a scientific aspect that would help to understand human and medical condition mechanics. The benefits of achieving a smart suture are plentiful but must outweigh individual disadvantages such as degradation rate, incidence of soiling, introduction of infection risks if secondary procedures are needed to retrieve implanted technology and the absence of an immunological response to the materials used within the wound site. The use of smart sutures may actually be within an implantation point of view rather than a helpful diagnostic tool for wound infections and thus this will reshape how a suture is seen as the suture can be used traditionally or be reconstituted as a 3D structure into an implanted feature. Smart sutures that require an external power source can be seen as less advantageous from a medical point of view with more risk of introducing infections with the addition of extra external structures to the susceptible wound site, but also for a patient point of view with the mobility impracticalities of being attached to an unknown size of power supply. Further studies are required in all aspects to enable the implementation of what could be a new novel promising monitoring and diagnostic tool.

References [1] P. Mostafalu, M. Akbari, K. Alberti, Q. Xu, A. Khademhosseini, S. Sonkusale, A toolkit of thread-based microfluidics, sensors, and electronics for 3D tissue embedding for medical diagnostics, Microsyst. Nanoeng. 2 (16039) (2016) 1e10. [2] A. Najafabadi, N. Tamayol, N. Annabi, M. Ochoa, P. Mostafalu, M. Akbari, M. Nikkhah, R. Rahimi, M. Dokmeci, S. Sonkusale, B. Ziaie, A. Khademhosseini, Biodegradable nanofibrous polymeric substrates for generating elastic and flexible electronics, Adv. Mater. 26 (2014) 5823e5830. [3] World Health Organisation, Global Guidelines for the Prevention of Surgical Site Infection, WHO, Switzerland, 2016. [4] D. Kim, S. Wang, H. Keum, R. Ghaffari, Y. Kim, H. Tao, B. Panilaitis, M. Li, Z. Ka, Thin, flexible sensors and actuators as ‘instrumented’ surgical sutures for targeted wound monitoring and therapy, Nano Micro Small 8 (21) (2012) 3263e3268. [5] X. Loh, A. Karim, C. Owh, Poly(glycerol sebacate) biomaterial: synthesis and biomedical applications, J. Mater. Chem. B 3 (39) (2015) 7641e7652. [6] A. Ghosh, A. Ali, S. Collie, Effect of wool karatin on mechanical and morphological characteristics of polycaprolactone suture fibre, J. Textil. Eng. 63 (11) (2017) 1e4. [7] J. Rouse, M. Van Dyke, A review of keratin-based biomaterials for biomedical applications, Materials 3 (2010) 999e1014.

Smart sutures

147

[8] S. McNeil, H. Griffiths, Y. Perrie, Polycaprolactone fibres as a potential delivery system for collagen to support bone regeneration, Curr. Drug Deliv. 8 (4) (2011) 448e455. [9] J. GFennisson, T. Baldeweck, M. Tanter, S. CAtheline, M. Fink, L. Sandrin, C. Cornillon, B. Querleux, Ferro-electrics and frequency control, IEEE Trans. Ultrasonics. 51 (2004) 980. [10] J. Melle, Early history of the ligature, S. Afr. Med. J. 8 (8) (1934) 290e292. [11] R. Ballweg, E. Sullivan, D. Brown, D. Vetrosky, Medical knowledge: clinical procedures, in: Physician Assistant: A Guide to Clinical Practice, fifth ed., Elsevier, Philaedelphia, 2013, pp. 127e157. [12] N. Davies, E-Textiles: smart sutures, AATCC Rev. 17 (3) (2017) 38e43. [13] I.-Y. Jeon, D. Chang, N. Kumar, J.-B. Baek, Functionalization of carbon nanotubes, in: S. Yellampalli (Ed.), Carbon Nanotubes - Polymer Nanocomposites, Intech, 2011, https://doi.org/10.5772/18396. [14] F. Saleh, B. Palmieri, D. Lodi, K. Al-Sebeih, An innovative method to evaluate the suture compliance in sealing the surgical wound lips, Int. J. Med. Sci. 5 (6) (2008) 354e360. [15] J. Leng, X. Lan, Y. Liu, S. Du, Shape-memory polymers and their composites: stimulus methods and applications, Prog. Mater. Sci. 56 (7) (2011) 1077e1135. [16] S. Dhivya, V. Padma, E. Santhini, Wound dressings e a review, Biomedicine 5 (4) (2015) 24e28. [17] G. Gethin, The significance of surface pH in chronic wounds, Wounds U. K. (3) (2007) 52e56. [18] J. Wu, M. Sailor, Chitosan hydrogel-capped porous SiO2 as a pH responsive nanovalve for triggered release of insulin, Adv. Funct. Mater. 19 (5) (2009) 733e741. [19] H. Lambers, S. Piessens, A. Bloem, H. Pronk, P. Finkel, Natural skin surface pH is on average below 5, which is beneficial for its resident flora, Int. J. Cosmet. Sci. 28 (5) (2006) 359e370. [20] N. Evans, R. Oreffo, E. Healy, P. Thurner, Y. Man, Epithelial mechanobiology, skin wound healing,and the stem cell niche, J. Mech. Behav. Biomed. Mater. 28 (2013) 397e409. [21] M. Zöllner, A. Tepole, A. Gosain, E. Kuhl, Growing skin: tissue expansion in pediatric forehear reconstruction, Biomech. Model. Mechanobiol. 11 (2011) 855e867. [22] M. Silva, M. Boyer, K. Ditsios, M. Burns, F. Harwood, D. Amiel, R. Gelberman, The insertion site of the canine flexor digitorum profundus tendon heals slowly following injury and suture repair, J. Orthopardic Res. 20 (3) (2002) 447e453. [23] M. Schönberger, M. Hoffstetter, 6- emerging trends, in: Emerging Trends in Medical Plastic Engineering and Manufacturing, Elsevier, Oxford, 2016, pp. 235e268. [24] H. Tao, S.-W. Hwang, B. Marellia, B. An, J. Moreaua, M. Yanga, M. Brencklea, S. Kimb, D. Kaplana, J. Rogers, F. Omenett, Silk-based resorbable electronic devices for remotely controlled therapy and in vivo infection abatement, Proc. Natl. Acad. Sci. U. S. A. 111 (49) (2014) 17385e17389. [25] E. Korol, K. Johnstyon, N. Waser, F. Sifakis, H. Jafri, M. Lo, M. Kyaw, A systematic review of risk factors associated with surgical site infections among surgical patients, PLoS One 8 (12) (2013) 83743. [26] E. Green, J. Mecsas, Bacterial secretion systems - an overview, Microbiol. Spectr. 4 (1) (2015) 1e32. [27] D. Sharp, P. Gladstone, R. Smith, S. Forsythe, J. Davis, Approaching intelligent infection diagnostics: carbon fibre sensor for electrochemical pyocyanin detection, Bioelectrochemistry 77 (2010) 114e119. [28] D. Sharp, S. Forsythe, J. Davis, Carbon fibre composites: integrated electrochemical sensors for wound management, J. Biochem. 144 (1) (2008) 87e93.

148

Advanced Technologies and Polymer Materials for Surgical Sutures

[29] M. Fernandez, Z. Upton, H. Edwards, K. Finlayson, G. Shooter, Elevated uric acid correlates with wound severity, Int. Wound J. 9 (2) (2012) 139e149. [30] T. Dargaville, B. Farrugiaa, J. Broadbent, S. Pace, Z. Upton, N. Voelcker, Sensors and imaging for wound healing: a review, Biosens. Bioelectron. 41 (2013) 30e42. [31] A. Hasmann, U. Gewessler, E. Hulla, K. Schneider, B. Binder, A. Francesko, T. Tzanov, M. Schintler, J. Van der Palen, G. Guebitz, E. Wehrschuetz-Sigl, Sensor materials for the detection of human neutrophil elastase and cathepsin G activity in wound fluid, Exp. Dermatol. 20 (6) (2011) 508e513. [32] J. Edwards, S. Gaston-Pierre, A. Bopp, W. Goynes, Detection of human neutrophil elastase with peptide-bound cross-linked ethoxylate acrylate resin analogs, Pept. Res. 66 (4) (2005) 160e168. [33] J. Edwards, S. Caston-Pierre, P. Howley, B. Condon, J. Arnold, A bio-sensor for human neutrophil elastase employs peptide-p-nitroanilide cellulose conjugates, Sens. Lett. 6 (4) (2008) 518e523. [34] M. Reches, K. Mirica, R. Dasgupta, M. Dickey, M. Butte, G. Whitesides, Thread as a matrix for biomedical assays, ACS Appl. Mater. Interfaces 2 (6) (2010) 1722e1728. [35] X. Li, J. tian, W. Shen, Thread as a versatile material for low-cost microfluidic diagnostics, Appl. Mater. Interfaces 2 (19) (2010) 1e6. [36] A. Tamayol, M. Akbari, Y. Zilberman, M. Comotto, E. Lesha, L. Serex, S. Bagherifard, Y. Chen, G. Fu, S. Ameri, W. Ruan, E. Miller, M. Dokmeci, S. Sonkusale, A. Khademhosseini, Flexible pH-sensing hydrogel fibers for epidermal applications, Adv. Healthc. Mater. 5 (2016) 711e719.

CHAPTER 7

Bioactive sutures: advances in surgical suture functionalization Rukhsar Shah1, Louise Taylor1, Morvarid Saeinasab1, Ximu Zhang2, 3, Wei Zhang4, Karthik Nair5 and Farshid Sefat1, 6 1

Department of Biomedical and Electronics Engineering, School of Engineering, University of Bradford, Bradford, United Kingdom; 2Chongqing Key Laboratory of Oral Disease and Biomedical Sciences, Chongqing Municipal Key Laboratory of Oral Biomedical Engineering of Higher Education, Stomatological Hospital of Chongqing Medical University, Chongqing, Sichuan Province, China; 3State Key Laboratory of Oral Diseases & National Clinical Research Center for Oral Diseases & Dept. of Preventive Dentistry, West China Hospital of Stomatology, Sichuan University, Chengdu, Sichuan Province, China; 4State Key Laboratory of Polymer Materials Engineering, Polymer Research Institute at Sichuan University, Chengdu, Sichuan Province, China; 5Summit Medical Ltd., Bourton on the water, United Kingdom; 6Interdisciplinary Research Centre in Polymer Science & Technology (Polymer IRC), University of Bradford, Bradford, United Kingdom

7.1 Introduction Bioactive sutures are sutures that have been tissue-engineered with the addition of a bioactive substance, which is organically found compounds that can come from metals, plants and marine organisms [1] that have a positive effect on the body [2]. There are currently several types of bioactive implants being used clinically to date such as implants for orthopedic injuries [3], ocular implants [4], middle ear implants [4], cancer treatments [4] and dentistry implants [4] with one of the first uses of medicinal bioactive compounds being calcium phosphate-coated dental implants that were introduced in the 1980s [5]. Bioactive sutures are aimed to be used within a medicinal outlook where the suture will be applied with the aim of optimizing healing alongside their essential function of a suture [6]. Bioactive compounds often form the basis of new drugs [7] and thus drug-eluting sutures can often be confused with bioactive sutures, however, these are two separate developments. Currently, there are no bioactive sutures within clinical practice [8], however, there are several types of bioactive sutures in development within various stages of progress.

7.2 Suture structure As a bioactive suture is an adapted suture, it is important to understand the makeup of each type of suture to be able to apply bioactive compounds Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00007-3

© 2023 Elsevier Ltd. All rights reserved.

149

150

Advanced Technologies and Polymer Materials for Surgical Sutures

effectively. There are several different types of suture structures that can be characterized in various ways however the most commonly used methods of characterization are either absorbable and nonabsorbable or Monofilament and multifilament [9]. 7.2.1 Absorbable sutures Sutures are classified as absorbable when the suture materials are degraded by hydrolysis of the ester linkages where the material is then excreted or metabolized by the body [10]. Although absorbable sutures can derive from natural and synthetic materials if synthetical, macromolecules of the material can be left in the wound site as residual debris from the degradation of the suture. These macromolecules can then cause the surrounding tissue to inflame due to the foreign materials present and thus a perfected synthetic absorbable suture is yet to be found [11]. An absorbable suture can be created by both monofilament and multifilament methods, each having its own strengths and absorption rates. To use absorbable sutures within a bioactive design, the original material is commonly chemically changed to be more advantageous such as adding NaOH to PLGA sutures, most commonly known as vicryl, which then alters the surface of the suture to become more hydrophilic, increased surface area, altered porosity for additional coatings and an increased degree of roughness to the surface [12]. 7.2.2 Nonabsorbable sutures As the name states, nonabsorbent sutures will not break down and thus are required to be physically removed if wished. Clinically there can be a need to ensure that the suture remains physically in place and thus the chosen suture needs to be nonabsorbable, more often internal sutures, or superficial skin closure [13] however the structure can be either mono or multi filamented to further suit the needs of the procedure. 7.2.3 Monofilament sutures This is a suture that is made from one continuous strand of material, which can be constructed from both natural and synthetic materials. It has been found that monofilament sutures have a reduced incidence of surgical site infection (SSI) than multifilament sutures due to the bacterial holding a weaker bond to the suture as the surface is smoother than that of a braided suture [14] alongside the reduced surface area than the braided suture touching the tissue fluid also reducing the risk [11]. Monofilament sutures are known to be harder to work with due to the characteristics of being weaker sutures in general with poor flexibility and knot strength [11].

Bioactive sutures: advances in surgical suture functionalization

151

7.2.4 Multifilament sutures Multifilament sutures can still be made up of just one material, although a combination can also be created, where multiple strands are braided together to form a stronger suture [11]. Sutures that are multifilament naturally have more spaces between the interwoven strands and thus can be seen as advantageous to use with coatings over the monofilament as more volume of coating can be enclosed within the suture alongside multifilament sutures can be improved with coatings, such as improving gripping [15].

7.3 Fabricating bioactive suture methods 7.3.1 Fiber level To ensure that the bioactive suture will engage with the tissue, it is essential that the body recognizes and responds to the structure of the suture. One already existing method within tissue-engineered implants is to create a structure, which mimics the natural extracellular matrix (ECM) components. This can be established by fabricating the structure of the implanted element at a submicron scale, which not only matches the ECM component but also creates a high surface area to volume ratio and porosity [16]. To apply this technique to bioactive sutures, it has been suggested by Abharri et al. [8] to mimic the collagen fibrils within the ECM, which is the main component of the extracellular matrix within tendons and ligaments, would one approach to creating a base of a bioactive suture, that can then be further developed for additional specific uses. Collagen fibrils within the ECM vary in diameter depending on age and have been found that the smaller the collagen fibrils present, the reduced amount of scarred tissue occurs. This will need to be taken into account when designing the collagen fibril based bioactive suture for the correct application for different targeted tissues and age groups. The same application of mimicking the ECM can be applied to soft tissue areas by identifying the important component of the ECM of the intended tissues and mimicking the structure onto the suture surface. This technique is particularly promising within stem cell biomaterial implantation, as this will help determine the final fate of the stem cell adaption to be the correct format for the proliferation of the tissue [17]. Alongside the application of the suitable fiber structure of a bioactive suture, the intention of the bioactive structure is to encourage and

152

Advanced Technologies and Polymer Materials for Surgical Sutures

accelerate the cell proliferation where the suture is situated and thus the structure requires to be porous. This then creates a difficult balance between mechanical strength and porosity, by controlling pore size and frequency, to ensure that this will be a successful bioactive suture [8]. 7.3.2 Cell and gene activators Further incorporating cell or gene activating sequences within a bioactive suture will boost the rate of healing due to the increased potential to stimulate the natural repair response that the body processes [8]. This method has been heavily trialed with varying levels of success, which will be further elaborated on within the chapter. 7.3.3 Stimuli responsive An alternative approach to bioactive sutures is to use a stimulus from the surrounding intended environment to trigger a release of bioactive material, which has been embedded into the suture [8]. Similar sutures that possess stimuli triggered response, such as the shape memory suture that is temperature stimulated, which the same approach can be tailored to the bioactive suture. A chemically programmed polymer would be used to hold the intended bioactive compound, which would be released when a stimulus, either external or internal, is exhibited [18]. Currently, there are several proposed stimuli that can be used in this style of suture including the external scope of the magnetic field, ultrasound, light and electrical fields, and internally through pH, temperature, mechanical loading and redox potential [8]. This is one of the approaches to eluting sutures where research can share bioactive and eluting approaches to using a suture as a physical carrier for improving the healing process, thus adding to the confusion of the term bioactive sutures being used for drug-eluting sutures. Currently, the author found no bioactive studies, which included the stimuliresponsive sutures as this is a more suitable technique for eluting sutures as there are several other techniques to achieve a bioactive suture. 7.3.4 Researched bioactive suture There are several different types of bioactive sutures that are currently being researched. The following studies are a selection of the various types representing the various application of the same type of suture, however as there are no active bioactive sutures currently in clinical use, there are various methods of constructing the sutures. Due to the high variations, this

Bioactive sutures: advances in surgical suture functionalization

153

has been grouped into the type of bioactive sutures as opposed to applications as the methods may not be comparable.

7.4 Cell based bioactive sutures 7.4.1 Stem cells The tissue wound healing process occurs within three stages of inflammation, proliferation and remodeling, which involves several different types of cells to carry out the interlinked mechanics of healing, however, it has been repeatedly identified that mesenchymal stem cells (MSCs) play a key role within promoting the healing process [19]. MSCs are a type of stem cell that the body produces within the bone marrow of long proximal bones, which can differentiate into different types of cells and therefore are highly sought after to be incorporated within tissue-engineered products. Their application within a suture would benefit the healing process in several medical fields and a precursor to a bioactive suture containing MSCs has already been trialed with fibrin spray [20]. 7.4.2 Stem cells for wound healing Falanga et al. [20] researched the use of MSCs being incorporated within the medium of a biocompatible spray to apply onto a wound to accelerate the healing of skin wounds. MSCs were selected due to their capabilities of differentiating into several different types of cells including fat, bone, muscle, cartilage and endothelium. The intended optimum use of this creation is to stimulate the tissue repair in chronic wounds where other treatments, including growth factors that are a more targeted approach, have not been able to stimulate the tissue repair process. Samples of bone marrow were collected from the recipient’s iliac crest within the pelvis of 35e50 mLs, which was then separated into 4 mL samples into conical tissue culture tubes that were then centrifuged. The mononuclear layer is then removed and plated into tissue culture flasks before being incubated for 48 h at a body temperature of 37 C. From this original culture, the cells are passaged every 3e4 days until the number of cells reaches the proposed level for the intended use. In this instance, the MSCs were intended to be incorporated into an already established fibrin spray for topical wound use. The study found that any MSC passages above 10e12 showed that cell morphology was altered and thus would be unusable alongside the doubling time increased within the greater number of passages.

154

Advanced Technologies and Polymer Materials for Surgical Sutures

In vivo experimentation of the fibrin, a spray used titration to identify the ideal concentration of the two elements of the fibrin spray to be used to form a suitable consistency of gel, which was identified to be 5 mg/mL fibrinogen and 25 U/mL of fibrin. Once this was established, murine wounds were then trialed with the new fibrin MSC spray and found that on full-thickness wounds, the speed of healing increased at the day 10 mark, and was seen on day 20 with the diabetic mouse, which has a slower healing process than a healthy mouse. Staining the cells showed that the added MSC although increasing the healing process, were not evident in longterm structures that were created within the healing process showing that although MSC will increase healing speed, the added cells are a stimulant rather than being used within the repair process as a structural element. Human trials were then carried out on acute surgical wounds and chronic wounds with a 2-week window given before the procedure to acquire and culture MSC cells in the acute wounds to allow application on day one of the wound. The application of the fibrin spray was up to three applications, a week apart, with no greater than 2 mL applied to the wound. Within the acute wounds, the full-thickness wound was resurfaced by the 6-week mark, which remained healed at the end of the trial at week 12. Chronic wounds within this trial were at least 1 year old and were positioned on the leg or foot. The treatment regime was the same as the acute wounds and showed that by weeks 16e20, chronic wounds had either healed or become smaller in size with all wounds bar one reducing size within weeks 2e4. The positive response to the application of MSCs within a fibrin spray allows the possibility that MSCs can be added to sutures made from preexisting collagen to increase the rate of healing in many types of skin wounds. This has been further studied by Reckhenrich et al. [19] who have developed a biodegradable suture, which contains adipose-derived MSCs (ASC).In this study, ASC original sample was taken from the patient’s abdomen, flanks and legs before being processed through digestion with the enzyme collagenase A at 37 C. Once the fat emulsion took place, the enzyme was neutralized and the remaining substance was then centrifuged. The pellet was then suspended in saline before being filtered through a 100 mm pore strainer. After a final centrifuge, the pellet was resuspended and transferred to culture flasks to start the proliferation of ASC cells. Once ASCs were cultured to the required numbers, a biodegradable suture, which was vicryl (polyglactin 910), was cut to 3 cm size and ASCs were

Bioactive sutures: advances in surgical suture functionalization

155

transferred via syringe. The ASC sutures were incubated overnight in a culture flask with a medium for the first 24 h before being separated into independent wells containing 1 mL of culture medium to continue cultivation. Assessment of the ASC sutures was carried out including tensile testing, which found that the ASC sutures were more flexible than the original viryl sutures, however, were reduced in tensile strength by 28%. The ASC suture was then trialed in an ex vivo model where human skin samples were used. Cell viability was assessed presuturing and postsuturing, which established that there was no significant decrease in metabolic activity of cells after suturing took place. This trial also involved frozen ASC sutures to assess the impact of storage in freezing, which will be required for clinical use of the suture, where the metabolic rate was reduced by 22% after thawing. Cell migration was evident through the ex vivo model suggesting that this mechanism of delivering ASCs through the center of a suture would be compatible with in vivo testing. Within this study, the assessment of the ASC success was measured in the several cytokine secretions including growth factors that have been proved to be involved within all three stages of the healing progress. Reckhenrich et al. suggest that this model could be suited to an in vivo trial however the next suggested point of interest would be to determine if the ASC’s could be allogeneic rather than donated from the recipient as this will decrease the time required to use such a suture and open up the uses from chronic wounds to various types of application. This model of an ASC biodegradable suture has the clinical potential to be successful, however, this has not been modeled with an in vivo testing to ensure that the cytokine secretions are impactful and do increase wound healing thus the future of this model needs further studies. 7.4.3 Stem cells e cardiovascular application Although previously mentioned studies were specifically aimed at stem cell application to accelerate wound healing, stem cells within sutures have other applications due to their adaptive abilities to become several types of specialized cells. Hansen et al. [21] have studied the use of MSCs within a suture with the aim of improving the functionality of the heart after suffering a myocardial infarction (MI). There have been several studies, which have documented the use of MSCs being delivered intravenously to patients after suffering MI’s through

156

Advanced Technologies and Polymer Materials for Surgical Sutures

various clinical trials have shown improvements in both cardiac and pulmonary scopes [22], however, these benefits are usually do not last longer than 1 year [23] and thus a more long term solution would be beneficial. Current MSC therapy deliveries have an insufficient grafting rate with intramyocardial injection resulting in approximately 10% retention and transendocardial delivery resulting in an approximately 19% cell retention. This, therefore, demands the initial cell population to be extensively large to enable the 10%e19% retention of MSCs to replace a billion myocytes that are lost due to MI [24]. Hansen et al. [21] developed a fibrin-based suture that was seeded with MSCs, using the findings of a previous study by Guyette et al. [25] where this style of suture delivery of MSCs was carried out on rats, which found an engraftment rate of 64%. To build on the previous work, Hansen et al. trialed three various concentrations of MSCs on rats with induced MI’s. Fibrin microthreads were created and then build into a 24-strand suture, which was seeded with MSC solutions with the concentrations of 25, 50 and 100 k cells per 4 cm of a suture. The findings from this study show that 100 k MSC concentration was most beneficial as cardiac analysis after the studies showed MSCs were present within other areas of the heart 1 week after the procedure showing that MSC cell migration can occur through the suture into damaged tissues. Using rats as the subject of this testing was disadvantaged due to typical ventricular remodeling after MI occurring at 4e6 weeks postMI, the study length was time constricted to 1 week following the rules of the animal studies for the area that the study took place. Although this study did not find statistical proof that MSC delivery via suture was beneficial to improving MI heart function, it did identify the most beneficial concentration of MSC and highlighted the fact that although MSCs are able to specialize and improve the healing process, alternative stem cells such as the pluripotent and embryonic stem cells, which have been demonstrated to develop into cardiomyocytes, which are identified to be ideal for cell replacement applications [26]. 7.4.4 Stem cells e tendon repair Tendon repairs are known to be a slow and complicated process, which are prone to becoming damaged early postoperatively alongside being weaker than healthy tendons and thus can continue to be weak and can become detached over prolonged use or excessive uses [27]. Developing methods in

Bioactive sutures: advances in surgical suture functionalization

157

which tendon repairs can heal sooner is extremely beneficial due to the reduction in early postoperative failure, alongside the possibility of strengthening the tendon overall prolonging the timescale of the repair. There have been several studies in which tendon repairs have received the addition of bone marrow-derived MSCs via various delivery methods of injection or tissue-engineered cell scaffolds. Both delivery methods of MSCs produce three positive processes within tendon repair, including the release of growth factors and cytokines, host cell recruitment and direct differentiation of stem cells [28]. Adams et al. [27] highlighted the opportunity to develop another method of MSC delivery via biological sutures due to the use of a woven suture being used within nearly all types of tendon repair. The use of a biological suture would be twofold by replacing the current woven suture by maintaining the mechanical stabilizer alongside the benefits of accelerating the healing process. Within the in vivo study, Adams et al. created groups of the rats to partake in one of three treatment paths: suture only, suture and injection and stem cell suture. The concentration of the stem cells was kept consistent at 1  106 alongside the same suture type throughout all groups providing a fair comparison of the treatment methods to establish the stem cell impact depending on the delivery method. Stem cells were acquired from an iliac crest bone marrow aspiration donated by a human donor, which is in line with previous stem cell studies on rat models. The stem cells were loaded into the center of the braided suture in a 2 cm portion that was marked to ensure the suture was used within the correct part of the suture during the tendon repair in the figure of eight repair suture and knot. Once treated, the tendons were analyzed blindly in two-time frames of 14 and 28 days with biomechanical tests of ultimate failure load and crosssectional areas and with histological grading. This histological grading was carried out according to the Hospital for Special Surgery (HSS) where the lower the grade, the closer to a healthy tendon the structure is. Findings showed that the ultimate failure load was increased significantly with both delivery methods of MSC however in the 28-day group, the ultimate failure load decreased within the injected group and not the stem cell suture. There were no notable observations with cross-sectional measurements, which can indicate scar tissue within the tendon, across all three groups in both time frames. Histologically, the stem cells suture group in the 14-day group had the best grading however all groups in the 28-day period were higher and therefore show a degrading in the tendon

158

Advanced Technologies and Polymer Materials for Surgical Sutures

structure, in particular with the collagen orientation and the number of fibroblasts present. Within Adams et al. [27] study model, rats were able to move without restrictions after the tendon repair thus stressing the repair immediately postoperative, which would not be the same as the human repair where immobilization is implemented during the initial healing process. The application of this style of suture based on the findings can be applied with two different approaches. These results find that the MSC suture has positive immediate effects on the tendon repair at 14 days of full movement, this finding indicates that this will be the same in human applications and have the possibility of being a prolonged time frame due to the immobility imposed postoperatively. The immobilization period after human tendon repair has shown that scar tissue can form and this, therefore, leads to a reduced strength or function of the repair, however, this MSC suture can accelerate the repair and reduce the immobilization time frame as opposed to the lengthening time frame of the accelerated healing. A previous study by Dines et al. [29] established that growth factors can be successfully delivered in tendon repairs using the external sutures as a delivery method and thus stem cell sutures within tendon repairs are a realistic goal. Despite the positive effects of MSC sutures, all tendon repair studies have been tested in vivo, which have had a short follow up and analysis time frame, which may not be representative of long-term results. Alongside this, there are no studies, which compare the results of the various MSC delivery methods currently being researched within tendon repair nor studies, which establish the long term effects of stem cell therapies is required before this can become a realistic option for clinical use. Lui [30] has identified that to further improve the stem cell therapy approach to tendon repairs there needs to be an increased understanding of the embryonic tendon development due to the fetal tendon healing process does not produce scarred tissue. 7.4.5 Stem cell suture conclusions Due to the body’s natural immune system, any cells that will be used within any type of implant must originate from the recipient’s own body and therefore any device, which uses MSCs require cell growth through cultures to enable the device to be effective. The process of obtaining the original MSC cells from the bone marrow involves steps to identify MSCs correctly from the combination of several cells from within the bone

Bioactive sutures: advances in surgical suture functionalization

159

marrow structure before cultivation can occur. This in turn inflicts a wait time in producing the sutures and this wait between recognizing the need for such a device and the fabrication of the suture, therefore, becomes a preplanned application and cannot be used in an acute or emergency setting. Although the proliferation of stem cells is not widely used in medicine to date, there are indications that the stem cell banks and allogenic MSCs can become a feasible option in the near future [31]. Stem cell therapy, irrelevant of whether this is delivered through sutures or another method, involves living cells, which introduces a new set of obstacles when it comes to the manufacturing and regulatory processes [30]. Specific problematic areas of cell transplantation processes include the scaling up of production, donor variability, maintaining structurally sound and viable cells through several culture passages within minimal mutations, and immunological responses to the alloantigens on transplanted cells [30]. However, there are several advantages to using MSCs in wound management as there are several cytokines and growth factors that are secreted as opposed to a singular biological factor of other bioactive suture designs [32]. Due to the involvement of a live cell, there will be variations within the regulation and classification of stem cell bioactive sutures dependent on the country that may cause delay in the introduction of such a suture into clinical practice [32]. 7.4.6 mRNA suture An alternative method for stimulating the natural healing process has been designed by Link et al. [33] where specifically encoded mRNA is used as a suture coating to trigger the natural production of growth factors. mRNA techniques are already founded through the production of vaccinations and have established production kits to aid in the building of the mRNA. This process includes the process of obtaining and passaging cells from the host and therefore has the same time limitations as stem cell-based sutures. This mRNA suture was tested in vitro with the forehead cells alongside additional tests with the sutures reaction to blood to ensure that coating does not affect the blood clotting response. Results from the in vitro testing showed that the mRNA suture has a transfection of growth factor success rate of 29.11%, which is an increased response to topic growth factor application where less than 10% of the dose is observed within the granulated tissue and 2% in deeper levels [34]. The study found that the amount of growth factor transfection by the mRNA suture was evident in

160

Advanced Technologies and Polymer Materials for Surgical Sutures

improved quantities however the amount of specific growth factor involved in the healing process, KGF, secretion levels were lower than if the mRNA was directly used for transfection. However, this process has advantages compared to DNA based genetic approach as the mRNA is a temporary change that is introduced to the host by suture and is removed and thus avoids the chance of the host obtaining a mutational response as the nucleic acid does not enter the nucleus. Alongside this, the tested length of suture was 1 cm, which is considerably smaller than the length that would be used within a clinical setting and thus the rate of growth factor production may be more substantial when longer lengths are used. Further testing and the possible adaption of the suture are needed on a skin model or ex vivo before this style of suture can be considered for any clinical use. This is a novel approach to accelerating wound healing with mRNA sutures and thus is at the early stages of the study. 7.4.7 Gene regulation Gene activation occurs when a stimulation, normally a physical-biological element, stimulates the natural proteins that are part of the genetic material of cells and is thus turning the genes “on” that were “off” and is therefore considered to be gene-regulating. Incorporating these elements into sutures will stimulate the body through natural means to accelerate the healing process and does not necessarily need a cellular component and can therefore be more appropriate for clinical use as the need for the device may not be planned like that cell therapy styles [8]. 7.4.8 Growth factor bioactive suture Growth factors are a peptide or hormone that is involved in the stimulation of cell division and multiplication [35] and are therefore a vital part of the healing process with different tissue structures having a specific growth factor for the area. The scope of using growth factors within bioactive sutures is vast providing that the most suitable growth factor can be identified in the targeted area. To date growth factors are used systemically using collagen sponges, catheters, injections or as nutritional supplements, which aid in the healing process of several structures including bone, cartilage and gastrointestinal tract [36]. Currently, the author found two areas where studies were concentrating on the growth factor of sutures on tendon repair and gastrointestinal surgery.

Bioactive sutures: advances in surgical suture functionalization

161

Linderman et al. [37] highlighted the lower rates of successful tendon repairs to intrasynovial tendons compared to extrasynovial tendons to be targeted for a bioactive suture to increase the rates of a successful repair. Further developing on previous attempts of improved rates through bioactive scaffolds and directly with both growth factors and MSc, the study trialed the use of suture as the carrier of growth factor. A nonabsorbable suture was made porous through a swelling and freeze-drying technique creating a 1 mm pore within the outer sheaths to hold the Connective Tissue Growth Factor (CTGF). The space created by the rearrangement of strands can be seen through the SEM images in Fig. 7.1 before (A,C) and after the procedure (B&D).

Figure 7.1 SEM images of the cross sections and side surfaces of both unmodified and porous sutures. Scale bars: 50 mm in panels (A, B) and 2 mm in panels (C, D). (Reproduced from Ref. [37] with permission from PMC, copyright 2018 e Open Access [37].)

162

Advanced Technologies and Polymer Materials for Surgical Sutures

Once the sutures had been prepared with the pores, sterilization by gas was then loaded with a heparin and fibrin delivery system of the growth factor CTGF. Sutures were then incubated at 37 C to ensure the population of CTGF reached the acquired number. This in vivo study used 10 female canine’s flexor tendon intrasynovial repairs, using two different digits per dog with a control group of five received an empty porous suture for the repair. Aftercare comprised of fully flexed and extended digits for 5 min, every day for 6 days a week to mimic the clinical scenario as completely as possible. The study lasted 14 days where cross section analysis was then carried out to assess the healing process and effects of the growth factor sutures. Histological analysis revealed that no inflammation was present and a higher number of cells within the growth factored sutures, particularly at the repair site. Staining also revealed that new collagen was elevated in the growth factor sutures than the control alongside the creation of blood vessels in this group was also present in the growth factor suture with a noted distance from the repair site. Biomechanically, ex vivo analysis shows that the adapted growth factor suture remained robust and was in comparison to the control group. Fuchs et al. [36] identified the need to establish a carrier for a growth factor that is bioabsorbable and biocompatible to minimize the immune response being triggered and thus reduce the healing process. The study then proceeded to design a coating for suture materials, which will be the growth factor carrier and be able to be applied to various types of bioabsorbable sutures and thus be beneficial to varying different types of surgeries. The basis of the coating is the biodegradable polymer Poly (D, L-lactide), known as PDLLA, which has been used in previous biological coating studies. The aim of this study was to assess the effects of a growth factor suture on colorectal resections in an animal model of rats. The chosen growth factor was insulin-like growth factor, which has been previously studied in rat models for the same colorectal resections. Postoperatively, the colon resection was assessed in three groups based on time on either day 1, 3 or 7. The analysis showed that the coated sutures increased the rate of healing compared to the control group through the biomechanical assessment where the untreated group achieved the same rate on day 7 as that on day 3. The application of a growth factor suture is not yet ready for clinical trials however is well on its way for become a valid suggestion for being clinically relevant. Both internal application of growth factor and external coating have been shown to be clinically successful and both tested in vivo.

Bioactive sutures: advances in surgical suture functionalization

163

The application of a suture will require time for the assembly of the suture as this needs to be incubated, similarly to cell therapy sutures, however, do not need a specific donor as can be bought as a separate product. However, due to the addition of more textures on the external of the suture, the suture has been noted to become more resistant to being sown. The use of growth factors is already established within a clinical setting and thus may allow the fabrication of the sutures to become more readily available than previous styles of cell therapy sutures.

7.5 Incorporated bioactive material Bioactivating materials is achieved by modifying the material to create new biological characteristics [38]. Because this can be done to a wide range of particles, there is a wide scope of varying types of bioactive materials. One group, in particular, is called bioactive glass, although the element of biological elements varies, was first established in 1970s with a silicate-based glass, which was then named 45S5, or commercially known as Bioglass [39]. 7.5.1 Chitin bioactive sutures Chitin is a natural element, which has biological characteristics have been increasing in scientific interest to be incorporated into tissue engineering structures. Chitin is recognized as the second most abundant polysaccharide after cellulose [40] and has a biological effect on genes that are intrinsic to healing, in particular with the inflammation stage [41]. Zhang et al. [11] developed a study that developed a Chitsin-based suture, which was created using a derivative of chitin called chitosan, which was strengthened with the addition of graphene oxide for mechanical strength alongside the future possibilities of utilizing further suture developments such as thermal and electronic properties. This suture was then used in rat model to assess the biocompatibility of the finished monofilament suture alongside in vitro assessments. In vitro cytotoxicity assessments showed that cell viability remained above 90% alongside the suture showing no inflammation responses further confirming that the suture has acceptable biocompatibility for its role. The designed chitsin suture showed that genetic expressions of inflammatory response remained in line with the normal healing process whereas antiinflammatory gene TGF-b showed a decrease within the 3e7 day window, which would suit the role of increasing healing rates as a suture.

164

Advanced Technologies and Polymer Materials for Surgical Sutures

An alternative approach to chitin sutures has also been studied by Altinel et al. [42] where chitosan was used as a suture coating to provide more strength in the healing process of colonic surgeries by preventing the common complications when an inflammation response occurs at the surgery site. This study compared the effects of chitosan on both monofilament and multifilament sutures where the sutures were used to close wounds created in the intestinal tract alongside a second comparison group of the application of chitosan directly to an induced bleed to the intestinal tract. Sutures were created by adding acetic acid to chitosan powder to create the solution that was then cured onto the sutures after 30 min of incubation. Histological analysis was carried out blindly of the in vivo studies showed that there were slight differences between the noncoated sutures and the chitosan sutures, which included the vascularization of the chitosan group was lower than the original sutures after 2 weeks. This coincides with the supportive finding that neovascularization was lowered in chitosan groups showing that chitosan is actively reducing the incidences of abnormal blood vessels but also impacts the body’s ability to form new vessels within the healing process. There were no noted differences histologically between the suture groups in terms of inflammation or collagen accumulation. The study also focused on the adhesion of the sutures and the effects the coating would have. Adhesion was measured using the diamond classification scale, which noted that there were no significant differences between the varying types of sutures observed. This process of coating a suture is easily replicable at a larger scale and thus would be an achievable product to become readily available alongside no additional storage requirements. Despite there being no significant difference in the addition of a chitosan coating in this application, the approach of coating a suture successfully where there was no degradation to the original characteristics of the suture. There may be a more suitable application of a chitosan-coated suture due to the recognized biological impacts chitin has. 7.5.2 Bioactive glass for antibacterial sutures Bioglass has been known to bond to both soft tissues alongside boney structures [43] and is therefore is a versatile biological that can be added to sutures that will provide a vast range of applications. Several variations of Bioglass sutures have been trialed with varying degrees of success and include differing additions to Bioglass.

Bioactive sutures: advances in surgical suture functionalization

165

Ciraldo et al. [43] focused on a study, which aided in comparing the various Bioglass suture techniques that have been established in the previous 10 years. The variations of Bioglass included Chitosan and Silver Mesoporous Bioglass (MBG), silver MBG, Chitosan and Zinc Bioglass and Zinc Bioglass. MBG is where the Bioglass substance has a structure, which contains pores between 2 and 50 mm. For each type of suture, the durability of the coating was tested using a knot test alongside the effect the coating has on the original suture characteristics. Using SEM analysis, all variations showed coating deterioration, in particular with the chitosan and Bioglass coating with the observation that the monofilament suture with Bioglass coating was less flexible and constricting knotting movement. Sutures were then tested for bioactivity, which was carried out in vitro with a Synthetic Body Fluid (SBF) for a week. This found that chitosan and Silver MBG were homogeneously covered by a layer of cells however the silver MBG and Zinc Bioglass did not form an ant layer of cells despite being submerged for 28 days. From an antibacterial point of view, the sutures were tested for antibacterial properties against both gram-positive and negative bacteria in comparison to their respective uncoated suture. As seen in Fig. 7.2 above, the chitosan and Silver MBG, and chitosan Zing BG provided a clear response to both types of bacteria, despite the Zinc BG being less bioactive than previously thought. Ultimately there is great potential for the use of Bioglass within sutures, however further testing will be required to completely understand the

Figure 7.2 Agar-disk diffusion test with (i) gram-positive e S. camosus and (ii) gramnegative e. coli. Suture segments are (A) VICRYL Plus, uncoated, (B) PCL/Ag-MBG, (C) chitosan/Ag-MGB, (D) chitosan/Zn-BG and (E) PCL/Zn-BG. (Reproduced from Ref. [43] with permission from MDPI, copyright 2019 e Open Access [43].)

166

Advanced Technologies and Polymer Materials for Surgical Sutures

correct selection of Bioglass additives alongside the concentration in the coating. Once perfected the applications of such a suture is wide, however, there may be different types of coatings more responsive within different types of body tissue.

7.6 Future developments of bioactive sutures 7.6.1 Surface architecture sutures There are promising initial investigations, which explore the varying effects that specifically designed shaping on a nanoscale will have on the body’s ability to increase healing rate, which could then be applied to implant structure, including sutures. Kim et al. [44] conducted a study where three varying ratios of nanogrooves inflicted on a scaffold of synthetic ECM to ensure minimal interference with the scaffold material. Nanogroove size was selected to mimic the natural size of collagen fibers, which occur during the healing process. Ex vivo studies were carried out with fibroblasts, which are involved in the majority of wound healing processes where various directions of grooving were carried out in the three ratios to ensure the most inclusive representation of various wound types. Results showed that grooves on a ratio of 1:2 spacing possessed the best wound healing properties including the greatest migrations speed and the formation naturally. This study creates a sturdy basis for further analysis or pre-existing and new implantation technology, which concerns the acceleration of healing alongside other applications to further increase the incorporation of an implant, including biological sutures to aid the body’s ability to accept the suture more readily and thus been effective sooner. This method of surface modification may not be solely beneficial on its own however in conjunction with another bioactive application, can further develop the suture design and results.

7.7 Conclusion There are various areas in which bioactive sutures are being explored however only a handful of designs are close to a randomized controlled trial process. Despite the scientific benefits of a suture design, the thought process must include the practicalities of a suture within the clinical application as cell therapy sutures are not readily available and thus may become less beneficial than anticipated within the practice. Another aspect to consider is the classification of some sutures may become more complex

Bioactive sutures: advances in surgical suture functionalization

167

and thus become more challenging to become a mainstream treatment choice as various countries have different classifications, which will also slow down the process of becoming in clinical use. Despite these barriers, there is great scope for a more effective and targeted approach for advanced bioactive sutures, which will have a great area of application when introduced into clinical practice.

References [1] G. Jones, Bioactive compounds in marine organisms, Bot. Mar. 51 (2008) 161e162. [2] H.-K. Biesalski, L. Dragstead, I. Elmadfa, R. Grossklaus, M. Müller, D. Schrenk, P. Walter, P. Weber, Bioactive compounds: definition and assessment of activity, Nutrition 25 (11e12) (2009) 1202e1205. [3] G. Kaur, Biomedical, Therapeutic and Clinical Application of Bioactive Glass, Elsevier, Cambridge, 2019. [4] F. Baino, S. Hamzehlou, S. Kargozsar, Bioactive glasses: where are we and where are we going? J. Funct. Biomater. 9 (1) (2018) 1e25. [5] D. Greenspan, Bioactive ceramic implant materials, Curr. Opin. Solid State Mater. Sci. 4 (4) (1999) 389e393. [6] F. Alshomer, A. Madhavan, O. Panthan, W. Song, Bioactive sutures: a review of advances in surgical suture functionalisation, Curr. Med. Chem. 24 (2) (2017) 215e223. [7] C. Katiyar, A. Gupta, S. Kanjilal, S. Katiyar, Drug discovery from plant sources: an integrated approach, J. Res. Ayurveda 33 (1) (2012) 10e19. [8] R. Abhari, J. Martins, H. Morris, P.-A. Mouthuy, A. Carr, Synthetic sutures: clinical evaluation and future developments, J. Biomater. Appl. 32 (3) (2017) 410e421. [9] A. Goel, Surgical sutures e a review, Delhi J. Opthalmol. 26 (3) (2016) 159e162. [10] S. Ruengdechawiwat, R. Molloy, J. Siripitayananon, R. Somsunan, P. Topham, B. Tighe, Synthesis, processing and tensile testing of a poly(l -lactide-co -caprolactone) monofilament fiber for use as an absorbable surgical suture, Macromol. Symp. 354 (1) (2015) 347e353. [11] W. Zhang, B. Yin, Y. Xin, L. Li, G. Ye, J. Wang, J. Shen, X. Cui, Q. Yang, Preparation, mechanical properties, and biocompatibility of graphene oxide-reinforced chitin monofilament absorbable surgical sutures, Mar. Drugs 17 (4) (2019) 210e224. [12] G. Park, M. Pattison, K. Park, T. Webster, Accelerated chondrocyte functions on NaOH-treated PLGA scaffolds, Biomaterials 26 (2005) 3075e3082. [13] A. Trott, Chapter 8 e instruments, suture materials and closure choices, in: Wounds and Lacerations, fourth ed., Elsevier, 2012, pp. 82e94. [14] J. Fowler, T. Perkins, B. Buttaro, A. Truant, Bacteria adhere less to barbed monofilament than braided sutures in a contaminated wound model, Clin. Orthop. Relat. Res. 471 (2013) 665e671. [15] C. Chu, Types and properties of surgical sutures, in: M. King, B. Gupta, Guidoin (Eds.), R Biotextiles as Medical Implants, Woodhead, Philadelphia, PA, 2013, pp. 231e273. [16] J. Venugopal, S. Ramakrishna, Applications of polymer nanofibres in biomedicine and biotechnology, J. Appl. Biochem. Biotechnol. 125 (3) (2005) 147e158. [17] A. Dolatshahi-Pirouz, M. Nikkhah, K. Kolind, M. Dokmeci, A. Khademhosseini, Micro- and nanoengineering approaches to control stem cell-biomaterial interactions, J. Funct. Biomater. 2 (2011) 88e106.

168

Advanced Technologies and Polymer Materials for Surgical Sutures

[18] S. Mura, J. Nicolas, P. Couvreur, Stimuli-responsive nanocarriers for drug delivery, J. Nat. Mat. 12 (2013) 991e1003. [19] A. Reckhendrich, B. Kirsch, E. Wahl, T. Schenck, F. Rezaeian, Y. Harder, P. Foehr, H.-G. Machens, J. Egaña, Surgical sutures filled with adipose-derived stem cells promote wound healing, PLoS One 9 (3) (2014) e91169. [20] V. Falanga, S. Iwamoto, M. Chartier, T. Yufit, J. Butmarc, N. Kouttab, D. Shrayer, P. Carson, Autologous bone marrow-derived cultured mesenchymal stem cells delivered in a fibrin spray accelerate healing in murine and human cutaneous wounds, Tissue Eng. 13 (6) (2007) 1299e1312. [21] K. Hansen, J. Favreau, J. Guyette, Z.-W. Tao, S. Coffin, A. Cunha-Gavidia, B. D’Amore, L. Perreault, J. Fitzpatrick, A. DeMartino, G. Gaudette, Functional effects of delivering human mesenchymal stem cell-seeded biological sutures to an infarcted heart, BioRes. Open Access 5 (1) (2016) 249e260. [22] J. Hare, J. Traverse, T. Henry, N. Dib, R. Strumpf, S. Schulman, G. Gerstenblith, A. DeMaria, A. Denktas, R. Gammon, J. Hermiller (Jr), M. Reisman, G. Schaer, W. Sherman, A randomized, double-blind, placebo-controlled, dose-escalation study of intravenous adult human mesenchymal stem cells (prochymal) after acute myocardial infarction, J. Am. Coll. Cardiol. 54 (24) (2009) 2277e2286. [23] G. Meyer, K. Wollert, J. Lotz, J. Steffens, P. Lippolt, S. Fichtner, H. Hecker, A. Schaefer, L. Arseniev, B. Hertenstein, A. Ganser, H. Drexler, Intracoronary bone marrow cell transfer after myocardial infarction: eighteen months’ follow-up data from the randomized, controlled BOOST (BOne marrOw transfer to enhance ST-elevation infarct regener, Circulation 113 (10) (2006) 1287e1294. [24] M. Laflamme, C. Murray, Regenerating the heart, Nat. Biotechnol. 23 (7) (2005) 845e856. [25] J. Guyette, M. Fakharzadeh, E. Burford, Z.-W. Tao, G. Pins, M. Rolle, G. Gaudette, A novel suture-based method for efficient transplantation of stem cells, J. Biomed. Mater. Res. 101 (2013) 809e818. [26] M. Laflamme, K. Chen, A. Naumova, V. Muskheli, J. Fugate, S. Dupras, H. Reinecke, C. Xu, M. Hassanipour, S. Police, C. O’Sulivan, L. Collins, Y. Chen, E. Minami, E. Gill, S. Ueno, C. Yuan, J. Gold, C. Murray, Cardiomyocytes derived from human embryonic stem cells in pro-survival factors enhance function of infarcted rat hearts, Nat. Biotechnol. 25 (9) (2007) 1015e1024. [27] S. Adams (Jr), M. Thorpe, B. Parks, G. Aghazarian, E. Allen, L. Schon, Stem cellbearing suture improves achilles tendon healing in a rat model, Am. Orthop. Foot Ankle Soc. 3 (3) (2014) 293e299. [28] X.-J. Wang, Z. Dong, X.-H. Zhong, R.-Z. Shi, S.-H. Huang, Y. Lou, Q.-P. Li, Transforming growth factor-beta1 enhanced vascular endothelial growth factor synthesis in mesenchymal stem cells, Biochem. Biophys. Res. Commun. 365 (2008) 548e554. [29] J. Dines, L. Weber, P. Razzano, R. Prajapati, M. Timmer, S. Bowman, L. Bonasser, S. Dines, D. Grande, The effect of growth differentiation factor-5-coated sutures on tendon repair in a rat model, J. Shoulder Elbow Surg. 16 (5) (2007) S215eS221. [30] P. Lui, Stem cell technology for tendon regeneration: current status, challenges, and future research directions, Stem Cell. Clon Adv. Appl. 8 (2015) 163e174. [31] L. Moroni, P. Fornasari, Human Mesenchymal stem cells: a bank perspective on the isolation characterization and potential of alternative sources from the regeneration of musculoskeletal tissues, J. Cell. Physiol. 28 (2013) 680e687. [32] J. Casado, R.E. Blazquez, I. Jorge, V. Alvarez, G. Gomez-Mauricio, M. OrtegaMunoz, J. Vazquez, Mesenchymal stem cell-coated sutures enhance collagen depositions in sutured tissues, Wound Repair Regen. 22 (2) (2014) 256e264.

Bioactive sutures: advances in surgical suture functionalization

169

[33] A. Link, H. Haag, T. Michel, M. Denzinger, H. Wendel, C. Schlensak, S. Krajewski, Development of a novel polymer-based mRNA coating for surgical suture to enhance wound healing, Coatings 9 (6) (2019) 374e391. [34] A. Singer, R. Clark, Cutaneous wound healing, N. Engl. J. Med. 341 (10) (1999) 738e746. [35] F. Martini, J. Nath, E. Bartholomew, The cellular level of organisation, in: Fundamentals of Anatomy & Physiology, Pearson Education Limited, Harlow, 2017, pp. 111e159. [36] T. Fuchs, C. Surke, R. Stange, S. Quandte, B. Wildermann, M. Raschke, G. Schmidmaier, Local delivery of growth factors using coated suture material, Sci. World J. 2012 (2012) 1e8 (Article ID 109216). [37] S. Linderman, H. Shen, S. Yoneda, R. Jayaram, M. Tanes, S. Sakiyama-Elbert, Y. Xia, S. Thomopoulos, R. Gelberman, Effect of connective tissue growth factor delivered by porous sutures on the proliferative stage of intrasynovial tendon repair, J. Orthop. Res. 36 (7) (2017) 2052e2063. [38] D. Williams, Definitions in Biomaterials, Elsevier, New York, NY, 1987. [39] M. Rahaman, D. Day, S. Bal, Q. Fu, S. Jung, L. Bonewald, A. Tomsia, Bioactive glass in tissue engineering, Acta Biomater. 7 (2011) 2355e2373. [40] D. Zhou, R. Yang, T. Yang, M. Xing, G. Luo, Preparation of chitin-amphipathic anion/quaternary ammonium salt ecofriendly dressing and its effect on wound healing in mice [Corrigendum], Int. J. Nanomed. 13 (2018) 5255e5256. [41] K. Azuma, S. Ifuku, T. Osaki, Y. Okamoto, S. Minami, Preparation and biomedical applications of chitin and chitosan nanofibers, J. Biomed. Nanotechnol. 10 (2014) 2891e2920. [42] Y. Altinel, S. Chung, G. Okay, N. Ugras, A. Isik, E. Ozturk, H. Ozguc, Effect of chitosan coating on surgical sutures to strengthen the colonic anastomosis, Turkish J. Trauma Emerg. Surg. 24 (5) (2018) 405e411. [43] F. Ciraldo, K. Schnepf, W. Goldman, A. Boccaccini, Development and characterization of bioactive glass containing composite coatings with ion releasing function for antibiotic-free antibacterial surgical sutures, Materials 12 (3) (2019) 423e431. [44] H. Kim, Y. Hong, M. Kim, K.-Y. Suh, Effect of orientation and density of nanotopography in dermal wound healing, Biomaterials 33 (34) (2012) 8782e8792.

CHAPTER 8

Engineering aspects of suture fabrication Smrithi Padmakumar and Deepthy Menon Amrita Centre for Nanosciences and Molecular Medicine, Amrita Vishwa Vidyapeetham, Kochi, Kerala, India

8.1 Introduction 8.1.1 Surgical sutures Surgical sutures are predominantly intended to promote wound closure, immobilize prosthesis, or repair the damage caused to tissues. They are inevitable surgical tools of clinicians owing to their role in holding tissues together in the course of the procedure, approximating body tissues to implants and ligation of blood vessels to restore blood flow. Moreover, sutures have to ensure firm, intact, tighter and hermetic joining of tissues and maintain their fixed position by imparting a constant compression throughout the course of wound healing, which also includes postoperative edema [1]. These functions emphasize the structural and functional role of sutures with respect to the restoration of body tissues, because of which they have been classified as Class II medical devices by the FDA. 8.1.2 The association of surgical sutures with wound healing cascade Surgical wound healing involves the process of matrix synthesis, which bridges the wound margins, supports the constituent cells and associated vascular cells to regenerate, thereby aiding the tissues to respond to functional stress in terms of resistance. Wound closure comprises of steps such as the elimination of dead space, along with a uniform distribution of tension throughout the suture lines [2]. There is a huge disparity between the cutaneous tensile strength possessed by normal skin and that of cut or wounded skin. The wound strength tends to increase gradually along the course of its healing cascade within weeks or months until it attains the original strength of the normal tissue [3]. Suturing plays an important role in this context wherein suture strands aid in approximating the cut tissues, without imparting excessive tension or trauma to tissues. The wound Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00013-9

© 2023 Elsevier Ltd. All rights reserved.

171

172

Advanced Technologies and Polymer Materials for Surgical Sutures

acquires 3%e5% of its inherent strength within 2 weeks postsuturing, which increases to 20% and 50% within 3 and 4 weeks respectively [4]. Therefore, the extent and success of tissue repair depend on the characteristics of sutured tissue, the material attributes of the suture as well as the technique employed for suturing. Optimal suture design and choice of the appropriate suture are therefore imperative for mitigating adverse and undesirable surgical outcomes [5]. Fig. 8.1 shows the utility of dermal and epidermal sutures in wound closure of the skin and the importance of balancing wound strength [6,7].

8.2 Why is the engineering of suture fabrication important? 8.2.1 Suture design parameters Suture design is an important attribute, which in fact governs the overall functionality of the suture. This is dependent on several parameters, which determine the overall structure, integrity, function and applicability of the suture (Fig. 8.2). The structural attributes comprise of suture size/dimensions, constituent geometry or configuration of the suture (mono or multifilament/braided), needle type, surface features like texture variations and presence of coatings, etc. The surface design conferring a smooth uniform texture to the suture is very important so as to ease its handling such that the suture easily passes or traverses through the tissue in a gentle

Figure 8.1 Steps involved in surgical wound closure with dermal and epidermal sutures. (Reprinted with permission from Ref. [7].)

Engineering aspects of suture fabrication

173

Figure 8.2 Various design parameters involved in the process of suture fabrication.

manner. The physical attributes with regard to suture design predominantly include the overall mechanical integrity of the suture constituted by straight pull tensile strength, knot pull strength, suture retention, force at break, flexibility, stiffness, capillarity and memory features. The rate at which suture degrades also determines its retention property within the body. Additionally, it is imperative for the suture material to possess adequate biocompatibility as well as sterility such that it does not cause any adverse effects, infections, inflammatory responses or undesirable reactions in tissues. The success towards the development of a strong and functional suture depends on the complete compliance with all of these factors [3,8]. The material properties of different tissues vary from each other with respect to the location, anatomy, function etc. For example, sutures used for orthopedic surgery would vary based on the tissue attributes of fascia, bone, tendon etc. and locations demanding repair such as superficial or deep regions. In such an instance, the balance between tendon repair and fracture fixation has to be maintained with respect to its correlation with elastic and rigid fixation aspects respectively [5]. Apart from these, the key factor behind choosing the appropriate suture is its utility or specific application. A suture intended for closure of incisions or lacerations is ideally supposed to conserve the actual positioning of the tissues and the structural integrity

174

Advanced Technologies and Polymer Materials for Surgical Sutures

until the healing cascade of tissue response restores its strength and stability [9]. Nonetheless, all of these design parameters, which can be considered as the engineering aspects of suture parameters are dependent on each other as shown in Fig. 8.3. 8.2.1.1 Structural attributes 8.2.1.1.1 Suture size The suture material and its constituent strand diameter determine the overall performance of the suture, which is directly dependent on absorbability, good knot security, ease in handling and minimal tissue reactivity [9]. The dimensions of sutures are based on the set of guidelines put up by the United States Pharmacopeia (Rockville, Maryland) wherein “suture size” refers to the diameter of the suture strand. Suture sizes were initially framed as those ranging from one to six, and then refined considering the requirement of very fine and slender sutures. The USP gauge size is generally denoted as zeroes (For example: #00 referred to as #2-0 or #2/0) to #000000 (#6-0 or #6/0) such that the resultant strand diameter is inversely proportional to the number of zeroes characterizing the suture size. Nevertheless, the currently available sutures range from the thicker ones such as that of #5 to thinner and finer ones, such as #11-0 [10].

Figure 8.3 The interdependence of engineering aspects in suture design.

Engineering aspects of suture fabrication

175

Table 8.1 USP suture sizes, diameters and metric gauges for collagen and synthetic sutures. Collagen sutures USP size

Metric size

#7 #6 #5 #4 #3 #2 #1 #0

Synthetic sutures

Diameter range

Metric size

Diameter range

Application in clinics

e e e 8 7 6 5 4

0.900e0.999 0.800e0.899 0.700e0.799 0.600e0.699 0.600e0.699 0.500e0.599 0.400e0.499 0.350e0.399

Orthopedic surgeries

0.800e0.899 0.700e0.799 0.600e0.699 0.500e0.599 0.400e0.499

9 8 7 6 6 5 4 3.5

#2-0

3.5

0.400e0.399

3

0.300e0.339

#3-0

3

0.300e0.339

2

0.200e0.249

#4-0

2

0.200e0.249

1.5

0.150e0.199

#5-0

1.5

0.150e0.199

1

0.100e0.149

#6-0

1

0.100e0.149

0.7

0.070e0.099

#7-0 #8-0 #9-0 #10-0

0.7 0.5 0.4 e

0.070e0.099 0.050e0.069 0.040e0.049

0.5 0.4 0.3 0.2

0.050e0.069 0.040e0.049 0.030e0.039 0.020e0.029

Abdominal walls, fascia, arterial walls, drain sites and orthopedic surgery Fascia, viscera, trunk and blood vessels Limbs, gut, trunk and blood vessels Mucosa, limbs, hand, neck, tendons and blood vessels Face, neck and blood vessels Face and blood vessels Ophthalmology and microsurgery

Table 8.1 provides a consolidation of USP suture sizes, their corresponding diameters and metric gauges for collagen and synthetic sutures. Most importantly, suture size is also one of the critical parameters, which governs the application of sutures such as microsurgical uses in ophthalmology, blood vessel repair, facial closures, drain stitches, tendon repairs, and abdominal wall closures etc. Each definite USP size has a range of

176

Advanced Technologies and Polymer Materials for Surgical Sutures

approved diameters due to which a disparity within the tensile strengths of sutures of the same USP size is not unusual [11]. The suture gauge size is also reported to influence the size of the knot, which is integral to the prevention of adverse tissue reactions. Larger the knot size, the greater the difficulty to bury them, which consequently leads to distortion of incisions and therefore inflammatory responses. This is particularly a challenge with regard to ophthalmic microsuturing, wherein these adverse effects occurring at the ocular surface might lead to astigmatism [9]. 8.2.1.1.2 Suture configuration/geometry Suture materials can be classified based on their geometrical structure or configuration into different types such as monofilament and multifilament sutures. Monofilament sutures possess single-stranded filaments (thread-like structures) whereas multifilament sutures are composed of several filaments, which are often twisted or braided upon one another (Fig. 8.4) [12,13]. Monofilament sutures pass smoothly through tissues and hence encounter less resistance as they traverse along tissues (lower tissue drag) causing lesser reactions or tissue scarring (Examples: polypropylene, catgut etc.) (Fig. 8.4). But they have low knot security and are relatively harder to handle owing to higher packaging memory. Although multifilament ones have higher tissue drag and cause higher tissue trauma/cutting owing to high coefficient of friction, they are softer, easier to handle and offers better knot security with fewer knots (Examples: silk (braided), steel (twisted), PGA e polyglycolic acid (coated/braided) etc.) (Fig. 8.4) [14]. But in comparison to monofilament ones, these have a higher risk of infections as their high capillarity enhances water absorption. Tensile strength is also dependent on the number of constituent filaments. Multifilament ones render higher mechanical properties, better tensile strength, and flexibility than monofilament ones [8]. But on the other hand, monofilament ones have a lesser

Figure 8.4 Different kinds of suture configurations. (Reprinted with permission from Ref. [13].)

Engineering aspects of suture fabrication

177

coefficient of friction, which makes their gliding easier through the smooth surfaces [15]. Multifilament braided sutures are often coated so as to enhance the tissue drag and decrease the possibility of infection risks [16]. PDS, maxon, nylon and polypropylene are monofilament sutures with high memory while monocryl, pronov, gore-tex and biosyn are monofilaments, which are exceptions in this regard. Vicryl, silk and chromic sutures belong to the category of multifilamented ones. Those sutures, which are constituted of a core-sleeve geometry wherein a braided core is contained within an extruded material sleeve are called pseudo monofilaments. The unique geometry of them provides lesser tissue drag due to which they offer better handling, flexibility and knotting properties [14]. A typical example is Resolon TwistÔ (manufactured by Resorba), nonabsorbable suture supplied as pseudofilament in large diameters, made of polyamide 6/polyamide 6.6 copolymer. 8.2.1.1.3 Needle type Just like the suture thread size, the physical characteristics of the suture needle are also an important attribute of suture design. Needles are chosen considering the factors such as rigidity, strength, ductility, malleability, surface features etc., but most importantly, with respect to the clinical attributes such as the thickness, flexibility and accessibility/location of tissue. The size of the suture strand associated with the needle should also be taken into consideration so as to attain an accurate stitch rendering better cosmetic results [17]. Needles can be manufactured from carbon steel or stainless steel and they can be electroplated or plated with silicon or aluminum. A surgical needle is constituted of three sections viz., point, body, as well as the swage (Fig. 8.5A) [18]. The sharpest portion meant for penetrating the tissue is referred to as “point” whereas the middle portion is considered as “body” of the needle. The swage or eye is the thickest portion attaching the suture to its needle. Needle design parameters such as the point, curvature length and diameter should be carefully chosen considering the suturing technique, type of tissue to be sutured and accessibility characteristics of the operative area. When straight suture needles favor skin closure for easily accessible tissues, curved needles ensure a better turnout dependent on the curvature and confinement of tissues. The needlepoint of the suture can be of different types such as conventional cutting, reverse cutting, taper point, vas cutting, spatula type etc. based on the requirement of skin closure and the associated area (Fig. 8.5B) [19].

178

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 8.5 (A) Anatomy of a suture needle (B) Different kinds of needles. (Reprinted with permission from Ref. [18].)

Suturing through dense, thick and more fibrous tissues requires cutting needles. The square body geometry of spatula (side cutting) needles makes them bending-resistant and yields more secured needle holder positioning, helping to mitigate injury to tissues by maximizing the ease of tissue penetration. Reverse cutting needles with triangular cross-sectioned body and apex cutting edge on exterior cause negligible trauma to tissues due to which they are very useful in ophthalmic operations, cosmetic surgeries and for approximating ligamentous or fibrous tissues. These are very strong and resistant to breakage and bending [20e22]. Conventional cutting needles with inner curvature oriented cutting edges are extensively used in general surgical skin closure procedures involved with subcutaneous tissues, plastic or reconstructive surgery and ophthalmic procedures. Taper point ones with no cutting edges are used for working with soft tissues such as gastrointestinal, vascular, facial tissues etc. 1/4 circle needles are good for convex surfaces and they possess shallow curvature. When 3/8 circular needle is preferred for ocular surgeries due to size constraints and tissue penetration aspects, the sharpness of the curvature of the half circle needle is imperative for sutures used for approximating orbital tissues. An increase in the confinement of the operative area of the body demands greater needle curvature. 5/8 needles favor deeper confinements [21]. Therefore, needle type is one of the important parameters, which can significantly alter upon

Engineering aspects of suture fabrication

179

the application of the suture. The construction of a strong and secure knot, which would not untie or slip away is also based on choosing the appropriate needle type [22]. The needle should possess an adequate amount of strength and ductility such that the repeated passages through the tissue neither cause any deformation nor breakage upon bending. Furthermore, the secure seating and positioning of the needle in the needle holder are based on the clamping moment, which is a measure of the interaction of needle holder jaws with the needle body. The suture should also possess adequate length such that it can be driven smoothly through the tissue using the needle holder without damaging the tissue. Summarizing the above-mentioned aspects an ideal surgical needle should be: (i) Necessarily rigid such that it can withstand bending and severe deformation (ii) Of adequate diameter so as to facilitate a pointed geometry and required sharpness to the cutting edge, imperative to bury the knot (iii) Of particularly good length favoring smooth handling via needle holder, which would ease the passage and retrieval through tissues [9]. 8.2.1.1.4 Surface features Sutures can again be classified according to surface features into smooth (standard) or barbed sutures. Traditional standard sutures have smooth nonbarbed surfaces, which rely on knots for anchoring tissues. It is widely known that a surgical knot acts as the point of highest foreign body material density, which is directly linked to the overall inflammatory reaction. Although wound healing demands minimizing such reactions, knots were considered to be inevitable for tissue approximation [23]. In this scenario, the advent of knotless barbed sutures, which possess nicks (barbs) along the surface helping in tissue penetration and locking them onto position thereby aiding a faster placement of sutures, mitigates the requirement of tying knots [16]. These sutures propagate tension all along the wound length, unlike traditional ones wherein, tension is elevated at the individual loops of suture, which eventually leads to higher ischemia risks [24]. Barbed sutures aid in lowering mechanical stress at the wound edges of the skin and this seems to reduce the complications relative to those imparted by smooth ones, including blood loss, which happens intraoperatively. Bidirectional barbed sutures are structurally unique wherein barbs are helically arranged in opposite directions on a midsegment, therefore conferring a design for even distribution of tension along the line of incision [25]. This aids in

180

Advanced Technologies and Polymer Materials for Surgical Sutures

saving time for placing sutures and skin closure owing to the absence of knot tying step and it is ideal for closure of wounds with high tension, simple and multi-layered closures etc. (Fig. 8.6) This consequently reduces the overall time taken for a surgical procedure, and therefore surgical expenses [24]. Quill bidirectional barbed polydioxanone suture was the first FDA approved barbed suture (2004) followed by V-Loc 180 barbed suture of Covidien, approved in the year 2009. The unidirectional and bidirectional barbed suture materials, which has needle on one and both ends respectively, are currently available include, poliglecaprone 25, nylon, polypropylene, PDO, glycomer 631 and polyglyconate. Nevertheless, unidirectional ones were reported to cause complications such as extrusion or migration owing to the absence of counterbalancing forces on the suture line [26]. However recent reports have indicated the utility of barbed sutures in the manipulation of superficial and deep fascia, tendon, cartilage etc., also in the field of plastic and cosmetic surgeries for facial rejuvenation, particularly owing to the ability to lift tissues during the procedure [24,27]. A randomized clinical trial in 2018 reported the advantage of knotless barbed sutures for uterine incision closure at caesarean delivery over standard sutures in reduction of closure time, obviating the requirement of more hemostatic sutures and decreased blood loss [28]. There have been similar reports emphasizing the applicability of such sutures in various procedures relating to obstetrics and

Figure 8.6 Bidirectional Barbed Suture for closure of simple wound (A) Suture ends pulled through tissue and needles aligned such that both ends are of equal lengths (B) Arcs completed through different arms of the device (C) Tissue approximation (D) Exiting through the skin. (Reprinted with permission from Ref. [25].)

Engineering aspects of suture fabrication

181

gynecology [23,26,29]. Multifilament barbed sutures also called “intertwined sutures” have been reported to confer better mechanical performance through improved pliability making them good candidates for closures of deeper wounds[8]. 8.2.1.1.5 Surface coatings The majority of sutures are often coated with dyes or coatings so as to improve their handling properties. This is particularly intended to reduce the tissue drag when the suture traverses through the needle and to improve the notability [14,30]. Coated sutures generally exhibit reduced flexural rigidity and a lesser coefficient of friction [31]. The application of dyes would also aid in enhancing the identification and suture recognition properties. Logwood extract, chromium-cobalt-aluminium oxide and ferric ammonium citrate are a few of the FDA approved dyes, which impart violet, blue and brown or green colors to sutures respectively [30]. The chemical properties of coating materials should be similar to that of the associated suture so as to improve the compatibility between them. Traditional coating materials such as paraffin wax, bee wax silicone, fluorocarbon, polytetrafluoroethylene etc. are nonabsorbable in nature. As the development of absorbable sutures demanded the usage of absorbable coatings, materials such as poloxamer 188 and calcium stearate were adapted as easily degradable, nonflaking, and adherent lubricant coatings. Poloxamer 188 is a water-soluble coating used over Dexon Plus suture, which dissolves easily in order to reveal the underlying uncoated suture postwound closure. At the same time, coated sutures can also lead to poor knot security wherein the number of square throws required to attain a strong secure knot for uncoated sutures are lesser than that for coated ones [31]. In comparison to water-insoluble calcium stearate coated Vicryl sutures, water-soluble poloxamer coated Dexon plus sutures exhibited an increase in knot security with respect to an increase in hydration time owing to the dissolution of coating [32]. Additionally, multifilament sutures are mostly coated in comparison to monofilament ones [30]. 8.2.1.2 Physical attributes 8.2.1.2.1 Absorbability of sutures The rate at which a suture degrades within the human body is an important attribute, which determines the overall performance of the suture. This measurement of in vivo degradation categorize sutures into two basic classes such as absorbable and nonabsorbable (Table 8.2) [33].

182

Advanced Technologies and Polymer Materials for Surgical Sutures

Table 8.2 Examples of absorbable and nonabsorbable sutures categorized as mono, multi and pseudofilaments. Suture type

Monofilament

Multifilament

Absorbable sutures

Maxon (polyglyconate) PDS II (Polydioxanone) Monocryl (Polyglecaprone) Biosyn Monosyn (25 glyconate) Ethilon (Polyamide 6/6) Prolene (polypropylene) Dermalon (Nylon) Surgilene (polypropylene)

Vicryl (Polyglactin 910) Vicryl rapide (Polyglactin 910) Dexon II (PGA)

Non-absorbable sutures

Pseudofilament

Catgut Chromic gut Ethibond (Polyester) Permahand (Silk)

Supramid (Polyamide 6) Resolon twist

Nurolon (Nylon)

The former constitutes those sutures, which rapidly degrade within tissues and lose tensile strength within 60 days, either by hydrolytic degradation or by enzymatic route followed by hydrolysis [14]. The rate at which degradation of suture occurs is based on the degradation profile of the constituent polymer and also the rate of wound healing. These absorbable sutures can either be natural or synthetic. A natural suture degrades by the mechanism of proteolysis whereas a synthetic one degrades by hydrolysis. Absorbable ones are typically used subcutaneously in order to minimize the dead space and tension usually observed on the edges of the wound, eventually helping in approximating them prior to placement of additional material for aiding wound closure such as staples, tissue adhesive or another suture at epidermal locations. Therefore they act as an option for providing additional temporary strength until the acquisition of the inherent tensile strength of the native tissue, which happens in the course of wound maturation. Moreover, absorbable sutures are also used for the repair of defects in deeper regions [16]. An ideal absorbable suture will have high tensile strength and knot security with a slow absorption profile and less reactivity. Water enters into the filament of strands of synthetic absorbable sutures such as Vicryl and Monocryl upon hydrolysis, which further leads to disruption of constituent polymeric chains. Vicryl was

Engineering aspects of suture fabrication

183

reported to retain 75% of its initial tensile strength at 14 days post in vivo implantation whereas this was 30%e40% for Monocryl at day 14. On the other hand, Polydioxanone based PDS II suture could retain about half of its original strength at 1-month postimplantation [19]. An initial study by Greenwald et al. investigated the tensiometric evaluation of 10 2-0 suture materials at time points t ¼ 0 and t ¼ 42 days post in vivo incubation in Sprague Dawley rats (Fig. 8.7) [34]. Assessment of elastic modulus

Figure 8.7 Changes in mechanical properties of 10 2-0 suture materials after in vivo incubation of 42 days (A) Changes in mean tangent modulus (B) changes in mean strength (mean  SEM). (Reprinted with permission from [34].)

184

Advanced Technologies and Polymer Materials for Surgical Sutures

confirmed the dependence of material property changes on the design of sutures (Fig. 8.7A). When monofilament and braided sutures displayed similar tangent elastic moduli, gut and chromic gut sutures behaved differently. Compliance was the highest for monofilament sutures (1.3  108 N/m2), intermediate for gut sutures (1.5  108 N/m2) and least for braided sutures (greatest modulus of elasticity: 2.6  108 NI/m2). The toughest and strongest sutures were reported to be the monofilament absorbable ones such as Maxon and PDS whereas silk showed the least strength and toughness (Fig. 8.7B) [34]. The rates of degradation or absorption are in fact dependent on so many clinical attributes such as: (1) Anatomical location of the sutures (2) Access to any kind of body fluid (3) Presence of infection or any sort of inflammation (4) Utility of the suture [35]. Nonabsorbable ones are those sutures, which have the ability to maintain their tensile strength for a time period >60 days. These sutures do not have the ability to dissolve by themselves and hence they have to be surgically removed. In comparison to many other nonabsorbable implanted sutures, silk suture was reported to lose 50% of its original tensile strength within a year, which was completely lost within 2 years whereas nylon was found to lose 25% in 2 years [36]. While polypropylene and nylon are the nonabsorbable ones preferred for dermatology associated surgeries, polydioxanone, polyglycolic acid, and polyglactin 910 are the absorbable ones recommended for the same application [16]. Nevertheless, synthetic absorbable sutures are mostly preferred over nonabsorbable ones due to their additional advantages [37]. 8.2.1.2.2 Engineering of mechanical performance An ideal suture is expected to have adequate mechanical integrity owing to its basic utility to close the wound and keep the skin intact during the process of healing. It should neither break in the course of suturing nor elongate with the edema associated with the wound [38]. The longitudinal changes in the mechanical properties of sutures are considered to be highly dependent on the material itself [39]. Additionally, elongation, stress and strain at failure as well as modulus are important tensile parameters, which have been tuned to fit the required utility of the suture [38]. Therefore, each of the mechanical properties should be sufficiently tuned with regard to the application of the suture and its desired implantation period.

Engineering aspects of suture fabrication

185

8.2.1.2.2.1 Tensile strength. Among the various mechanical properties of sutures, tensile strength is the most popular one owing to its correlation to cross-sectional area of the suture as well as normalization with respect to its dimensions. Hence it can be used as an indicator to compare sutures of different sizes. Nevertheless, the force at break doesn’t take into account the diameter of the suture stand and therefore comparisons using this mechanical attribute should be made among candidates of similar sizes and physical forms [11]. Therefore tensile strength of different sutures has been reported to vary with the size and is also dependent on the weight required to disintegrate the suture matrix [8]. 8.2.1.2.2.2 Knot strength. A decent and reliable but firm approximation of sutures tissues can be warranted by the associated strength properties of sutures. However, a decrease in knot strength is a characteristic feature of most sutures owing to which knot tightening step is a crucial step during the overall suturing procedure [1]. This step is crucial to safeguard the security of the suture and it is dependent on mechanical attributes such as coefficient of friction, stiffness, plasticity and elasticity [15]. The ease of handling owing to which a surgeon can make knots effortlessly with sutures possessing good fastening ability are additional perquisites for an ideal suture [40]. Hence knot strength measures the resistance capability of suture strands to withstand the effect of being pulled apart from one another and it is an important parameter of suture strength [38]. Knot security, which maintains knot strength thereby preventing suture slippage is inversely proportional to suture memory, the inherent ability of the suture to maintain its original shape or form. Those sutures possessing high memory will have lesser pliability, which makes their handling difficult. The capability of the suture to retain the tightness of suture loops upon tying the knots is referred to as loop security. Higher the abrasive nature of suture materials, the higher the friction between their loops, and thus greater the loop security [15]. A lesser coefficient of friction results in a higher tendency of suture wherein its knot can loosen easily upon tying [13]. Although knot strength can be enhanced by modulating the thickness of suture strands, this concurrently increases the weights of both sutures as well as the knot. This is also dependent on the number of loops made for tying the knot, which is integral to balancing the elevated tissue pressure in the course of postoperative edema [1]. 8.2.1.2.2.3 Stiffness and flexibility. Variations in the stiffness property of sutures can have significant implications in their clinical utilities particularly

186

Advanced Technologies and Polymer Materials for Surgical Sutures

for those applications with younger tissues [15]. Pliability is a measure of the ease with which a suture can bend without obstructions. A more pliable (flexible) suture can ease the handling, knotting and suturing processes [10]. 8.2.1.2.2.4 Elasticity and plasticity. An ideal suture should possess a good level of elasticity, which facilitates its stretching along with the progression of wound associated edema and its return back to the initial dimensions upon subsiding of swelling. An elastic suture can attain an optimum level of wound approximation without cutting through skin tissue during swelling. In contrast, plastic sutures would not be able to return back to their original dimensions when the swelling subsides. Therefore, they become loose and often hamper wound stitching without being able to approximate the edges [16,17]. 8.2.1.2.2.5 Coefficient of friction. This value determines the ease of suture to pass or slide through tissue. A higher coefficient of friction would guarantee better suturing and knotting. 8.2.1.2.2.6 Capillarity. It measures the fluid absorption or transportation through the length of the suture strand. It is an inherent attribute of multifilament sutures owing to their design and availability of interstitial space. Hence, it also increases the propensity of risk of infections as it can enhance the transport and spread of microorganisms. Monofilament sutures do not possess capillarity and therefore they don’t have to pose higher chances of infections. For example, a braided nylon suture has the potential to promote microorganism entry three times that of its monofilament form. As synthetic sutures are more hydrophobic than natural ones, they have the lesser capability of fluid absorption [17]. Therefore, engineering multifilament sutures can be tricky with regard to the balance of mechanical properties and biocompatibility. 8.2.1.2.2.7 Memory. Memory is the capability of the suture matrix to return back to its initial shape after undergoing the process of deformation [4]. Sutures possessing good memory are generally strong, but their knot security will be lesser than other sutures, which makes their handling a little difficult. Monofilament sutures belong to this category. 8.2.1.2.2.8 Comparisons of mechanical performances of few sutures. With regard to orthopedic surgery, suture material will be subjected to extensive periods of a significant amount of mechanical load, associated with the slow healing of tissues such as ligaments, tendon, fascia etc.

Engineering aspects of suture fabrication

187

Therefore, mechanical parameters such as failure strain, failure load, suture stiffness and hysteresis have to be integrated according to the engineering aspects of such orthopedic sutures [39]. It is imperative for the suture material to possess the endurance to tensile forces, which render stability to the ligaments of joints [40]. A study conducted by Muller et al. compared six different sutures used in orthopedic surgery such as Vicryl, Maxon, Vicryl Rapide, Moncoryl and PDS II (absorbable sutures) and Ethibond, a nonabsorbable one for different mechanical parameters though the course of 56-day long in vitro incubation at physiological conditions such as 37.0  0.02 C; pH 7.4  0.2. Apart from load to failure and strain under maximum load (elongation normalized with original initial length), stress-strain curve dependent attributes, such as stiffness (load to displacement ratio on the linear proportion region) and hysteresis (area under the curve) were also quantified using suture loops with the material testing machine. The overall results recommended the use of PDS sutures for tolerating clinically important loads such as that for applications pertaining to tendon, bone and ligaments. Situations compromising the demand for higher load or increased mechanical strength include facial repairs and tendon sheath closure for which Vicryl and Maxon could be used. Properties of Ethibond were comparatively lesser than that of others over the entire study duration. Vicryl, PDS and Maxon were found to withstand higher loads during the initial 14 days of study. PDS and Maxon were found to sustain the elastic characteristics despite their loss of tensile strength owing to degradation (Fig. 8.8) [39].

Figure 8.8 Comparison of mechanical parameters of different sutures at various time points after in vitro incubation. (Reprinted with permission from Ref. [39].)

188

Advanced Technologies and Polymer Materials for Surgical Sutures

The prominence of Dacron 2-0 sutures in the retention of tensile strength and knot pull breaking strength over 2 week period postsubcutaneous implantation in rats is established [41]. A study by Karabulut et al. made a comparison of the tensile strength and durability of various sutures in different in vivo conditions with respect to pH and anatomical variations. Although polypropylene nonabsorbable sutures were reported to have the highest durability and stability in all conditions, its probability to induce the formation of biliary stones upon stability preservation hampers its utility. Therefore, Maxon, Vicryl, and Biosyn, which also possess relatively better durability could be alternatives for biliary surgery. Vicryl was shown to lose its tensile strength in all conditions such as that of stomach, bile duct, and intestine except in intravesical conditions. Caprosyn and Biosyn are ideal for urological procedures as they showed the maximal tensile loss in urine and intravesical conditions and complete dissolution, which would decrease the probability of infections and stone formation [42]. All of these emphasize the importance of engineering mechanical aspects of suture design with respect to its exact utility and desired implantation period. 8.2.1.3 Biological attributes 8.2.1.3.1 Capillarity, biofilms and bacterial attacks Surgical sutures materials must be biocompatible and should not have properties of capillarity and wicking. Most importantly, they have to retain all of these characteristics during storage [1]. Additionally, even the first and foremost step of skin piercing in the entire suture technique can actually pave the opening route for opportunistic bacteria to infiltrate the probable incision site. Herein, the presence of foreign material such as a suture further adds to the vulnerability of tissue to contamination [37]. Although braided sutures offer better manipulation with respect to suturing, they impart a sawing effect upon tissues, which are often emphasized by high capillarity and wicking property. Wicking is often manifested as wound exudate spreading along with pores present between the suture fibers [1]. The open water channels formed within the polymeric matrix of suture materials can enclose biofilms constituted by complex cell populations [43]. These biofilms as well as the interstices n braided sutures would act as homing niches for bacteria to reside evade the immune cell attack and further trigger a proliferation cascade to spread infection in the neighboring tissues. These bacteria would have evaded the attack of immune cells and a wound exudate enriched environment further favors the further growth

Engineering aspects of suture fabrication

189

and incubation of pathogens, leading to an enhanced rate of infections [1]. Antimicrobial coatings provided on the surface of braided sutures are intended to prevent such kinds of infections [37]. 8.2.1.3.2 Tissue responses and adhesions Sutures being fabricated from foreign materials can certainly induce injury by the passage of the needle and the suture thread, which causes some extent of tissue reaction [44]. Although this is a lot dependent on the configuration or the geometry of the suture material, the tissue response is unavoidable. Tissue reactions are higher when a considerable length of suture is used for implantation, particularly, at the knotted regions and the additional throws to knot [43]. Additionally, cells could also grow within these braids of sutures in the course of wound healing and such tissue ingrowth culminates in the adhesion of tissues to the surface of sutures, a condition called “postsurgical adhesions” [45]. This is also one of the complications, which affects wound healing. The length of the suture, composition and constituent geometry of the suture can also influence the rate of tissue adhesion. Hence, it is imperative to consider the structural attributes of suture design with respect to its biocompatibility and propensity to cause infections during the engineering phase. There have been instances wherein certain suture materials were reported to induce allergic responses in patients. One typical example is the catgut suture, to which chromic salt is added for modulation of degradation profiles, which can consequently trigger an allergic response in chromatesensitive patients [46]. Other examples are allergies caused by nylon sutures during ophthalmic surgeries, and allergies caused by silk sutures [47,48]. Although sericin coated silk fibers impart mechanical strength to the sutures, they are reported to cause inflammatory reactions by stimulating macrophage adhesion. Therefore the addition of coating materials and other constituents intended to improve the performance of the suture should not hamper the biocompatibility of the suture. 8.2.1.3.3 Influence of pH of body fluid The retention of tensile strength is also based on the pH variations of the body fluid, which encounters the site of suture implantation. The rate of degradation of synthetic absorbable sutures is reported to be dependent on alkaline pH to a greater extent (polydioxanone being an exception). Among the nonabsorbable ones, silk is very susceptible to pH variations. When polypropylene can sustain its strength over several pH ranges, nylon

190

Advanced Technologies and Polymer Materials for Surgical Sutures

shows a decrease in tensile properties in acidic environments [43]. All of these factors mention the importance of tuning the biological attributes related to suture fabrication such that it remains biocompatible throughout the course of action.

8.3 Broadening the functionality of sutures Apart from the general utility of wound closures and tissue repair, the applicability of sutures can now be further extended as drug-eluting depots. Drug-eluting sutures are now considered to be the next generation surgical sutures, owing to their synergistic role in wound closure/tissue repair along with elution of the drug near the vicinity of implantation [49]. The nature of the biological agent incorporated within the suture material determines the therapeutic utility of the suture matrix. Reduction in the occurrence of surgical site infections and postoperative issues as well as accelerating the rate of wound healing and tissue regeneration are a few of the practical applications of drug-eluting sutures. Such an approach can also possibly mitigate the usage of additional supplementation of drugs. Additionally, the sustained release of a specific therapeutic agent to a particular area can aid in prolonging the therapeutic local concentrations further enhancing its therapeutic utility as drug depots [50]. This could be a good strategy to retain the concentrations of antibiotics, antiinflammatory drugs, analgesics, antimicrobials, extracellular matrix proteins, peptides, growth factors, cytokines etc. One of the first commercial examples justifying this concept was the approval of ‘Plus sutures’ with triclosan for anti-bacterial protection to mitigate surgical site infections and these sutures were manufactured by Ethicon (part of Johnson & Johnson Medical Devices Companies) [51]. It is composed of biodegradable polymer polyglactin and coated with the antibacterial agent triclosan, which is proved to be effective against the leading causative bacteria of surgical site infections such as staphylococcus aureus, staphylococcus epidermidis and methicillin-resistant strains of staphylococcus (MSRA and MRSE) [52]. On the other hand, a modified Coated VICRYL* Plus suture is composed of a copolymer of glycolide and lactide, alone with a small proportion of triclosan and calcium stearate, which is intended for approximating soft tissues and for ligation purposes, except for those tissues related to ophthalmic, neurological, and cardiovascular applications. These sutures are absorbed completely within a period of 56e70 days. Similarly, PDS Plus and Monocryl Plus are triclosan coated monofilament sutures composed of polydioxanone polyester and

Engineering aspects of suture fabrication

191

copolymer of glycolide with epsilon-caprolactone, respectively [53]. These sutures, which are proved to be successful for subcuticular skin and facial closures are incorporated with triclosan and absorbed within 91e119 days and 182e238 days respectively [54,55]. Subsequently, there have been numerous studies investigating different techniques of manufacturing drugeluting sutures, which include dip coating, coaxial electrospinning, blend electrospinning, grafting, melt extrusion etc. [49,56e59]. These approaches were based on two strategies such as the incorporation of the therapeutic agent into the material in the course of the fabrication process and the addition of those agents to a prefabricated suture [60]. The methods falling under the former category are advantageous as they are relatively less expensive, one-step processes, making use of the same equipment and hence rendering massive yields. The exact method should be opted with regard to the utility of suture, location in the body, duration of desired drug release profile etc. 8.3.1 Engineering drug-eluting sutures Dip coated sutures often render a shorter and relatively uncontrolled release profile, but they do not compromise on mechanical strength as drugs are basically coated onto conventional sutures [50]. Therefore, they are ideal for releasing antibiotics, antiinflammatory drugs etc., which are required for shorter durations, typically, a few days [61e63]. On the contrary, engineering drug-eluting sutures fabricated by other techniques is challenging, as they demand an optimum balance between mechanical strength and drug release. It is imperative for the suture to release the optimum concentration of the drug good enough to induce a favourable therapeutic outcome, without compensating the mechanical characteristics. Blend electrospinning makes use of those solutions wherein a solvent compatible for dissolving both drug and polymer is selected so as to ensure homogeneity, which is required to attain uniform drug distribution throughout the suture matrix. Nevertheless, drug release occurs by the combinatorial effect of both drug diffusion and polymer degradation because of which deterioration of the mechanical tensile properties in a time-dependent manner also occurs. Hence, engineering drug-eluting sutures requires precise control of several parameters [50,64]. 8.3.1.1 Choosing the right technique of fabrication It is important to choose the right technique for fabricating the suture according to the requirement.

192

Advanced Technologies and Polymer Materials for Surgical Sutures

Melt extrusion is one of the conventional techniques of suture fabrication, which is easy to implement and solvent-free. It ensures the homogenous distribution of the therapeutic agent along the cross section go the matrix [65]. Nonetheless, thermo-labile or heat-sensitive drugs cannot be incorporated with those polymers possessing high melting temperature. Blend electrospinning can aid in fabricating sutures with good and efficient drug loading, but the overall mechanical strength is reported to be inversely proportional to loading content [50,64]. Coaxial electrospinning could be a better alternative wherein the properties of core and shell constituents can be better modulated considering the specific characteristics of polymer and drug [59]. Recent reports indicate the usage of a modified electrospinning strategy, wherein drug-loaded electrospun sheath was deposited onto a mechanically strong core, thereby attaining a controlled release profile and good mechanical strength [64]. Techniques such as melt extrusion and electrospinning show greater avenues for altering the suture diameters, morphology, porosity etc., as these can be tuned according to the diameter of the spinneret and spinning conditions. Techniques like coating and grafting can help in achieving high initial drug loadings, but they concurrently enhance the overall diameter of the suture too. The former technique favors the adsorption of the drug onto the suture surface, which could also lead to an undesirable burst effect and poor control over the regulation of coating thickness. On the other hand, direct functionalization of fiber surface by radiation grafting method can resolve these issues upon tuning the conditions such as concentration and proportions or monomers, choosing the right temperature, reaction time and irradiation dose [49,66]. Soaking is considered to be the simplest method of drug loading as it facilitates drug impregnation into sutures by soaking it in a drug solution [67]. This would ensure the diffusion of the drug into the matrix and its retention by polymer-drug interactions. Usage of toxic organic solvents is a limitation of techniques such as electrospinning and coating. In this context, the use of supercritical carbon dioxide (scCO2) has been shown to mitigate organic solvent associated issues, which can sob in numerous polymers facilitating their swelling [68]. This technology also helps in the manufacture of sutures free of solvent residues by a simple depressurization step and also favors the impregnation of thermo labile drugs [10,49]. 8.3.1.2 Choosing the right polymer The flexibility offered in the selection of a range of polymers possessing different physicochemical characteristics is an advantage in this regard. As

Engineering aspects of suture fabrication

193

mentioned earlier, sutures can be fabricated from a wide range of degradable and nondegradable polymers. Degradable polymers include poly lactic co glycolic acid (PLGA), poly caprolactone (PCL), poly L-lactic acid (PLLA), polydioxanone (PDS) etc., which mostly dissolve by bulk degradation. The incorporation of drugs into such matrices will ensure uniform dispersion as well as homogeneity throughout the matrix. This is also dependent on the entrapment efficiency of the drug within the polymeric matrix, which is based on the affinity of the drug with respect to polymer, their interactions, hydrophilicity/hydrophobicity attributes etc. Moreover, the rate of drug release will be completely dependent on the rate of polymer degradation, which happens by hydrolysis, enzymatic degradation etc. For a majority of these polymers, the incorporation of the drug during the manufacturing process is often accompanied by a burst release of the drug. This can also be prevented by providing a layer or sheath over the suture matrix such that it offers an additional rate-limiting step in the course of core polymer degradation and consequently drug release. At the same time, the possibility of a negative influence on mechanical properties owing to this additional layer cannot be ruled out. Hence, tuning the thickness of the coating layer is an important step with regard to the engineering perspective. 8.3.1.3 Choosing the right technique of postprocessing Various postprocessing techniques such as heating, stretching, twisting or plying etc. have demonstrated an increase in mechanical strength of suture matrices. High tensile properties can be attained by the synergistic effect of uniaxial orientation and elevated crystallinity imparted onto suture strands by such postprocessing conditions, which tend to influence the alignment and microstructure of the polymeric chains [64,69,70]. Therefore, choosing a semicrystalline polymer can offer advantages such as high strength as well as sustained release profile owing to the time taken for complete degradation. Braiding or plying the overall suture was found to enhance the force at break, thereby forming a mechanically stable drug-eluting suture.

8.4 Conclusions Successful performance of an ideal suture depends on the effective implementation of the right engineering aspects of suture design. The different parameters to be considered prior to the selection of an appropriate technique of suture fabrication have been classified according to the

194

Advanced Technologies and Polymer Materials for Surgical Sutures

structural uniqueness, mechanical integrity and biological relevance of sutures. Although they have been associated with multiple factors, these attributes are dependent on each other, based on the specific application and clinical utility of suture. Structural constraints pertaining to suture size, dimensions, geometry and surface features have an important role in the mechanical performance of sutures. Additionally, mechanical parameters such as tensile, knot pull strength, flexibility, elasticity etc. would vary for different sutures with respect to the anatomical location of their placement within the human body and the specific characteristics of tissues associated with the application. It is imperative to tune the biocompatibility of the suture matrix such that the tissue responses would be minimal, without evoking serious infections or adhesions. Therefore, careful consideration of the right design parameters will ensure the development of a strong and stable suture. Considering the recent progress in the development of drugeluting sutures, an adequate modification of the engineering aspects is required in order to attain an optimal drug release profile from the suture depot, without compromising its mechanical strength.

References [1] V.A. Zhukovskii, Problems and prospects for development and production of surgical suture materials, Fibre Chem. 40 (2008) 208e216, https://doi.org/10.1007/s10692008-9039-0. [2] S.V. Khiste, V. Ranganath, A.S. Nichani, Evaluation of tensile strength of surgical synthetic absorbable suture materials: an in vitro study, J. Periodontal Implant Sci. (2013) 130e135. [3] M. Byrne, A. Aly, The surgical suture, Aesthet, Surg. J. 39 (Supplement_2) (2019) S67eS72, https://doi.org/10.1093/asj/sjz036. [4] M.H. Kudur, S.B. Pai, H. Sripathi, S. Prabhu, Sutures and suturing techniques in skin closure, Indian J. Dermatol. Venereol. Leprol. 75 (2009) 425e434, https://doi.org/ 10.4103/0378-6323.53155. [5] S. Najibi, R. Banglmeier, J. Matta, M. Tannast, Material properties of common suture materials in orthopaedic surgery, Iowa Orthop. J. 30 (2010) 84e88. [6] M.T. Ho, C.E. Saunders, Current emergency diagnosis & treatment, Adv. Emerg. Nurs. J. 12 (4) (1990) 72. [7] G.M. Doherty (Ed.), Current Diagnosis & Treatment: Surgery, Lange Medical Books/ McGraw-Hill, 2010. [8] C. Dennis, S. Sethu, S. Nayak, L. Mohan, Y. Morsi, G. Manivasagam, Suture materials - current and emerging trends, J. Biomed. Mater. Res. 104 (2016) 1544e1559, https:// doi.org/10.1002/jbm.a.35683. [9] J.H. Smith, M.S. Macsai, Needles, sutures, and instruments, Ophthal. Microsurg. Suturing Tech. (2007) 9e20, https://doi.org/10.1007/978-3-540-68041-3_2. [10] A. Arora, G. Aggarwal, J. Chander, P. Maman, M. Nagpal, Drug eluting sutures: a recent update, J. Appl. Pharm. Sci. 9 (7) (2019) 111e123, https://doi.org/10.7324/ JAPS.2019.90716. [11] S. Dumitriu (Ed.), Polymeric Biomaterials, Marcel Dekker, New York, NY, 1994.

Engineering aspects of suture fabrication

195

[12] J.R. Ajmeri, C.J. Ajmeri, Surgical sutures: the largest textile implant material, Med. Text. Biomater. Heal (2006) 432e440, https://doi.org/10.1533/ 9781845694104.7.432. Woodhead Publ. [13] A. Goel, Surgical sutures - a review, Delhi J. Ophthalmol 26 (2016) 159e162, https:// doi.org/10.7869/djo.161. [14] K.P. Chellamani, D. Veerasubramanian, R.S.V. Balaji, Surgical sutures : an overview, J. Acad. Ind. Res. 1 (12) (2013) 778e782. [15] E. Acan, O. Hapa, F.A. Barber. Mechanical Properties of Suture Materials, In Knots in orthopedic surgery: Open and arthroscopic techniques, in: U. Akgun, M. Karahan, P.S. Randelli, J. Espregueira-Mendes (Eds.), Springer, Berlin, Heidelberg, 2018, https://doi.org/10.1007/978-3-662-56108-9. [16] C.G. Regula, C. Yag-Howard, Suture products and techniques: what to use, where, and why, Dermatol. Surg. 41 (2015) S187eS200, https://doi.org/10.1097/ DSS.0000000000000492. [17] J. Hochberg, K.M. Meyer, M.D. Marion, Suture choice a nd other methods of skin closure, Surg. Clin. NA. 89 (2009) 627e641, https://doi.org/10.1016/ j.suc.2009.03.001. [18] L. J.A., T. S., P. F., P. M.J., Fundamentals of sutures, needles, knot tying, and suturing technique, Palazzo F. Fundam. Gen. Surgery (2018) 1e350, https://doi.org/10.1007/ 978-3-319-75656-1. Springer, Cham. [19] Wound Closure Manual, DL. Dunn (contributing editor), Ethicon, Inc., Somerville, N.J. (2005). [20] D. Ratner, Suturing Techniques Periprocedural Care, In: Drugs & Diseases > Clinical Procedures, Medscape, Accessed 22 February 2020 https://emedicine.medscape.com/ article/1824895-periprocedure. [21] S. Baran, E. Johnson, M. Perret-Gentil, Understanding and Selecting Surgical Suture and Needle, ALN (2013) 1e9. [22] R.R. Szarmach, J. Livingston, G.T. Rodeheaver, J.G. Thacker, R.F. Edlich, An innovative surgical suture and needle evaluation and selection program, J. Long Term Eff. Med. Implants 12 (2002) 211e229. [23] J.A. Greenberg, R.H. Goldman, Barbed suture : a review of the obstetrics and gynecology, Rev. Obstet. Gynecol. 6 (2013) 107e115, https://doi.org/10.3909/ riog0231. [24] K. Aliano, M. Trostler, I.M. Fromm, A. Dagum, S. Khan, D. Bui, A comparison of barbed sutures and standard sutures with regard to wound cosmesis in panniculectomy and reduction mammoplasty patients, Plast. Surg. Int. 2016 (2016) 1e5, https:// doi.org/10.1155/2016/7590396. [25] M.D. Paul, Bidirectional barbed sutures for wound closure: evolution and applications, J. Am. Col. Certif. Wound Spec. 1 (2009) 51e57, https://doi.org/10.1016/ j.jcws.2009.01.002. [26] J.A. Greenberg, The use of barbed sutures in obstetrics and gynecology, Rev. Obstet. Gynecol. 3 (2010) 82e91, https://doi.org/10.3909/riog0124. [27] A.D. Rosen, New and emerging uses of barbed suture technology in plastic surgery, Aesthet. Surg. J. 33 (3_Supplement) (2013) 90Se95S, https://doi.org/10.1177/ 1090820X13500048. [28] D.P. Facog, R.S. Ahmad, S.L. Warsof, N. Marcus-braun, Y. Sciaky-tamir, I. Ben Shachar, A randomized clinical trial of knotless barbed suture vs conventional suture for closure of the uterine incision at cesarean delivery, Am. J. Obstet. Gynecol. 218 (2018) 343.e1e343.e7, https://doi.org/10.1016/j.ajog.2018.01.043. [29] B. Gardella, A. Daniela, What is the role of barbed suture in laparoscopic myomectomy? A meta-analysis and pregnancy outcome evaluation, Gynecol. Obstet. Invest. 83 (6) (2018) 521e532, https://doi.org/10.1159/000488241. [30] C.C. Chu, Types and properties of surgical sutures, in: B.S. Gupta, M. King, R. Guidoin (Eds.), Biotextiles as medical implants. Woodhead Publishing Series in Textiles, Woodhead Publishing, Cambridge, U, 2013, pp. 231e273.

196

Advanced Technologies and Polymer Materials for Surgical Sutures

[31] G.T. Rodeheaver, J.G. Thacker, J. Owen, M. Strauss, T. Masterson, R.F. Edlich, Knotting and handling characteristics of coated synthetic absorbable sutures, J. Surg. Res. 35 (1983) 525e530, https://doi.org/10.1016/0022-4804(83)90043-4. [32] C.-C. Chu, Classification and General Characteristics of Suture Materials. in: C.C. Chu, J.A. von Fraunhofer, H.P. Greisler (Eds.), Wound Closure Biomaterials and Devices. CRC Press (Boca Raton), 1996 pp. 43. [33] R. Edlich, K.D. Gubler, A.G. Wallis, J.J. Clark, J.J. Dahlstrom, W.B. Long III, Scientific basis for the selection of skin closure techniques, J. Environ. Pathol. Toxicol. Oncol. 29 (2010). [34] D. Greenwald, S. Shumway, P. Albear, L. Gottlieb, Mechanical comparison of 10 suture materials before and after in vivo incubation, J. Surg. Res. 56 (1994) 372e377. [35] J.W.W. Van, J.C. Hastings, E. Barker, D. Hines, W. Nichols, Effect of suture materials on healing skin wounds, Surg. Gynecol. Obstet. 140 (1975) 7e12. [36] R.W. Postlethwait, Long-term comparative study of nonabsorbable sutures, Ann. Surg. 171 (1970) 892e898, https://doi.org/10.1097/00000658-197006010-00010. [37] R.S. Rengasamy, S. Ghosh, Technical Sewing Threads in Technical Textile Yarns, Woodhead Publishing Series in Textiles, 2010, pp. 495e533. [38] S.E. Naleway, W. Lear, J.J. Kruzic, C.B. Maughan, Mechanical properties of suture materials in general and cutaneous surgery, J. Biomed. Mater. Res. B Appl. Biomater. 103 (2015) 735e742, https://doi.org/10.1002/jbm.b.33171. [39] D.A. Müller, J.G. Snedeker, D.C. Meyer, Two-month longitudinal study of mechanical properties of absorbable sutures used in orthopedic surgery, J. Orthop. Surg. Res. (2016) 1e7, https://doi.org/10.1186/s13018-016-0451-5. [40] L.C. Gomide, D. de O. Campos, C.A. Araújo, G.L. Menegaz, R.S. Cardoso, S.C. Saad Júnior, Mechanical study of the properties of sutures used in orthopedics surgeries, Rev. Bras. Ortop. 54 (2019) 247e252, https://doi.org/10.1055/j.rbo.2018.02.001. [41] J.B. Herrmann, Changes in tensile strength and knot security of surgical sutures in vivo, Arch. Surg. 106 (1973) 707e710, https://doi.org/10.1001/archsurg.1973.013501700 71018. [42] R. Karabulut, K. Sonmez, Z. Turkyilmaz, B. Bagbanci, A.C. Basaklar, N. Kale, An in vitro and in vivo evaluation of tensile strength and durability of seven suture materials in various pH and different conditions: an experimental study in rats, Indian J. Surg. 72 (2010) 386e390, https://doi.org/10.1007/s12262-010-0158-5. [43] A.J. Tsugawa, F.J. Verstraete, Suture Materials and Biomaterials, in: Oral and Maxillofacial Surgery in Dogs and Cats, Elsevier Ltd., 2012, pp. 69e78. [44] R.W. Postlethwait, D.A. Willigan, A.W. Ulin, Human tissue reaction to sutures, Ann. Surg. 181 (1975) 144e150, https://doi.org/10.1097/00006534-197508000-00082. [45] R.P.G. Ten Broek, N. Kok-Krant, E.A. Bakkum, R.P. Bleichrodt, H. Van Goor, Different surgical techniques to reduce post-operative adhesion formation: a systematic review and meta-analysis, Hum. Reprod. Update 19 (2013) 12e25, https://doi.org/ 10.1093/humupd/dms032. [46] R.J. Engler, C.B. Weber, R. Turnicky, Hypersensitivity to chromated catgut sutures: a case report and review of the literature, Ann. Allergy 56 (1986) 317e320. [47] C. Yip, K. Bowen, B.K. Chew, A report of rare adverse tissue reaction to Ethilon® Nylon Suture, J. Surg. Case Rep. 3 (2018) rjy037. [48] S. Kurosaki, H. Otsuka, M. Kunitomo, M. Koyama, R. Pawankar, K. Matumoto, Fibroin allergy: IgE mediated hypersensitivity to silk suture materials, Reports Exp. Clin. Cases. 66 (1999) 41e44. [49] M. Champeau, J. Thomassin, T. Tassaing, Expert opinion on drug delivery current manufacturing processes of drug-eluting sutures, Expet Opin. Drug Deliv. 00 (2017) 1e11, https://doi.org/10.1080/17425247.2017.1289173.

Engineering aspects of suture fabrication

197

[50] C.B. Weldon, J.H. Tsui, S.A. Shankarappa, V.T. Nguyen, M. Ma, D.G. Anderson, D.S. Kohane, Electrospun drug-eluting sutures for local anesthesia, J. Contr. Release 161 (2012) 903e909. [51] Ethicon Plus Sutures Become First and Only Sutures With Antibacterial Protection Recommended for Use in NHS by NICE Medical, Bloomberg, US edition, 2021. Accessed 20 May 2022, https://www.bloomberg.com/press-releases/2021-06-29/ ethicon-plus-sutures-become-first-and-only-sutures-with-antibacterial-protection-rec ommended-for-use-in-nhs-by-nice-medical. [52] S. Rothenburger, D. Spangler, S. Bhende, D. Burkley, In vitro antimicrobial evaluation of Coated VICRYL_ plus antibacterial suture (coated polyglactin 910 with triclosan) using zone of inhibition assays, Surg. Infect. 3 (2002) s79es87. [53] Plus Sutures for preventing surgical site infection, Medical technologies guidance. NICE.org., The National Institute for Health and Care Excellence. (Accessed 20 May 2022). https://www.nice.org.uk/guidance/ [54] X. Ming, M. Nichols, S. Rothenburger, In vivo antibacterial efficacy of MONOCRYL plus antibacterial suture (poliglecaprone 25 with triclosan), Surg. Infect. 8 (2007) 209e214, https://doi.org/10.1089/sur.2006.004. [55] X. Ming, S. Rothenburger, D. Yang, In vitro antibacterial efficacy of MONOCRYL plus antibacterial suture (Poliglecaprone 25 with Triclosan), Surg. Infect. 8 (2007) 201e207, https://doi.org/10.1089/sur.2006.005. [56] C.L. He, Z.M. Huang, X.J. Han, Fabrication of drug-loaded electrospun aligned fibrous threads for suture applications, J. Biomed. Mater. Res. 89 (2009) 80e95. [57] W. Hu, Z.-M. Huang, Biocompatibility of braided poly(L-lactic acid) nanofiber wires applied as tissue sutures, Polym. Int. 59 (2010) 92e99. [58] X. Chen, D. Hou, L. Wang, Q. Zhang, J. Zou, G. Sun, Antibacterial surgical silk sutures using a high-performance slow-release carrier coating system, ACS Appl. Mater. Interfaces 7 (2015) 22394e22403, https://doi.org/10.1021/acsami.5b06239. [59] W. Qian, D.G. Yu, Y. Li, Y.Z. Liao, X. Wang, L. Wang, Dual drug release electrospun core-shell nanofibers with tunable dose in the second phase, Int. J. Mol. Sci. 15 (2014) 774e786. [60] M. Champeau, J. Thomassin, T. Tassaing, C. Jérôme, Drug loading of polymer implants by supercritical CO 2 assisted impregnation: a review, J. Contr. Release 209 (2015) 248e259, https://doi.org/10.1016/j.jconrel.2015.05.002. [61] A. Obermeier, J. Schneider, S. Wehner, F.D. Matl, M. Schieker, R. von EisenhartRothe, A. Stemberger, R. Burgkart, Novel high efficient coatings for anti-microbial surgical sutures using chlorhexidine in fatty acid slow-release carrier systems, PLoS One 9 (2014) 1e7. [62] Y. Li, K.N. Kumar, M. Dabkowski, M. Corrigan, R.W. Scott, K. Nu, G.N. Tew, New bactericidal surgical suture coating, Langmuir 28 (2012) 12134e12139. [63] Z.X. Wang, C.P. Jiang, Y. Cao, Y.T. Ding, Systematic review and meta-analysis of triclosan-coated sutures for the prevention of surgical-site infection, Br. J. Surg. 100 (2013) 465e473, https://doi.org/10.1002/bjs.9062. [64] S. Padmakumar, J. Joseph, M.H. Neppalli, S.E. Mathew, S. V Nair, S.A. Shankarappa, D. Menon, Electrospun polymeric core-sheath yarns as drug eluting surgical sutures, ACS Appl. Mater. Interfaces 8 (2016) 6925e6934, https://doi.org/10.1021/ acsami.6b00874. [65] O. Catanzano, S. Acierno, P. Russo, M. Cervasio, M. Del Basso De Caro, A. Bolognese, G. Sammartino, L. Califano, G. Marenzi, A. Calignano, D. Acierno, F. Quaglia, Melt-spun bioactive sutures containing nanohybrids for local delivery of anti-inflammatory drugs, Mater. Sci. Eng. C. Mater. Biol. Appl. 43 (2014) 300e309, https://doi.org/10.1016/j.msec.2014.07.012. [66] B. Gupta, Rachna Jain, Harpal Singh, Preparation of antimicrobial sutures by preirradiation grafting onto polypropylene monofilament, Polym. Adv. Technol. 19 (2008) 1698e1703, https://doi.org/10.1002/pat.

198

Advanced Technologies and Polymer Materials for Surgical Sutures

[67] R. Zurita, J. Puiggalí, A. Rodríguez-Galán, Loading and release of ibuprofen in multiand monofilament surgical sutures, Macromol. Biosci. 6 (2006) 767e775, https:// doi.org/10.1002/mabi.200600084. [68] R.D. Weinstein, K.R. Muske, S.A. Martin, D.D. Schaeber, Liquid and supercritical carbon dioxide-assisted implantation of ketoprofen into biodegradable sutures, Ind. Eng. Chem. Res. 49 (2010) 7281e7286, https://doi.org/10.1021/ie901913x. [69] S. Ozbek, D.H. Isaac, Carbon fiber processing: effects of hot stretching on mechanical properties, Mater. Manuf. Process. 9 (1994) 199e219. [70] Z. Song, X. Hou, L. Zhang, S. Wu, Enhancing crystallinity and orientation by hotstretching to improve the mechanical properties of electrospun partially aligned polyacrylonitrile (PAN) nanocomposites, Materials 4 (2011) 621e632.

CHAPTER 9

Revisiting the properties of suture materials: an overview Ashwin Kumar Narasimhan1, Thella Shalem Rahul1 and Saravanan Krishnan2 1

Department of Biomedical Engineering, SRM Institute of Science and Technology, Kattankulathur, Chennai, Tamil Nadu, India; 2Creative Carbon Labs Pvt Ltd., Chennai, Tamil Nadu, India

9.1 Introduction The success of surgery depends on how the tissues are connected through proper suturing. Sutures are the combination of individual fibers that form larger fibrous threads and are used for connecting the blood vessels and tissues in animals and humans’ postsurgery. Suturing requires two key components namely, metallic needles and sutures [1]. The development of novel sutures and associated suturing techniques have evolved over time to improve the wound healing process. During postsurgery, the four major wound healing mechanisms are arrest of bleeding, blood clotting, tissue regeneration and tissue strengthening [2]. For this reason, surgeons have been using sutures and a variety of biomaterials for connecting the separated tissues. Toward managing wound healing, Egyptians have used suture materials such as cotton, tendons, fibers, silk threads, linen, hair and grass. A famous Indian physician, Susruta, in the book named Samhita, has documented eight important classifications of surgery along with surgical procedures for better treatment planning [3]. For the suturing process, he has elaborated on the use of biological surgical materials such as tendon, hair, bark and silk for wound closure [4]. The use of sutures for the fast recovery of wounds is growing at a rapid pace. Suture materials have different properties related to its intrinsic features such as biosorption, tensile strength, elasticity, static friction, plasticity, ease of handling and memory. Material properties of the sutures also depend on sizes, diameters, and type of materials used. Physicians usually choose the sutures depending on the type of surgery to be performed [5].

Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00011-5

© 2023 Elsevier Ltd. All rights reserved.

199

200

Advanced Technologies and Polymer Materials for Surgical Sutures

9.1.1 Characteristics of sutures Ideally, suture materials possess the characteristics such as bio-inertness, handling compatibility, unfavorable atmosphere for bacterial growth, withstand sterilization process, easy to knot and tissue absorption [6]. Some of the suture characteristics are based on factors like patient age, physician experience, wound exposure to the environment, wound location, length and shape [7]. Bacterial infection within the surgically sutured tissue area is a major concern. To overcome this, drug-loaded sutures have been developed depending on the patient type and wound location [8]. As compared to aged people, suture-based wound healing is effective for younger patients. Hence, it is important to choose suture materials, which avoid a scar on the sutured area and also exhibit suture thickness, healing capacity and degradation capacity as similar to tissues. Practically, these ideal characteristics represented here may not be available in a single suture material. Therefore, the need for a new type of composite sutures is the need of the hour.

9.2 Types of suture materials and examples Surgical sutures can be classified based on the types of material used, suture size and physical formation of threads. Sutures materials are also classified into two types namely absorbable and non-absorbable sutures, which are discussed in the next section. 9.2.1 Absorbable sutures Absorbable sutures are defined as those materials that lose their mechanical properties within the span of 90 days and lay beneath the tissue structures. Absorbable sutures are classified into natural and synthetic. Earlier, the most widely used absorbable suture material is Catgut prepared from the small intestine of sheep or cattle, which is mainly composed of collagen [8]. However, the inherent problem with catgut is its rapid degradation rate i.e., losing the mechanical properties within 4e5 days and knot security for only 2 weeks. For this reason, Catgut-based sutures are used only for fast-healing wounds and preferably, in surface sutures. By tanning with chromium, the degradation rate of Catgut sutures was increased from 80 to 120 days. Also, the tensile strength of the plain gut is increased from 7 to 14 days [9]. Considering the inherent problems of catgut and chromic acid dipped catgut, synthetic absorbable polymer-based sutures are being developed.

Revisiting the properties of suture materials: an overview

201

Synthetic absorbable polymeric sutures such as polyglycolic acid, Polyglactic acid, Polydioxanone, and Polytrimethylene carbonate, Barbed sutures have been developed. Some of the commercially available synthetic absorbable polymers are Polyglycolide e Dexon®, Poly glactin e Vicryl®, Polydioxanone e PDS®, Di- (or) Tri- (or) Teri-block copolymers of Polytrimethylene carbonate with polycaprolactone, polylactide and polyglycolide e Maxon®, Biosyn®, Monosyn®, Caprosyn®, and Poly-L-lactide e Orthodek® [10]. In 2007, FDA approved a new absorbable suture prepared from poly-4-hydroxybuytrate (P4HB)- TephaFLEX®, through a fermented process involving recombinant DNA technology. P4HB is a biocompatible material with thermoelastic properties [11]. Other than Catgut, polyglycolic acid (PGA) based absorbable sutures were developed in 1970. PGA is commercially known as Dexon and exists in two forms viz., monofilament and braided form. The property of PGA makes it easier to handle the sutures while the coated PGA has ease of knot and good binding performance with tissues [12]. Compared to Catgut, PGA has a tensile strength of 63% at the time-lapse of 14 days and gets absorbed into tissues within 120 days via hydrolysis mechanism. PGA possess biological properties such as lesser tissue reactivity, and no inflammatory response. Dexon-II is a commercially available PGA suture, which is coated with smooth material, easy to knot and protects from bacteria [13]. Vicryl (polyglactic acid) is another example of a synthetic absorbable copolymer composed of glycolic acid and lactic acid (Fig. 9.1). The composition of Vicryl polymer is 90% glycolide and 10% lactide. The physical structure of vicryl is in braided form having retention of 65%

Figure 9.1 Commercially available FDA approved absorbable sutures (A) Polydiaxonone (PDS II e Violet Monofilament) and (B) Polyglactin 910 (Vicryl e Violet Braided). (Adapted from Ref. [14].)

202

Advanced Technologies and Polymer Materials for Surgical Sutures

tensile strength for about 2 weeks. This bioabsorbable polymer degrades via hydrolysis reaction within a time period of 90 days. Also, the tissue reactivity and inflammatory response are poor as compared to polydioxanone [15]. Reactivity with skin tissues shows initial lump formation and it gradually resolves with time. Ethicon Inc. has developed a vicryl polymer, which displays an increased absorption rate and is sterilized using gamma radiation. This vicryl polymer is generated under the trademark Vicryl Rapide®, which absorb completely within 5 days and exhibits no tensile strength after 2 weeks. Triclosan, an antimicrobial drug is loaded into Vicryl Rapide and the antimicrobial suture thus developed inhibits pain after surgery in pediatrics [16]. In 1982, the next absorbable suture material named polydioxanone (PDS®) was developed by Ethicon incorporation. PDS is composed of paradioxanone monomer, otherwise known as polyester form a monofilament structure. This type of suture possesses characteristics such as low tensile strength, ease of passage within tissues, stiffer material and affects bacterial growth in the sutures [17]. Some of the challenges like material stiffness and handling capability of the PDS have been met by the recent version, PDS-II (Fig. 9.1). However, PDS-II sutures are costlier than Dexon and Vicryl. In addition, tensile strength was only 25% after 6 weeks and bioabsorbed within a span of 6 months longer than absorbable suture materials like Polyglycolic and Polyglactic acid. This type of suture is mainly required to prolong the wound healing for more than 180 days. Other polyester-based copolymer sutures are Caprosyn and Biosyn [18]. Polytrimethylene carbonate (PTMC) is a commercially known synthetic absorbable suture named as Maxon developed in the year 1985. PTMC is composed of copolymers of trimethylene carbonate and glycolide with monofilament structured sutures. Among suture materials like polyglactin and PDS II, Maxon possesses characteristics like knot safety, handling capability, and good tensile strength before implantation. Postsurgery, the tensile strength of Maxon reduces to 81%, 59% and 30% at the lapse of 14, 28 and 42 days respectively. These synthetic absorbable sutures have similar degradation properties through hydrolysis and are usually absorbed within 180 days [19]. 9.2.2 Non-absorbable sutures To enable longer-term tissue approximation, non-absorbable sutures are utilized. It can be used directly on the skin and removed later, or can be

Revisiting the properties of suture materials: an overview

203

used inside the body and left as such. Non-absorbable sutures are commonly used for vascular repair/anastomosis, bowel repair, tendon repair, and skin closure. Non-absorbable sutures elicited a considerably lower amount of immune response and cause less scarring in surgeries performed with an improved cosmetic outcome. Depending on the severity of the incision, these sutures may be left in place or removed after a stipulated period of time. Non-absorbable sutures, which are typically used to close skin wounds, may or may not have coatings to improve their overall performance. After a few weeks of posttreatment, the suture is removed [19]. Table 9.1 shows the examples of non-absorbable sutures and their advantages and disadvantages. 9.2.2.1 Silk suture Silk Suture is a sterile, non-absorbable surgical suture made of Fibroin, an organic protein. This protein is extracted from domesticated silkworm Bombyx mori which belongs to Bombycidae family. Silk is colored black to make it more visible and braided to make it easy to handle. The black braided suture is further coated with a unique wax mixture. As shown in Fig. 9.2, a black braided silk suture is used for soft tissue approximation and/ or ligation in a variety of operations, including cardiovascular, ophthalmic, and neurological systems. Due to the fact that it is a natural product, it is more prone to bacterial growth. Of all the sutures currently available in the market, silk suture has the lowest tensile strength. Because of its softness and pliability, it is being utilized as a temporary suture for retraction during surgery or surgical procedures involving mucosal or intertriginous areas. It can retain its tensile strength for up to 50% even after 1 year [20]. 9.2.2.2 Nylon In the early 1930s, nylon was introduced as the primary synthetic suture, which is a strong non-absorbable material. Nylon sutures are available in both monofilament (Ethilon, Monosof) and multifilament variants (Nurolon, Surgilon, Supramid). Monofilament is significantly more prevalent and less prone to infection as compared to multifilament sutures. The commercial name of Nylon sutures is Ethilon®, which is the most extensively used non-absorbable suture material available in monofilament form (Fig. 9.2). Nylon sutures provide high knot security and are easily removable without causing tissue adhesion and infection-resistant sutures [21]. Nylon sutures are commonly utilized in general surgery, cutaneous surgery, and cosmetic surgery. In cardiovascular surgery, nylon sutures are

204

Suture type

Raw material

Strand

Pros

Cons

Silk suture Nylon

Silk Polyamide thermoplastics

Polypropylene

Polymer of propylene

Braided Mono/ multifilament Monofilament

Good handling Inert, good elasticity, good knot security Good tensile strength, low tissue reactivity and high plasticity

Polybutester

Copolymer of polyglycol terephthate and polybutylene terephthate The alloy of iron, chromium and molybdenum

Monofilament

Excellent strength, elastic, good handling characteristics, good knot security Biologically inert and the strongest suture material

Low tensile strength Low handling characteristics Poor knot security and handling characteristics Tissue reactivity

Stainless steel suture

Monofilament or multifilament

Poor handling properties cause injury to the tissue

Advanced Technologies and Polymer Materials for Surgical Sutures

Table 9.1 Examples of non-absorbable sutures and their pros and cons.

Revisiting the properties of suture materials: an overview

205

Figure 9.2 Commercially available FDA approved non-absorbable sutures (A) Silk suture, (B) Nylon, (C) Polypropylene, (D) Polybutester, (E) Stainless Steel suture. (Adapted from Ref. [14].)

less recommended for the attachment of artificial prostheses. Nylon sutures possess properties such as good durability, low tissue reactivity, high elasticity and cost-effective. For nearly 11 years, monofilament forms retain its 2/3rd of its original strength. Nylon hydrolyses slowly, and loses w15% to 20% of its enduringness every year [22e24]. 9.2.2.3 Polypropylene Prolene sutures (colored and undyed) are sterile surgical sutures prepared from an isotactic crystalline stereoisomer of polypropylene, a synthetic linear polyolefin [25]. Prolene sutures have a long track record of strength, highly dependable, adaptable and biologically inert. To improve the visibility, these sutures are coated with blue colored dyes. As shown in Fig. 9.2, it is an excellent suture for skin because of its durability without causing a substantial loss in tensile strength, resistance to infection, and tissue reactivity. Polypropylene sutures are used in a variety of soft tissue surgeries, including cardiovascular, ophthalmology, and neurosurgical procedures for ligation [26]. Since there is no progressive absorption in prolene, this suture is preferred for long-term skin surgeries. Compared to other materials, cross-hatching is less prevalent in this case because it accommodates tissue expansion. On contrary, nylon suture gets loosened once the wound heals. Due to stiffness and memory, knot security is poor with nylon sutures.

206

Advanced Technologies and Polymer Materials for Surgical Sutures

Usually, clinicians perform thermocautery on nylon sutures to fuse the knots. When it is stretched, it loses its original shape. It has a very smooth surface, which results in a low coefficient of friction. Unfortunately, the low pliability of prolene combined with its great memory resulted in lower knot security. If the knots are tight enough, the polypropylene would flatten where the strands cross, increasing its knot stability and strength. 9.2.2.4 Polybutester e novafil Polybutester is a more recent form of polyester suture that comes in monofilament form with molecular weight ranging from 500 to 3000 Da [22]. Polyglycol terephthalate and polybutylene terephthalate are the components that make up this copolymer. Compared to previous polyester sutures, polybutester has a reduced coefficient of friction and tissue reactivity. Polybutester is a superior suture material used for the closure of skin wounds due to its good handling properties and can be easily removed from healed wound. Since it is made of monofilament, they are less-prone to bacterial infection. Under modest loads, polybutester possesses a high degree of elastic stretch, and thus it is superior to other non-absorbable sutures. Once the swelling subsided, it would return to its previous strength. Because of its high elasticity, it is less likely to cut through the skin during swelling and thus well suited for good issue approximation throughout the healing process. With 25% of load, these sutures can stretch up to 50%, which is known as its breaking point. 9.2.2.5 Stainless steel non-absorbable sutures The use of iron steel sutures started as early as 1666. Stainless steel is a chromium-molybdenum alloy that can be used as a monofilament and multifilament suture [26]. It is biologically inert and possesses the highest tensile strength of any suture material. Only 304, 316, and 316 L wrought stainless steels are used as sutures. Series 300 stainless steel has high chromium and nickel content with an austenitic structure and is characterized by high corrosion-resistance and ductility [27]. The letter “L” in 316L stainless steel indicates low carbon content. Stainless steel sutures are mostly employed in orthopedic surgery, and it has been utilized to mend tendons and ligaments as well. In animals, the stainless steel type of sutures is used to treat the contaminated and infected wounds in horses [28]. For example, abdominal wall hernias after wound infection in horses are connected through this type of suture material. Sutures made of stainless steel have a

Revisiting the properties of suture materials: an overview

207

tendency to pierce through tissue and are difficult to handle [29]. Tissue necrosis can occur while moving over the knot ends [30]. 9.2.3 Emerging alternatives to conventional sutures 9.2.3.1 Staples Tissue staples are a convenient and rapid alternative to manual suturing, readily employed in gastrointestinal surgery, liver and lung resection, and reproductive surgery [31]. Individual rectangular-shaped sutures are stapled in a single row on skin, but staples used in other tissues have at least a double row of staples in the form of a staggered “B” arrangement [26]. Stapling devices reduce tissue handling without causing damage, surgical time, and contamination, while it also ensures visceral, vascular, and cutaneous closures and maintains the blood flow [24]. If staples were employed on thick tissues, it might be inefficient, and massive blood vessel hemostasis may be ineffective [32]. These staples are further classified as absorbable and non-absorbable sutures. 9.2.3.2 Absorbable staples Absorbable staples are made of a polylactide-based copolymer with a little amount of polyglycolide. These subcuticular staples offer a new type of skin closure that is quick and eliminates the need for non-absorbable staples or sutures to puncture the epidermis [33]. It loses about 60% of holding power in 2 weeks, has a half-life of 10 weeks, and gets absorbed in a complex manner over several months. Compared to subcuticular sutures or stainless steel skin staples, it generates low-grade inflammation during 7e21 days following surgery [32]. It works effectively for polluted wounds by reducing the infection rate as compared to vicryl sutures. 9.2.3.3 Non-absorbable staples These staples are made up of stainless steel and have higher tensile strength than other sutures. An additional advantage is that the tissue reactivity was low [22]. Metal staples come in two sizes, regular (4.8e6.1 mm) and wide (6.5e7 mm), discharged from lightweight cartridges. When compared to sutures, it closes faster and evert incision margins without strangling the tissue, and lowers the risk of crosshatch scarring. On the other hand, removal of non-absorbable staples may take a longer time as compared to sutures. In rare cases, staples may produce a better cosmetic effect and incur lesser risk of complications in infected wounds.

208

Advanced Technologies and Polymer Materials for Surgical Sutures

9.2.3.4 Tissue adhesives Tissue adhesives have been used for skin closure in a diverse types of surgical procedures for many years. Octyl-2-cyancorylate (OCA) was approved for use in the year 1998 [34]. The stabilizer in the adhesive is neutralized by partially ionized water molecules on the skin, allowing for the polymerization of molecules and adhesion to adjacent tissue within 10 s. A novel higher viscosity OCA has been developed to reduce the adhesive migration and subsequently improve the bonding and cosmesis. The breaking strength of OCA is approximately 5 times of that of nylon sutures, but as reepithelialization develops over 5e10 days, the adhesive frequently sloughs [35]. Tissue adhesives can be expensive, but when used properly, it reduces infection as a result of its antibacterial properties. It saves time, provides good cosmetic effects, and allows for a faster return to sports work without the need for sutures. It is non-carcinogenic and safe to use in youngsters and patients with a history of severe scarring. Furthermore, it is also effective in facial wounds and used widely for various surgical conditions depending on the need for these tissue adhesives [36].

9.3 Suture materials and their properties: recent advances From materials perspective, the performance of the suture depends on its physical, chemical, mechanical, structural and biological properties (antimicrobial, wound healing, and few others) as indicated in Fig. 9.3. Physical

Figure 9.3 Outline of the properties of suture materials.

Revisiting the properties of suture materials: an overview

209

properties are measurable parameters of the material without undergoing any changes. Few examples are density, configuration, rigidity, and conductivity. Chemical properties are those characteristics identified from particular reaction with the material. Some of the chemical properties are pH, surface tension, corrosion resistance, and so on. Structural properties are related to the features attributed to the structure of materials e.g. size, shape and texture. Mechanical properties of sutures are similar to physical properties but the difference is that the former is measured only after the application of known load or forces. Classic examples of mechanical properties are knot pull strength, tensile strength, suture retention, elasticity, stiffness, knot security, shape-memory and few others. Biological properties are those functional features of the suture materials like biocompatibility, wound healing, antimicrobial nature, tissue absorbtion and reactivity, which made them suitable for biomedical applications. Considering the societal need and commercial market, different types of suture materials are prepared and evaluated for the aforementioned properties, which are discussed below. 9.3.1 Silk-based sutures Silk sutures are those developed based on natural silk proteins and contain braided structure. Due to this feature, it has excellent handling characteristics. Chen et al. have developed a silk-based suture with an inbuilt slowrelease property of levofloxacin hydrochloride and poly (ε-caprolactone) by loading them before braiding using two-dipping-two-rolling method [37]. More interestingly, the developed suture has passed the minimum criteria of knot-pull strength as set by U.S. Pharmacopoeia. Also, the prepared sutures have slightly different physical and handling properties such as bending stiffness and thread-to-thread friction. This suture also demonstrated good antibacterial activity and sustained-drug release pattern. As per USP, the sutures used in clinical management should possess a minimum strength of 14.1 MPa while the mean radio-opacity values should be a minimum of 162.9
17.01. As shown in Fig. 9.4, radioopaque silk sutures have been developed by sequential crosslinking reactions involving 2,5-dimethoxy-2,5-dihydro-furan (DMDF)iodine employing higher temperature and pressure [38]. In vivo studies revealed that modified silk sutures possess a mean radio-opacity value of 213
19.46 under computed tomography after 28 days. At the same time, the developed sutures also showed appreciable mechanical properties (tensile strength and knot strength).

210

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 9.4 Outline of the preparation of silk sutures through crosslinking with iodine to produce antimicrobial and radio-opacity properties. (Reprinted from Ref. [38].)

Braided silk sutures are prone to microbial growth and hence antibacterial agents are added to impart antimicrobial properties. Chen et al. studied the physical and chemical properties of the antibacterial braided silk sutures (18/3000/2) such as knot-pull tensile strength, bending stiffness, tissue-drag friction resistance, and pullout friction resistance [39]. This study’s findings revealed that knot pull strength was decreased but still higher than the standard readings while the suture-to-tissue friction and bending stiffness was increased by nearly 50%. Through adjusting the coating parameters, the physio-chemical characteristics of the antibacterial silk sutures can be determined. Similarly, in a report, the biomechanical evaluation and handling test on surgical silk sutures after modification with antibacterial compounds was studied [40]. Apparently, a 50% increase in the bending stiffness and suture-to-tissue friction was seen with these treated sutures and the results are comparable to synthetic ones. Also, these sutures have good knot-pull tensile strength. Among the sutures investigated in this study, 18/3000/2 antibacterial silk sutures were found to be optimum for clinical applications. Franco et al. reported the preparation of antibacterial silk sutures involving the bio-engineered spider silk proteins and further coated with AMP (6mer-HNP1) [41]. The incorporation of these silk proteins into

Revisiting the properties of suture materials: an overview

211

Perma-Hand® sutures did not change their mechanical properties. As compared to untreated sutures, a significant reduction of biofilmproducing methicillin-resistant Staphylococcus aureus (w1.5 log) and E. coli (2 log) was observed with coated sutures. The use of non-mulberry silk scaffold named as Spidrex® in tendon regeneration was evaluated through in vitro studies and compared with fiberwire®, Bombyx mori silk scaffold, and a 3D collagen gel [42]. Spidrex® facilitated the adhesion of primary human and rat tenocytes and also promoted its cellular proliferation ad expression of tenocyte-related genes namely tenomodulin, fibromodulin and scleraxis. Immunogenicity analysis revealed that the dendritic cell maturation favored by Spidrex® was similar to that of the fiberwire®. An upregulated signals of proinflammatory cytokines indicate that Spidrex® stimulate early immune response and thus assists in tendon repair at an accelerated rate. The biocompatibility of silk fibroin sutures was improved by depositing the polysaccharides such as chitosan or hyaluronic acid as surface layers [43]. These type of suture materials finds their applications in general surgery and compression plastic surgery. 9.3.2 Poly(ε-caprolactone) based sutures Multi-fluid electrospinning technique was employed to fabricate the coresheath suture yarns composed of polyethylene glycol, polylactic acid, and polycaprolactone, which is further impregnated with curcumin [44]. The developed suture material possesses improved mechanical properties and controlled release of a drug. By varying the structural parameters of the yarns, the drug quantity and time of release can be significantly controlled. Suture made of fine fibers as sheath matrix displayed rapid decomposition and burst drug-release pattern. In a similar study, polymeric compositebased sutures made of PEG/PCL/chitosan-keratin blends were evaluated for their physical, thermal stability and mechanical properties while the diclofenac potassium-loaded sutures were tested for drug-release characteristics [45]. Optimum polymeric suture formulation with appreciable tensile properties was noticed at a ratio of 80/19/1 w/w of PEG/PCL/ chitosan-keratin. The fibers produced by this technique demonstrated good homogeneity and cell viability. In order to improve the mechanical properties of poly (ε-caprolactone), polyglycolic acid (PGA) suture yarn was added as a reinforcing material during the 3D printing process using fused deposition modeling [46]. For

212

Advanced Technologies and Polymer Materials for Surgical Sutures

the incorporation of 22 wt% of suture yarn in a 3D printed PCL matrix, an increase in the elastic modulus and tensile strength of up to 775% and 374% was found respectively. Compared to non-reinforced samples, reinforced samples demonstrated good mechanical strength for up to 4 weeks. Similarly, the impact of the incorporation of ethyl ester L-Lysine Triisocyanate on mechanical properties of poly(lactide)/poly(ε-caprolactone) polyester blends was studied by Visco and his group [6]. By altering the composition of these polyesters’ blends, the mechanical and physical properties can be selectively maneuvered. Both these polymers contribute synergistically to the overall mechanical characteristics. Interestingly, an increase in the ductility of the final polymer mixture was seen. In addition, the mechanical strength of the blend was significantly improved as reflected by tensile strength and yield stress. Owing to these features, the prepared polymer blends can be exploited as reabsorbable suture threads. Alternatively, a layer-by-layer assembly approach was employed to develop N-Halamine modified polyglycolide-based antibacterial sutures [40]. Even though the chlorination treatment slightly reduced the tensile properties of the prepared surgical suture, the biocompatibility was not compromised as evaluated through in vitro cytotoxicity and hemolytic studies. 9.3.3 Polyamide-based sutures Synthesis of polyamide suture materials with good shape memory effects and stereo-regulated thermomechanical properties was reported [47]. Here, the stereochemistry of the alkene delineates the material properties such as tensile strength, glass transition and modulus. Interestingly, a, b  unsaturated carbonyl moiety in the polymer backbone is formed as a result of the nucleophilic thiol-yne reaction. Through fine-tuning the reaction conditions, the alkene cis-content in the range of 35%e82% can be achieved. Moreover, the developed materials exhibited excellent in vitro and in vivo biocompatibility. As a non-resorbable biomaterial, polyamides demonstrated superior mechanical properties as compared to nylon and thus it is more suited for the development of stable biomedical devices. Braided sutures are widely used in dermatologic surgery. However, they lack antibacterial activity, which often leads to postoperative complications. Debbabi et al. have developed an effective antibacterial polyamide-based braided suture, which involves a three-layered biopolymer coating made of 1% chitosan and 10% citric acid using PAD-dry process [48]. This coating process did not hamper the mechanical properties of native braided

Revisiting the properties of suture materials: an overview

213

sutures and possess excellent antibacterial activity against four widespread clinical isolates. Similarly, the same group has developed a PET basedbraided antibacterial suture using citric acid as a crosslinking agent, which connects to the biofilm and the chitosan acts as an antibacterial agent [49]. An in vitro study demonstrated that rinsing the non-absorbable sutures namely silk and polyamide with a chemical adjunct, hyaluronic acid (HA) did not change their native tensile properties [50]. However, polyamide sutures displayed better tensile strength as compared to silk, which could be due to the stabilizing role of HA. In order to control incision site infections, the sutures are protected by antimicrobial coatings. In a study by Massod et al. polyethylene terephthalate and polyamide, threads were coated using a formulation containing naturally available materials like hydrolyzed chitosan, clove oil and turmeric powder [51]. Owing to the addition of these coatings, there is an improvement in both tensile and knot strength of the suture threads and the antibacterial activity against S. aureus was observed. The performance of polyamide and silk sutures on healing postdental extraction in a split-mouth study is investigated [52]. Experimental studies revealed that polyamide sutures performed better as compared to silk sutures, based on the socket reduction and healing index. Polyamide multifilament sutures are commonly used in flexor tendon repair. However, their mechanical properties could vary depending on the size of the knot and suture material used. Netscher et al. have found out that the forces occurring due to tendon movements are resisted by the two-knot threeloop polyamide sutures [53]. Apparently, the two-knot four looped suture constructs failed to resist these tendon motion forces. Hernandez et al. developed a new family of biocompatible and biodegradable amino acidbased poly(ester amide) sutures i.e., Phe-PEA and Arg-Phe-PEA [54]. This polymer-coated silk or plain gut sutures demonstrated a noticeable decrease in inflammatory response both in vitro and in vivo. 9.3.4 Collagen-based sutures Recently, collagen-based absorbable sutures from tannery solid waste were successfully developed [55]. The prepared suture has a tensile strength of 43.16 MPa and storage stability of 6 months. Besides, this suture possesses good biocompatibility and wound healing properties. Dasgupta et al. assessed the role of crosslinking chemistry in improving the mechanical properties of collagen microfibers-based sutures prepared using the microfluidic extrusion method [56]. In this study, glyoxal was employed as

214

Advanced Technologies and Polymer Materials for Surgical Sutures

a collagen crosslinker. Improved mechanical features of collagen microfibers include tensile strength (300 MPa), and young’s modulus (3 GPa). Also, the cytocompatibility and biocompatibility of collagen-based sutures were significantly improved after the crosslinking step. Furthermore, the M2 macrophage response was strongly elicited. Collagen-based scaffold (CS) membranes have shown improved mechanical properties in the presence of silk fibroin and it is well suited for applications in the field of corneal tissue engineering [57]. CS10 membrane, which contains 10 wt% silk fibroin demonstrated excellent mechanical properties. Importantly, developed CS membranes possess high suture retention strength. Moreover, these scaffolds demonstrated in vitro biocompatibility and also support the proliferation of human corneal endothelial cells. 9.3.5 Polyurethane-based sutures Thermoplastic polyurethane/poly(ε-caprolactone) blends were prepared and tested for their shape memory properties [58]. Sutures prepared with a percentage ratio of TPU and PCL of 1:3 yielded excellent shape recovery ratio and shape fixing ratio of 90% and 98% respectively. When stretched, the crystalline portion present in the TPU undergoes deformation and comes back to a temporary shape. Upon heating, the rubbery part of TPU assists in restoring the shape. Under hot water bath conditions, TPU25% could introduce a knit by itself. These blends showed good viability with 3T3 fibroblast cell lines. In a similar study, thermoplastic poly (carbonate) urethane (TPU) based synthetic sutures were compared against nonelastic polypropylene (PP) sutures for the mechanical strength and biocompatibility [59]. These TPU sutures demonstrated reduced accumulation of macrophages and improved host response. Mechanical evaluation of these TPU sutures confirmed the better tension curve and retained good tensile properties even after 30 min of external tension. Monofilaments prepared from TPU showed appreciable shore hardness (95A), and good elasticity with a >88% elongation rate. Xu and Chen, have reported the two-step shape memory properties of the polyurethane suture materials based on disulfide bonds and shape memory effects of poly(epsilon-caprolactone) of molecular weight (MW) 3000 [60]. These distinct shape-memory features are attributed due to the crystalline soft segment of polyurethane made of materials having MW > 2000. Owing to these properties, improved healing efficiency in

Revisiting the properties of suture materials: an overview

215

lesser healing time was seen. Similarly, Joo et al. reported that 30% polymer blends of polyurethane/polycaprolactone yielded superior shape-memory properties [61]. Here, PCL (highly crystalline) acts as a hard segment while the PU (heat stable) acts as a soft segment. Biocompatibility of these polymer blends was assessed with MC3T3-E1 cells. Results showed that these blends are non-toxic and also biodegradable in nature. Inspired by the features of collagen and elastin in blood vessels, triplelayered vascular grafts (Fig. 9.5) were fabricated using polyurethane, polyacrylamide (PAM) hydrogel and braided silk [46]. In this case, silk fibers mimic the characteristics of collagen while the thermoplastic polyurethane and PAM hydrogel contribute to the elasticity as that of elastin. Importantly, these grafts exhibit suture retention strength and burst pressure required during surgical operations. On the whole, the developed vascular grafts resemble the non-linear mechanical property of blood vessels. Javaid et al. devised a methodology to prepare biocompatible suture material using chitosan and polyurethane and various diisocyanates [62]. The incorporation of chitosan improves the biological behavior of polyurethane and exhibits diminished cytotoxicity, low hemocompatibility, and non-mutagenic.

Figure 9.5 Mechanistic view of nonlinear mechanical property of the triple-layered vascular grafts. (Reprinted from Ref. [46].)

216

Advanced Technologies and Polymer Materials for Surgical Sutures

9.3.6 Polypropylene sutures The radiation grating method was employed in the production of silverloaded antibacterial polypropylene (PP) sutures [39]. For functionalization of silver onto PP, grafting of either hydroxyethyl methacrylate (HEMA) or N-isopropyl acrylamide (NIPAAm) is employed. The composite PP monofilaments prepared with Ag contents of 3e5 wt% exhibited good control over the pathogens like E. coli, S. aureus and appreciable biocompatibility. Likewise, Razumov et al. described a method to develop monofilament polypropylene (PP) based suture materials with enhanced anti-inflammatory properties [63]. In this study, aspirin and indomethacin are the anti-inflammatory agents functionalized on the surface of PP. Initially, sutures were first treated using hydrogen peroxide for activation and subsequently anti-inflammatory agents were anchored on the PP surface. In vivo studies revealed excellent healing of surgical skin wounds, which was attributed to the anti-inflammatory substances present. Due to the beneficial features of this approach such as ease of implementation and cost-effective, these functional materials can be exploited in surgical applications. 9.3.7 Chitosan-based sutures The use of chitosan (CS) filaments as absorbable suture materials has gained more attention in recent years. In a separate study, the introduction of N-Acetyl-D-Glucosamine (GlcNAc) into the chitosan filaments did meet the minimum mechanical resistance values (1.7 N) set forth by the US pharmacopeia for suture number 6.0 (100e149 mm dia.) [64]. The surface morphology of these CS filaments remains unaltered even after GlcNAc incorporation. Moreover, the rate of biodegradation of CS filaments was rapid in the presence of N-Acetyl-D-Glucosamine. These CS/ GlcNAc suture filaments were non¼toxic to L929 cells. In a study by Mohammadi et al. chitosan and hyaluronic acid were sequentially incorporated onto nylon monofilament using a bilayer coating approach to improve its physical and biological properties [44]. Importantly, this suture composite possesses roughness of 164
129 nm, the friction coefficient of 0.26 and good antibacterial activity. A two-step dip-coating process employing chitosan and silver-doped mesoporous bioactive glass is reported for the preparation of antibacterial Vicryl® plus sutures [65]. The developed sutures showed antimicrobial activity against both Gram-positive S. carnosus and Gram-negative E. coli.

Revisiting the properties of suture materials: an overview

217

Vicryl plus, a triclosan-loaded polyglactin suture is the most commonly used antibacterial absorbable suture used to reduce surgical site infections in orthopedic surgery. However, due to the rising concerns associated with the toxicity and clinical safety of triclosan, there is an urgent need to find potential sutures for these applications. As a potential alternative, Yang et al. developed an antibacterial suture by coating the hydroxypropyl trimethyl ammonium chloride chitosan, a derivative of chitosan onto the surface of vicryl suture [66]. The prepared suture displayed excellent antibacterial performance and biofilm inhibition both in vitro and in vivo. Importantly, these sutures exhibited good cytocompatibility with human skin-derived fibroblast cells. Similarly, the performance of two absorbable suture materialspolyglactin 910 6-0 (Vicryl) and polydioxanone 6-0 (PDS-II), on the facial scar spread during the revision procedure by elliptical excision and intradermal suturing was comparatively investigated [67]. The findings of the controlled trial with the help of Vancouver Scar Scale indicated that the scar spread and quality were better in the case of polydioxanone as compared to polyglactin 910. The effect of chitosan coating over Vicryl and PDS sutures for intestinal anastomosis was investigated [68]. In vitro experiments revealed that an increase in tensile strength was observed in chitosan-coated Vicryl as compared to Vicryl on 7th and 14th day while it decreases upon chitosan coating of PDS. On contrary, in vivo studies demonstrated that chitosan coating of PDS provided clinical benefits over anastomosis. Moreover, the vascularization of anastomosis follows the sequence, Chitosan@Vicryl > Vicryl > Chitosan@PDS sutures. Shao et al. studied the feasibility of using the diacetyl chitin (DAC) as an absorbable suture in wound closure [69]. Their mechanical properties were almost comparable to that of synthetic polymeric sutures. The potential use of DAC sutures in wound healing was supported by the experiments in a linear incisional wound model, which showed it gets completely absorbed within 42 days, causes a negligible adverse reaction in vivo, and promoted rapid tissue reconstruction and healing. 9.3.8 Bio-based sutures Degummed fiber of ramie (Boehmeria nivea) plant was reportedly used to prepare suture material, which demonstrates good tensile properties, biocompatibility and in vivo would healing activity [70]. These results were comparable to commercial BMSF sutures. Similarly, the use of animal skin

218

Advanced Technologies and Polymer Materials for Surgical Sutures

fibers (buffalo) and cotton fibers in the construction of multifilament sutures was recently reported [71]. The prepared suture material was robust and biocompatible in nature with 85.25 Mpa tensile strength, knot strength of 7.59 kgf, and elongation at a break of 81.9%. In vivo studies demonstrated its usefulness in accelerated wound healing. A novel suture material with remarkable mechanical properties such as tensile strength (4 N) and swelling ratio (165.55%e183.23%) under saline conditions was prepared from gellan gum, which is commonly referred as gellan-polylysine polyion complex fibers (GPF) [72]. Interestingly, these sutures showed neither in vitro cytotoxicity effects with CCK8 cells nor hemolytic behavior with mice blood. More importantly, in vivo mice models failed to show any adverse reactions or infections.

9.4 Properties of suture materials: comparative analysis 9.4.1 Physico-mechanical properties Polyglyconate or polydioxanone resorptive surgical materials are commonly used in gastro-intestinal surgery and several manufacturers’ produce these sutures for the commercial market. Experimental investigations revealed that polydioxanone (Maxon®) sutures possess higher tensile strength than that polyglyconate sutures [73]. After 10 days of implantation of these sutures within the mucosa of the stomach, polydioxanone (Surgicryl®) displayed higher tensile strength retention (91.4%) and initial elasticity and higher initial young modulus. However, up to 55% of initial tensile strength is only retained by the polydioxanone (PDX®). Similarly, in a study, by Dueñas-Garcia et al. different suture materials (polyglyconate, Polyester, Polytetrafluoroethylene, 0-polydioxanone, 00polydioxanone, Polypropylene) and knot techniques were explored for pelvic reconstructive procedure [74]. Compared to robotic knots, hand-tied knots showed better tensile strength and low elongation. Importantly, polyglyconate suture displayed the greatest tensile properties while the OO-polydioxanone showed the lowest strength. Knot integrity was almost similar between the slip knot and flat square knots. A group of researchers from Oregon have investigated the tensile strength of 13 suture materials used in cutaneous surgery, as per US Pharmacopeia [75]. Among the absorbable and non-absorbable sutures studied, a high value of initial modulus was recorded for rapid polyglactin 910 (9320 MPa) and silk (8701 MPa). Compared to absorbable sutures

Revisiting the properties of suture materials: an overview

219

(Glycomer 631, polyglyconate), non-absorbable sutures namely nylon and polypropylene exhibited relatively higher failure load under straight and knotted configurations. Handling characteristics of the absorbable (Poliglecaprone & Polyglactin 910) and non-absorbable (nylon and polyester) surgical sutures of both braided and multifilament nature could be influenced by the mechanical parameters such as bending rigidity and suture-tosuture coefficient of friction [76]. Polyglactin 910, polypropylene, and polydioxanone are the suture materials used for abdominal fascial closure. Segura-Ibarra et al. have examined the mechanical properties of these sutures after the application of a known amount of tensile strain [77]. Among the suture materials studied, polypropylene, and polydioxanone did not show significant changes in the maximum tensile forces while polyglactin 910 displayed an increase in young modulus. All these sutures have a maximum tensile force and elastic limit as high to withstand the pressure in the abdominal region. Tribological interactions between the surgical suture and human skin is an important factor, which is characterized by structure and surface topography. Zhang et al. studied the tribological interaction of three different suture materials viz. nylon, silk and polyglactin using the capstan method [78]. Amongst, vicryl (polyglactin-based synthetic suture) exhibited the lowest apparent coefficient of friction. This measured quantity increases with the increase in sliding velocity or decrease in applied load. Regier et al. reported a comparative study of the intradermal closure using two different types of sutures-monofilament versus unidirectional barbed sutures in dogs [79]. Fluid leakage, maximum load, and stiffness were the parameters monitored in this study. There is no significant difference in stiffness between both these sutures, the maximum load tolerated by polyglyconate monofilament was 2.5-fold stronger than that of Quill Monoderm barbed suture. However, barbed suture provides low fluid release and hence provides a relatively better watertight seal. Sutures used in surgeries are also affected by the type of treatment regimen. It is important to investigate the tensile properties of the suture used in various types of surgery. In a separate study, the effect of (a) heat and (b) chemotherapy on the biomechanical properties of sutures was investigated [80]. Here, the elongation rate and tensile breaking force (TBF) of four different types of absorbable sutures namely Biosyn, Dexon II, Maxon, Monocryl, PDS II, and Vicryl Plus were not significantly affected due to heat and chemotherapy. Among these sutures, Maxon remains the strongest type with TBF of 59.6 N. Elongation rate was higher for the monofilament sutures than the multifilament sutures. Muller et al. performed the

220

Advanced Technologies and Polymer Materials for Surgical Sutures

mechanical evaluation of the different types of absorbable suture materials namely Vicryl, Maxon, Monocryl, PDS II, Vicryl rapide and a nonabsorbable suture, Ethibond used in orthopedic surgery [81]. For the initial 2 weeks, higher loads were sustained only by absorbable sutures e Maxon, Vicryl, PDS II. In spite of degradation and loss of tensile effects, elasticity properties were retained only by Maxon and PDS II. Similarly, Johnson et al. investigated the effect of instrumentation on knot pullout strength and tensile strength of five different suture materials such as FiberWire, Ethibond, Vicryl, Prolene and Monocryl [82]. Results showed that there is no significant difference between instrumented and noninstrumented groups. In a particular study by Silver et al., the knot security of four different suture materials (chromic gut, nylon, silk, and vicryl) was tested for its application in oral and maxillofacial surgery [83]. Results showed that the knot security follows the order: vicryl > chromic gut > nylon > silk. Moreover, knot security increases with an increase in the number of throws up to five. Suture materials such as polytetrafluoroethylene, silk, and polyglactin 910 are commonly used in periodontal surgery. Tensile properties of the suture materials may be affected through regular contact with the ingredients present in the mouthwash or oral rinse formulations. Myrrh is a type of plant resin, which is used in mouthwash formulations. Alshehri et al. showed that the tensile strength was reduced upon exposure to high concentrations of myrrh [84]. Considering this effect, silk sutures are the least preferred choice among others. Tensile strength of different types of suture materials namely silk, e-PTFE and polyamide (Supramid®, monofilament) was investigated [85]. For the 5 mm traction applied, e-PTFE 4-0 showed better resistance without breaking or untying. Among the 5-0 sutures, tensile strength follows the order: Supramid® > monofilament > silk. In a similar study by Abellan et al., where the five different suture materials namely polyamide 6/66, glycolide-ecaprolactone copolymer, polyglycolic acid, polytetrafluoroethylene and silk were physically examined on three-knot configurations [86]. Mechanical properties such as knot strength and elongation depending on the suture materials being tested. The failure load of the suture materials is as follows, polyglycolic acid > glycolide-e-caprolactone copolymer >> polytetrafluoroethylene. The knot failure load and knot configuration are directly related. Expanded politetrafluorene (ePTFE) sutures are used in mitral valve repair as chordae substitutes. Caimmi et al. reported the mechanical properties of ePTFE sutures [87]. Here, an increase in stiffness was

Revisiting the properties of suture materials: an overview

221

significant for 3 cm long, which is the requirement of artificial chordae. An increase in length of chordae not more than 7 cm increases the stiffness during the transapical approach but the resistance gets compromised in this case. For the braided-polyethylene sutures, different arthroscopic knot types such as were studied to evaluate the mechanical properties [88]. Among the knot types, dines knot displayed high reliability with the highest loop and knot security. In rotator cuff repair, the suture-tendon interface is the crucial point where most of the time, the suture fails. To encounter this, abrasive and strong sutures is the need of the hour. Williams et al. investigated the abrasiveness of nine different braided sutures namely FiberWire, Collagen Coated FiberWire, Orthocord, MaxBraid, Force Fiber, ULTRABRAID, Phantom Fiber BioFiber, Ti-Cron and Surgipro, which are utilized in rotator cuff surgery [89]. Mean displacement rate and abrasiveness are the parameters monitored in this study. Amongst, Collagen Coated FiberWire was found to be the most abrasive braided suture while a mean displacement rate of >0.150 mm/cm was common with sutures such as Ti-Cron, Phantom Fiber BioFiber, FiberWire, and Collagen Coated FiberWire. Lipatov et al. have described a methodology to choose the suture material used in liver surgery [90]. Physico-mechanical properties of the three suture threads namely catgut (monofilament), capron (twisted), and polyglycolide (braided) were evaluated in this study. For experimental research on the liver, chosen suture materials should have the ability to withstand the high axial load. Two parameters namely extent of pulling of thread (Lu), and maximal effort applied (Fmax) of the suture materials were measured. Among the suture threads tested, monofilament is the least preferred. 9.4.2 Biological properties Postoperative patient care is an important criterion in health care management. Surgical site infections are one of the most frequent problems encountered in our day-to-day life. Dixit et al. reported the clinical safety and efficacy of MITSUÔ Polyglactin 910 Suture for closure of surgical incision [91]. The clinical performance of this test suture was comparable to the reference product, coated Vicryl® Polyglactin 910 suture. The efficacy of the Monocryl® Plus and silk sutures toward the surgical removal of impacted lower third molars (I3M) was investigated [92]. Here, the critical parameters such as surgical site bacterial infections, wound bleeding and degree of discomfort are taken into account. Even after 72 h, good antibacterial activity was observed with Monocryl® Plus sutures.

222

Advanced Technologies and Polymer Materials for Surgical Sutures

In a study by Asher et al. four kinds of sutures viz. polyester, silk, nylon, and coated polyglactin were mainly tested for bacterial accumulation postdental surgery [93]. Amongst, nylon (monofilament) sutures showed fewer bacterial counts. In all cases, antibiotic treatment and surgery type (periodontal/implant) did not completely retard the colonization of bacteria over sutures. After caesarean delivery, wound complications associated with surgical sutures used are quite frequent in pregnant women. In order to mediate the subcuticular skin closure, the suture materials used should be carefully chosen. In a study by Tuuli et al. monofilament (Monocryl) and multifilament braided (Vicryl) synthetic sutures were critically examined for subcuticular closure of skin incisions seen after caesarean [94]. Results indicated that there is no significant difference in infections and wound complications associated with surgical sites between these two synthetic sutures. A group of researchers from Germany have studied the adhesion of bacteria onto the suture materials-barbed Quill, monofilament Ethilon II, braided sutures Vicryl and triclosan-coated Vicryl Plus in wound closure [95]. Results indicated that barbed sutures are recommended for aseptic surgery while careful monitoring is required during septic surgery. The influence of different suture materials like PDS, Vicryl, and Prolene for abdominal wall closure was investigated in male Wistar rats [42]. Based on histological outcomes, these sutures were compared to provide better fascia healing. Considering the number of macrophages and foreign body giant cells accumulated, the preference of suture materials for abdominal wall surgery is as follows, PDS > Prolene > Vicryl. Wang et al. studied the utilization of PDS (absorbable) versus nylon (non-absorbable) materials on nasal width for applications of cinch sutures in orthognathic surgery [96]. Alar curvature (Ac) width and the alar base (Al) are the parameters of the nasal width, measured from preoperative and postoperative 3D craniofacial images. Results showed that there is no significant difference between PDS and nylon in maintaining the peri-operative nasal widening. However, nylon has a few clinical limitations such as nasal irritation and suture exposure, which could be easily met by the use of PDS sutures. Morris et al. investigated the effect of three different absorbable sutures -monofilament, barbed monofilament and braided, on bacterial adhesion and tissue reactivity of S. aureus contaminated wound mouse model [97]. On the whole, between monofilament (PDS II) and barbed monofilament (Quill) sutures, there was no significant differences in bacterial adhesion and tissue reactivity. However, braided sutures (Vicryl) may be the least preferred based on these results.

Revisiting the properties of suture materials: an overview

223

As a step toward preventing post-operative wound complications, two different suture materials namely polydiaxanone and polypropylene were investigated for closure of midline laparotomy incision employing the interrupted X suture technique [98]. Among the two materials, polydiaxanone is found superior in preventing wound problems. As compared to polypropylene, both the wound-associated complications such as (suture sinus, burst abdomen) and infection rate were lesser in the case of polydiaxanone. The performance of resorbable sutures (5/0 irradiated Polygalactic acid) and non-resorbable sutures (polypropylene) for nasal transcolumellar incision closure in rhinoplasty was examined [99]. Importantly, less patient discomfort was seen with resorbable sutures as compared to non-resorbable ones. However, no risk for post-operative infection and scar visibility was observed with both these groups. In a clinical study by Dragovic et al. suture materials of the absorbable and non-absorbable types under two different categories of materials viz. monofilament and multifilament were investigated for oral wound healing and its clinical safety [100]. Results have shown that soft tissue healing observed with the monofilament sutures was relatively better. Overall, polypropylene suture (non-absorbable) showed appreciable clinical features, including soft tissue healing and low inflammation. On the other hand, non-absorbable silk sutures received the least preference due to poor tissue healing, higher biofilm formation, and adverse inflammatory effects. In a recent report, two different types of absorbable sutures e knotless barbed and monofilament materials were compared for wound closure in the oral mucosa of cats [101]. The average wound closure time was 3.6 min lesser for the barbed sutures than that of conventional monofilament. These data suggest that knotless barbed sutures were not less superior to the interrupted monofilament sutures with a similar pattern of dehiscence and swelling-like postoperative complications. Lambertz et al. have investigated the in vivo biocompatibility of polyvinylidene fluoride (PVDF) based monofilamental suture material against five different commercial sutures in male rats [43]. Although qualitatively, the foreign body reaction was almost the same between the sutures. Among all, PVDF-based sutures showed decreased granuloma size and comet tail-like infiltrate as compared to the other established suture materials.

224

Advanced Technologies and Polymer Materials for Surgical Sutures

9.5 Micro and nanotechnology-enabled suture materials Most of the current existing sutures release drugs or antimicrobials based on a passive diffusion mechanism. The development of smart sutures, which have active control over the drug release is the need of the hour. Toward this, Morelli et al. have reported the construction of electronic sutures, which is capable of metered dosing of drugs [102]. In the design shown in Fig. 9.6, gold microwires were loaded in the drug encapsulated polymeric

Figure 9.6 Graphical outline of the controlled drug release pattern of electronic sutures. (Reprinted from Ref. [102].)

Revisiting the properties of suture materials: an overview

225

suture composed of cellulose acetate phthalate. Upon introduction of suitable redox potential at the gold surface, pH was increased and thus causes the polymers to dissolute and release the active agents. The drug release process can be regulated by controlling the magnitude of potential applied or the time over which it is applied. One of the major limitations of this technique is that the drug should not carry any reducible functional group e.g., nitro moiety, which may be probably reduced due to the imposed redox potential. To increase the practical applicability of the microfibers, microfluidic spinning fibers have been reportedly used as ophthalmology suture material [103]. Interestingly, this microfluidic spinning method has exercised good control over the mechanical properties of the fibers produced. This can be also achieved through coating the biomolecules over the fiber, controlling the number of the single microfibers strands that tend to form fibers, and in situ initial mixing of the biomolecules with fibers. In vivo experiments of bundle fiber in porcine eyes exhibited smoother surfaces during postoperative management. Infections associated with surgical sutures are very common in postoperative care. To encounter this, antimicrobial sutures are currently being developed. With the advent of nanotechnology, nanostructured materials are used as antimicrobial agents. Serrano et al. have described a method to produce antibacterial sutures where the nanostructures are coated on the surface of commercially available sutures to prevent the adherence of bacteria and inhibit the biofilm formation [104]. For instance, silk sutures were enriched by Cefixime nanoparticles incorporated polymer hydrogel [105]. An improvement in mechanical strength was observed between the untreated and treated silk sutures. As compared to the untreated sutures, a reduction in SSI against resistant and non-resistant bacteria was seen in the reinforced sutures. Wu et al. reported the preparation of triclosan-loaded L-lysine based nanogel grafted silk sutures (TLNGSs), which demonstrated antibacterial activity against S. aureus and E. coli [106]. TLNGSs showed good biocompatibility and in vitro cytocompatibility. Moreover, this suture showed a sustained release of triclosan in trypsin solution. This suture possesses relatively good mechanical properties and meets the minimal USP requirements following ASTM D2256 standards. To overcome the antibiotic resistance, antibacterial surgical sutures were developed based on the deposition of in situ formed silver clusters [107]. The silver coating on the sutures renders antibacterial properties with no adverse cytotoxicity and also promotes wound healing. Syukri et al. have

226

Advanced Technologies and Polymer Materials for Surgical Sutures

reported the preparation of antibacterial silk sutures by depositing the biosynthesized silver nanoparticles on its surface [108]. The addition of nanoparticles increases the overall elasticity and reduced the surface roughness of sutures. These sutures possess excellent bactericidal properties against Gram-negative pathogens such as E. coli, K. pneumoniae, A. baumannii, P. aeruginosa and this property is retained even up to 12 weeks. Bioactive polyester surgical sutures were prepared by immobilization of chlorhexidine and nanosilver nanogels on the surface of poly (ethylene terephthalate) [109]. Particles found within the nanogel are of diameter 10e50 nm. These sutures displayed good antimicrobial and would healing properties without causing inflammation. Gallo et al. described a methodology to produce absorbable PLGA sutures with both antibacterial and tissue regenerative activity [110]. Here, the sutures were functionalized with silk sericin and silver nanoparticles are also in situ deposited to provide wound healing properties. Malafeev et al. developed a polymeric nanocomposite where the chitin nanofibrils were loaded onto the polylactide based monofilament fibers [111]. The incorporation of 1% chitin-biolignin and CN-PEG makes the composite fibers brittle while the mechanical properties did not change. Importantly, strength in a single knot is one of the critical parameters, which determines the characteristic of suture yarn and it remained the same value even after the addition of chitin nanofibrils within the PLA matrix. In a similar study, Zhang et al. reported that there is a significant improvement in mechanical properties such as knot strength, knot-pull strength and tensile strength of chitin-based sutures after the addition of 1.6 wt% of graphene-oxide within the pure chitin matrix [112]. Moreover, these suture materials exhibit biocompatibility both in vitro and in vivo. Also, the expression of key genes involved in mediating inflammation remains unaffected in the presence of graphene-oxide. In a recent study by Wu et al. addition of bacterial cellulose nanocrystals (BCNC) into regenerated chitin (RC) fibers tend to drastically improve the mechanical strength of the BCNC/RC filament shown in Fig. 9.7 [113]. However, the biodegradability of this suture material depends on the amount of BCNC loaded into the yarn. Knot-pull tensile strength of 9.8 N was observed with a yarn, which is built using 30 BCNC/RC filament fibers. Moreover, the developed BCNC/RC fibers did not show cytotoxicity, facilitate cell proliferation, and wound healing. The synthesis of smart bio-based sutures is an emerging area of research. Duarah et al. have reported the preparation of surgical sutures based on

Revisiting the properties of suture materials: an overview

227

Figure 9.7 Proposed scheme of bacterial cellulose nanocrystals/regenerated chitin fibers-based sutures for in vivo wound healing applications. (Reprinted from Ref. [113].)

starch-modified polyurethane-reduced carbon dot nanocomposites [114]. The incorporation of 2 wt% carbon dots in hyperbranched-polyurethane imparted remarkable thermal (286 C) and mechanical properties such as toughness (439.28 MJ m3), tensile strength (32.14 MPa), and elongation (1576%). Most importantly, these nanocomposites possess noncontact selftightening and shape memory properties (15 s at 37 C) and were biocompatible, hemocompatible, and biodegradable. The degradation properties of poly (lactic acid) sutures can be controlled by the addition of nanomaterials. This was indeed proved through in vitro experiments by the incorporation of carbon nanotubes (CNT) into PLA sutures [115]. The strength valid time of these PLA/CNT composite sutures was 26.6 weeks while that of plain PLA sutures was only 13.5 weeks. Interestingly, the complete degradation of these composite sutures based on mass loss measurements took 63e73 weeks and that of PLA sutures took about 49 weeks. Initially, slow mass loss is recorded and rapid loss is observed at a later stage. The total degradation period was found to vary with an increase in CNT i.e., 63 weeks (0.5 wt% CNT/PLA), 73 weeks (1 wt% CNT/PLA) and 65 weeks (2 wt% CNT/PLA).

9.6 Conclusions and future outlook Initially, the surgeons have the option of choosing either an absorbable or non-absorbable suture. But, at present, factors like rapid wound-healing, antimicrobial property, shape-memory effect, elasticity and many others are in practice. For instance, Vicryl Plus with TriclosaneEthicon, is an antimicrobial suture, which prevents surgical site infections and also promotes

228

Advanced Technologies and Polymer Materials for Surgical Sutures

wound healing to a certain extent. NovafileTyco Healthcare/Kendall Animal Health, a suture can be readily used in rapid healing tissues and proper closure of linea alba. This chapter discusses the recent advances in the properties of suture materials. Amongst, the mechanical and biological properties of suture materials are extensively studied. Generation of novel suture materials with shape-memory, elastic, electronic metered-drug release properties and stereochemical regulated systems are the significant findings. From an application perspective, tissue approximation, wound closure, and negligible surgical site infections are the important benefits of bio-active sutures. The incorporation of nanostructured materials onto suture materials imparts antibacterial behavior and thus significantly controls the surgical site infections and also provides a better microenvironment for wound healing. Apparently, most of the properties of suture materials have been substantially improved due to their conjunction with micro and nanotechnology-enabled solutions. However, only limited literature exists to develop novel smart sutures at the bio-nano interface. At present, the lack of clinical validation of these nano-engineered suture systems is the main bottleneck in the life cycle of suture-like product development. For a careful selection of suture materials for wound closure, it is important to study the parameters like bio-material interactions, wound tissue configuration and mechanical properties. Many sutures have been developed so far. However, only a few sutures have been tested for their safety and biocompatibility aspects. Clinical evaluation of bio-active sutures is the most essential and rate-determining step in the commercialization process. However, it is still in its infancy. Recently, a lot of attention has been gained on the design and development of bio-based sutures involving the use of plant or animal fibers as raw materials, which impart material characteristics like biocompability and mechanical strength. It is important to develop novel surgical threads inbuilt with native biological components such as stem cells, growth factors or specific proteins, which favors the regeneration of tissues and wound healing at a rapid rate. Taking into account the surgical procedures and the forces held, one should test the prepared sutures for their characteristic properties to yield better clinical outcomes. In order to improve the surgical outcomes, researchers may pave the attention to regulatory issues associated with the development of suture materials. Surgical sutures are slowly evolved from being a physical material to a better theragnostic agent capable of delivering drugs or cells to the targeted site.

Revisiting the properties of suture materials: an overview

229

References [1] C. Dennis, S. Sethu, S. Nayak, L. Mohan, Y. Morsi, G. Manivasagam, Suture materials e current and emerging trends, J. Biomed. Mater. Res. A. 104 (6) (2016) 1544e1559. [2] T.E. Bucknall, Factors affecting wound-healing problems in general surgery 6 (2) (1989) 194e219. [3] K. Bath, S. Aggarwal, V. Sharma, Sushruta: father of plastic surgery in benares, J. Med. Biogr. 27 (1) (2019) 2e3. [4] K. Wu, P. Tang, The history of knots and surgical suturing, in: P. Tang, K. Wu, Z. Fu, H. Chen, Y. Zhang (Eds.), Tutorials in Suturing Techniques for Orthopedics, Springer, Singapore, 2021, https://doi.org/10.1007/978-981-33-6330-4_1. [5] C.C. Chu, Mechanical properties of suture materials: an important characterization, Ann. Surg. 193 (3) (1981) 365. [6] X. Chen, D. Hou, X. Tang, L. Wang, Quantitative physical and handling characteristics of novel antibacterial braided silk suture materials, J. Mech. Behav. Biomed. Mater. 50 (2015) 160e170. [7] M. Pensalfini, S. Meneghello, V. Lintas, K. Bircher, A.E. Ehret, E. Mazza, The suture retention test, revisited and revised, J. Mech. Behav. Biomed. Mater. 77 (2018) 711e717. [8] X. Deng, M. Qasim, A. Ali, Engineering and polymeric composition of drug-eluting suture: a review, J. Biomed. Mater. Res. A. 109 (2021) 2065e2081. [9] S. Sheik-Ali, W. Guets, Absorbable vs non absorbable sutures for wound closure. Systematic review of systematic reviews, Wound Med. 23 (2018) 35e37. [10] C.C. Chu, Types and properties of surgical sutures, in: W.K. Martin, S.G. Bhupender, G. Robert (Eds.), Biotextiles as Medical Implants, Woodhead Publishing, Cambridge, UK, 2013 Jan 1, pp. 231e273. [11] B. Yadav, A. Pandey, L.R. Kumar, R.D. Tyagi, Bioconversion of waste (water)/ residues to bioplastics-a circular bioeconomy approach, Bioresour. Technol. 298 (2020) 122584. [12] K.B. Lee, K.R. Yoon, S.I. Woo, I.S. Choi, Surface modification of poly (glycolic acid)(PGA) for biomedical applications, J. Pharmacol. Sci. 92 (5) (2003) 933e937. [13] K. Budak, O. Sogut, U.A. Sezer, A review on synthesis and biomedical applications of polyglycolic acid, J. Polym. Res. 27 (8) (2020) 1e9. [14] A.J. Dart, C.M. Dart, Suture material: conventional and stimuli responsive, in: P. Ducheyne (Ed.), Comprehensive Biomaterials, Elsevier Science, Amsterdam, Netherlands, 2011, pp. 573e587. [15] H.R. Ford, P. Jones, B. Gaines, K. Reblock, D.L. Simpkins, Intraoperative handling and wound healing: controlled clinical trial comparing coated VICRYL® Plus antibacterial suture (coated polyglactin 910 suture with triclosan) with Coated VICRYL® suture (coated polyglactin 910 suture), Surg. Infect. 6 (3) (2005) 313e321. [16] J. Conn, R. Oyasu, M. Welsh, J.M. Beal, Vicryl (polyglactin 910) synthetic absorbable sutures, Am. J. Surg. 128 (1) (1974) 19e23. [17] J.L. Myers, D.B. Campbell, J.A. Waldhausen, The use of absorbable monofilament polydioxanone suture in pediatric cardiovascular operations, J. Thorac. Cardiovasc. Surg. 92 (4) (1986) 771e775. [18] I. Manavitehrani, A. Fathi, H. Badr, S. Daly, A. Negahi Shirazi, F. Dehghani, Biomedical applications of biodegradable polyesters, Polymers 8 (1) (2016) 20. [19] C.C. Chu, Materials for absorbable and nonabsorbable surgical sutures, in: Biotextiles as Medical Implants, Woodhead Publishing, Cambridge, UK, 2013, pp. 275e334.

230

Advanced Technologies and Polymer Materials for Surgical Sutures

[20] R.W. Postlethwait, D.A. Willigan, A.W. Ulin, Human tissue reaction to sutures, Ann. Surg. 181 (2) (1975) 144. [21] S. Sahu, S. Mishra, S. Lenka, R. Banerjee, S. Pachisia, S. Ghosh, Comparison between N-butyl cyanoacrylate tissue adhesive and Ethilon nylon sutures in extraoral maxillofacial incisions: a randomized prospective study, J. Oral Biol. Craniofacial Res. 9 (3) (2019) 173e178. [22] R.L. Moy, B. Waldman, D.W. Hein, A review of sutures and suturing techniques, J. Dermatol. Surg. Oncol. 18 (9) (1992) 785e795. [23] B.S. Bloom, D.J. Goldberg, Suture material in cosmetic cutaneous surgery, J. Cosmet. Laser Ther. 9 (1) (2007) 41e45. [24] R.H. Tan, R.J. Bell, B.A. Dowling, A.J. Dart, Suture materials: composition and applications in veterinary wound repair, Aust. Vet. J. 81 (3) (2003) 140e145. [25] H.W. Boothe, Suture materials, tissue adhesives, staplers, and ligating clips, Textbook Small Anim. Surgery 1 (2003) 235e244. [26] J. Niles, J. Williams, Suture materials and patterns, Practice 21 (6) (1999) 308e320. [27] E.J. Sutow, Iron-based alloys, in: D.F. Williams (Ed.), Concise Encyclopedia of Medical and Dental Materials, Pergamon Press, New York, NY, 1990, p. 232. [28] V. Arma, M.W. Lumb, L.W. Johnson, H.L. Ferguson, Sutures in infected experimental wounds, Am. J. Vet. Res. 42 (1981) 571e574. [29] D.D. Smeak, Selection and use of currently available suture materials, in: M.J. Bojrab, G.W. Ellison, B. Slocum (Eds.), Current Techniques in Small Animal Surgery, Williams and Wilkins, Philadelphia, PA, 1998, pp. 19e27. [30] C.R. Bellenger, Sutures: part 1. The purpose of sutures and available suture materials, Compend. Continuing Educ. Pract. Vet. 4 (1982) 507e515. [31] R.E. Doran, D. Allen, The use of stapling devices in equine gastrointestinal surgery, Compend. Continuing Educ. Pract. Vet. 9 (1987) 854e860. [32] C.T. Ammirati, Advances in wound closure material, in: W.D. James (Ed.), Advances in Dermatology, 18, Mosby, St. Louis, 2002, pp. 313e318. [33] J.L. Fick, R.E. Novo, N. Kirchhof, Comparison of gross and histologic tissue responses of skin incisions closed by use of absorbable subcuticular staples, cutaneous metal staples, and polyglactin 910 suture in pigs, Am. J. Vet. Res. 66 (2005) 1975e1984. [34] J.W. Beam, Tissue adhesives for simple traumatic lacerations, J. Athl. Train. 43 (2008) 222e224. [35] P.J. Sniezek, H.W. Walling, J.R. Debloom, M.J. Messingham, M.J. Vanbeek, C.D. Kreiter, D.C. Whitaker, C.J. Arpey, A randomized control trial of high-viscosity 2-octyl cyanoacrylate tissue adhesive versus sutures in repairing facial wounds following mohs micrographic surgery, Dermatol. Surg. 33 (2007) 966e977. [36] G.R. Scott, C.L. Carson, G.L. Borah, Dermabond skin closures for bilateral reduction mammaplasties: a review of 255 consecutive cases, Plast. Reconstr. Surg. 120 (2007) 1460e1465. [37] X. Chen, D. Hou, L. Wang, Q. Zhang, J. Zou, G. Sun, Antibacterial surgical silk sutures using a high-performance slow-release carrier coating system, ACS Appl. Mater. Interfaces 7 (40) (2015) 22394e22403. [38] N.K. Francis, H.S. Pawar, P. Ghosh, S. Dhara, In situ iodination cross-linking of silk for radio-opaque antimicrobial surgical sutures, ACS Biomater. Sci. Eng. 2 (2) (2016) 188e196. [39] F. López-Saucedo, G.G. Flores-Rojas, J. López-Saucedo, B. Magariños, C. AlvarezLorenzo, A. Concheiro, E. Bucio, Antimicrobial silver-loaded polypropylene sutures modified by radiation-grafting, Eur. Polym. J. 100 (2018) 290e297.

Revisiting the properties of suture materials: an overview

231

[40] M.M. Umair, Z. Jiang, W. Safdar, Z. Xie, X. Ren, N-halamine-modified polyglycolide (PGA) multifilament as a potential bactericidal surgical suture: in vitro study, J. Appl. Polym. Sci. 132 (2015) 42483. [41] A.R. Franco, E.M. Fernandes, M.T. Rodrigues, F.J. Rodrigues, M.E. Gomes, I.B. Leonor, D.L. Kaplan, R.L. Reis, Antimicrobial coating of spider silk to prevent bacterial attachment on silk surgical sutures, Acta Biomater. 99 (2019) 236e246. [42] S. Van Steensel, L.C. van den Hil, A. Bloemen, M.J. Gijbels, S.O. Breukink, J. Melenhorst, K. Lenaerts, N.D. Bouvy, Prevention of incisional hernia using different suture materials for closing the abdominal wall: a comparison of PDS, Vicryl and Prolene in a rat model, Hernia 24 (1) (2020) 67e78. [43] A. Lambertz, K.M. Schröder, D.S. Schöb, M. Binnebösel, M. Anurov, U. Klinge, U.P. Neumann, C.D. Klink, Polyvinylidene fluoride as a suture material: evaluation of comet tail-like infiltrate and foreign body granuloma, Eur. Surg. Res. 55 (1e2) (2015), 1-1. [44] H. Mohammadi, F. Alihosseini, S.A. Hosseini, Improving physical and biological properties of nylon monofilament as suture by Chitosan/Hyaluronic acid, Int. J. Biol. Macromol. 164 (2020) 3394e3402. [45] E. Sharifisamani, F. Mousazadegan, R. Bagherzadeh, M. Latifi, PEG-PLA-PCL based electrospun yarns with curcumin control release property as suture, Polym. Eng. Sci. 60 (7) (2020) 1520e1529. [46] H.Y. Mi, Y. Jiang, X. Jing, E. Enriquez, H. Li, Q. Li, L.S. Turng, Fabrication of triple-layered vascular grafts composed of silk fibers, polyacrylamide hydrogel, and polyurethane nanofibers with biomimetic mechanical properties, Mater. Sci. Eng. C 98 (2019) 241e249. [47] J.C. Worch, A.C. Weems, J. Yu, M.C. Arno, T.R. Wilks, R.T. Huckstepp, R.K. O’Reilly, M.L. Becker, A.P. Dove, Elastomeric polyamide biomaterials with stereochemically tuneable mechanical properties and shape memory, Nat. Commun. 11 (2020) 3250. [48] F. Debbabi, S. Gargoubi, M.A. Hadj Ayed, S.B. Abdessalem, Development and characterization of antibacterial braided polyamide suture coated with chitosan-citric acid biopolymer, J. Biomater. Appl. 32 (3) (2017) 384e398. [49] F. Debbabi, M.A. Ayed, S.B. Abdessalem, Development of Antibacterial Braided PET Suture Using ChitosandLaurus nobilis Essential Oil Emulsion, in: EuroMediterranean Conference for Environmental Integration, Springer, Cham, 2017 Nov 20, pp. 1485e1487. [50] S.R. Varma, M. Jaber, S.A. Fanas, V. Desai, A.M. Al Razouk, S. Nasser, Effect of hyaluronic acid in modifying tensile strength of nonabsorbable suture materials: an in vitro study, J. Int. Soc. Prev. Community Dent. 10 (1) (2020) 16. [51] R. Masood, T. Hussain, M. Umar, Azeemullah, T. Areeb, S. Riaz, In situ development and application of natural coatings on non-absorbable sutures to reduce incision site infections, J. Wound Care 26 (3) (2017) 115e120. [52] E. Quaglia, L. Moscufo, M. Corsalini, D. Coscia, P. Sportelli, F. Cantatore, C. De Rinaldis, B. Rapone, M. Carossa, S. Carossa, Polyamide vs silk sutures in the healing of postextraction sockets: a split mouth study, Oral Implant. 11 (2) (2018) 115e120. [53] D.T. Netscher, J.J. Badal, J. Yang, Y. Kaufman, J. Alexander, P. Noble, Biomechanical evaluation of double-strand (looped) and single-strand polyamide multifilament suture: influence of knot and suture size, Hand 10 (3) (2015) 417e424. [54] K.A. Hernandez, R.C. Hooper, T. Boyko, A.R. Golas, M. van Harten, D.Q. Wu, A. Weinstein, C.C. Chu, J.A. Spector, Reduction of suture associated inflammation after 28 days using novel biocompatible pseudoprotein poly (ester amide) biomaterials, J. Biomed. Mater. Res. B Appl. Biomater. 103 (2) (2015) 457e463.

232

Advanced Technologies and Polymer Materials for Surgical Sutures

[55] S. Rethinam, S.T. Parvathaleswara, G. Nandhagobal, T. Alagumuthu, B. Robert, Preparation of absorbable surgical suture: novel approach in biomedical application, J. Drug Deliv. Sci. Technol. 47 (2018) 454e460. [56] A. Dasgupta, N. Sori, S. Petrova, Y. Maghdouri-White, N. Thayer, N. Kemper, S. Polk, D. Leathers, K. Coughenour, J. Dascoli, R. Palikonda, Comprehensive collagen crosslinking comparison of microfluidic wet-extruded microfibers for bioactive surgical suture development, Acta Biomater. 128 (2021) 186e200. [57] K. Long, Y. Liu, W. Li, L. Wang, S. Liu, Y. Wang, Z. Wang, L. Ren, Improving the mechanical properties of collagen-based membranes using silk fibroin for corneal tissue engineering, J. Biomed. Mater. Res. B. 103 (3) (2015 Mar) 1159e1168. [58] X. Jing, H.Y. Mi, H.X. Huang, L.S. Turng, Shape memory thermoplastic polyurethane (TPU)/poly (ε-caprolactone)(PCL) blends as self-knotting sutures, J. Mech. Behav. Biomed. Mater. 64 (2016) 94e103. [59] R.R. Vogels, A. Lambertz, P. Schuster, S. Jockenhoevel, N.D. Bouvy, C. DisselhorstKlug, U.P. Neumann, U. Klinge, C.D. Klink, Biocompatibility and biomechanical analysis of elastic TPU threads as new suture material, J. Biomed. Mater. Res. B Appl. Biomater. 105 (1) (2017) 99e106. [60] Y. Xu, D. Chen, Shape memory-assisted self-healing polyurethane inspired by a suture technique, J. Mater. Sci. 53 (14) (2018) 10582e10592. [61] Y.S. Joo, J.R. Cha, M.S. Gong, Biodegradable shape-memory polymers using polycaprolactone and isosorbide based polyurethane blends, Mater. Sci. Eng. C 91 (2018) 426e435. [62] M.A. Javaid, K.M. Zia, R.A. Khera, S. Jabeen, I. Mumtaz, M.A. Younis, M. Shoaib, I.A. Bhatti, Evaluation of cytotoxicity, hemocompatibility and spectral studies of chitosan assisted polyurethanes prepared with various diisocyanates, Int. J. Biol. Macromol. 129 (2019) 116e126. [63] M. Razumov, O. Gornukhina, O. Golubchikov, I. Vershinina, A. Vashurin, Polypropylene suture material with anti-inflammatory action, Iran. Polym. J. (Engl. Ed.) 27 (9) (2018) 629e634. [64] M.C. da Silva, H.N. da Silva, R.D. Alves Leal Cruz, S.K. Sagoe Amoah, S.M. de Lima Silva, M.V. Lia Fook, N-acetyl-d-glucosamine-loaded chitosan filaments biodegradable and biocompatible for use as absorbable surgical suture materials, Materials 12 (11) (2019) 1807. [65] F.E. Ciraldo, K. Schnepf, W.H. Goldmann, A.R. Boccaccini, Development and characterization of bioactive glass containing composite coatings with ion releasing function for antibiotic-free antibacterial surgical sutures, Materials 12 (3) (2019) 423. [66] Y. Yang, S.B. Yang, Y.G. Wang, S.H. Zhang, Z.F. Yu, T.T. Tang, Bacterial inhibition potential of quaternised chitosan-coated VICRYL absorbable suture: an in vitro and in vivo study, J. Orthop. Translat. 8 (2017) 49e61. [67] D. Gupta, U. Sharma, S. Chauhan, S.A. Sahu, Improved outcomes of scar revision with the use of polydioxanone suture in comparison to polyglactin 910: a randomized controlled trial, J. Plast. Reconstr. Aesthetic Surg. 71 (8) (2018) 1159e1163. [68] Y. Altınel, S.S. Chung, G. Okay, N. Ugras, A.F. Isık, E. Öztürk, H. Özgüç, Effect of chitosan coating on surgical sutures to strengthen the colonic anastomosis, Ulus Travma Acil Cerrahi Derg 24 (5) (2018) 405e411. [69] K. Shao, B. Han, J. Gao, Z. Jiang, W. Liu, W. Liu, Y. Liang, Fabrication and feasibility study of an absorbable diacetyl chitin surgical suture for wound healing, J. Biomed. Mater. Res. B Appl. Biomater. J 104 (1) (2016) 116e125. [70] R. Kandimalla, S. Kalita, B. Choudhury, D. Devi, D. Kalita, K. Kalita, S. Dash, J. Kotoky, Fiber from ramie plant (Boehmeria nivea): a novel suture biomaterial, Mater. Sci. Eng. C 62 (2016) 816e822.

Revisiting the properties of suture materials: an overview

233

[71] S. Rethinam, G. Nallathambi, S. Vijayan, B. Basaran, A. Mert, O. Bayraktar, A new approach for the production of multifilament suture-in vitro and in vivo analysis, Int. J. Polym. Mater. 70 (18) (2021) 1306e1315. [72] X. Peng, G. Liu, L. Zhu, K. Yu, K. Qian, X. Zhan, In vitro and in vivo study of novel antimicrobial gellanepolylysine polyion complex fibers as suture materials, Carbohydr. Res. 496 (2020) 108115. [73] M. Kreszinger, B. Toholj, A. Acanski, S. Balos, M. Cincovic, M. Pecin, M. Lipar, O. Smolec, Tensile strength retention of resorptive suture materials applied in the stomach wall-an in vitro study, Vet. Arh. 88 (2) (2018) 235e243. [74] O.F. Dueñas-Garcia, G.M. Sullivan, K. Leung, K.L. Billiar, M.K. Flynn, Knot integrity using different suture types and different knot-tying techniques for reconstructive pelvic floor procedures, Int. Urogynecol. J. 29 (7) (2018) 979e985. [75] S.E. Naleway, W. Lear, J.J. Kruzic, C.B. Maughan, Mechanical properties of suture materials in general and cutaneous surgery, J. Biomed. Mater. Res. B Appl. Biomater. J. 103 (4) (2015) 735e742. [76] M.D. Roy, S. Ghosh, A. Yadav, S.D. Roy, Effect of coefficient of friction and bending rigidity on handling behaviour of surgical suture, J. Inst. Eng. Electron. 100 (2) (2019) 131e137. [77] V. Segura-Ibarra, J.I. de Elguea-Lizarraga, J. Vazquez-Armendariz, A.L. GarciaGarcia, J.A. Diaz-Elizondo, C.A. Rodríguez, E.F. Villalba, In-vitro biomechanical evaluation of suture materials used for abdominal fascial closure, Mater. Res. Express 8 (2021) 115401. [78] G. Zhang, T. Ren, X. Zeng, E. Van Der Heide, Influence of surgical suture properties on the tribological interactions with artificial skin by a capstan experiment approach, Friction 5 (1) (2017) 87e98. [79] P.J. Regier, D.D. Smeak, K.C. McGilvray, Ex vivo comparison of intradermal closures with conventional monofilament suture vs unidirectional barbed suture in dogs, Vet. Surg. 48 (8) (2019) 1399e1405. [80] S. Lapointe, F. Zhim, L. Sidéris, P. Drolet, S. Célestin-Noël, P. Dubé, Effect of chemotherapy and heat on biomechanical properties of absorbable sutures, J. Surg. Res. 200 (1) (2016) 59e65. [81] D.A. Müller, J.G. Snedeker, D.C. Meyer, Two-month longitudinal study of mechanical properties of absorbable sutures used in orthopedic surgery, J. Orthop. Surg. Res. 11 (1) (2016) 1e7. [82] P.C. Johnson, A.D. Roberts, J.M. Hire, T.L. Mueller, The effect of instrumentation on suture tensile strength and knot pullout strength of common suture materials, J. Surg. Educ. 73 (1) (2016) 162e165. [83] E. Silver, R. Wu, J. Grady, L. Song, Knot security-how is it affected by suture technique, material, size, and number of throws? J. Oral Maxillofac. Surg. 74 (7) (2016) 1304e1312. [84] M.A. Alshehri, J.K. Baskaradoss, A. Geevarghese, R. Ramakrishnaiah, D.N. Tatakis, Effects of myrrh on the strength of suture materials: an in vitro study, Dent. Mater. J. (2015) 2013e2317. [85] A. González-Barnadas, O. Camps-Font, D. Espanya-Grifoll, A. España-Tost, R. Figueiredo, E. Valmaseda-Castellón, In vitro tensile strength study on suturing technique and material, J. Oral Implantol. 43 (3) (2017) 169e174. [86] D. Abellán, J. Nart, A. Pascual, R.E. Cohen, J.D. Sanz-Moliner, Physical and mechanical evaluation of five suture materials on three knot configurations: an in vitro study, Polymers 8 (4) (2016) 147. [87] P.P. Caimmi, M. Sabbatini, L. Fusaro, A. Borrone, M. Cannas, A study of the mechanical properties of ePTFE suture used as artificial mitral chordae, J. Card. Surg. 31 (8) (2016) 498e502.

234

Advanced Technologies and Polymer Materials for Surgical Sutures

[88] M.H. Baums, C. Sachs, T. Kostuj, K. Schmidt-Horlohé, W. Schultz, H.M. Klinger, Mechanical testing of different knot types using high-performance suture material, Knee Surg. Sports Traumatol. Arthrosc. 23 (5) (2015) 1351e1358. [89] J.F. Williams, S.S. Patel, D.K. Baker, J.M. Schwertz, G. McGwin, B.A. Ponce, Abrasiveness of high-strength sutures used in rotator cuff surgery: are they all the same? J. Shoulder Elbow Surg. 25 (1) (2016) 142e148. [90] V.A. Lipatov, D.A. Severinov, A.A. Denisov, S.V. Lazarenko, N.N. Grigor’yev, Research of physical and mechanical characteristics of suture material in experiment in operations on liver, IP Pavlov Russian Med. Biol. Herald 28 (2) (2020) 193e199. [91] A. Dixit, P. Nadkarni, V. Shah, B. Patel, P.K. Turiya, A. Thakkar, Evaluation of safety and efficacy of polyglactin 910 suture in surgical incision closure: clinical study protocol for a randomized controlled trial, Int. J. Clin. Trials 5 (1) (2018) 80. [92] S. Sala-Pérez, M. López-Ramírez, M. Quinteros-Borgarello, E. Valmaseda-Castellón, C. Gay-Escoda, Antibacterial suture vs silk for the surgical removal of impacted lower third molars. A randomized clinical study, Med. Oral Patol. Oral Cirugía Bucal 21 (1) (2016) e95. [93] R. Asher, T. Chacartchi, M. Tandlich, L. Shapira, D. Polak, Microbial accumulation on different suture materials following oral surgery: a randomized controlled study, Clin. Oral Invest. 23 (2) (2019) 559e565. [94] M.G. Tuuli, M.J. Stout, S. Martin, R.M. Rampersad, A.G. Cahill, G.A. Macones, Comparison of suture materials for subcuticular skin closure at cesarean delivery, Am. J. Obstet. Gynecol. 215 (4) (2016) 490-e1. [95] J. Dhom, D.A. Bloes, A. Peschel, U.K. Hofmann, Bacterial adhesion to suture material in a contaminated wound model: comparison of monofilament, braided, and barbed sutures, J. Orthop. Res. 35 (4) (2017) 925e933. [96] P.F. Wang, D.C. Pascasio, S.H. Kwon, S.H. Chen, P.Y. Chou, C.F. Yao, Y.A. Chen, C.H. Lin, Y.R. Chen, The effect of absorbable and non-absorbable sutures on nasal width following cinch sutures in orthognathic surgery, Symmetry 13 (8) (2021) 1495. [97] M.R. Morris, C. Bergum, N. Jackson, D.C. Markel, Decreased bacterial adherence, biofilm formation, and tissue reactivity of barbed monofilament suture in an in vivo contaminated wound model, J. Arthroplasty 32 (4) (2017) 1272e1279. [98] C.T. Kailas, A randomised control trial comparison the efficacy between delayedabsorbable polydioxanone and non-absorbable suture material in abdominal wound closure, Int. Surg. J. 4 (1) (2016) 153e156. [99] B. Alinasab, P.O. Haraldsson, Rapid resorbable sutures are a favourable alternative to non-resorbable sutures in closing transcolumellar incision in rhinoplasty, Aesthetic Surg. J. 40 (4) (2016 Aug) 449e452. [100] M. Dragovic, M. Pejovic, J. Stepic, S. Colic, B. Dozic, S. Dragovic, M. Lazarevic, N. Nikolic, J. Milasin, B. Milicic, Comparison of four different suture materials in respect to oral wound healing, microbial colonization, tissue reaction and clinical featuresdrandomized clinical study, Clin. Oral Invest. 24 (4) (2020 Apr) 1527e1541. [101] C.L. Durand, Comparison of knotless barbed suture versus monofilament suture in the oral cavity of cats, J. Vet. Dent. 34 (3) (2017) 148e154. [102] F. Morelli, A. Anderson, A. McLister, J.J. Fearon, J. Davis, Electrochemically driven reagent release from an electronic suture, Electrochem. Commun. 81 (2017) 70e73. [103] D. Park, I.S. Yong, K.J. Cho, J. Cheng, Y. Jung, S.H. Kim, S.H. Lee, The use of microfluidic spinning fiber as an ophthalmology suture showing the good anastomotic strength control, Sci. Rep. 7 (1) (2017), 1-1. [104] C. Serrano, L. García-Fernández, J.P. Fernández-Blázquez, M. Barbeck, S. Ghanaati, R. Unger, J. Kirkpatrick, E. Arzt, L. Funk, P. Turón, A. del Campo, Nanostructured medical sutures with antibacterial properties, Biomaterials 52 (2015) 291e300.

Revisiting the properties of suture materials: an overview

235

[105] A. Alirezaie Alavijeh, M. Dadpey, M. Barati, A. Molamirzaie, Silk suture reinforced with Cefixime nanoparticles using polymer hydrogel (CFX@ PVA); preparation, bacterial resistance and mechanical properties, Nano Res. J. 3 (3) (2018) 133e139. [106] D.Q. Wu, H.C. Cui, J. Zhu, X.H. Qin, T. Xie, Novel amino acid based nanogel conjugated suture for antibacterial application, J. Mater. Chem. B 4 (15) (2016) 2606e2613. [107] A.L. Gallo, F. Paladini, A. Romano, T. Verri, A. Quattrini, A. Sannino, M. Pollini, Efficacy of silver coated surgical sutures on bacterial contamination, cellular response and wound healing, Mater. Sci. Eng. C 69 (2016) 884e893. [108] D.M. Syukri, O.F. Nwabor, S. Singh, J.C. Ontong, S. Wunnoo, S. Paosen, S. Munah, S.P. Voravuthikunchai, Antibacterial-coated silk surgical sutures by ex situ deposition of silver nanoparticles synthesized with Eucalyptus camaldulensis eradicates infections, J. Microbiol. Methods 174 (2020) 105955. [109] S. Anjum, A. Gupta, D. Sharma, S. Kumari, P. Sahariah, J. Bora, S. Bhan, B. Gupta, Antimicrobial nature and healing behavior of plasma functionalized polyester sutures, J. Bioact. Compat. Polym. 32 (3) (2017) 263e279. [110] A.L. Gallo, M. Pollini, F. Paladini, A combined approach for the development of novel sutures with antibacterial and regenerative properties: the role of silver and silk sericin functionalization, J. Mater. Sci. Mater. Med. 29 (8) (2018) 1e3. [111] K.V. Malafeev, O.A. Moskalyuk, V.E. Yudin, P. Morganti, E.M. Ivan’Kova, E.N. Popova, Y.U. Elokhovskii, Biodegradable polylactide/chitin composite fibers: processing, structure, and mechanical properties, J. Appl. Cosmetol. 35 (2017) 163e173. [112] W. Zhang, B. Yin, Y. Xin, L. Li, G. Ye, J. Wang, J. Shen, X. Cui, Q. Yang, Preparation, mechanical properties, and biocompatibility of graphene oxidereinforced chitin monofilament absorbable surgical sutures, Mar. Drugs 17 (4) (2019) 210. [113] H. Wu, G.R. Williams, J. Wu, J. Wu, S. Niu, H. Li, H. Wang, L. Zhu, Regenerated chitin fibers reinforced with bacterial cellulose nanocrystals as suture biomaterials, Carbohydr. Polym. 180 (2018) 304e313. [114] R. Duarah, Y.P. Singh, P. Gupta, B.B. Mandal, N. Karak, Smart self-tightening surgical suture from a tough bio-based hyperbranched polyurethane/reduced carbon dot nanocomposite, Biomed. Mater. 13 (4) (2018) 045004. [115] S. Liu, G. Wu, X. Chen, X. Zhang, J. Yu, M. Liu, Y. Zhang, P. Wang, Degradation behavior in vitro of carbon nanotubes (CNTs)/poly (lactic acid)(PLA) composite suture, Polymers 11 (6) (2019) 1015.zz.

CHAPTER 10

Suture materials, emerging trends Hemand Aravind Aromatic and Medicinal Plant Research Station, Kerala Agricultural University, Thrissur, Kerala, India

10.1 Introduction The word suture is derived from the Latin word sutura, “a sewn seam” [1]. In Ancient times, materials including linen, cotton, horsehair, animal tendons and intestines, and wire from precious metals have been used to hookup wounds and body openings. Several adaptations over time have led to the highly sophisticated products doctors use in their current practice. A surgical suture is one of the important medical devices used by doctors during surgical procedures. Generally, surgical sutures are used to facilitate closure and healing of surgical or trauma-induced wounds by perpetuating tissues together to facilitate the healing process. Majority of people confused the word suture with stitch. A suture is a medical device used by the doctor to repair the wound, while stitching is simply the technique the doctor uses to close the wound [2]. Sutures are mainly classified as absorbable and nonabsorbable. There has been progress in the classes of suture materials based on their properties and capabilities to improve tissue approximation and wound closure. In this chapter, we discuss the emerging trends in suture technology with a special mention of barbed knotless sutures, antimicrobial sutures, bio-active sutures such as drug-eluting and stem cells seeded sutures, and smart sutures including elastic, and electronic sutures. These impressive models expand the resourcefulness of sutures from being used just as a physical thing for suturing or to a more biologically active component enabling the delivery of drugs and cells to the target site with immense application potential in both therapeutics and diagnostics [3].

10.2 Taxonomy of sutures Sutures are generally classified as absorbable and nonabsorbable sutures based on their material nature. Among them absorbable sutures are peculiar Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00005-X

© 2023 Elsevier Ltd. All rights reserved.

237

238

Advanced Technologies and Polymer Materials for Surgical Sutures

in their action, they will break down harmlessly in the body by the action of hydrolytic enzymes in the tissues. Hence doctors do not need to remove them from the wound after surgery. However, the nonabsorbable sutures are usually removed after a few days of the surgery. Sometimes doctors permanently leave the sutures in the body of the patient, by considering the type of surgery that the patient has undergone. Sutures can be monofilament or multifilament or braided. A monofilament suture encompasses a single thread that allows the suture to pass through the tissues easily and they have minimal tissue reaction and dragging [4]. On the other hand, a multifilament suture consists of many small threads which are braided together and because of its typical structural arrangement a greater chance of infections may be predicted. Sutures are also categorized into synthetic or natural material depending on the origin of the material. A detailed description of suture materials helps to know them better.

10.3 Absorbable and nonabsorbable suture materials One of the major challenges in internal surgery is it requires reopening for the removal of suture materials used inside the body. An externally applied suture material can be easily removed by the doctors without reopening the wound. Such a ruinous situation demands absorbable sutures to be often used internally and nonabsorbable sutures externally. Some of the major surgeries like heart and urinary bladder repair may require sutures with highly specialized and durable materials to perform their role well; usually, such sutures have some unique material properties and are often nonabsorbable to minimize the risk of degradation.

10.4 Monofilament, multifilament sutures and barbed sutures brands The material structure of sutures determines its category, that is, monofilament sutures, multifilament or braided sutures and knotless monofilament sutures or barbed sutures. Monofilament sutures show lower tissue reaction compared to braided sutures [5]. However braided sutures exhibit greater mechanical properties than nonbraided sutures.

10.5 Categories of absorbable sutures 10.5.1 Catgut sutures In some countries, absorbable sutures are also made of gut material obtained from cow or sheep intestines and is termed as catgut sutures. Two varieties

Suture materials, emerging trends

239

of catgut are plain and chromic [6]. In the 1860s, the physician Joseph Lister invented a technique for sterilizing catgut and it was perfected finally in 1906 [7]. The naturally derived catgut suture with monofilament structure and absorbable nature provides them good tensile strength to hold tissues together. A highlight of catgut suture is its smooth and flexible nature with good knotting ability, and can completely vanish between 60 and 120 days. This eventual disintegration process makes it a good choice for surgeons to use in tissues healing rapidly. 10.5.2 Chromic gut sutures They are developed from animal intestinal submucosa, particularly from sheep intestine and bovine intestinal serosa and are classically treated with formaldehyde. Systematic treatment of these sutures with chromic salt helps them to last longer in the body than the plain kind. This also helps to increase strength and decreases tissue reaction. In the body, it is absorbed by phagocytosis and can be rapidly absorbable. Usually, it loses 33% tensile strength in 7 days and 67% in 28 days. The impetuous loss of strength is found in infected wounds which have enormous digestive enzyme secretion (e.g., stomach), in highly vascular tissues, and patients with poor protein content. Compared with synthetic materials chromic gut sutures are economical and have relatively good handling characteristics, good knot security and stiffness. It is not recommended to use them in the closure of the abdominal wall because of rapid, unpredictable loss of strength. 10.5.3 Polyglycolic acid sutures Polyglycolic acid sutures are made up of a glycolic acid polymer material. It is rapidly absorbable by hydrolysis; most of its tensile strength is lost within 2 weeks of implantation and can be removed from the body within 3e4 months. It is widely used in urinary bladder procedures. Compare to chromic gut, polyglycolic sutures are stronger, have more predictable strength loss, are less reactive and can be removed completely from the body. 10.5.4 Polydioxanone sutures This type of synthetic monofilament suture is widely used to handle various kinds of soft-tissue cuts and for some abdominal closures. In pediatric and cardiac procedures, polydioxanone sutures are commonly used by surgeons. 10.5.5 Poliglecaprone sutures The Poliglecaprone suture is a synthetic monofilament suture having a combination of a copolymer of glycolide and epsilon-caprolactone,

240

Advanced Technologies and Polymer Materials for Surgical Sutures

generally used to repair soft tissues. It’s negligible reaction on the tissue surface and soft material structure helps it to be used for subcuticular dermis closures on a patient’s face and as a ligature. Overall, the scar-free, esthetic healing ability of poliglecaprone sutures makes them highly preferred for plastic surgery procedures. In the case of vascular anastomosis procedures, this suture material is a better choice to connect blood vessels. 10.5.6 Polyglactin sutures Polyglactin suture is a member of braid type suture emerged by the combination of two copolymers of glycolide and epsilon-caprolactone, in general, soft tissue approximations and in specific lacerations, on the face and hands, it is the. For vascular anastomosis procedures, poliglecaprone suture and polyglactin suture are the most preferred options. Instead of using a catgut suture, polyglactin sutures are a good alternative for soft tissue surgeries because the absorption level of this suture is more predictable and also exhibits little tissue reactions.

10.6 Slowly absorbable sutures 10.6.1 Polydioxanone (PDS II) Polydioxanone sutures are composed of paradioxanone polymers, which are absorbed by hydrolysis and retains 50% strength after 4 weeks of the procedure. It takes up to 6 months for complete removal from the body. The suture maintains its integrity in alkaline environment makes and hence is suitable for urinary bladder surgery. 10.6.2 Polyglyconate (Maxon) The best composition of glycolic acid and trimethylene carbonate enables polyglycolide to be a good slowly absorbable suture. About 50% of its tensile strength is lost only within 3 weeks and for complete absorption, 6e7 months are required. Overall strength is similar to polydioxanone but is more economical than PDS. The handling characteristics limits its use in small animals but could be selected for soft tissues. 10.6.3 Nonabsorbable sutures Nonabsorbable sutures are mainly silk or synthetic polyesters, polypropylene or nylon. Nonabsorbable sutures may or may not comprise coatings that augment their performance characteristics and are

Suture materials, emerging trends

241

characteristically used to close skin wounds. It can be removed after a couple of weeks. Classically these sutures are used in highly sensitive and complicated surgeries like vascular anastomosis procedures of the heart (due to the constant movement and pressure on the heart). Because of its much lower level of immune response, nonabsorbable sutures usually cause less scarring, which is why they are also used in surgical procedures where the cosmetic outcome is momentous. Depending on the intensity of the wound, these sutures may be left in the body permanently or removed after a specific period. 10.6.4 Silk suture They are derived from silkworm cocoons and technically come under the nonabsorbable category. In a slow manner they lose tensile strength over 6 months and may be completely absorbed after 2 years. Its has excellent handling properties with good knot security. It is an economical multifilament suture but is not used in infected or contaminated areas. 10.6.5 Polymerized caprolactum suture (Supramid) Supramid is composed of synthetic twisted multifilament polyamide wrapped in a smooth sheath, which makes it too strong and rather stiff to handle. When used on skin, it is highly reactive and should not be buried because of the high rate of draining tract formation. Comparatively, it is economical and primarily used for skin closure. 10.6.6 Polyester suture (Mersilene, Ethibond) Polyester sutures are synthetic multifilament sutures obtained from ethylene glycol and terephthalic acid. They are available as both coated and uncoated ones. It has high tensile strength and retains them to a larger extent when implanted. The uncoated sutures generally show slight tissue reaction. The Mersilene polyester sutures are generally used for soft tissue approximation, cardiovascular, ophthalmic, and neurological procedures. 10.6.7 Nylon (Dermalon or Ethilon) The Nylon sutures are generally available as both monofilament and multifilament forms, and is composed of polyamine. It is an inert suture with minimal tissue reaction and relatively poor knot security, so care is necessary to ensure secure knots.

242

Advanced Technologies and Polymer Materials for Surgical Sutures

10.6.8 Polybutester (Novafil) Polybutester is a copolymer composed of polyglycolterephthate and polybutylene terephthate. Novafil sutures are somewhat slippery to manage and possess high flexibility and elasticity than polypropylene or nylon. It also exhibits good handling characteristics and knot security. Since it is monofilament, lesser tissue reaction is only elicited. 10.6.9 Polypropylene (Prolene) Polypropylene suture is a nonabsorbable suture having excellent tensile strength. It possesses minimal tissue reaction and can be used as a permanent implant suture for prosthetic devices. 10.6.10 Structurally coated and un-coated sutures To improve the quality and performance of sutures, scientists introduced specially coated materials for suture manufacturing which enhance properties like knotting, easy passage through tissue and reduce tissue reaction. This upgraded technology is applied to braided sutures rather than monofilament sutures and mostly the coating materials are chromium salt, silicon, wax, PTFE, polycaprolactone, calcium stearate etc. Compared to conventional coating materials like chromium salts, beeswax, paraffin and gelatine, polymeric coating materials are known to be more biocompatible. Monofilament and multifilament sutures with antibacterial and antimicrobial properties have high application value in the suture industry. Stem cell coated sutures are its advanced version and they exhibit improved healing properties on surgical sites. To minimize surgical site infection and to increase healing ability antimicrobial coatings like chlorhexidine, triclosan, and silver ions may be given to any suture as a surface coating in addition to the regular coating materials. Some examples of coated and uncoated sutures are given below. 10.6.10.1 Coated sutures include PGA sutures, Catgut Chromic, Polyglactin 910, silk and polyester sutures, braided or twisted nylon, Poliglecaprone and Polydioxanone sutures. 10.6.10.2 Un-coated sutures include Polypropylene sutures, Nylon, PVDF, stainless steel, and PTFE sutures.

Suture materials, emerging trends

243

10.6.11 Application-based suture categories Sutures are organized into various categories based on their common application and use, in a way we can group them into general sutures, valve sutures, cardiovascular sutures, dental sutures, ophthalmic sutures, orthopedic sutures, gynaec, cosmetic surgery sutures, veterinary sutures, etc.

10.7 New trends in sutures 10.7.1 Knotless barbed sutures Knotless sutures are comparatively a new trend-setting suture that has been widely used in both skin and deeper structures. In its monofilament structure the barbs are oriented in the opposite direction to the needle. Surgical knots usually reduce the tensile strength of all sutures by thinning and stretching the material. On the other hand, the novel barbs on the ligatures make the suture grab the tissue, without allowing the suture to slide back and also reduce the operation time. Recently knotless barbed sutures are widely used in orthopedics surgeries, particularly in total knee arthroplasty and hip surgery [8]. 10.7.2 Antibacterial sutures Usually after surgery, the wounds can be contaminated leading to delayed wound healing. All types of sutures either absorbable or nonabsorbable, synthetic or natural represent an artificial implant; a 90 cm length material with a total surface area of 130 cm [9]. It is highly evident that the presence of a suture increases the frequency of a site-specific infection [10]. Sutures which is coated or impregnated with a wide spectrum of biocide have a high antimicrobial potential against biofilm production around the suture site and can successfully prevent infection rate. The advantage of using an antiseptic as opposed to an antibiotic for example triclosan has valuable benefits as reported earlier [11]. For instance, polymerized cyclodextrin (pCD) coated surgical suture interestingly introduces both antiinflammatory and antimicrobial properties throughout the phases of acute and chronic healing [12]. Silk sutures incorporated with 4-Hexylresorcinol (4HR) is a suitable example of the antimicrobial suture. It exhibits antimicrobial activity and has better degradation properties as shown in Fig. 10.1. The silk fibroin protein is attached with 4HR a

244

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 10.1 Antimicrobial suture with an attached 4HR suture degradation mechanism [13]. Diagrammatic representation of antimicrobial silk suture with attached 4HRmicelles. Firmly bound 4HR(4-hexylresorcinol) in a hydrophobic domain is exposed during the wound remodeling phase by macrophages. The released 4HR induces MMPs (Matrix metalloproteinases) in macrophages and produces MMPs that can digest silk suture material. (Reproduced with permission from Springer Nature.)

resorcinolic lipid and has been used as an antiseptic and food ingredient for preventing melanosis [13]. The positive side of local delivery of drugs by sutures at the wound is that it can improve the healing rate and reduce related systemic side effects and drug resistance. 10.7.3 Stem cell seeded suture Another promising development in suture manufacturing is the development of stem cell-delivering sutures with an ability to release cells on the sites. It was shown that suture-based cell engraftment in rat heart muscles is more stable than intramuscular injection [14]. In another work, the authors described that MSC (mesenchymal stem cells)-coated sutures can improve wound healing and tissue remodeling in vivo [15]. A major limitation of these studies is that most cells are delivered at the initial insertion site and the systemic mechanical stress induced by the suturing itself may seriously diminish the viability of the cells. Some recent also studies check

Suture materials, emerging trends

245

the feasibility of using ASC (adipose-derived mesenchymal stem cells)-filled sutures as a novel therapeutic methodology for wound healing. However, the basic mechanisms of action of such cell-coated sutures have not been yet been described [16]. 10.7.4 Smart sutures: electronic/elastic sutures Smart sutures are an advanced form of sutures which have sensors attached on them that could monitor wounds and speed up healing. Structurally they are devices with silicon membranes and gold electrodes and wires that are just a few hundred nanometres thick. Electronic sutures can precisely measure temperature at the wound site and in case of an infection-related temperature rise, they can deliver heat to a wound site along with electrical stimulation, which is known to aid healing. Some other electronic sutures can release drugs from them in a programmed way [17]. The electronic sutures, with a special structural design of ultrathin silicon sensors integrated on polymer or silk strips, can be legitimately threaded through needles, and in animal tests, the scientists were able to lace them through the skin, pull them tight, and knot them without disturbing the devices [18]. The studies related to next generation of smart sutures still on the pipeline, rely on silicon-based devices that exhibit significant flex and stretchability. Drugreleasing electronic suture system (DRESS) provides a wider outlook on multifunctional sutures as shown in Fig. 10.2 and agrees with an

Figure 10.2 A multifunctional electronic suture for continuous strain monitoring and on-demand drug release [17]. (Reproduced with permission from the Royal Society of Chemistry.)

246

Advanced Technologies and Polymer Materials for Surgical Sutures

opportunity for additional therapeutic and diagnostic options in clinical applications.

10.8 Conclusion Advancement of suture technology improves surgical procedures and reduce their comorbidities. A variety of natural and synthetic sutures, absorbable and nonabsorbable are available commercially. In the midst, biodegradable sutures have received incomparable attention because they open up a large slot of scientific opportunities in targeted drug delivery mechanisms. In the current scenario especially, coated sutures are an effective approach for the delivery of antibacterial agents or antiinflammatory drugs to the surgical site. A lot of promising developments in the direction of suture material selection have come up for improvisation of tissue approximation and wound closure. The current and new trends in suture technology comprising knotless barbed sutures, antimicrobial sutures, drug releasing sutures and various stem cells seeded sutures, and smart sutures including elastic, and electronic sutures have been well discussed in this chapter. In this context, more randomized clinical trials are needed to better elucidate its full potential.

References [1] B. Miriam, A. Al, The surgical suture, Aesthetic Surg. J. 39 (2) (2019) 67e72. [2] https://www.merillife.com/blog/medtech/>types-of-surgical-sutures-and-their-uses. [3] D. Christopher, S. Swaminathan, N. Sunita, M. Loganathan, Y.M. Yosry, M. Geetha, Suture materials - current and emerging trends, J. Biomed. Mater. Res. 104 (6) (2016) 1544e1559. [4] https://wcvm.usask.ca/vsac205/Lab4/>suture-materials.php#Rapidlyabsorbable. [5] A.J. Dart, C.M. Dart, Suture material: conventional and stimuli responsive, Comprehensive Biomater. II 7 (2017) 746e771. [6] L. Crampton, Stitches or sutures: wound closure today and in history, A-Stitch-inTime-Medical-and-Surgical-Sutures-Today-and-in-History 6 (3) (2021). https:// youmemindbody.com/health-care-industry/. [7] R. Andrew, History of Sutures vol 4, 2016. https://www.flushinghospital.org/ newsletter/history-of-sutures/. [8] L. Yifei, L. Sike, H. Jin, D. Liang, The efficacy and safety of knotless barbed sutures in the surgical field: a systematic review and meta-analysis of randomized controlled trials, Sci. Rep. 6 (2016) 23425. [9] Health Protection Agency, Surveillance of Surgical Site Infections in NHS Hospitals in England, London: HPA, 2012. [10] S.D. Elek, P.E. Conen, The virulence of Staphylococcus pyogenes for man. A study of the problems of wound infection, Br. J. Exp. Pathol. 38 (1957) 573e586. [11] D. Leaper, P. Wilson, O. Assadian, C. Edmiston, M. Kiernan, A. Miller, G. BondSmith, J. Yap, 99 (6) (2017) 439e443, https://www.ncbi.nlm.nih.gov/pmc/articles/ PMC5696981/.

Suture materials, emerging trends

247

[12] R.M. Haley, V.R. Qian, G.D. Learn, H.A. Von Recum, Use of affinity allows antiinflammatory and anti-microbial dual release that matches suture wound resolution, J. Biomed. Mater. Res. 107 (7) (2019) 1434e1442. [13] Y.Y. Jo, H. Kweon, D.W. Kim, Accelerated biodegradation of silk sutures through matrix metalloproteinase activation by incorporating 4-hexylresorcinol, Sci. Rep. 2 (13) (2017) 42441. [14] J.P. Guyette, M. Fakharzadeh, E.J. Burford, Z.W. Tao, G.D. Pins, A novel suturebased method for efficient transplantation of stem cells, J. Biomed. Mater. Res. 101 (2012) 809e818. [15] H.T. Georgiev, A.M. Garcia, G.I. Garcia, M.A. Garcia-Cabezas, J. Trebol, Sutures enriched with adipose-derived stem cells decrease the local acute inflammation after tracheal anastomosis in a murine model, Eur. J. Cardio. Thorac. Surg. 42 (2012) 40e47. [16] K.R. Ann, M.K. Bianca, A.W. Elizabeth, L.S. Thilo, R. Farid, H. Yves, F. Peter, M. Hans-Günther, T.E. José, Surgical sutures filled with adipose-derived stem cells promote wound, Healing 3 (13) (2014). [17] L. Yeontaek, K. Hwajoong , K. Yeonju, N. Seungbeom, C. Beomsoo, K. Jinho, P.Charnmin, C. Minyoung, P.Kijun, L.Jaehong, S. Jungmok, A multifunctional electronic suture for continuous strain monitoring and on-demand drug release, Nanoscale. https://pubs.rsc.org/en/content/articlelanding/2021/nr/d1nr04508c#! divCitation/. [18] P. Prachi, Smart sutures that detect infections, MIT Technol. Rev. 8 (24) (2012). https://www.technologyreview.com/2012/08/24/184113/smart-sutures-that-detectinfections/.

Further reading [1] M. D Paul, Bidirectional barbed sutures for wound closure: evolution and applications, J. Am. Col. Certif Wound Spec. 1 (2009) 51e57. [2] J. Charnley, N. Eftekhar, Postoperative infection in total prosthetic replacement arthroplasty of the hip-joint, Br. J. Surg. 56 (1969) 641e649.

CHAPTER 11

Biocompatibility and cytotoxicity of polymer sutures Smitha Vijayan1, Teena Merlin1 and M.S. Jisha2 1

School of Bioscience, Mar Athanasios College for Advanced Studies, Tiruvalla, Kerala, India; 2School of Bioscience, Mahatma Gandhi University, Kottayam, Kerala, India

11.1 Introduction Suture materials are the most frequently used wound closure materials. They play a major role in fixation of injury. Sutures were familiarized about 4000 years ago. At that time the forefathers of medicine used natural fibers like linen, silk, horse tail, etc. for wound closure [1]. Selection of suture material is the vital step, and it depends on the physical and natural properties of the material, evaluation of the injury, rate of healing of various tissues, and health status of patient. The current advancement in the improvisation of suture materials increased the types of available suture materials in the industry [2]. At the same time, staples and tissue adhesives are also considered as an effective substitute to sutures for wound closure. The main role of suture in a wound is to initiate the healing process without infection and other complications. Now a days the main challenge faced by a surgeon is the selection of an ideal suture material from the enormous number of available materials. To enhance wound healing and retardation of scar formation, it is very important to select the suture wisely. Polymeric materials are versatile and have wide range of applications in the medical field especially in wound closure [3]. Ideal sutures should possess some specific characteristics such as good tensile strength, easy handling, ability to form knots, the capacity to stretch in order to lodge wound edema and to shrink to achieve the original size with wound retraction capacity, cost-effective, visible, easily sterilizable, biocompatibility, nontoxic, and their size [4].It is very difficult to find all these characters in one suture material [5]. Recent research activities enormously helped in the development of better suture materials [6]. Advanced surgical sutures made up of polymeric substances are available in the industry and nowadays artificial sutures are replaced with natural sutures and that enhanced the healing process [3]. Advanced technologies are now available to increase the Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00009-7

© 2023 Elsevier Ltd. All rights reserved.

249

250

Advanced Technologies and Polymer Materials for Surgical Sutures

therapeutic value of polymer sutures, even though, the cytotoxicity and biocompatibility are considered to be the major concerns in the field [7]. The latest and more advanced type of natural sutures are considered to be made up of natural carbohydrates. This article discusses about different types of sutures and its biocompatibility and cytotoxicity aspects.

11.2 Classification of sutures Sutures can be classified on the basis of their origin, type of material, size, and physical configuration. Fig. 11.1 clearly depicts the classification of sutures based on different criteria [8]. 11.2.1 Origin based classification Based on the source of material from which the suture material is derived, we can classify them into natural and synthetic. Suture materials derived directly from animal or plant sources are named as natural sutures. Ancient Egyptians used plant fibers, hair, tendons, and wool threads for wound closure. Samhita, written by Indian surgeon Susrutha explains the usage of natural sutures like bow strings (derived from the sheep upper small

Figure 11.1 Classification of surgical sutures.

Biocompatibility and cytotoxicity of polymer sutures

251

intestine) for wound closure in rhinoplasty, tonsillectomy, amputation, and anal fistulae repair [9]. Natural sutures differ from synthetic sutures in their degradation. Natural sutures are proteolyzed and biodegradable in nature, whereas synthetic sutures degraded by hydrolysis [10]. Proteolysis of natural sutures cause adverse inflammatory response at the site of suture but the synthetic sutures doesn’t cause much inflammation. Silk fibers are widely used natural sutures because of their attractive properties such as easy handling and knot security. Natural sutures are highly susceptible to infection because of their animal and plant origin. Most of the natural sutures require prior treatment to increase their properties. For instance, catgut sutures are treated with aldehyde solution and chromium trioxide to strengthen the material [11]. Advanced synthetic sutures with upgraded features and low immunogenic reactions are widely accepted. Postlethwait et al. [12], studied the reaction of human tissues to various suture materials and where nylon sutures exhibited minimum tissue reactions. At the present time, absorbable synthetic sutures with reduced tissue reactions are developed by scientists and are absorbed by hydrolysis [13]. 11.2.2 Material based classification As per the definitions by US Pharmacopeia; Absorbable sutures are materials which lost their complete or partial tensile strength in a period of 1e3 months. Nonabsorbable sutures retain most of their tensile strength even after 3 months. Absorbable suture materials may be synthetic or natural in nature. First absorbable natural suture was catgut, which undergoes enzymatic degradation and phagocytosis after 7e10 days of implantation in vivo. However, polyglycolic acid considered as the foremost absorbable synthetic suture, which is absorbed within 100 days of implantation in vivo by hydrolysis [3]. PDS II (Polydioxanone) is popular absorbable synthetic suture material among plastic surgeons. It is a polymeric material, which gradually absorbed after in vivo implantation. The complete absorption takes about 6 months (25% absorbed in 6 weeks and 70% in 3 months). They are predominantly used in pediatric, cardiovascular, and ophthalmic surgeries due to its minimum tissue reactions. Advanced synthetic absorbable sutures offer expectable degradability in a controlled environment [4]. Nonabsorbable sutures are again classified into natural and synthetic sutures. Natural fibers such as silk, cotton, and linen are examples for

252

Advanced Technologies and Polymer Materials for Surgical Sutures

nonabsorbable sutures. They remain at the site of implantation forever, as long as they are not removed using external devices or forces. Yet considered as nonabsorbable, they gradually undergo proteolytic cleavage and absorbed slowly within 1e3 years. Synthetic nonabsorbable sutures exhibit low immunogenic response and very cheaper and easy to handle [14]. Nonabsorbable synthetic sutures include enormous variety of polymeric materials such as polypropylene, polyamide, polyester, poly (ether ester), polytetrafluoroethylene, polyvinylidene fluoride, and stainless steel [8]. 11.2.3 Classification based on size Sutures are also classified on the basis of size. Two standards are currently employed for the classification; USP (United States Pharmacopeia) and EP (European Pharmacopeia). Commonly employed type of classification is USP standard and where the classification is denoted by a combination of two Arabic numerals. For instance, 2e0 (or 2/0). Finer will be the material, when the first number is higher. In EP standard of classification, the size ranges from 0.1 to 10. USP classification varies according to the type of material but the type of material is not considered in EP standard of classification [15]. 11.2.4 Classification based on physical configuration On the basis of structure, the sutures are classified into multifilament, monofilament, and pseudomonofilament. Twisted and braided multifilament and monofilament sutures are also there based on the structural organization. Rebuilt collagen fibers, catgut, and cotton are considered as twisted multifilament sutures and Dexon, vicryl monson, polysorb silk, polyamide, and polyester-based beads are examples of braided multifilament configuration [16]. Monofilament sutures consist of only one filament and multifilament sutures are made up of several strands. Fig. 11.2 depicts the structure of monofilament and multifilament sutures [14]. Fig. 11.3

Figure 11.2 (A) Monofilament thread (nylon); (B) multifilament thread, braided.

Biocompatibility and cytotoxicity of polymer sutures

253

Figure 11.3 Scanning Electron Microscope images of Monofilament (A) and Multifilament sutures (B).

illustrates the electron micrograph of monofilament and multifilament physical configuration of polymer sutures [8]. Compared to the monofilament sutures, multifilament sutures show high mechanical property with substantial elasticity and flexibility [17]. On the basis of their surface organization, they can be classified into smooth and barbed sutures. Barbed sutures are more effective than smooth ones because of their sharp projections on the surface. Multifilament barbed sutures are known as intertwined sutures and are employed widely in closure of deep wounds because of their improved mechanical properties [16]. Smooth sutures require knots to be held in position, and these knots have hostile effects on tissues. Moreover, the knots are difficult to be placed in deeper surgical regions. Barbed sutures replaced these knots and are widely used in complex reconstructive surgeries [18].

11.3 Necessary characteristics of suture materials The ideal suture material should possess some necessary characteristics. The characteristics are explained in the schematic diagram Fig. 11.4. The characteristics are as follows: 1. Tensile strength 2. Elasticity 3. Knot strength and security 4. Biodegradability 5. Tissue reaction 6. Packaging memory 7. Noncapillary transport and antimicrobial [19].

254

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 11.4 Essential characteristics of suture materials.

All these characteristics are interrelated to each other and depends on four main properties such as • Physical and mechanical properties • Handling properties • Biological properties • Biodegradation properties 11.3.1 Physical and mechanical properties Physical and mechanical properties are very crucial for the sutures because they play major role in wound closure activities. These properties directly related to strength, stiffness, viscoelasticity, coefficient of friction, compliance, size, absorption, transport, and so on. Tensile strength is the widely studied mechanical property of the suture material. It measures the capacity of the suture to resist damage. Tensile strength is important, and it determines the ability of the suture to hold the tissue until complete healing and also to withstand the force of tying knots. Whereas viscoelasticity, bending stiffness and compliance are not studied extensively [20,21]. Knot strength and security are also vital properties of sutures and there are certain external and internal factors which influence knot strength and security.

Biocompatibility and cytotoxicity of polymer sutures

255

The internal factors contribute to the knot strength and security are chemical composition of sutures, treatment of sutures before implantation, and friction coefficient of sutures. Extrinsic factors are size of suture, type of knot, number of throws on knot, and force excreted on tying knots [8]. 11.3.2 Handling properties Pliability, packaging memory, knot tie down, knot slippage, knot security, and tissue drag are the main properties to be considered. Unbiased evaluation of handling properties is very difficult. These properties are interrelated to each other and with the other physical and mechanical properties [22]. Pliability is directly proportional to flexibility of the suture and indirectly related to the coefficient of friction. Table 11.1 is provided with some suture materials with desirable and undesirable properties. 11.3.3 Biological properties Sutures are foreign substances, which can initiate an immune response in the human body after implantation. Exploration of biological properties of suture, materials will help us to know in detail about the inflammation cascade caused by the material, degradation, allergic response initiated by body, abrasion, and degradation, etc. Different factors are there which influence the biological properties [24]. Tissue response also depends on these factors (Table 11.2). Selection of appropriate sutures based on the biological properties is important for enhanced wound healing properties, low inflammatory reactions, reduced scar formation etc. [25]. Biocompatibility and cytotoxicity are the major biological properties, which require great attention.

11.4 Biocompatibility of sutures Biocompatibility is a quite common word used in case of biomaterials but still a great deal of clarification is required about the mechanisms that contribute to this property. Biocompatibility is crucial in case of devices or materials which are intended to remain in the body for a long period of time [26]. Donaruma defined [27] biocompatibility as the “ability of a material to perform with an appropriate host response in a specific situation.” Biocompatibility and associated mechanisms vary according to physiological conditions human body. Varied biocompatibility is mainly due to factors such as age, sex, health status, prolonged disease, lifestyle features, and pharmacological factors [28].

256

Advanced Technologies and Polymer Materials for Surgical Sutures

Table 11.1 Examples of sutures with desirable and undesirable properties [23]. Sl No

Suture type

Desirable properties

Undesirable properties

1

Nylon

2

Silk

High tensile strength, low tissue reaction, high elasticity, cost effective Easy to handle Good knot security High flexibility Good for mucosal surfaces.

3

Polyglyconate

4

Polydioxanone (PDS II)

Poor knot security, difficulty in handling, high memory Low tensile strength High friction coefficient High tissue reaction High capillarity High coefficient of friction High knot splitting Stiff and difficult to handle Poor knot security Memory

5

Polyglactin 910

6

Chromic (Chromic Gut)

7

Stain less steel

8

Polypropylene

High tensile strength, low tissue reaction Easy to handle High and extended tensile strength Easy to handle Low tissue reactivity Maintain its integrity during infection Antibacterial Minimal inflammatory reaction Good handling characteristics Strong with high tensile strength Used in rapidly healing tissues like Subcutaneous tissue Variable absorption rate Moderate inflammatory reactions Slow healing Use in infected tissues High tensile strength Biologically inert Good tensile strength Negligible tissue Reactivity Least thrombogenic activity High plasticity

May cut through crumbly tissues

Poor knot security Poor handling characteristics Tendency to fray Fast absorption High capillarity Not useful in highly mobile tissue sites Poor handling characteristics Nonabsorbable suture Scar free or good cosmetic concern Knot less

Biocompatibility and cytotoxicity of polymer sutures

257

Table 11.2 Factors which affect tissue reactions of sutures. Factors

Material of the suture Physical configuration of the suture Picks per inch in braided suture Twist angle Size Type of suture

More tissue reactions

Less tissue reaction

Natural Braided

Synthetic Monofilament

Less Low Thicker Absorbable

More High Thinner Nonabsorbable

Ng et al. [29], studied biocompatibility and mechanical stiffness of FDA approved medical sutures comprised of silk, polytetrafluoroethylene, and polybutester as a function of sterilization/disinfection mode. Cellular response and enzyme histochemistry analysis are the two basics methods to study biocompatibility. Cellular response is the common and widely used method to study biocompatibility of sutures. Saxena et al. [30], studied the biocompatibility aspect of polypropylene sutures by in vitro and in vivo methods. In vitro studies were carried out on fibroblast cell line, and in vivo analysis was on Swiss albino mice. The study revealed the significant compatibility of the suture material. 11.4.1 Measuring biocompatibility Suture materials should be screened extensively before implantation into humans. Biological acceptance of such materials can be assessed by various testes. These tests can be classified as in vitro, in vivo and usage tests [31]. 11.4.1.1 In vitro tests Cell culture or some other biological systems are used for the in vitro analysis. Direct or indirect method can be used for the biocompatibility analysis by in vitro method. In direct analysis the material directly contacts with the cell system without any barriers. Direct test is conducted based on direct physical contact with the cells or in a manner that the extract from the material will be in contact with cell. Subdivision of in vitro test is based on two main concepts i.e., measurement of cytotoxicity or cell growth and quantifying some other metabolic functions. The crucial factor for in vitro analysis is the standardization of methods and cell lines used for the tests. In vitro tests possess the advantages such as low cost, easy performance, experiment can be standardized easily, availability of large-scale screening,

258

Advanced Technologies and Polymer Materials for Surgical Sutures

good experimental control is possible and excellent interaction mechanism. The major disadvantage of the method is their relevance to the in vivo system [31]. 11.4.1.2 In vivo tests This includes animal tests which involves mammals such as mice, rats, hamsters and guinea pigs. In vivo analysis has many disadvantages such as extremely expensive, time-consuming, many ethical and legal issues, difficult to control, harmful to test animals, and also need skilled person to evaluate the results. This method possesses many advantages also. It is more inclusive than in vitro analysis and provides more relevant results. This method helps for more complex interaction with the material and tissue, which provides significant outcomes [24]. Biocompatibility can be evaluated by several in vivo methods such as sensitization, irritation intracutaneou s reactivity, systemic toxicity, subacute and subchronic toxicity, genotoxicity, hemocompatibility, carcinogenicity, reproductive toxicity, biodegradation studies, immunotoxicity, and toxicokinetics [32]. 11.4.1.3 Usage tests The test is conducted in an environment which is identical to the actual situation. The main advantage of this test is that the relevance to use of material is assured. They are very expensive, time-consuming, and difficult to interpret and control. The best method to assess the biocompatibility is to use the combination of in vitro, in vivo, and usage tests because no single method is adequate to characterize the complete biocompatibility characters of materials [33]. Previously the testing strategy was focused only on toxicity. Now the biocompatibility testing strategy focuses on specific toxicity, unspecific toxicity, and clinical trials. Future testing strategies are also developed and which focus on continuous evaluation and assessment of biocompatibility even after introduction to the market. This testing method also helps to evaluate the efficacy and accuracy of existing testing methods. This scheme is based on the aspect that biocompatibility of suture materials should be assessed continuously [31]. 11.4.1.4 Standards that regulate the measurement of biocompatibility Standardization of test method is difficult and lengthy process because development of a uniform test method for every material is a tough task. In

Biocompatibility and cytotoxicity of polymer sutures

259

1976, a bill was passed for biological testing for all medical devices. The committee that accepted the mill also recognized the need for a standardized method for continuous assessment of biocompatibility of medical devices and materials. In 1980s, several organizations initiated an effort to constitute certain international standards for the biomedical devices and materials. In 1992, International Organization for Standardization published ISO 10993 standard for biological evaluation of medical devices. This standard guideline provides perfect guidance to industries and Food and Drug Administration (ISO:10993). Later in 2002, FDA proposed special guidelines for the biological evaluations of medical devices. FDA requires the biocompatibility tests to be conducted according to the Good Manufacturing Practices.

11.5 Cytotoxicity of sutures Measurement of cytotoxicity is also related to biocompatibility. It is considered as one method to establish the compatible nature of the suture material. Cytotoxicity test measures the cell death caused by a specific material or the percentage of viability before and after exposure to the specific material. Membrane permeability test can also be used for the assessment of cytotoxicity. In vitro test which measures biosynthetic or enzymatic activity are also used for the cytotoxicity test. Measurement of DNA synthesis or protein synthesis is some common example of these kinds of test. Familiar method of cytotoxicity measurement such as MTT [3-(4,5dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide] test, as well as the NBT [nitro blue tetrazolium], XTT [ 2, 3-Bis-(2 methoxy-4-nitro-5sulfophenyl)] -2H-tetrazolium-5-carboxanilide salt] and WST (a watersoluble tetrazolium). All are colorimetric assays based on tetrazolium dyes. These tests allow continuous monitoring of cytotoxicity. Vijayan et al. [34], studied the cytotoxicity of chitosan using MTT assay and the results were promising. Javaid et al. [35], studied the cytotoxicity and biocompatibility of chitosan assisted polyurethanes and the results indicated enhanced biocompatibility of polyurethanes as suture material when assisted with chitosan. The study used hemolytic activity, mutagenic, and cytotoxicity to evaluate the biocompatibility of the suture material. Vijayan et al. [34] studied the cytotoxicity of chitosan uing MTT assay and the results were promising (Fig. 11.5). Graphene oxide-reinforced chitin monofilament absorbable surgical sutures were tested for their biocompatibility by in vitro and in vivo

260

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 11.5 Showing dose dependent cytotoxicity of chitosan toward normal L929 fibroblast cell line (A) 25, (B) 50, (C) 100, (D) 125, (E) 150 mg/mL and (F) control.

methods. MTT assay on L929 fibroblast cell lines was the in vitro method used for the study. For the in vivo studies, the subcutaneous implantation of graphene oxide reinforced with chitin sutures was implanted on dorsal skin. The studies were promising and the suture materials showed excellent biocompatibility with negligible tissue reactions [36]. Ho et al. [37], in their studies proved that long-term coatings with antimicrobial silver nanoparticles can increase the efficacy of surgical sutures made up of polyglycolic acid and it was tested for the cytotoxicity on L929 mouse fibroblast cell lines. The cytotoxicity studies were promising and showed no toxicity. Chen et al. [38], studied the in vitro cytotoxicity of surgical sutures made up of silk and were evaluated using the ISO 10993e5:2009 standards. The evaluation was conducted by measuring the viability of cells using liquid extract test method. The cytotoxicity of silk sutures was acceptable with reference to ISO standards. Cyanoacrylate(CA)-based tissue adhesives were extensively studied for the cytotoxicity and biocompatibility by Pascual et al. [39], and they used both in vitro and in vivo methods to evaluate the cytotoxicity. In surgical treatment of hernia CA is currently used by clinical practioners but (Glubran (n-butyl) and Ifabond (n-hexyl)) and a longerchain CA (OCA (n-octyl)) is not in the clinical usage. The studies indicated that CA exhibits good tissue integration with short-term biocompatibility, whereas OCA induces slightest seroma and macrophage response. Souza et al. [40], compared the biocompatibility of ethyl-cyanoacrylate (ECA) and octylcyanoacrylate (OCA) wound closures to conventional

Biocompatibility and cytotoxicity of polymer sutures

261

sutures by in vivo study. The study was conducted on rat’s skin. The study showed the biocompatibility of ECA without any necrosis, allergic reactions or infections. ECA have lower cost, and fewer complications compared to OCA and other sutures. Considering the cytotoxicity and biocompatibility studies, advanced novel suture materials derived from carbohydrate polymers are currently arising in the field, they exhibit excellent biocompatibility and have immense potential as suture materials. Cellulose, chitin, and alginate are extensively studied and employed to create advanced and novel suture materials [1]. A multitude of research and evaluation are going on with suture materials from the time of their discovery onwards. Now a days advanced absorbable sutures with low cytotoxicity and high biocompatibility is available to surgeons. New and advanced polymers play a major role in the development of advanced suture materials. Biocompatibility and cytotoxicity of these sutures are major concerns and are directly related to each other.

11.6 Conclusion Biocompatibility and cytotoxicity of the polymer sutures depends on the composition and mode of interaction of the suture with the tissue materials. Interaction of the material and the body is an important factor in determination of biocompatibility and cytotoxicity. Choosing the exact method and elucidation of biocompatibility and cytotoxicity is very important before using the suture materials. Higher compatibility and lower toxicity are the suitable property in case of suture materials. The methods used to study biocompatibility and cytotoxicity are based on certain guidelines set by FDA and ISO. Standardization and optimization of such tests are difficult and lengthy but it is very important to adhere with the standards.

References [1] K.M. de la Harpe, P.P.D. Kondiah, T. Marimuthu, Y.E. Choonara, Advances in carbohydrate-based polymers for the design of suture materials: a review, Carbohydr. Polym. 261 (2021) 117860, https://doi.org/10.1016/j.carbpol.2021.117860. [2] W. Zhou, P. Tan, X. Chen, Y. Cen, C. You, L. Tan, H. Li, M. Tian, Berberineincorporated shape memory fiber applied as a novel surgical suture, Front. Pharmacol. 0 (2020), https://doi.org/10.3389/fphar.2019.01506.

262

Advanced Technologies and Polymer Materials for Surgical Sutures

[3] B. Joseph, A. George, S. Gopi, N. Kalarikkal, S. Thomas, Polymer sutures for simultaneous wound healing and drug delivery e a review, Int. J. Pharm. 524 (2017) 454e466, https://doi.org/10.1016/j.ijpharm.2017.03.041. [4] M. Byrne, A. Aly, The surgical suture, Aesthetic Surg. J. 39 (2019) S67eS72, https:// doi.org/10.1093/asj/sjz036. [5] R.L. Moy, B. Waldman, D.W. Hein, A review of sutures and suturing techniques, J. Dermatol. Surg. Oncol. 18 (1992) 785e795, https://doi.org/10.1111/j.15244725.1992.tb03036.x. [6] G. Molea, F. Schonauer, G. Bifulco, D. D’Angelo, Comparative study on biocompatibility and absorption times of three absorbable monofilament suture materials (Polydioxanone, Poliglecaprone 25, Glycomer 631), Br. J. Plast. Surg. 53 (2000) 137e141, https://doi.org/10.1054/bjps.1999.3247. [7] F. Jummaat, E.B. Yahya, H.P.S. Abdul Khalil, A.S. Adnan, A.M. Alqadhi, C.K. Abdullah, A.K. Atty Sofea, N.G. Olaiya, M. Abdat, The role of biopolymerbased materials in obstetrics and gynecology applications: a review, Polymers 13 (2021) 633, https://doi.org/10.3390/polym13040633. [8] C.C. Chu, 10 e types and properties of surgical sutures, in: M.W. King, B.S. Gupta, R. Guidoin (Eds.), Biotextiles as Medical Implants, Woodhead Publishing Series in Textiles. Woodhead Publishing, Cambridge, UK, 2013, pp. 231e273. [9] T.M. Muffly, A.P. Tizzano, M.D. Walters, The history and evolution of sutures in pelvic surgery, J. R. Soc. Med. 104 (2011) 107e112, https://doi.org/10.1258/ jrsm.2010.100243. [10] J. Rose, F. Tuma, Sutures and needles, in: J. Rose, F. Tuma (Eds.), StatPearls, StatPearls Publishing, Treasure Island, FL, 2021. [11] D.E. Firestone, A.J. Lauder, Chemistry and mechanics of commonly used sutures and needles, J. Hand Surg. 35 (2010) 486e488, https://doi.org/10.1016/j.jhsa.2009. 10.036, quiz 488. [12] R.W. Postlethwait, D.A. Willigan, A.W. Ulin, Human tissue reaction to sutures, Ann. Surg. 181 (1975) 144, https://doi.org/10.1097/00000658-197502000-00003. [13] W. Kromka, M. Cameron, R. Fathi, Tie-over bolster dressings vs basting sutures for the closure of full-thickness skin grafts: a review of the literature, J. Cutan. Med. Surg. 22 (2018) 602e606, https://doi.org/10.1177/1203475418782152. [14] A. Schneider, H. Feussner, Chapter 6 e classical (open) surgery, in: A. Schneider, H. Feussner (Eds.), Biomedical Engineering in Gastrointestinal Surgery, Academic Press, Cambridge, MA, 2017, pp. 221e267, https://doi.org/10.1016/B978-0-12803230-5.00006-3. [15] A.J. Dart, C.M. Dart, 6.636 e Suture material: conventional and stimuli responsive, in: P. Ducheyne (Ed.), Comprehensive Biomaterials, Elsevier, Oxford, 2011, pp. 573e587, https://doi.org/10.1016/B978-0-08-055294-1.00245-2. [16] C. Dennis, S. Sethu, S. Nayak, L. Mohan, Y. Morsi, G. Manivasagam, Suture materials d current and emerging trends, J. Biomed. Mater. Res. A 104 (2016) 1544e1559, https://doi.org/10.1002/jbm.a.35683. [17] B. Busse, Selecting materials for wound closure, in: B. Busse (Ed.), Wound Management in Urgent Care, Springer International Publishing, Cham, 2016, pp. 19e23, https://doi.org/10.1007/978-3-319-27428-7_4. [18] D. Kirsch, S. Marczyk, Multifilament Barbed Suture, US8414612B2, 2013. [19] C.C. Chu, Suture Materials. Kirk- Othmer Encyclopedia of Chemical Technology, Cornell University, Ithaca, NY, USA, 2017, pp. 1e35. [20] A. Vasanthan, K. Satheesh, W. Hoopes, P. Lucaci, K. Williams, J. Rapley, Comparing suture strengths for clinical applications: a novel in vitro study, J. Periodontol. 80 (2009) 618e624, https://doi.org/10.1902/jop.2009.080490.

Biocompatibility and cytotoxicity of polymer sutures

263

[21] R. Karabulut, K. Sonmez, Z. Turkyilmaz, B. Bagbanci, A.C. Basaklar, N. Kale, An in vitro and in vivo evaluation of tensile strength and durability of seven suture materials in various pH and different conditions: an experimental study in rats, Indian J. Surg. 72 (2010) 386e390, https://doi.org/10.1007/s12262-010-0158-5. [22] J. Tjandra, G.J. Clunie, A.H. Kaye, J.A. Smith (Eds.), Textbook of Surgery, third ed., 2008 [WWW Document], eBooks.com, https://www.ebooks.com/en-us/book/ 351492/textbook-of-surgery/joe-tjandra/ (Accessed 17. November 2021). [23] A.J. Dart, C.M. Dart, 7.38 suture material: conventional and stimuli responsive, in: P. Ducheyne (Ed.), Comprehensive Biomaterials II, Elsevier, Oxford, 2017, pp. 746e771, https://doi.org/10.1016/B978-0-12-803581-8.10135-3. [24] O. Basçı, U. Akgun, F.A. Barber, Biological properties of suture materials, in: U. Akgun, M. Karahan, P.S. Randelli, J. Espregueira-Mendes (Eds.), Knots in Orthopedic Surgery: Open and Arthroscopic Techniques, Springer, Berlin, Heidelberg, 2018, pp. 11e20, https://doi.org/10.1007/978-3-662-56108-9_2. [25] M.M. Al-Qattan, H. Kfoury, A delayed allergic reaction to polypropylene suture used in flexor tendon repair: case report, J. Hand Surg. 40 (2015) 1377e1381, https:// doi.org/10.1016/j.jhsa.2015.03.004. [26] D.F. Williams, On the mechanisms of biocompatibility, Biomaterials 29 (2008) 2941e2953, https://doi.org/10.1016/j.biomaterials.2008.04.023. [27] L.G. Donaruma, Definitions in biomaterials, D. F. Williams, Ed., Elsevier, Amsterdam, 1987, 72 pp, J. Polym. Sci. C Polym. Lett. 26 (1988), https://doi.org/10.1002/ pol.1988.140260910, 414e414. [28] J.A. Porter, J.A. Von Fraunhofer, Success or failure of dental implants? A literature review with treatment considerations, Gen. Dent. 53 (2005) 423e432. [29] J.L. Ng, V.D.L. Putra, M.L. Knothe Tate, In vitro biocompatibility and biomechanics study of novel, Microscopy Aided Designed and ManufacturEd (MADAME) materials emulating natural tissue weaves and their intrinsic gradients, J. Mech. Behav. Biomed. Mater. 103 (2020) 103536, https://doi.org/10.1016/j.jmbbm.2019.103536. [30] S. Saxena, A.R. Ray, A. Kapil, G. Pavon-Djavid, D. Letourneur, B. Gupta, A. Meddahi-Pellé, Development of a new polypropylene-based suture: plasma grafting, surface treatment, characterization, and biocompatibility studies, Macromol. Biosci. 11 (2011) 373e382, https://doi.org/10.1002/mabi.201000298. [31] Chapter 6 e biocompatibility and tissue reaction to biomaterials, in: R.L. Sakaguchi, J.M. Powers (Eds.), Craig’s Restorative Dental Materials, thirteenth ed., Mosby, Saint Louis, 2012, pp. 109e133, https://doi.org/10.1016/B978-0-323-08108-5.10006-4. [32] J.M. Anderson, 4.402 e Biocompatibility and the relationship to standards: meaning and scope of biomaterials testing, in: P. Ducheyne (Ed.), Comprehensive Biomaterials, Elsevier, Oxford, 2011, pp. 7e26, https://doi.org/10.1016/B978-0-08-0552941.00002-7. [33] D. Pappalardo, T. Mathisen, A. Finne-Wistrand, Biocompatibility of resorbable polymers: a historical perspective and framework for the future, Biomacromolecules 20 (2019) 1465e1477, https://doi.org/10.1021/acs.biomac.9b00159. [34] S. Vijayan, K. Divya, M.S. Jisha, In vitro anticancer evaluation of chitosan/biogenic silver nanoparticle conjugate on Si Ha and MDA MB cell lines, Appl. Nanosci. 10 (2019) 715e728, https://doi.org/10.1007/s13204-019-01151-w. [35] M.A. Javaid, K.M. Zia, R.A. Khera, S. Jabeen, I. Mumtaz, M.A. Younis, M. Shoaib, I.A. Bhatti, Evaluation of cytotoxicity, hemocompatibility and spectral studies of chitosan assisted polyurethanes prepared with various diisocyanates, Int. J. Biol. Macromol. 129 (2019) 116e126, https://doi.org/10.1016/j.ijbiomac.2019.01.084. [36] W. Zhang, B. Yin, Y. Xin, L. Li, G. Ye, J. Wang, J. Shen, X. Cui, Q. Yang, Preparation, mechanical properties, and biocompatibility of graphene oxide-reinforced

264

[37]

[38] [39]

[40]

Advanced Technologies and Polymer Materials for Surgical Sutures

chitin monofilament absorbable surgical sutures, Mar. Drugs 17 (2019) 210, https:// doi.org/10.3390/md17040210. C.H. Ho, E.K. Odermatt, I. Berndt, J.C. Tiller, Long-term active antimicrobial coatings for surgical sutures based on silver nanoparticles and hyperbranched polylysine, J. Biomater. Sci. Polym. Ed. 24 (2013) 1589e1600, https://doi.org/10.1080/ 09205063.2013.782803. X. Chen, D. Hou, L. Wang, Q. Zhang, J. Zou, G. Sun, Antibacterial surgical silk sutures using a high performance slow-release carrier coating system, ACS Appl. Mater. Interfaces 7 (2015), https://doi.org/10.1021/acsami.5b06239. G. Pascual, S. Sotomayor, M. Rodríguez, B. Pérez Köhler, A. Kühnhardt, M. Fernandez Gutierrez, J. San Roman, J. Bellón, Cytotoxicity of cyanoacrylate-based tissue adhesives and short-term preclinical in vivo biocompatibility in abdominal hernia repair, PLoS One 11 (2016) e0157920, https://doi.org/10.1371/journal.pone. 0157920. S.C. de Souza, W.L. de Oliveira, D.F.O.S. de Soares, C. Briglia, P. Athanazio, M. Cerqueira, P. Guimarães, M.C. Carreiro, Comparative study of suture and cyanoacrylates in skin closure of rats, Acta Cir. Bras. 22 (2007), https://doi.org/10.1590/ S0102-86502007000400013.

CHAPTER 12

Shape memory polymers as sutures Haritha R. Das1, Arya Uthaman1, Hiran Mayookh Lal1, Allan Babu1 and Sabu Thomas2 1 2 School of Energy Materials, Mahatma Gandhi University, Kottayam, Kerala, India; International and Inter University Centre for Nanoscience and Nanotechnology, Mahatma Gandhi University, Kottayam, Kerala, India

12.1 Introduction Shape memory refers to the capability of a material to regain or recall its previous shape after subjecting to a deformation due to mechanical force or by heating and cooling. A structural phase change is at the root of this behavior. Shape memory polymers (SMPs) retype of polymer that can switch between a temporary and a permanent shape essentially endlessly. The material attains a temporary shape by applying a mechanical deformation of a material with permanent shape under certain conditions, such as heating above the glass transition or melting temperature, and then maintaining their temporary shape while the deformation is sustained (e.g., using crystallization, chemical crosslinking, or supramolecular interactions). The polymer will hold its temporary shape after stress is released until an appropriate trigger is provided, such as light, increased temperature, contact with particular chemicals, and so on, causing the permanent shape recovery [1e4]. The thermal behavior of SMPs is illustrated in Fig. 12.1 [5].

Figure 12.1 Thermal response of shape-memory polymers [5]. (Reproduced with permission from John Wiley and Sons.) Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00001-2

© 2023 Elsevier Ltd. All rights reserved.

265

266

Advanced Technologies and Polymer Materials for Surgical Sutures

However, shape memory metal alloys (SMAs), such as FeNiAl and CuZnAl alloys, were the first materials to exhibit these properties. Their martensitic transformation occurs when there is a change in the structural phase of these materials. Furthermore, the SMPs are being advanced and capable to replace or supplement SMAs because they are lighter, have a higher shape recovery capacity, are easier to handle, and are less expensive than SMAs [6] Some of the typical applications of these SMPs are demonstrated in Fig. 12.2 [7]. The common advantages and disadvantages of the SMPs compared with shape memory alloys [8] are represented in Table 12.1. Recently, SMPs for biomedical applications are tremendously gaining attention [5,9,10]. Some applications such as vascular stents, medical guidewires, sutures, orthodontic wires, and other applications have been typically proposed for these materials [11]. Moreover, incorporating the porous materials (due to the permeability, high porosity of porous materials) [12] and the shape memory effect can enhance the properties, especially in tissue engineering applications [13]. Shape-memory polymers are exciting options for developing tissue-engineered constructions because of their potential to understand complicated shapes that are different from their permanent shape. SMPs are phase-separated linear block copolymers with a hard section and a soft section. The hard section is usually crystalline and has a known melting point; on the other hand, the soft section is amorphous and has a specified glass transition temperature [14]. The shape-memory effect is one of the

Figure 12.2 Common applications of shape memory polymers [7]. (Reproduced with permission from Elsevier.)

Shape memory polymers as sutures

267

Table 12.1 Merits and demerits of SMPs compared with SMAs. Advantages

Disadvantages

Low cost Good processability Light weight High recoverable deformability

Low recovery stress Low thermal resistance Short lifetime Low mechanical strength

properties of polymers. However, it is not a polymer’s inherent feature and necessitates additional processing stages [15]. Entropic elasticity lies at the heart of the shape-memory process. The firmly coiled shape of an amorphous chain segment of the polymers is the entropically most advantageous state. The randomly arranged polymer chains will elongate and arrange in the direction of the applied mechanical force due to elastic deformation, limiting the number of active configurations and hence the entropy. The relaxation of the tension causes the polymers to revert to their original, more entropically favorable, coiled configuration because chain entanglements of polymer chains high molar mass prevent the whole polymer chain from moving and functioning as a kind of physical crosslinks. The polymer can be said to have a “memory” of its original, undeformed condition. However, applying the stress over longer timeframes will cause slippage of the polymer chains, causing the polymer to forget its initial structure [5]. The shape-memory effect requires a constant polymer chain network and an adequate reversible transition of the polymer chains. The crystalline phase, molecule entanglement, an interpenetrated network, a chemical cross-linking, could all be used to create stable network structures. Melting/ crystallization transitions, glass transitions, isotropic/anisotropic transitions, reversible molecular cross-linking, and supramolecular disassociation/association are all reversible switching transitions [16]. During the shape recovery cycle, SMPs are first deformed. Because of the poor mobility of the macromolecules, the distorted polymer cannot recover if the switch is locked. The SMP recovers as a result, along with the release of internal stress. The internal tension can also be partially stored in the cross-linking network since the locks control the macromolecules’ mobility. Finally, the stimuli open the reversible locks when they are stimulated. Other tactics that can considerably modify the movement of the SMP that may initiate the shape recovery and to unlock the reversible locks. Moisture, ultrasonic field, light, magnetic field, electric impulses, and ionic concentration activate the shape-memory effect [8].

268

Advanced Technologies and Polymer Materials for Surgical Sutures

Both the amorphous and crystalline areas of a polymer contribute to the shape memory property of the polymer. The reversible crosslinking or the thermal transitions in the polymer chains can fix the temporary shape depending on the SMP type. These components can be act as additional crosslinks, allowing the polymer to become distorted, entropically less advantageous while storing the deformation energy. When an appropriate trigger is applied, the temporary crosslinking is undone, softening the polymer chains and regaining their permanent shape [17]. Programming is used to create this effect in a polymer. The first step involves the treatment of the polymer so that it can take on its final shape. The polymer then deformed, and the temporary shape is fixed by heating, deforming, cooling, or drawing the sample at a low temperature. Then the sample transforms to a temporary shape while the permanent shape is preserved. The shape memory effect in polymers can be induced by heating the SMP over its transition temperature. As a result, the permanent shape could recover, called thermo responsive SMPs [18]. Although the temperature is a popular trigger for SMPs, it has limitations in providing shape-memory switching under specific conditions. SMP is commonly filled with particles to provide shape-changing effects under stimuli with mechanical, electrical, and magnetic properties [18]. Several stimuli driven SMPs is shown in Table 12.2. Table 12.2 Several stimuli driven shape memory polymers. Polymer types

Driven methods

Transition temperature

Polyurethane/poly(methacrylic acid) Polyurethane /poly(methacrylic acid) Epoxy based on shape memory polymer Epoxy resin/PCL fibers polymer(methylmethacrylate)/ polyethylene glycol Crosslinked poly(methyl Methacrylate-co-butyl Acrylate copolymer Chemically cross-linked poly(3-caprolactone) Composite

Thermal stimuli

50
C/80
C

Water/pH UV and temperature

60
C

Temperature Temperature

40
C/80
C 35
C/50
C/ 70
C/110
C 47
C

High-intensity focused ultrasound Magnetic field, electric field and temperature

About 40
C

Shape memory polymers as sutures

269

SMPs can evolve in novel biomedical applications, particularly devices for minor invasive surgery, based on active reaction and deformation. SMP has numerous advantages over traditional materials, including lightweight, low cost, colossal deformation, and tunable glass transition temperature. Stress generation, shape change, vapor permeability change, and modulus change, are the four kinds of SMP applications depending on the applications. The applications of SMPs have been extended toward various fields such as smart textiles and apparel, smart medical instruments and auxiliaries, heat shrinkable packaging, selfhealing, microelectro-mechanical systems, etc. Biodegradability and biocompatibility are two main aspects of SMPs in biomedical fields. Suture threads, drug delivery vehicles, stents, and tissue engineering are among the many biological uses for which biodegradable SMPs have been developed. Electrospinning and electrospraying processes produce nanostructured fibers with tunable release kinetics that can be used in various biological applications. Drug-eluting sutures result in fewer surgical site infections, faster wound healing, fewer postoperative problems, and, most importantly, fewer supplement medications [19]. Moreover, for preparing SMPs nanofiber structures, the functional materials typically used are carbon nanotubes, graphene oxide, cellulose nanoparticles, while electrospinning to empower the fibers having multi-shape memory property [20]. Furthermore, Fe3O4, epoxy resins are used to reinforce with fibers to achieve better properties since epoxy resins have high strength [21e23], curing or self-healing performance, conductivity. In addition, silver nanoparticles were considered for attaining more excellent antimicrobial properties [24e26]. Iregui et al. [27] achieved an SME fiber structure of diglycidyl ether of bisphenol A (DGEBA) epoxy resin with PLA through electrospinning and UV radiation. Additional research is being done to manage the chemistry of the structure of SMPs to modify the glass transition temperature according to the applications. This considerably expands the spectrum of biomedical applications that could be developed. This chapter mainly focuses on the biomedical application of shape memory properties of SMPs, especially for sutures.

12.2 Sutures A suture is a natural or synthetic biomaterial, which is mainly applied to ligate blood vessels and approximate tissues after surgery or injury. It refers to the method used for mechanical wound closure. The main objective of the wound closure comprises eradication of dead space, even the uniform

270

Advanced Technologies and Polymer Materials for Surgical Sutures

transfer of tension along deep suture lines, preservation of tensile strength across the wound until the tissue strength is requisite, and about the epithelial portion of the closure [28]. While alternate techniques, such as surgical glue made from polyalkylcyanoacrylates, may be effective in some situations, the traditional procedure still involves a lengthy process of stitching and surgical knots to unite the edges of a wound. It is critical to knot a suture with the optimum amount of tension because if the knot is too tight, the surrounding tissues may undergo necrosis, and if it is too loose, the incision may not heal properly, resulting in the formation of scar tissues. Traditional sutures have the potential to promote infection in the incision. They are categorized as nonabsorbable and absorbable surgical sutures based on their biostability, multifilament (braided), and monofilament surgical sutures based on their design. Sutures that are commonly employed have several flaws: • They cannot wholly close pierced tissue. • Multifilament threads increase the risk of bacterial migration. • Tissue reactivity is increased owing to their design. Bottaro and Larsen invented electrospin sutures to provide anesthesia to the wounded area. This method is used to provide an analgesic agent, a critical component of practically all surgeries, to reduce postoperative discomfort, to reduce the need for additional systemic medications such as opioid drugs that impair the patient’s function, and to avoid the need for further procedures (such as local or regional anesthetic nerve blocks) [29]. Polylactic-co-glycolic acid (PLGA). PLGA is a biodegradable and biocompatible polymer. PLGA fibers bundled in a single stranded formulation with different dosages of local anesthetic bupivacaine hydrochloride were used in this study. During the electrospinning process, a charge is applied to the polymer-drug mixture dissolved in the volatile organic solvent, which expels the solution as a thin stream into a grounded collector where the polymer microfibers are deposited. Drug-containing stitches. This method can be used to make fibers from practically any polymer, which does not require high operating temperatures and can damage heat-labeled pharmaceuticals. Antibiotics, anticancer drugs, proteins, and other therapies can all be added due to the diversity of matrix material selection [30]. 12.2.1 SMP sutures The shape-memory characteristic has the added benefit of allowing not only biodegradable but also self-tightening stitches. This allows surgeons to

Shape memory polymers as sutures

271

loosen a wound before using a trigger to activate and close the shapememory effect. SMP-based absorbable surgical sutures have the potential to address current unrestricted needs. Among other features, they may allow for more accessible applications and secure security features. Surgical sutures made from biodegradable SMPs can provide permanent and unified restorative strength, making them “smart” surgical sutures [29]. Since when biodegradation is a desired attribute of a material, it should ideally leave no residue, and the degradation products should be harmless. In addition, the degradation speed of the polymer must be compatible with the application; since new tissue formation is often compatible with polymer degradation, polymer degradation should be slower or equivalent to new tissue formation. Degradation products must not be harmful to the human body and must have the proper speed and mechanical properties. Local acidity can occur as a result of the breakdown of polylactic acid polymers. This can lead to auto-accelerating deterioration due to acidity, which might trigger an inflammatory response. Thus, the addition of hexamethylene disocyanate and butane diamine to the polylactic SMPs added a hard segment to the SMP for better shape memory and increased the alkaline group content. As a result, the alkali group can control the rate of degradation [29]. Furthermore, SMPs can be made biodegradable by incorporating fragile and hydrolyzable bonded polymer segments that split under physical conditions. Biodegradable SMP sutures are helpful because they degrade after wound healing and eliminate the necessity for a further surgical procedure to take out the sutures. Common biodegradable SMPs for medical applications are poly (D, L-lactoid-co-glycolide), cross-linking poly (e-caprolactone) dimethycrylate, and n-butyl acrylate. 12.2.1.1 Main factors for SMPs sutures In the case of a biomedical context, shape-memory polymers must meet general requirements as well as application-specific requirements. Since shape-memory polymers are commonly extracted from traditional polymeric materials, these needs are generally in line with the requirements of conventional biomaterials, and there are no special requirements for shapememory materials. The most crucial factor is that its toxicity as it is a nontoxicity biomaterial. As a result, this refers to the noncarcinogenic, nonallergic, blood compatible, and noninflammatory materials. Also, do not emit or release any molecules unless the biomaterial is designed. Another important factor is that it must be biocompatible, and work with a

272

Advanced Technologies and Polymer Materials for Surgical Sutures

host response that is compatible with its specific application (i.e., the material must be nonmucogenic). Compared to synthetic polymers, biopolymers have an advantage because they generally have lower inflammatory reactions. For example, polymers with cardiovascular applications may contact the blood and thrombus may not form on the polymeric surface. In addition, there should be no absorption of proteins on the polymer surface. To evaluate this, a biomaterial must be subjected to a hemocompatibility study. Polymers previously used to meet this requirement include poly (caprolactone) (PCL), polylactic acid (PLA), polyglycolic acid (PGA), and polydioxenone with various polyurethanes. Another significant factor is that a polymer used as a biomaterial must not provoke a foreign body response. In addition, the biomaterials must match the mechanical properties of the implant material to the surrounding tissue. This is usually necessary to avoid modification of the in vivo biomechanics in the target area. The mechanical requirements are application-specific and depend on the function of the material and the implantation site. In addition, biodegradation also plays an important role. Although degradation of an implant material is required, how quickly degradation should occur depends on the application and the time it takes for an implant to perform its function. Therefore, biodegradable sutures are relevant because they may be destroyed after the wound has dried, thereby preventing the need for suture removal. The main advantage of the shape-memory characteristic in this case is the ability to make not only biodegradable, but also self-tight stitches. As a result, it allows surgeons to stitch the wound loosely, thereby activating the shape-memory effect and triggering a wound closure. Decomposition is not desirable when the polymer material remains permanently in the body. While biodegradation is a required property of a material, degradation products must be nontoxic. PGA, PLA, and PCL and their derivatives are generally considered biodegradable polymers and biopolymers such as cellulose or gelatin. In addition, the degradation speed of the polymer must be tuned to the application, which often corresponds to the formation of new tissue, so the polymer degradation must be slow or equal to the formation of new tissue. Moreover, another factor is the sterilization capacity of the polymer; since nonsterile polymers are not allowed to serve in vivo applications, they cannot be used for in vitro evaluation of cellular response, so these materials are not suitable for biomedical applications [5].

Shape memory polymers as sutures

273

12.2.1.1.1 Importance of polyurethane based sutures The terms “smart” or “intelligent” are used for shape memory polyurethanes because they are able to regain a certain temporary deformed shape back to its original form in the presence of external stimuli such as temperature. Normal polyurethanes cannot fully recover their shape when heated. However, due to the micro-Brownian motion of the polymer chains, when heated to 10e20
C above their glass transition temperature (Tg) or crystal melting temperature (Tm), the shaped memory polyurethane can temporarily recover the deformed form. These extensive deformation restorations upon temperature stimuli could be helpful for many potential applications. Sutures made with SMPs help reduce scarring, speeds healing, and reduce the risk of infection by knotting a suture with the right amount of tension in the influence of a body temperature. Sewlike-shaped memory polyurethane can be bonded together into a complete bundle due to their shape memory nature, which is activated by human body temperature. In addition, it can shrink and form a bundle of optimal tension in the surrounding tissues [31]. 12.2.1.1.2 Applications of SMPs sutures A major challenge associated with endoscopic surgery is the use of instruments and sutures to close a wound or open a lumen. Necrosis of the surrounding tissue can occur when the joint is tightened with strong force. It also leads to the formation of hernias if the knot is weak. Moreover, it is challenging to evaluate the suture to press the wound lips under the correct stress. The Smart Surgical Suture design may have a temporary shape, obtained by stretching the fiber with controlled pressure, which will solve this problem, and the suture can be applied loosely in its temporary form. As the temperature rises above the trans, the seam shrinks, the knot tightens, and the optimal force is applied [32]. Andreas Lendlien and Robert Langer have developed a smart seam with biodegradable, elastic shaped memory polymers that are heat reactive. In their opinion, the lesion in rats showed a contraction of the sample-shaped memory suture, while the temperature to close the wound increased from 20 to 41
C, as shown in Fig. 12.3 [33]. From past years, a compression process has been established for wound sutured anastomosis. For sutureless anastomosis, a spring made of shape memory polylactic acid was created. The spring expanded to some level before being put into the junction of the two tubes [34] as illustrated in Fig. 12.4. When the spring reached body temperature, it progressively

274

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 12.3 Degradable shape-memory suture for wound closure [33]. (Reproduced with permission from Science.)

Figure 12.4 Analysis of a spring based on shape memory PLA for sutureless anastomosis [34]. (Reproduced with permission from Elsevier.)

tightened and frapped the two tubes together. Furthermore, because of its degradable action, the options to use it in bleeding control are substantially increased [34]. Light-triggered SMPs also attract attention because of the possibility of remote activation with high spatial and temporal resolutions using a large spectrum of light [from UV-Vis to Infrared (IR)]. Toncheva et al. Developed light-reactive nanocomposites based on biocompatible polycaprolactone (PCL-SMP) containing complex nanofillers such as biosourced CNCs adorned with AGNPs. For self-tightening suture behavior and fast shape recovery, IR illumination was applied. The thermal heating of the light source material did not exceed more than 37
C. This technology has many advantages over the previous functions of thermally responsive SMPs. The nanofillers also enhanced the thermo-mechanical properties as well as showed good antibacterial properties [35].

Shape memory polymers as sutures

275

Arpan Biswas reported the synthesis of polyurethane nanohybrid through in situ polymerization. In this work, a polymer is grown inside the biologically modified layers of nano clay, which allows the molecular sheets to be better bonded and the crystallites to be distributed within the polymer matrix, which significantly increases the hardness. The limited flipping of the hard segment/nanoclay reduces the stacking density of the hard segment in the HPL-NH nano hybrid in the presence of large interaction of the nanoclay with the polymer network. Conversely, temperature-induced flipping of distributed hard segments takes place in pure polyurethane, which increases the stacking of the hard segment in pure HPL. Decreasing the melting point of the soft segment and decreasing the stacking pattern of the hard segment with temperature results in a significant increase in the shape memory characteristic of HPL-NH compared to pure HPL. The enhanced ratio of recovery at 40
C with sustained flexibility generates an opportunity for HPL-NH to be used as a self-tightening suture [36]. Bai et al. A new bio-friendly SMP based on ethyl cellulose (EC) and PCL was developed. EC was added as the backbone of the polymeric network to adjust the Tm and mechanical strength of the PCL materials. By changing the ratio between EC and e-CL, a series of EC-PCL SMPs (ECPCL SMP) with varying molecular weights of graft PCL were generated. Linear EC backbones constructed the polymeric network structure joined by grafted PCL chains, and the results showed that the polymer had exceptional mechanical strength and shape memory properties. The ECPCL SMP also exhibited shape memory effect actuated by body temperature. The introduction of EC to polymer network greatly improved the mechanical strength along with its shape memory property. Therefore, ECSMP can be applied as biomedical sutures [37]. In a study after 21 days in rabbits, the suture material revealed considerably low inflammatory cells and better collagen than sutures fabricated using polypropylene. Polyurethane sutures developed from shapedmemory polymers with increased elasticity and tensile strength, as well as self-tightening bindings, expand their therapeutic applications. Many polymers have a soft/hard segment structure similar to ethylene-vinyl acetate (EVA) and polyurethane (PU), while specific polymers and polymer blends/hybrids have an inclusion-matrix microstructure. From this, it can be seen that if the heating-responsive SME of these polymers is elastic in the hard segment or matrices temperature range of interest and in the soft segment or inclusion, its hardness can vary considerably when heated by glass transition. Melting. The elastic segment as the elastic component,

276

Advanced Technologies and Polymer Materials for Surgical Sutures

while the soft segment as the transition component. The elastic component holds elastic energy after programming at high or low temperatures. The plastically deformed and rehardened transition component, on the other hand, acts as a restriction to inhibit shape recovery at low temperatures [38]. The limitation can only be removed when the transition component is heated to soften, and the polymer can return to its original shape using the elastic energy stored in the elastic component. Newly designed thermoplastic polyurethane elastic sutures with promising tensile properties are possible and safe for midline laparotomy wound closure, preventing postoperative problems such as burst abdomen during abdominal surgery [5]. Sutures with drug-eluting properties are a more advanced suture that delivers medicine to a specific region during surgery. Various novel ways are being developed to improve the efficacy of sutures as a physical entity for obtaining an enhanced biologically active component that allows the delivery of different types of desirable medications and cells to the affected spot. Modified sutures that are ideal must keep their mechanical stability during the wound healing process, and deliver the medications contained in them in a controlled manner. The research group from Maitland created a thrombectomy device that is useful for stroke patients using shape memory thermoset polyurethanes. The thrombectomy device is initially directed into the desired location to thrombus accumulation through a tube. The device returns to the coil configuration after puncturing the clot, which gathers the thrombus and removes it. The strategy of the thromboectomy device was examined 10 times under neurovascular pressure. The Mytland Group later developed a hybrid thromboectomy device using a nickeltitanium alloy in a thermosetting polyurethane shell. The hybrid device has greater recovery power than the pure SMP thromboectomy device. An SMP-Nitinol hybrid wire is straightened throughout the working process and prevents SMP alloy recovery in the glass. Both shape memory effects were activated, and the NieTi alloy wire was heated to restore the initial appearance. The device was tested on a rabbit model of acute artery occlusion, and a quarter of its results showed that the clot had been removed [39] is shown in Fig. 12.5. John Rogers invented smart sutures consisting of plastic or silk threads with temperature sensors and microheaters to detect infections. They started by slicing an ultrathin sheet of silicon from a silicon wafer with chemicals. They took off the nanomembranes with a rubber stamp and transferred them to polymeric or silk strips. The metal electrodes and wires are then deposited on top, and the entire device is encased in an epoxy

Shape memory polymers as sutures

277

Figure 12.5 SMP-based clot microactuator [39].

coating. On the sutures, they installed two types of temperature sensors. The first is a silicon diode that changes its current output as temperature changes, while the second is a platinum nanomembrane resistor that changes its resistance as temperature changes. Meanwhile, the microheaters are just gold filaments that heat up as current travels through them. The electronic sutures, which feature ultrathin silicon sensors embedded on polymer or silk strips, can be threaded through needles, and researchers were able to lace them through the skin, pull them tight, and knot them without causing the electronics to degrade in animal tests. The sutures can accurately assess temperature (high temperatures indicate infection) and provide heat to a wounded location, which is believed to help heal [40]. Houshyar et al. suggested that hybrid smart PCL/CS suture was designed as a multifunctional suture that can monitor in situ temperature, promote mammalian cell development, and be easily coated with

278

Advanced Technologies and Polymer Materials for Surgical Sutures

antibiotics to avoid wound infection. The authors observed that the antibiotics quickly coat the surface of the sutureþ, reducing the likelihood of bacterial infection in the incision [41].

12.3 Conclusion SMPs offer current and future opportunities in many medical applications. When the right amount of pressure is applied to the tissues, the SMP sutures can tighten themselves into a perfect bundle. Some sutures are biodegradable, capable of being absorbed into the human body, eliminating the need for a second surgery. The mechanical properties and stress of the suture applied to the tissue can be adjusted by modifying the programming process and modifying the molecular structure of the shape-memory sutures. Surgical sutures play an important role as a medical tool in wound management, and recent advances in this area may have contributed to technological advances in material science. Polymers are important because of their high flexibility, which results in a wide variety of suture materials with good mechanical and thermal properties. Also, biodegradable polymers as sutures are widely applied as it is easy to broke and removed from the body without any surgery. These substances are also widely known for their ability to transport drugs, proteins, stem cells, peptides, DNA, antibodies, nanoparticles, to the desired area, enhancing the healing properties of sutures. The physicochemical properties of sutures affect their effectiveness and function. When modified or coated with bioactive agents and sensors it is important to maintain these properties with excellent handling properties and the necessary modifications; the stitches should be cancer-free, nontoxic, and bio-compatible. For some applications, the mechanical strength and recovery pressure of SMPs are still low, and the material is much softer than changing temperatures. The properties of shape memory alloys are excellent in these respects, but the main advantage of SMPs is that they are not rigid and can be biodegradable. Recent works mainly focus on the incorporation of fillers in these polymers. Mainly metal based and carbon nanomaterials would enhance their mechanical and thermal properties, and also, the nanofillers show a good antibacterial property, which is an advantage for the sutures. The evolution of new materials as the sutures as an outcome of scientific research will have a major impact on wound management and surgery.

Shape memory polymers as sutures

279

References [1] H.Y. Jiang, S. Kelch, A. Lendlein, Polymers move in response to light, Adv. Mater. 18 (2006) 1471e1475, https://doi.org/10.1002/adma.200502266. [2] B. Yang, W.M. Huang, C. Li, L. Li, Effects of moisture on the thermomechanical properties of a polyurethane shape memory polymer, Polymer 47 (2006) 1348e1356, https://doi.org/10.1016/J.POLYMER.2005.12.051. [3] J. Lee, S.-K. Kang, Principles for controlling the shape recovery and degradation behavior of biodegradable shape-memory polymers in biomedical applications, Micromachines 12 (2021) 757, https://doi.org/10.3390/mi12070757. [4] H. Tobushi, K. Hoshio, S. Hayashi, N. Miwa, Shape memory composite of SMA and SMP and its property, Eng. Plast. Its Appl. 340 (2007) 1187e1192, https://doi.org/ 10.4028/www.scientific.net/KEM.340-341.1187. Trans Tech Publications Ltd. [5] J. Delaey, P. Dubruel, S. Van Vlierberghe, Shape-memory polymers for biomedical applications, Adv. Funct. Mater 1909047 (2020) 1e23, https://doi.org/10.1002/ adfm.201909047. [6] J.H. Alcamo, Surgical Suture, 1964, p. 3123077. US PATENT. 3,123,077. [7] F. Pilate, A. Toncheva, P. Dubois, J.M. Raquez, Shape-memory polymers for multiple applications in the materials world, Eur. Polym. J. 80 (2016) 268e294, https://doi.org/ 10.1016/J.EURPOLYMJ.2016.05.004. [8] H. Meng, G. Li, B. Rouge, B. Rouge, Shape-memory and self- reinforcing polymers as sutures, Shape Memory Polym. Biomed. Appl. 22 (2015), https://doi.org/10.1016/ B978-0-85709-698-2.00014-3. [9] C.M. Yakacki, K. Gall, Shape-memory polymers for biomedical applications, in: A. Lendlein (Ed.), Shape-Memory Polym, Springer, Berlin, Heidelberg, 2009, https:// doi.org/10.1007/12_2009_23. [10] D.A. Tirrell, R. Langer, Materials for biology and medicine, Nature 428 (2012) 25. [11] H.M. Wache, D.J. Tartakowska, A. Hentrich, M.H. Wagner, Development of a polymer stent with shape memory effect as a drug delivery system, J. Mater. Sci. Mater. Med. 14 (2003) 109e112, https://doi.org/10.1023/A:1022007510352. [12] A. Uthaman, S. Thomas, T. Li, H. Maria (Eds.), Advanced Functional Porous Materials: From Macro to Nano Scale Lengths, first ed., Springer Nature Switzerland AG, Cham, 2021 https://doi.org/10.1007/978-3-030-85397-6. [13] M. Zare, N. Parvin, M.P. Prabhakaran, J.A. Mohandesi, S. Ramakrishna, Highly porous 3D sponge-like shape memory polymer for tissue engineering application with remote actuation potential, Compos. Sci. Technol. 184 (2019) 107874, https:// doi.org/10.1016/j.compscitech.2019.107874. [14] J.C. Worch, A.C. Weems, J. Yu, M.C. Arno, R.T.R. Huckstepp, R.K.O. Reilly, et al., Elastomeric polyamide biomaterials with stereochemically tuneable mechanical properties and shape memory, Nat Commun (2020) 1e11, https://doi.org/10.1038/ s41467-020-16945-8. [15] K.A. Cavicchi, M. Pantoja, M. Cakmak, Shape memory ionomers, J. Polym. Sci., Part B: Polym. Phys. 54 (2016) 1389e1396, https://doi.org/10.1002/polb.24052. [16] A. Lendlein, M. Behl, B. Hiebl, C. Wischke, A. Lendlein, M. Behl, et al., Shapememory polymers as a technology platform for biomedical applications, Expert Rev. Med. Devices 7 (3) (2010) 357e379, https://doi.org/10.1586/erd.10.8. [17] A. Lendlein, S. Kelch, Shape-memory polymers, Angew. Chem. Int. Ed. 41 (2002) 2034e2057. [18] W. Zhao, L. Liu, F. Zhang, J. Leng, Y. Liu, Shape memory polymers and their composites in biomedical applications, Mater. Sci. Eng. C 97 (2018). [19] A. Arora, G. Aggarwal, J. Chander, P. Maman, M. Nagpal, Drug eluting sutures: a recent update, J. Appl. Pharmaceut. Sci. 9 (2019) 111e123, https://doi.org/10.7324/ JAPS.2019.90716.

280

Advanced Technologies and Polymer Materials for Surgical Sutures

[20] M. Zare, P. Davoodi, S. Ramakrishna, Electrospun shape memory polymer micro-/ nanofibers and tailoring their roles for biomedical applications, Nanomaterials 11 (2021), https://doi.org/10.3390/nano11040933. [21] H.M. Lal, G. Xian, S. Thomas, L. Zhang, Z. Zhang, H. Wang, Experimental study on the flexural creep behaviors of pultruded unidirectional carbon/glass fiber-reinforced hybrid bars, Materials 13 (2020) 11e13, https://doi.org/10.3390/ma13040976. [22] A. Uthaman, G. Xian, S. Thomas, Y. Wang, Q. Zheng, X. Liu, Durability of an epoxy resin and its carbon fiber-reinforced polymer composite upon immersion in water, acidic, and alkaline solutions, Polymers 12 (2020), https://doi.org/10.3390/ polym12030614. [23] H.M. Lal, A. Uthaman, C. Li, G. Xian, S. Thomas, Combined effects of cyclic/sustained bending loading and water immersion on the interface shear strength of carbon/ glass fiber reinforced polymer hybrid rods for bridge cable, Construct. Build. Mater. 314 (2022) 125587, https://doi.org/10.1016/j.conbuildmat.2021.125587. [24] A. Uthaman, H.M. Lal, S. Thomas, Fundamentals of silver nanoparticles and their toxicological aspects, in: H.M. Lal, et al. (Eds.), Polym. Nanocomposites Based Silver Nanoparticles, first ed., Springer Nature Switzerland AG, Cham, 2021, pp. 1e24, https://doi.org/10.1007/978-3-030-44259-0_1. [25] H.M. Lal, S. Thomas, T. Li, H.J. Maria, Polymer Nanocomposites Based on Silver Nanoparticles: Synthesis, Characterization and Applications, Springer Nature Switzerland AG, Cham, 2021. [26] H.M. Lal, A. Uthaman, S. Thomas, Silver Nanoparticle as an Effective Antiviral Agent, in: H.M. Lal, T. Li, H.J. Maria (Eds.), Polymer Nanocomposites Based on Silver Nanoparticles, Engineering Materials, Springer, Cham, 2021, https://doi.org/ 10.1007/978-3-030-44259-0_10. [27] A. Iregui, L. Irusta, O. Llorente, L. Martin, T. Calvo-Correas, A. Eceiza, et al., Electrospinning of cationically polymerized epoxy/polycaprolactone blends to obtain shape memory fibers (SMF), Eur. Polym. J. 94 (2017) 376e383, https://doi.org/ 10.1016/J.EURPOLYMJ.2017.07.026. [28] C. Krishna, S. Pillai, C.P. Sharma, Review paper: absorbable polymeric surgical sutures: chemistry, production, properties, biodegradability, and performance, J. Biomater. Appl. 25 (2010) 291e366, https://doi.org/10.1177/0885328210384890. [29] New Plastic Products | Technologies for Processing Plastic (plastemart.com), Biocompatible Shape Memory Polymers in medical applications. atozplastics.com/PrintFile.asp? REF¼/webtech/upload/literature/Biocompatible-Shape-Memory-Polymers-SMPmedical-applications.asp& 1/2 1930:1e2. atozplastics.com/PrintFile.asp?REF¼/ webtech/upload/literature/Biocompatible-Shape-Memory-Polymers-SMP-medicalapplications.asp& 1/2. [30] L.B. Bottaro, Bone 23 (2008) 1e7, https://doi.org/10.1016/j.jconrel.2012. 05.021.Electrospun. [31] S. Mondal, Temperature responsive shape memory polyurethanes, Polym. Technol. Mater. 00 (2021) 1e28, https://doi.org/10.1080/25740881.2021.1906903. [32] N.C. Hodgson, R.A. Malthaner, T. Ostbye, The search for an ideal method of abdominal fascial closure: a meta-analysis, Ann. Surg. 231 (2000) 436e442, https:// doi.org/10.1097/00000658-200003000-00018. [33] A. Lendlein, R. Langer, Biodegradable , elastic shape-memory polymers for potential biomedical applications, Science (2013) 1673, https://doi.org/10.1126/ science.1066102. [34] W.M. Huang, C.L. Song, Y.Q. Fu, C.C. Wang, Y. Zhao, H. Purnawali, et al., Shaping tissue with shape memory materials, Adv. Drug Deliv. Rev. 65 (2013) 515e535, https://doi.org/10.1016/j.addr.2012.06.004.

Shape memory polymers as sutures

281

[35] A. Toncheva, F. Khelifa, Y. Paint, M. Voué, Fast IR-actuated shape-memory polymers using in situ silver nanoparticle-grafted cellulose nanocrystals, ACS Appl. Mater. Interfaces 10 (2018), https://doi.org/10.1021/acsami.8b10159. [36] A. Biswas, A.P. Singh, D. Rana, V.K. Aswal, P. Maiti, Biodegradable toughened nanohybrid shape memory polymer for smart biomedical applications, Nanoscale 10 (2018) 9917e9934, https://doi.org/10.1039/c8nr01438h. [37] Y. Bai, C. Jiang, Q. Wang, T. Wang, A novel high mechanical strength shape memory polymer based on ethyl cellulose and polycaprolactone, Carbohydr. Polym. 96 (2013), https://doi.org/10.1016/j.carbpol.2013.04.026. [38] X. Wu, W.M. Huang, Y. Zhao, Z. Ding, C. Tang, J. Zhang, Mechanisms of the shape memory effect in polymeric materials, Polymers (2013) 1169e1202, https://doi.org/ 10.3390/polym5041169. [39] J. Hartman, W. Small, T.S. Wilson, J. Brock, P.R. Buckley, W.J. Benett, Technical note embolectomy in a rabbit acute arterial, Technology (2007) 872e874. [40] A. Bhatnagar, S. Wagh, B. Singh, A. Rr, F. Khan, Smart materials-a review, Ann. Dent. Spec. 4 (2016). [41] S. Houshyar, A. Bhattacharyya, A. Khalid, A. Rifai, C. Dekiwadia, G.S. Kumar, et al., Multifunctional sutures with temperature sensing and infection control, Macromol. Biosci. 2000364 (2021) 1e11, https://doi.org/10.1002/mabi.202000364.

CHAPTER 13

Drug release kinetics of sutures Neethu Ninan Clinical and Health Sciences, University of South Australia, Adelaide, South Australia, Australia

13.1 Introduction Drug release is a vital characteristic of a therapeutic system controlled by dissolution, diffusion or both, by which drug solutes migrate from the therapeutic system to the outside release medium [1,2]. The ultimate aim of the therapeutic system is to sustain the concentration of the drug in the target tissues or blood at a required value, having control over the duration and release rate of the drug. The purpose of release kinetics is to regulate the drug level in the target tissue or blood within the therapeutic window, amid the minimum toxic concentration (MTC) and minimum effective concentration (MEC) (Fig. 13.1). Suppose the drug is taken as a single large dose, then the drug level is elevated above the MTC, causing critical toxic impacts and soon quickly drops below MEC. Multiple dosing after a particular interval may decrease the variation of drug levels in plasma but the patient may face serious issues. Hence, it is required to design therapeutic systems that deliver sustained or controlled release of a drug with a small dosing frequency. In a controlled release system, a fraction of the dose is released initially to attain the MEC, called burst release [3]. Then the level

Figure 13.1 Release kinetics aims to control the drug level in the target tissue or blood within the therapeutic window, between MTC and MEC. (Reproduced from Ref. [2] with permission from American Chemical Society, 2011.) Advanced Technologies and Polymer Materials for Surgical Sutures ISBN 978-0-12-819750-9 https://doi.org/10.1016/B978-0-12-819750-9.00010-3

© 2023 Elsevier Ltd. All rights reserved.

283

284

Advanced Technologies and Polymer Materials for Surgical Sutures

is maintained for the desired period. Thus, the drug is released slowly for an extended time frame. Drug release is affected by several driving forces including release medium, properties of the drug solutes, nature of the drug carrier or the therapeutic system, etc [4]. Mathematical modeling intends to make the complex drug release process simpler and to make us understand the release mechanisms of a therapeutic system (Fig. 13.2). Thus, a mathematical model is based on one or two leading driving forces. There are various mathematical models for drug release including zero-order kinetics, firstorder kinetics, Hixson Crowell model, Korsmeyer Peppas model, Higuchi model, etc (Table 13.1) [11]. A zero-order kinetic model is used for drugs that do not disaggregate and release the drug slowly, as in the case of transdermal systems or low soluble drugs in coated forms [12]. In first-order release kinetics, the drug release is proportional to the amount of remaining drug and is related to water-soluble drugs [13]. In Higuchi kinetics, drug release is proportional to the square root of time and is applicable for both

Figure 13.2 Different release kinetic models for the release of curcumin. (Reproduced from Ref. [5], with permission from Elsevier, 2021.)

Drug release kinetics of sutures

285

Table 13.1 Release kinetics model and their equations. Release kinetics model

Equation

Abbreviations

Zero-order kinetics

Ft ¼ K0t

First order kinetics

lnQt ¼ lnQ0 þ Qt t

Higuchi kinetics

Ft ¼ Kht1/2

Hixson Crowell kinetics

1/3 ¼ Kst W1/3 0 Wt

Korsmeyer Peppas kinetics

Mt/M2 ¼ Ktn

Ft is a fraction of drug dissolved in time t, k0 is zero-order release constant [6] Q0 is the drug released at zero hours, Qt is the drug released at time t, Kt is the first-order release constant [7] Ft is a fraction of drug released at time t, Kh is the Higuchi release rate constant [8] W0 is the initial amount of drug present in the matrix, Wt is the amount of drug released at time t, and Ks is the drug release constant [9] M1 is the amount of drug released at time t, M2 is the amount of drug released at infinite time, K is release rate constant, n is release exponents [10]

water-soluble and less soluble drugs in solid or semisolid matrixes (diffusion matrix systems) [9]. In Hixson Crowell’s kinetics, the release rate is restricted by the dissolution rate of drug particles and not affected by diffusion across matrix such as in erodible matrix formulations [14]. Korsmeyer Peppas model is applicable for swellable polymeric devices where many release phenomena could be involved [10].

13.2 Surgical sutures Surgical sutures are devices used by medical professionals to close the wound after surgery. There are various types of sutures, namely, absorbable, nonabsorbable, ophthalmic, cardiovascular, veterinary, dental, etc., [15]. Drug-eluting sutures are fabricated using electrospinning, dip coating, grafting, etc [16]. Based on the drugs incorporated, they can show antiinflammatory effects, antibacterial effects, antithrombotic effects, and also accelerate wound healing. They can provide the required therapeutic concentration over a long time without creating toxicity by their sustained release profile (Fig. 13.3). The biggest trial in the manufacture of drug-

286

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 13.3 The concept of a drug-eluting suture. (Reproduced from Ref. [17], with permission from Elsevier, 2014.)

eluting sutures is to attain the needed concentration and desired impact of the drug without losing the mechanical properties of the suture, which can be attained with faster polymer degradation and precise release approaches. This chapter discusses the recent progress in drug release kinetics of sutures loaded with various kinds of drugs.

13.3 Drug release from antiinflammatory sutures Pain from the injured site is an unavoidable consequence for patients needful of stitches, after surgery or major wounds [18]. To combat this pain, nonsteroidal antiinflammatory drugs (NSAIDs) are often provided [19]. Nevertheless, these drugs are not administered during the entire wound healing process, to decrease their systemic exposure. Hence, they do not show any effects throughout the whole wound healing window. Recently, there has been much research on the manufacture of antiinflammatory sutures (drug-coated sutures and drug embedded sutures), which would

Drug release kinetics of sutures

287

provide pain relief locally and improve wound healing. Drug-coated sutures are obtainable for sale with good biodegradability, high mechanical strength, and low tissue reaction [20]. Nevertheless, their main disadvantage is the trouble in achieving competent drug loading and poor release kinetics. To subjugate this challenge of surface coating, another rational option is to integrate drugs within the polymeric matrix of the suture. There are several techniques to load drugs within sutures such as dip coating [21], supercritical carbon dioxide impregnation [22], melt spinning [23], electrospinning [24], nanoparticle-controlled drug delivery [21], layer-bylayer technique [25], etc. The supercritical carbon dioxide impregnation process is a method by which drugs and polymers are pressurized in a vessel using carbon dioxide [26]. Here carbon dioxide is used as a plasticizer. This technique can be used to incorporate ketoprofen (antiinflammatory drug) in poly-L-lactic acid (PLLA) fibers. The process enabled possible tuning of the release profile of ketoprofen by varying temperature and pressure [22]. The release of ketoprofen was influenced by the free volume of the polymer matrix and the polymer’s degradation rate. When more ketoprofen is loaded in the matrix, the degradation rate of PLLA is increased leading to a higher release of ketoprofen. This is because the carboxylic acid groups in the drug, catalyze the hydrolysis of ester bonds of polymer. Various PLLA sutures that could release ketoprofen from 3 days to 3 months were fabricated using this process. Compressed carbon dioxide can also be used as a solvent to swell poly-lactic-co-glycolic acid (PLGA) sutures and load ketoprofen [27]. The study showed that the drug attached to the suture surface was easily released compared to the embedded drugs which were released only after the suture dissolved. Melt spinning is used to synthesize commercial sutures in which polymers are processed at high temperatures. Melt extruded polyethylene glycol (PEG)/polycaprolactone (PCL)/chitosan/keratin sutures loaded with diclofenac sodium demonstrated rapid and sustained release of drug [23]. The intention of hydrophilic/phobic polymers was to effectively improve the drug solubility. For instance, lipophilic ingredients of drugs show increased solubility in hydrophobic polymer while hydrophilic ingredients of drugs dissolve in a hydrophilic polymer. Also, as the amount of drugloaded increased, the drug release rate augmented due to the formation of cavities in the polymer matrix at higher drug loading. PCL sutures loaded with diclofenac or nanohybrids of diclofenac intercalated with synthetic hydrotalcite were also synthesized by melt-spinning [17]. There was an

288

Advanced Technologies and Polymer Materials for Surgical Sutures

initial burst release of diclofenac in all the samples. The control samples released all the drugs in 14 days, whereas the sutures loaded with nanohybrids released the whole drug in 55 days. The drug release kinetics from the suture depends on drug dissolution and diffusion in the matrix polymer. However, using this technique, heat-sensitive drugs cannot be incorporated in the polymer blend as they get degraded at high temperatures. Electrospinning can overcome this issue to produce drug-loaded sutures using micro- or nanosized fibers. Menon et al. have fabricated electrospun yarn with a central core made of PLLA and a drug-eluting PLGA sheath incorporating drugs such as aceclofenac and insulin [24]. The release of aceclofenac in phosphate-buffered saline at 37 C showed an initial burst release within the first 24 h followed by a sluggish and sustained release for 10 days. Similarly, insulin showed a similar release pattern with a sustained and slow release for 4 days. These sutures were found to reduce cellularity and epidermal hyperplasia in an animal model with skin inflammation. In another study, diclofenac loaded strand of PLGA was braided with a clinically available surgical suture (VICRYLTM) [28]. This approach enabled sustained release of diclofenac which cannot be achieved by the normal dipping process. The suture was able to release diclofenac for 10 days without compromising its mechanical strength. Nanoparticle controlled drug delivery is another method to increase the antiinflammatory effect of the suture. At the site of inflammation, increased interstitial fluid flow and lymphatic drainage clear drugs quickly. However, the clearance of nano-sized drug-loaded particles takes a longer time compared to small molecule drugs. Along with that, polymeric nanocarriers facilitate sustained release of loaded drugs and the release rate can be tuned by changing component ratios. Again, nanoparticles functionalized with ligands can release drugs at target sites. Kim et al. used nanoparticle controlled drug delivery to attain sustained release, tissue retention for a longer period and selective delivery of drugs to macrophages to impart an antiinflammatory effect [21]. The suture was incorporated with mannose/ PEG/diclofenac nanoparticles by the dip-coating method. The polyvinyl alcohol (PVA) layer in which nanoparticles were incorporated facilitated their sustained release that could effectively reduce the immune reactions around the suture. The layer-by-layer technique is a convenient method to develop sutures with multifunctional films. The technique includes successive adsorption of oppositely charged groups in a precise manner on a substrate that is charged. The advantage of this technique is that films can be deposited on any

Drug release kinetics of sutures

289

complicated structure. Normally, the surfaces of sutures are hydrophobic and need to be surface modified by interfacial chemical reactions or plasma, before layer-by-layer deposition. Wang and coworkers have tried to directly deposit such films on hydrophobic surfaces without any surface modification [25]. They demonstrated that a microgel constituting polyallylamine hydrochloride and dextran can be deposited consecutively with hyaloplasm acid on sutures to develop multi-layer films. They further incorporated Ibuprofen (an antiinflammatory drug) in these multilayer films constructed on sutures. The driving force behind the loading of ibuprofen is the electrostatic interaction with protonated amine groups of microgel and carboxylate groups of ibuprofen along with hydrogen bond interactions between the two. There was the rapid release of the drug in the initial 4 h followed by a slow and sustained release in 10 days. Neointimal hyperplasia is the main cause of failure of bypass graft surgery of arteries, veins, and other prosthetics. Several factors cause this failure such as compliance mismatch, local blood vessel injury, the reaction of the wall of the blood vessel to the material of suture etc. The immune cells in the wall of the blood vessel secrete proinflammatory cytokines, mediators, metalloproteinase, adhesion molecules that will enhance the migration and proliferation of smooth muscle cells causing anastomosis narrowing. To prevent this proliferation, several strategies are used such as antiproliferative drugs, ionizing radiation, external stenting, gene therapy, etc. Tacrolimus (an immunosuppressant drug)-chitosan-coated sutures were used in vivo in a rat model to evaluate their potential in inhibiting neointimal hyperplasia [29]. The in vitro release studies for 1 month showed that there was no initial burst release of Tacrolimus. The release of Tacrolimus was quicker for coated sutures compared to chitosan-Tacrolimus sutures as chitosan aided in the prolonged release of Tacrolimus. Thus, antiinflammatory drug-loaded sutures with controlled-release behavior have lots of potential in alleviating inflammation at the site of injury.

13.4 Drug release from growth factor embedded sutures In surgery, impaired wound healing is a serious problem that can cause mortality and morbidity. The factors affecting impaired wound healing include diabetes mellitus, ischemia, immunosuppression, etc. Several approaches for accelerating wound healing are undergoing preclinical and

290

Advanced Technologies and Polymer Materials for Surgical Sutures

clinical assessment. In this scenario, the use of growth factors has been proven to promote cell growth, differentiation, and synthesis of proteins. However, the systemic application of growth factors often leads to unwanted systemic side effects. To vanquish these limitations, the growth factors are incorporated into the suture for effective wound healing. The controlled release of growth factors such as platelet-derived growth factor (PDGF), vascular endothelial growth factor (VEGF), basic fibroblast growth factor (bFGF) to wound sites is vital for wound healing and tissue regeneration. Electrospun multifunctional fibrous sutures were fabricated using PCL and collagen incorporated with bFGF to achieve controlled drug release and mechanical strength [30]. The release profile showed an initial burst release with a release rate of 4.3 ng/h in the first 24 h followed by a slow release of 2.1 ng/h from day 2 to day 21. The initial burst release was due to bFGF present on the surface of the suture, and the slow release was from bFGF buried inside the fibers. In another study, nanoparticle coated sutures were used for gene delivery for healing tendons [31]. Here, plasmids (bFGF and VEGF) and PLGA nanoparticles were mixed to form their complexes which were attached to polydopamine modified sutures. There was sustained release with no burst release of plasmids from these sutures. In an in vitro study, around 78% of genes loaded were released in 28 days. The release was mainly affected by the degradation of PLGA. A commercial suture material made of polydioxanone was coated with poly-L-lactide and VEGF [32]. The suture could release biologically effective VEGF within 5 days, whose potential was evaluated in vivo. In another study, electrospun core-sheath PLGA sutures were loaded with transforming growth factor-b1 (TGF-b1) for promoting wound healing [33]. PLGA microfibers were the core, and PLGA nanofibers prepared by electrospinning were the sheath covering the microfibers. The core yielded mechanical strength, whereas the sheath provided biocompatibility due to its similarity to the extracellular matrix. The sheath was able to absorb many drugs due to capillary action and kept them at interspace between sheath and core. The release profile showed an initial burst release of TGF-b1 within 24 h that stayed on the surface. This was followed by the slow release of TGF-b1 from the PLGA core-sheath till 7 days after which there was no release. The released TGF-b1 could synthesize collagen during wound healing. However, excess TGF-b1 can result in scar formation. This can be controlled by TGF-b inhibiting drug called Tranilast. In this context, Choi and coworkers made sutures loaded with Tranilast using electrospinning which was braided with clinically used

Drug release kinetics of sutures

291

PLGA-based surgical sutures [34]. There was an initial burst release of 64%, followed by sustained release of Tranilast in 14 days, due to the presence of PLGA outer layers that acted as a drug diffusion barrier. Thus, this section covered explanations for recent progress in the design of growth factor embedded sutures and the mechanism involved in their controlled release.

13.5 Drug release from antithrombotic sutures Vascular anastomosis is one of the applications of medical sutures in which two tubular structures are joined within the vascular system. Suture-based vascular anastomosis is an important part of both bypass surgery and vascular graft implantation. At the site of anastomosis, an injury to the vascular endothelium triggers a coagulation cascade, activating proinflammatory response causing several complications like thrombosis (clot formation) and stenosis (narrowing of blood vessels). Fabrication of bioactive antithrombotic sutures can be a possible solution to inhibit thrombosis. Such sutures also help in transporting tissues from one location to another in the case of various surgical fields including orthopedic, otolaryngology, oral, plastic, maxillofacial surgery etc. The main cause of graft failure in such tissue transportation is thrombotic occlusion. Drugs are used to prevent these complications but they often cause serious side effects. Hence, a different tactic to carry these drugs at the wound site in a sustained manner is essential. Heparin eluting sutures were fabricated by the electrostatic interaction of negatively charged heparin with positively charged electrospun nanofibers composed of PLGA, polyethylene oxides , poly(lactide-co-glycolide)-graft-polyethyleneimine (PgP) [35]. The release studies disclosed an initial burst release, followed by a sluggish and sustained release of heparin for 20 days (Fig. 13.4). The surface of these yarns was unaffected after 20 days demonstrating their mechanically robust nature. The strong electrostatic interaction was responsible for the slow release of heparin from the stents.

13.6 Drug release kinetics of antibacterial sutures Sutures can come in contact with bacteria growing in hair follicles or subcutaneous tissues after penetrating through the skin barrier. These bacteria develop biofilm on the suture surface leading to several infections. So, it is very important to fabricate antibacterial sutures especially in case of risky procedures involving placing a central venous catheter. Medical

292

Advanced Technologies and Polymer Materials for Surgical Sutures

Figure 13.4 Release of heparin in vitro from nanofiber yarns for the first (A) 8 h and for (B) 20 days (C) Field Emission Scanning Electron Microscope images of nanofiber yarns at 20 days postrelease (scale bars: 200 and 50 mm). (Reproduced from Ref. [35], with permission from American Chemical Society, 2018.)

sutures with antibacterial characteristics can efficiently prevent pathogens, thus circumventing the incidence of surgical site infection and decreasing the reoccurrence of patients ensuing in postoperative death [36,37]. Due to the increase in the development of bacterial resistance, research should focus on the development of sutures incorporated with various types of bactericides. Wu et al. developed a scalable antibacterial suture by braiding silk filaments with silk fibroin films impregnated with various percentages of Berberine (antimicrobial agent) [38]. Sustainable and slow drug release was observed for 7 days with an initial burst release after 12 h, followed by a stable release for 96 h. The release behaviors of the Berberine (drug) were

Drug release kinetics of sutures

293

controlled by its characteristics and the features of the drug incorporated matrix. The tighter packing of silk filaments, their molecular weight, longer diffusion path, and average pore size could retard the release rate of the drug (Fig. 13.5). Also, the drug release rate was higher in acidic conditions than alkaline conditions. Although there are no ionizable groups in Berberine, the upsurge in solubility instigated by phosphate can be due to its interaction with drugs, making it a more soluble complex [39]. Nanofiber yarns, the thread-like structures synthesized by bundling of nanofibers concurrently all through the electrospinning process have received attraction among researchers recently. These sutures not only secure tissues together but also deliver drugs locally at the site of the wound reducing infections at surgical sites. Electrospun curcumin loaded PLLA nanofiber yarns were found to provide a biomimetic milieu for drug encapsulation which increased the drug’s bioavailability [40]. They showed an initial burst release of curcumin in the first 12 h followed by sustained release with the highest release of 78% in 72 h. The released curcumin

Figure 13.5 Antibacterial mechanism of coated Silk filament/Berberine suture before and after coating. (Reproduced from Ref. [38], with permission from American Chemical Society, 2021.)

294

Advanced Technologies and Polymer Materials for Surgical Sutures

showed antibacterial, antiplatelet effects, and promoted cell migration in vitro. Antibacterial functionalization of sutures with nanoparticles is another promising approach. Silver nanoparticle (AgNPs) coating on sutures has been reported as an efficient antibacterial strategy due to their antibacterial and antiinflammatory effect [41e43]. Nylon monofilament surgical sutures were coated using in situ deposition of biogenic AgNPs [44]. The advantage of nylon monofilaments is that they show less susceptibility to pathogens at the surgical site. The release studies of nylon/AgNPs showed first-order release kinetics with the concentration-dependent release of Agþ. In another study, nonabsorbable silk sutures were incorporated with AgNPs (synthesized using microbes) [45]. Despite the higher Agþ release from suture’s degradation, the amount of silver measured using ICP-MS was lower than the toxicity limits stated in the literature. Another tactic to cure postoperative infections is to tether antimicrobials on the surface of the suture through a hydrogel. Sutures were made from silk waste fibers through a five-loop method. Onto the surface of sutures, a hydrogel base made of aloe vera and Gum acacia were coated incorporated with antimicrobials and growth factors. These sutures demonstrated continuous release for 6 days with an initial burst release followed by stable release after 96 h. The natural hydrogel base enabled continuous release in different pH conditions such as 6.3,6.6, and 7.7. It was found that release was highest in alkaline pH [46]. Lee et al. introduced an electronic suture to monitor suture integrity and promote tissue regeneration by activated drug release [47]. Its thermoresponsive layer consisting of a grafted layer of PVA onto poly(Nisopropyl acrylamide) aided in the release of the drug on demand via Joule heating. The potential of the suture was evaluated both in vitro and ex vivo. Vicryl Plus (polyglactin suture coated with triclosan) has been used to treat infections at surgical sites. However, the side effects and toxicity of triclosan are a huge concern. So, chitosan-coated Vicryl Plus incorporating gentamycin (antibacterial drug) has been investigated in vivo. There was an initial burst release followed by a slow release of Gentamycin in 144 h [48]. In another study, polypropylene sutures grafted with acrylic acid were loaded with the antibiotic drug, Vancomycin. The sutures sustained the release of Vancomycin for 2 h without any burst release [49]. Antimicrobial sutures have gained importance in ocular surgery. Universally, more than 12 million ocular surgeries in a year use conventional nylon sutures for treating ocular wounds. In keratoplasty, these nylon

Drug release kinetics of sutures

295

sutures stay in the eyes for 1e2 years, which leads to bacterial infections. Thus, it is essential to provide for local antibacterial effect while suturing implants in the eyes. An antibacterial suture should be fine (20e50 mm) with high strength for a longer period. Kunal et al. developed sutures made of PCL and levofloxacin (ophthalmic antibiotic). Its pharmacokinetics was evaluated in a rat model of bacterial keratitis [50]. There were high levels of antibiotics in harvested aqueous humor and cornea in the first 15 min. The presence of antibiotics was detected after 30 days at reduced levels. In another study, polyglycolide-b-poly(glycolide-co-trimethylene carbonateco-ε-caprolactone)-b-polyglycolide based single filaments were incorporated with chlorhexidine (antiseptic) and biguanide (antiseptic) [51]. The drugs were embedded separately on the surface of the coating and copolymer used for coating. The release characteristics of both drugs were different. In the case of biguanide, the coating could prevent burst release. On the other hand, chlorhexidine was released at a fast pace and the rate was affected by the amount of drug loaded. Some of the drugs are retained on the suture surface due to strong interactions with the matrix polymer. Antibacterial sutures were also synthesized using braided silk coated with PCL and levofloxacin hydrochloride (antibiotic) using two coating processes [52]. Slow and continuous drug release was observed in 5e6 days with an initial burst release in the first 12 h, followed by stable release after 96 h. The rate of drug release was higher in an alkaline environment than in an acidic environment [52]. Employing layer by layer assembly, chitosan and hyaluronic acid were applied on the surface of nylon monofilament sutures to impart good antibacterial and frictional properties [53]. An anionic and cationic drug models Acid Blue 80 and Astrazon blue was loaded in these sutures. Singlelayer coated nylon filament could release 80% of drugs after 24 h due to the repulsion effect. The nylon filament coated with chitosan and hyaluronic acid allowed the sustained release of drugs for 7 days due to strong ionic interaction between polymers and drugs. Thus, the controlled release of antibiotics from sutures helps in sustaining the concentration of drugs within a required level, decreases toxic side effects and increases the bioavailability of the drug.

13.7 Oxygen release from sutures The rupture of blood vessels occurs in trauma and surgery, and subsequently, wound healing involves low levels of oxygen, nutrients, etc., in

296

Advanced Technologies and Polymer Materials for Surgical Sutures

the renewing tissue. Thus, oxygen is a critical element for wound healing due to its key role in cell signaling and metabolism [54]. Oxygen also helps in cell proliferation, immune response, collagen production, regeneration of tissues, etc. Recently photosynthetic sutures were synthesized for the distribution of oxygen and recombinant growth factors in wounds [55]. In this strategy, genetically engineered microalgae were incorporated in commercially available surgical threads. These photosynthetic cells used light to trigger water photolysis and produce oxygen upon light stimulation. These cells could also deliver growth factors (VEGF, SDF, PDGF) to the wound site. The oxygen production was visualized as gas bubbles formed on the suture and quantified within 14 days (Fig. 13.6). They found that a rise in oxygen tension was adequate to produce a physiological impact in hypoxic fibroblasts by decreasing the expression of hypoxiainducible growth factor (HIF-1a).

Figure 13.6 (A) Photosynthetic sutures incorporated with different concentrations of genetically engineered microalgae (B) The oxygen production was visualized as gas bubbles formed on the suture (C) Release of oxygen quantified for 14 days (D) Reduction of HIF-1a as a result of oxygen production. (Reproduced from Ref. [55], with permission from Elsevier, 2018.)

Drug release kinetics of sutures

297

Bowel anastomosis is another clinical issue associated with hypoxia leading to anastomotic failure in urology. Although additional systemic oxygen was found to raise tissue oxygen tension in perianastomotic tissue, there is no clear idea on the impact of local use of oxygen on healing such severe clinical consequences related to bowel anastomosis. In this context, oxygen-producing biomaterials have attained lots of interest. Inglin et al. have developed oxygen releasing suture material incorporated with calcium peroxide nanocrystals and covered with a synthetic amino acid, poly (D, Llactide-co-glycolide) [56]. The release studies showed that a bulk of oxygen was produced within the first 24 h in contact with water. In 72 h, 90 mL of oxygen was released by the coated suture. Thus, these sutures could increase the level of oxygen in hypoxic conditions, leading to higher tissue oxygen saturation and mechanical integrity in rat models with colonic anastomosis. Thus, oxygen releasing sutures with controlled-release behavior can promote chronic wound healing.

13.8 Conclusion Drug release kinetic models that precisely envisage the release kinetics of drugs from sutures would significantly improve our understanding of drugloaded sutures and empower quick optimization of such therapeutic systems. Despite the increased level of complexity, real advancement has been made in trying to recognize the mechanisms leading to the release of drugs from drug-eluting sutures and also in recognizing the binding and transport properties of drugs at target tissues. Both in vitro and in vivo models are used to understand the drug release kinetics of sutures. Some researchers have tried to understand the complex binding process of drug and matrix while others have focused on the effect of driving forces such as diffusion, dissolution, etc. While one-dimensional models can give a rough idea about the problem, three-dimensional models give insight into the full complex geometry of suture and its surroundings. Further investigations are needed to develop improved suture platform strategies for optimized drug delivery and ideal pharmacokinetic profile.

References [1] J. Li, D.J. Mooney, Designing hydrogels for controlled drug delivery, Nat. Rev. Mater. 1 (12) (2016) 16071. [2] N. Kamaly, B. Yameen, J. Wu, O.C. Farokhzad, Degradable controlled-release polymers and polymeric nanoparticles: mechanisms of controlling drug release, Chem. Rev. 116 (4) (2016) 2602e2663.

298

Advanced Technologies and Polymer Materials for Surgical Sutures

[3] J.H. Lee, Y. Yeo, Controlled drug release from pharmaceutical nanocarriers, Chem. Eng. Sci. 125 (2015) 75e84. [4] J.K. Patra, G. Das, L.F. Fraceto, E.V.R. Campos, M.d.P. Rodriguez-Torres, L.S. Acosta-Torres, L.A. Diaz-Torres, R. Grillo, M.K. Swamy, S. Sharma, S. Habtemariam, H.-S. Shin, Nano based drug delivery systems: recent developments and future prospects, J. Nanobiotechnol. 16 (1) (2018) 71. [5] D. Sharma, B.K. Satapathy, Understanding release kinetics and collapse proof suture retention response of curcumin loaded electrospun mats based on aliphatic polyesters and their blends, J. Mech. Behav. Biomed. Mater. 120 (2021) 104556. [6] Z. Khan, S.A. Al-Thabaiti, F.M. Al-Nowaiser, A.Y. Obaid, A.O. Al-Youbi, M.A. Malik, Kinetics of silver nanoparticle growth in aqueous polymer solutions: 1st nano update, Arab. J. Chem. 5 (4) (2012) 453e459. [7] H. Chavda, C. Patel, Chitosan superporous hydrogel composite-based floating drug delivery system: a newer formulation approach, J. Pharm. BioAllied Sci. 2 (2) (2010) 124e131. [8] T.K. Giri, K. Kumar, A. Alexander, Ajazuddin, H. Badwaik, D.K. Tripathi, A novel and alternative approach to controlled release drug delivery system based on solid dispersion technique, Bull. Fac. Pharm. Cairo Univ. 50 (2) (2012) 147e159. [9] J. Siepmann, N.A. Peppas, Higuchi equation: derivation, applications, use and misuse, Int. J. Pharm. 418 (1) (2011) 6e12. [10] C.S. Brazel, N.A. Peppas, Modeling of drug release from swellable polymers, Eur. J. Pharm. Biopharm. 49 (1) (2000) 47e58. [11] S. Dash, P.N. Murthy, L. Nath, P. Chowdhury, Kinetic modeling on drug release from controlled drug delivery systems, Acta Pol. Pharm. 67 (3) (2010) 217e223. [12] K.T. Savjani, A.K. Gajjar, J.K. Savjani, Drug solubility: importance and enhancement techniques, ISRN Pharm. 2012 (2012) 195727. [13] S. Loew, A. Fahr, S. May, Modeling the release kinetics of poorly water-soluble drug molecules from liposomal nanocarriers, J. Drug Deliv. 2011 (2011) 376548. [14] D.J. Hines, D.L. Kaplan, Poly(lactic-co-glycolic) acid-controlled-release systems: experimental and modeling insights, Crit. Rev. Ther. Drug Carrier Syst. 30 (3) (2013) 257e276. [15] A.L. Tajirian, D.J. Goldberg, A review of sutures and other skin closure materials, J. Cosmet. Laser Ther. 12 (6) (2010) 296e302. [16] M. Champeau, J.-M. Thomassin, T. Tassaing, C. Jérôme, Current manufacturing processes of drug-eluting sutures, Expet Opin. Drug Deliv. 14 (11) (2017) 1293e1303. [17] O. Catanzano, S. Acierno, P. Russo, M. Cervasio, M. Del Basso De Caro, A. Bolognese, G. Sammartino, L. Califano, G. Marenzi, A. Calignano, D. Acierno, F. Quaglia, Melt-spun bioactive sutures containing nanohybrids for local delivery of anti-inflammatory drugs, Mater. Sci. Eng. C 43 (2014) 300e309. [18] N. Ninan, A. Forget, V.P. Shastri, N.H. Voelcker, A. Blencowe, Antibacterial and anti-inflammatory pH-responsive tannic acid-carboxylated agarose composite hydrogels for wound healing, ACS Appl. Mater. Interfaces 8 (42) (2016) 28511e28521. [19] D. Narayanan, N. Ninan, R. Jayakumar, S.V. Nair, D. Menon, O-carboxymethyl chitosan nanoparticles for controlled release of non-steroidal anti-inflammatory drugs, Adv. Sci. Eng. Med. 6 (5) (2014) 522e530. [20] Y.-Y. Jo, H. Kweon, D.-W. Kim, M.-K. Kim, S.-G. Kim, J.-Y. Kim, W.-S. Chae, S.P. Hong, Y.-H. Park, S.Y. Lee, J.-Y. Choi, Accelerated biodegradation of silk sutures through matrix metalloproteinase activation by incorporating 4-hexylresorcinol, Sci. Rep. 7 (1) (2017) 42441. [21] H. Kim, B.H. Kim, B.K. Huh, Y.C. Yoo, C.Y. Heo, Y.B. Choy, J.-H. Park, Surgical suture releasing macrophage-targeted drug-loaded nanoparticles for an enhanced antiinflammatory effect, Biomater. Sci. 5 (8) (2017) 1670e1677.

Drug release kinetics of sutures

299

[22] M. Champeau, I.T. Coutinho, J.-M. Thomassin, T. Tassaing, C. Jérôme, Tuning the release profile of ketoprofen from poly(l-lactic acid) suture using supercritical CO2 impregnation process, J. Drug Deliv. Sci. Technol. 55 (2020) 101468. [23] X. Deng, M. Gould, M.A. Ali, Fabrication and characterisation of melt-extruded chitosan/keratin/PCL/PEG drug-eluting sutures designed for wound healing, Mater. Sci. Eng. C 120 (2021) 111696. [24] S. Padmakumar, J. Joseph, M.H. Neppalli, S.E. Mathew, S.V. Nair, S.A. Shankarappa, D. Menon, Electrospun polymeric coreesheath yarns as drug eluting surgical sutures, ACS Appl. Mater. Interfaces 8 (11) (2016) 6925e6934. [25] L. Wang, D. Chen, J. Sun, Layer-by-Layer deposition of polymeric microgel films on surgical sutures for loading and release of ibuprofen, Langmuir 25 (14) (2009) 7990e7994. [26] C. Potter, Y. Tian, G. Walker, C. McCoy, P. Hornsby, C. Donnelly, D.S. Jones, G.P. Andrews, Novel supercritical carbon dioxide impregnation technique for the production of amorphous solid drug dispersions: a comparison to hot melt extrusion, Mol. Pharm. 12 (5) (2015) 1377e1390. [27] R.D. Weinstein, K.R. Muske, S.-A. Martin, D.D. Schaeber, Liquid and supercritical carbon dioxide-assisted implantation of ketoprofen into biodegradable sutures, Ind. Eng. Chem. Res. 49 (16) (2010) 7281e7286. [28] B.K. Huh, B.H. Kim, S.-N. Kim, C.G. Park, S.H. Lee, K.R. Kim, C.Y. Heo, Y.B. Choy, Surgical suture braided with a diclofenac-loaded strand of poly(lactic-coglycolic acid) for local, sustained pain mitigation, Mater. Sci. Eng. C 79 (2017) 209e215. [29] K. Ak, E. Ak, O. Dericioglu, T. Canak, J. Akbuga, N. Ozkan, S. Cetinel, S. Isbir, S. Arsan, A. Cobanoglu, Tacrolimus-eluting suture inhibits neointimal hyperplasia: an experimental in vivo study in rats, Eur. J. Vasc. Endovasc. Surg. 53 (3) (2017) 431e437. [30] J. Hu, Y. Song, C. Zhang, W. Huang, A. Chen, H. He, S. Zhang, Y. Chen, C. Tu, J. Liu, X. Xuan, Y. Chang, J. Zheng, J. Wu, Highly aligned electrospun collagen/ polycaprolactone surgical sutures with sustained release of growth factors for wound regeneration, ACS Appl. Bio Mater. 3 (2) (2020) 965e976. [31] Y.L. Zhou, Q.Q. Yang, Y.Y. Yan, L. Zhang, Q.H. Wang, F. Ju, J.B. Tang, Geneloaded nanoparticle-coated sutures provide effective gene delivery to enhance tendon healing, Mol. Ther. 27 (9) (2019) 1534e1546. [32] C. Bigalke, F. Luderer, K. Wulf, T. Storm, M. Löbler, D. Arbeiter, B.M. Rau, H. Nizze, B. Vollmar, K.-P. Schmitz, E. Klar, K. Sternberg, VEGF-releasing suture material for enhancement of vascularization: development, in vitro and in vivo study, Acta Biomater. 10 (12) (2014) 5081e5089. [33] Z. Gu, H. Yin, J. Wang, L. Ma, Y. Morsi, X. Mo, Fabrication and characterization of TGF-b1-loaded electrospun poly (lactic-co-glycolic acid) core-sheath sutures, Colloids Surf. B Biointerfaces 161 (2018) 331e338. [34] S.Y. Choi, B.H. Kim, B.K. Huh, W. Jeong, M. Park, H.J. Park, J.-H. Park, C.Y. Heo, Y.B. Choy, Tranilast-delivery surgical sutures to ameliorate wound healing by reducing scar formation through regulation of TGF-b expression and fibroblast recruitment, J. Ind. Eng. Chem. 67 (2018) 469e477. [35] S. Bae, M.J. DiBalsi, N. Meilinger, C. Zhang, E. Beal, G. Korneva, R.O. Brown, K.G. Kornev, J.S. Lee, Heparin-eluting electrospun nanofiber yarns for antithrombotic vascular sutures, ACS Appl. Mater. Interfaces 10 (10) (2018) 8426e8435. [36] N. Ninan, M. Muthiah, N.A.B. Yahaya, I.-K. Park, A. Elain, T.W. Wong, S. Thomas, Y. Grohens, Antibacterial and wound healing analysis of gelatin/zeolite scaffolds, Colloids Surf. B Biointerfaces 115 (2014) 244e252.

300

Advanced Technologies and Polymer Materials for Surgical Sutures

[37] T.A. Gillam, C.K. Goh, N. Ninan, K. Bilimoria, H.S. Shirazi, S. Saboohi, S. AlBataineh, J. Whittle, A. Blencowe, Iodine complexed poly(vinyl pyrrolidone) plasma polymers as broad-spectrum antiseptic coatings, Appl. Surf. Sci. 537 (2021) 147866. [38] Q. Wu, C. He, X. Wang, S. Zhang, L. Zhang, R. Xie, Y. Li, X. Wang, Z. Han, Z. Zheng, G. Li, Sustainable antibacterial surgical suture using a facile scalable silkfibroin-based berberine loading system, ACS Biomater. Sci. Eng. 7 (6) (2021) 2845e2857. [39] S.K. Battu, M.A. Repka, S. Maddineni, A.G. Chittiboyina, M.A. Avery, S. Majumdar, Physicochemical characterization of berberine chloride: a perspective in the development of a solution dosage form for oral delivery, Am. Assoc. Pharm. Sci. 11 (3) (2010) 1466e1475. [40] A.S. Richard, R.S. Verma, Bioactive nano yarns as surgical sutures for wound healing, Mater. Sci. Eng. C 128 (2021) 112334. [41] L.E. González García, M.N. MacGregor, R.M. Visalakshan, N. Ninan, A.A. Cavallaro, A.D. Trinidad, Y. Zhao, A.J.D. Hayball, K. Vasilev, Self-sterilizing antibacterial silverloaded microneedles, Chem. Commun. 55 (2) (2019) 171e174. [42] N. Ninan, B. Joseph, R.M. Visalakshan, R. Bright, C. Denoual, P. Zilm, Y.B. Dalvi, P.V. Priya, A. Mathew, Y. Grohens, N. Kalarikkal, K. Vasilev, S. Thomas, Plasma assisted design of biocompatible 3D printed PCL/silver nanoparticle scaffolds: in vitro and in vivo analyses, Mater. Adv. 2 (20) (2021) 6620e6630. [43] N. Ninan, N. Goswami, K. Vasilev, The impact of engineered silver nanomaterials on the immune system, Nanomaterials 10 (5) (2020) 967. [44] D.M. Syukri, O.F. Nwabor, S. Singh, S.P. Voravuthikunchai, Antibacterial functionalization of nylon monofilament surgical sutures through in situ deposition of biogenic silver nanoparticles, Surf. Coating. Technol. 413 (2021) 127090. [45] T. Baygar, N. Sarac, A. Ugur, I.R. Karaca, Antimicrobial characteristics and biocompatibility of the surgical sutures coated with biosynthesized silver nanoparticles, Bioorg. Chem. 86 (2019) 254e258. [46] H. Kalita, A. Hazarika, S. Kalita, R. Kandimalla, R. Devi, Antimicrobials tethering on suture surface through a hydrogel: a novel strategy to combat postoperative wound infections, RSC Adv. 7 (52) (2017) 32637e32646. [47] Y. Lee, H. Kim, Y. Kim, S. Noh, B. Chun, J. Kim, C. Park, M. Choi, K. Park, J. Lee, J. Seo, A multifunctional electronic suture for continuous strain monitoring and ondemand drug release, Nanoscale 13 (2021) 18112e18124. [48] Y. Yang, S.-B. Yang, Y.-G. Wang, S.-H. Zhang, Z.-F. Yu, T.-T. Tang, Bacterial inhibition potential of quaternised chitosan-coated VICRYL absorbable suture: an in vitro and in vivo study, J. Orthop. Translat. 8 (2017) 49e61. [49] M. García-Vargas, C. González-Chomón, B. Magariños, A. Concheiro, C. AlvarezLorenzo, E. Bucio, Acrylic polymer-grafted polypropylene sutures for covalent immobilization or reversible adsorption of vancomycin, Int. J. Pharm. 461 (1) (2014) 286e295. [50] K.S. Parikh, R. Omiadze, A. Josyula, R. Shi, N.M. Anders, P. He, Y. Yazdi, P.J. McDonnell, L.M. Ensign, J. Hanes, Ultra-thin, high strength, antibiotic-eluting sutures for prevention of ophthalmic infection, Bioeng. Trans. Med. 6 (2) (2021) e10204. [51] Y. Márquez, T. Cabral, A. Lorenzetti, L. Franco, P. Turon, L.J. Del Valle, J. Puiggalí, Incorporation of biguanide compounds into poly(GL)-b-poly(GL-co-TMC-co-CL)b-poly(GL) monofilament surgical sutures, Mater. Sci. Eng. C 71 (2017) 629e640. [52] X. Chen, D. Hou, L. Wang, Q. Zhang, J. Zou, G. Sun, Antibacterial surgical silk sutures using a high-performance slow-release carrier coating system, ACS Appl. Mater. Interfaces 7 (40) (2015) 22394e22403.

Drug release kinetics of sutures

301

[53] H. Mohammadi, F. Alihosseini, S.A. Hosseini, Improving physical and biological properties of nylon monofilament as suture by Chitosan/hyaluronic acid, Int. J. Biol. Macromol. 164 (2020) 3394e3402. [54] A. Forget, C. Staehly, N. Ninan, F.J. Harding, K. Vasilev, N.H. Voelcker, A. Blencowe, Oxygen-releasing coatings for improved tissue preservation, ACS Biomater. Sci. Eng. 3 (10) (2017) 2384e2390. [55] C. Centeno-Cerdas, M. Jarquín-Cordero, M.N. Chávez, U. Hopfner, C. Holmes, D. Schmauss, H.G. Machens, J. Nickelsen, J.T. Egaña, Development of photosynthetic sutures for the local delivery of oxygen and recombinant growth factors in wounds, Acta Biomater. 81 (2018) 184e194. [56] R.A. Inglin, L.E. Brügger, D. Candinas, B.S. Harrison, D. Eberli, Effect of oxygenproducing suture material on hypoxic colonic anastomoses in an experimental model, Br. J. Surg. 3 (6) (2019) 872e881.

Index ‘Note: Page numbers followed by “f ” indicate figures and “t” indicate tables.’

A Abrasion, 86 Abrasive resistance, 87e88 Abrasive testing apparatus, 88f Absorbable polymers, 96f, 101t Absorbable staples, 207 Absorbable sutures, 2e3, 47, 97e100, 150, 251 biological properties, 223 catgut sutures, 200, 238e239 chromic gut sutures, 239 definition, 200 functionalization antibiotics, 36e37 caprolactam sutures gentamicin/silver loaded, 38 chitin sutures, 37e38 drug-loading, 38e39 polyglycolic acid sutures, 36e37 silver, 37 oliglecaprone sutures, 239e240 polydioxanone sutures, 202, 239 polyglactin sutures, 240 polyglycolic acid (PGA), 201, 239 polytrimethylene carbonate (PTMC), 202 slowly, 240e243 synthetic, 200e201 vicryl (polyglactic acid), 201e202 Adipose-derived MSCs (ASC), 154e155 Agar-disk diffusion test, 165f Antibacterial sutures, 7, 225e226, 243e244 Antimicrobial sutures antimicrobial agent incorporation, 57e58, 57f bioactive glasses (BGs), 58 chitin, 57e58 disadvantage, 57 proteins, 59e60

silver nanoparticles (Ag NPs), 58e59 surface incorporation, 58e60 zwitterionic polymers, 57e58 bacterial attack, 56 Approximation loop, 77e78 Arduino-based microprocessor, 141f Azoles antimicrobial activity imidazolium group formation, 32e33 MeI treatment, 33 polypropylene suture functionalization, 32e33

B Bacterial cellulose nanocrystals (BCNC), 11 Bacterial cellulose nanocrystals/regenerated chitin (BCNC/RC), 226 Bacterial detection sensors, 143 Barbed sutures, 179e180, 253 Barb suture, 244f Berberine incorporated electrospun polyurethane fibers, 79 Bidirectional barbed sutures, 179e180, 180f Bioactive glasses (BGs), 58 Bioactive sutures, 101e103 antibacterial sutures, 115e116 bioactive growth factors, 116 bioactive implants, 149 cell based gene regulation, 160 growth factor bioactive suture, 160e163 mRNA suture, 159e160 stem cells. See Stem cell based bioactive sutures cell seeded biological sutures, 116e117 drug eluting sutures, 115e116 fabrication

303

304

Index

Bioactive sutures (Continued) cell and gene activators, 152 fiber level, 151e153 stimuli responsive, 152 incorporated material bioactive glass, 164e166 chitin bioactive sutures, 163e164 researched, 152e153 structure absorbable sutures, 150 monofilament sutures, 150 multifilament sutures, 151 nonabsorbable sutures, 150 surface architecture sutures, 166 tissue engineering and regenerative medicine strategies, 116e117 Bio-based sutures, 217e218 Biocompatibility. See also Cytotoxicity test cellular response, 257 definition, 255 measurement standardization, 258e259 usage tests, 258 in vitro tests, 257e258 in vivo tests, 258 mechanical stiffness, 257 polypropylene sutures, 257 Biodegradable sutures, 64e65 Bioglass sutures, variations, 165 Biohazard, 19 Bioinspired tissue adhesives, 113e114 Biological evaluation of medical devices, 100e101 Biological properties, 255 absorbable sutures, 223 nylon (monofilament) sutures, 222 PDS sutures, 222 polydiaxanone and polypropylene, 223 surgical site infections, 221 Biomedical materials, 46e47 Biomimetic sutures, 117e118 Biopolymers absorbable sutures, 2e3 biodegradation, 1e2 biomedical applications, 2e3 categories, 1e2, 2f

cellulose materials, 4 chitin, 4 collagen, 4 eco-friendly approach, 3e4 nanoparticles, 3e4 natural silk fibers, 4 nonabsorbable sutures, 2e3 poly(lactic acid) (PLA), 4 polyhydroxyalkanoates (PHA), 4 for sutures cellulose, 11e12 chitin and chitosan, 9 collagen, 4e5 polyhydroxyalkanoate (PHA), 10 polylactic acid (PLA), 5e6 silk, 6e8 Biosyn, 188 Black braided silk suture, 203 Blend electrospinning, 192 Bowel anastomosis, 297 Breaking strength, 82

C Capillarity, 85e86, 186 Caprosyn, 188 Carbon nanotubes (CNT), 133 CATGUT CHROMICTM, 51e52 Catgut sutures, 103e104, 200, 238e239 Cefotaxime sodium (CFX-Na), 107e109 Cell and gene activators, 152 Cellulose materials, 4, 11e12 Chemical initiators, 26 Chemically programmed polymer, 152 Chemical modification antimicrobial sutures, 56 chemical composition, 56 molecular conformation, 56 suture design, 56 Chitin sutures, 4, 9 absorbable suture functionalization, 37e38 bioactive biological characteristics, 163 chitosan effects, 164 chitsin suture, 163 coating process, 164

Index

histological analysis, 164 Chitosan-based sutures, 9 diacetyl chitin (DAC), 217 N-Acetyl-D-Glucosamine (GlcNAc), 216 properties, 216e217 two-step dip-coating process, 216e217 Vicryl, 217 Chlorhexidine, 61e62, 107e109, 109f Chromic gut sutures, 103e104, 239 loss of strength, 239 systematic treatment, 239 Chromium-molybdenum alloy, 206e207 Clinical evaluation, 118e119 Coated Vicryl suture, 106e107 Coaxial electrospinning, 192 Coefficient of friction, 186 Collagen, 4e5 Collagen-based sutures, 213e214 Collagen fibrils, 151 Collagen nanofibrils (CoNF), 4e5 Collagens, 131 Colorimetric smart sutures, 144e145, 144f Commercially available FDA approved absorbable sutures, 201f Commercial suture, 98te99t contraindications, 51e52 material choice, 50e51 MERSILENETM, 51 PERMA-HANDTM, 51 PRONOVATM, 51 Core-sleeve geometry, 177 Cotton, 132 Cyanoacrylate (CA)-based tissue adhesives, 259e260 Cytotoxicity test cell death, 259 cyanoacrylate (CA)-based tissue adhesives, 259e260 dose dependent, 260f ethyl-cyanoacrylate (ECA) and octylcyanoacrylate (OCA) wound closures, 260e261

305

graphene oxide-reinforced chitin monofilament absorbable surgical sutures, 259e260 membrane permeability test, 259

D Dacron 2-0 sutures, 188 Diacetyl chitin (DAC), 217 Diglycidyl ether of bisphenol A (DGEBA) epoxy resin, 269 Drug delivery sutures biodegradable sutures, 64e65 chlorhexidine, 61e62 controlled release, 60 dip-coated sutures, 61f direct/indirect coating method, 60e61 electrospinning, 62 erosion-controlled drug delivery systems, 64e65 grafting modification, 62e63, 63f grafting polymer incorporation, 63 heparin release, 63 levofloxacin, 61e62 localized release, 60 stimuli sensitive polymers, 64 triclosan-coated PLGA sutures, 61 Drug-eluting sutures, 115e116, 269, 285e286, 286f engineering aspects blend electrospinning, 191 dip coated sutures, 191 electrospinning, 192 fabrication, 191e192 melt extrusion, 192 polymer selection, 192e193 postprocessing techniques, 193 soaking, 192 Drug-loaded electrospun sheath, 192 Drug release kinetics antibacterial sutures, 291e295 antibacterial mechanism, 292e293, 293f electronic suture, 294 hydrogel base, 294 layer by layer assembly, 295 nanofiber yarns, 293 ocular surgery, 294e295

306

Index

Drug release kinetics (Continued) silver nanoparticle (AgNPs), 294 antiinflammatory sutures drug-coated sutures, 286e287 layer-by-layer technique, 288e289 melt spinning, 287e288 nanoparticle controlled drug delivery, 288 neointimal hyperplasia, 289 nonsteroidal antiinflammatory drugs (NSAIDs), 286e287 supercritical carbon dioxide impregnation proces, 287 antithrombotic sutures, 291 growth factor embedded sutures basic fibroblast growth factor (bFGF) release, 289e290 electrospun core-sheath PLGA sutures, 290e291 electrospun multifunctional fibrous sutures, 289e290 growth factors, 289e290 Higuchi kinetics, 284e285 mathematical model, 284e285 minimum effective concentration (MEC), 283e284 minimum toxic concentration (MTC), 283e284 model and equations, 285t oxygen release, 295e297 therapeutic system, 283e284 Drug-releasing electronic suture system (DRESS), 245e246

E Elasticity, 84, 186 Electrodeposition, 25e26 Electronic sutures, 245e246 Electrospinning process, 62, 192, 269e270 Electrospraying process, 269 Electrospun nanofibers, 117e118 Electrospun surgical sutures, 78 Engineering aspects, suture fabrication, 172f drug-eluting sutures blend electrospinning, 191

dip coated sutures, 191 electrospinning, 192 fabrication, 191e192 melt extrusion, 192 polymer selection, 192e193 postprocessing techniques, 193 soaking, 192 suture design absorbability, 181e184 biological attributes, 188e190 capillarity, 186 coefficient of friction, 186 design parameters, 172e173, 173f incision closure, 173e174 interdependence, 174f knot strength, 185 mechanical performance, 184e188 memory, 186 needle type, 177e179 orthopedic surgery, 173e174 pH influence, 189e190 physical attributes, 172e173, 181e188 pliability, 185e186 stiffness property, 185e186 structural attributes, 174e181 surface coatings, 181 surface design, 172e173 surface features, 179e181 suture configuration/geometry, 176e177, 176f suture size, 174e176, 175t tensile strength, 185 tissue responses and adhesions, 189 Entropic elasticity, 266e267 Erosion-controlled drug delivery systems, 64e65, 64f EthilonÒ, 203e205 Ethyl-cyanoacrylate (ECA), 260e261 Extracellular matrix (ECM), 54, 117e118 Extracted chitosan (EC), 9

F Fiber coating definition, 23 dip-coating, 24e25

Index

fiber-structured materials, 24 Fiber dimensions, 54e55 Functionalized sutures biocompatibility studies, 22e23 functionalization absorbable sutures, 36e39 electrodeposition, 25e26 fiber coating, 23e26 grafted sutures, 26e29 nonabsorbable sutures, 30e36 stimuli-responsive polymers, 29e30

G

Gellan-polylysine polyion complex fibers (GPF), 217e218 Gentamicin, 24e25 Glucose smart sutures carbon and CNT threads, 139e140 glucose levels monitoring, 139 glucose sensor, 139f in vivo studies, 140 Grafted sutures free-radical copolymerization, 26 functionalization, 26 hydrophobic grafted copolymers, 26e27, 27f initiators, 26 solvent, 28e29 temperature of reaction, 29 vinyl monomers grafting, 27e28 Graphene oxide-reinforced chitin monofilament absorbable surgical sutures, 259e260 Growth factor bioactive suture carrier, 162 coating design, 162 ex vivo analysis, 162 growth factors, 160 sterilization, 162 tendon repairs, 161 in vivo study, 162 Gut, 103e104

H 4-Hexylresorcinol (4HR), 104e105 Hixson Crowell’s kinetics, 284e285 Hydrogel, 136

307

Hydroxyl propyl trimethyl ammonium chloride (HACC) chitosan, 9

K Keratin, 131 Knotless barbed sutures, 179e180, 243, 245f Knot-pull tensile strength, 82e83 Knot strength, 81e82, 185 Korsmeyer Peppas model, 284e285

L Levofloxacin, 61e62 Light-triggered shape memory polymers (SMPs), 274 Low critical solution temperature (LCST), 30, 67e68

M Material based suture classification, 251e252 Mathematical model, 284e285 Maxon, 202 Mechanical properties abrasion, 86 abrasive testing, 87e88 antibacterial coating, 88e89 berberine incorporated electrospun polyurethane fibers, 79 bioactive photosynthetic suture, 79 breaking strength, 82 capillarity, 85e86 characterization techniques, 87e88 custom tensile testing, 79, 80f elasticity, 84 gentamicin/pluronic F127-silver/PCL nanofiber, 79 knot-pull tensile strength, 82e83 knot strength, 81e82 memory, 84e85 plasticity, 84 pliability, 85 short-term restoring, 79 tensile strength, 80e81 wound breaking strength, 83e84 Melt extrusion, 192 Melt spinning, 287e288

308

Index

Membrane permeability test, 259 Memory, 84e85, 186 MERSILENETM, 51 Mesoporous bioglass (MBG), 165 Microelectronics, 129 Microfluidic system analysis, 140e141 Micro & nanotechnology-enabled suture materials antibacterial surgical sutures, 225e226 drug release process, 224e225 electronic sutures, 224e225 gold microwires, 224e225 infections, 225 microfluidic spinning method, 225 polymeric nanocomposite, 226 smart bio-based sutures, 226e227 Minimum effective concentration (MEC), 283e284 Minimum toxic concentration (MTC), 283e284 MONOCRYLTM, 51e52 Monofilament sutures, 22, 100f, 150, 176e177, 244f, 252e253, 252f mRNA suture, 159e160 Mulberry silkworms, 4 Multifilament sutures, 22, 100f, 151, 176e177, 244f Multifunctional surgical bactericidal nanofibers, 96e97

N Nano-engineered suture systems, 228 Nanoparticle controlled drug delivery, 288 Natural polymers, 97e100 Natural silk fibers, 4 Natural sutures, 77e78, 250e251 Needle type, suture design anatomy, 177, 178f 3/8 circular needle, 178e179 conventional cutting needles, 178e179 5/8 needles, 178e179 physical characteristics, 177 reverse cutting needles, 178e179 square body geometry, 178e179 taper point ones, 178e179 Neointimal hyperplasia, 289

Neutrophil sensor, 143e144 Nonabsorbable polymers, 96f Nonabsorbable staples, 207 Nonabsorbable sutures, 2e3, 102t, 150, 240e241, 251e252 advantages and disadvantages, 202e203, 204t functionalization degree, 31f heart and blood vessels, 31 high energy source, 31 irritation/inflammation, 31 mechanical properties, 30 modified silk sutures, 35e36 polypropylene sutures, 31e35 modulus lines, 81f nylon, 203e205 polybutester, 206 polypropylene, 205e206 silk suture, 203, 241 stainless steel, 206e207 Novafil (Medtronic), 112, 206, 242 NUROLONTM, 52e53 Nylon, 110e111, 111f, 203e205, 241 Nylon 6,6, 110e111, 111f

O Octylcyanoacrylate (OCA), 47, 208, 260e261 Oliglecaprone sutures, 239e240 Origin based suture, 250e251 Oxidized cellulose, 11e12 Oxygen release, 295e297

P

PANACRYLTM, 52e53 PERMA-HANDTM, 51 PGA-PCL blend, 105e106 PGA-PLA blend, 106e109 Phase-separated linear block copolymers, 266e267 pH-responsive polymers, 30, 66e67, 67f pH sensor acidity levels, 135e136 alkaline environments, 135e136 coating method, 136 microfluidic approach, 136

Index

potentiometric approach, 136 swelling mechanism, 136 in vivo testing, 136e137 Physical configuration, 252e253 Physically perforating materials, 95 Physico-mechanical properties absorbable and nonabsorbable sutures, 218e219 knot strength and elongation, 220e221 mean displacement rate and abrasiveness, 221 polydioxanone, 218 polyglactin 910, 218e219 polyglyconate suture, 218 stiffness, 220e221 suture-tendon interface, 221 tensile properties, 219e221 tribological interactions, 218e219 Plain gut, 103e104 Plasticity, 84, 186 Platelet-derived growth factor-BB (PDGF-BB), 4e5 Pliability, 85, 185e186 Poliglecaprone suture, 239e240 Poly(3-hydroxybutyrate) (PHB), 10 Poly(4-hydroxybutyrate) (P4HB), 10, 109e110 Poly(ethylene terephthalate) (PET), 112 Poly(lactic acid) (PLA), 4 Poly(vinylidene fluoride) (PVDF), 113 Polyamide suture, 110e111 braided sutures, 212e213 coating process, 212e213 postdental extraction, 213 stereochemistry, 212 synthesis, 212 Poly(e-caprolactone) based sutures ethyl ester L-lysine triisocyanate incorporation, 211e212 layer-by-layer assembly approach, 211e212 multi-fluid electrospinning technique, 211 polymeric suture formulation, 211 yarn incorporation, 211e212 Polybutester, 112, 206, 242 Polybutylene terephthalate (PTFE), 113

309

Polycaprolactone (PCL), 131 Polydioxanone (PDS), 110, 201f, 202, 223, 239, 251 Polydioxanone (PDS II), 240 Polyesters, 241 Polyglactin 910, 201f, 218e219 Polyglactin sutures, 240 Polyglycerol sebacate (PGS), 131 Polyglycolic acid (PGA), 201, 239, 251 Polyglyconate, 240 Polyhydroxyalkanoate (PHA), 4, 10, 109e110 Polylactic acid (PLA) biocompatibility and cellular affinity, 6 degradation, 5e6 and polycaprolactone blends, 6 Polylactic-co-glycolic acid (PLGA), 270 Polymeric surgical sutures natural polymers gut, 103e104 poly(4-hydroxybutyrate) (P4HB), 109e110 polydioxanone (PDS), 110 silk, 104e105 synthetic and absorbable polymers PGA-PCL blend, 105e106 PGA-PLA blend, 106e109 synthetic and nonabsorbable polymers nylon, 110e111, 111f poly(ethylene terephthalate) (PET), 112 poly(vinylidene fluoride) (PVDF), 113 polybutester, 112 polybutylene terephthalate, 113 polypropylene (PP), 112 Polymerized caprolactum suture, 241 Polypropylene sutures, 112, 205e206, 223, 242 functionalization acrylonitrile (AN), 31e32 azoles, 32e33 gamma radiation, 32 gamma rays, 31e32 silver effect, 33e35 properties, 216

310

Index

Polytrimethylene carbonate (PTMC), 202 Polyurethane, 133 Polyurethane-based sutures healing efficiency, 214e215 shape memory polymers (SMPs), 273, 275e276 shape memory properties, 214 triple-layered vascular grafts, 215, 215f two-step shape memory properties, 214e215 Premarket Approval (PMA), 118e119 PROLENETM, 51e52 PRONOVATM, 51

Q Quill bidirectional barbed polydioxanone suture, 180e181

R Researched bioactive suture, 152e153 Resorbable smart sutures advantages, 141 electrical conductivity, 141e142 electrical resistance, 141e142 external power source, 142 silk, 141 silver, 141e142 Reverse cutting needles, 178e179

S Self-adjusting sutures, 69 Shape memory polymers (SMPs), 117 amorphous and crystalline areas, 268 applications, 266f auto-accelerating deterioration, 271 biodegradable, 271 biodegradation, 272 biomedical applications, 266e267, 269 cardiovascular applications, 271e272 clot microactuator, 277f compression process, 273e274 degradation speed, 271e272 drug-eluting properties, 276 drug-eluting sutures, 269 electrospin sutures, 270 endoscopic surgery, 273

entropic elasticity, 266e267 ethyl cellulose (EC) and PCL, 275 FeNiAl and CuZnAl alloys, 265e266 hard section, 266e267 hybrid smart PCL/CS suture, 277e278 internal tension, 267 light-triggered, 274 merits and demerits, 267t microheaters, 277 polyurethane based sutures, 273 polyurethane nanohybrid, 275 polyurethane sutures, 275e276 porous materials incorporation, 266e267 shape-changing effects, 268 shape-memory effect, 266e267, 270e271 shape recovery cycle, 267 silver nanoparticles, 269 Smart Surgical Suture design, 273 smart sutures, 276e277 SMP-Nitinol hybrid wire, 276 spring based, 273e274, 274f stimuli driven, 268t temporary shape, 265 thermal response, 265f Silk suture, 104e105 antibacterial silk sutures, 210e211 with antibiotics, 7 bacterial attachment, 6e7 biocompatibility, 211 biomechanical evaluation, 210e211 black braided silk suture, 203 braided silk sutures, 210 characteristics, 241 electron microscope images, 7f functionalization, 35e36 handling test, 210e211 knot-pull strength, 209 nonmulberry silk scaffold, 211 radio-opaque silk sutures, 209, 210f silver nanoparticles, 7e8 surface modification, 8 tetracycline hydrochloride (TCH) treated, 7 Silver absorbable sutures functionalization, 37

Index

polypropylene suture functionalization adhesion, 34 grafted PP monofilaments, 34e35 inhibition patterns, 35 intermediates, 34 ionic salts, 34 photoreduction, 34 silver nitrate, 33e34 Silver-cellulose hybrids, 3e4 Silver-doped bioactive glass powder (AgBG), 106e107 Silver nanoparticles (AgNPs), 3e4, 58e59, 107e109 Slowly absorbable sutures, 240e243 Smart bio-based sutures, 226e227 Smart Surgical Suture design, 273 Smart sutures, 117, 245e246 bacterial detection sensors, 143 base material, 130e133 carbon nanotubes (CNT), 133 cotton, 132 paper, 130e131 PCL/PGS blend, 132 polycaprolactone (PCL), 131 polyglycerol sebacate, 131 polyurethane, 133 wicking process, 133 categories and parameters, 130t colorimetric smart sutures, 144e145 glucose smart sutures carbon and CNT threads, 139e140 glucose levels monitoring, 139 glucose sensor, 139f in vivo studies, 140 microelectronics, 129 microfluidic analysis, 140e141 neutrophil sensor, 143e144 pH sensor acidity levels, 135e136 alkaline environments, 135e136 coating method, 136 microfluidic approach, 136 potentiometric approach, 136 swelling mechanism, 136 in vivo testing, 136e137 resorbable sutures, 141e142 strain smart sutures electrical resistance, 138 Fe3O4 spot, 137e138 strain and stress, 137

311

strain-controlled modulation, 138 strain sensor design, 138 wirelessly assess, 137e138 suture type, 129e130 temperature sensors accuracy test, 135 localized skin temperature abnormalities, 133e134 temperature sensitive suture, 133e134 tensile stress tests, 135 thermal based smart suture, 134e135 thread-based diagnostics, 134 variations tests, 135 vivo tests, 135 Smooth sutures, 253 Solvent, grafted sutures, 28e29 Spider silk chimeric proteins, 88e89 SpidrexÒ, 211 Stainless steel non-absorbable sutures, 206e207 Stainless steel sutures, 206e207 Staples, 115 absorbable, 207 nonabsorbable, 207 stapling devices, 207 tissue adhesives, 208 Stem cell based bioactive sutures cardiovascular application, 155e156 cell transplantation processes, 159 mesenchymal stem cells (MSCs), 153 tendon repair bone marrow-derived MSCs, 156e157 growth factors, 158 histological grading, 157e158 iliac crest bone marrow, 157 immobilization period, 158 woven suture, 157 wound healing adipose-derived MSCs (ASC), 154e155 bone marrow samples, 153 cell migration, 155 cell viability, 155 fibrin MSC spray, 154 human trials, 154 MSC passages, 153 Stem cell seeded suture, 244e245 Stiffness property, 185e186

312

Index

Stimuli-responsive polymers classification, 29e30 low critical solution temperature (LCST), 30 pH-responsive polymers, 30 swelling responsiveness, 30 Stimuli responsive sutures, 152 Stimuli responsive systems pH-responsive polymers, 66e67, 67f shape-memory characteristics, 69 smart grid networks, 66 swelling capacity, 65e66 thermo-responsive polymers, 67e68 Strain smart sutures electrical resistance, 138 Fe3O4 spot, 137e138 strain and stress, 137 strain-controlled modulation, 138 strain sensor design, 138 wirelessly assess, 137e138 Structurally coated and uncoated sutures, 242 Structural modification extracellular matrix (ECM), 54 fiber dimensions, 54e55 suture microstructure and topography, 55e56 Supercritical carbon dioxide impregnation proces, 287 Supramid, 241 Surgical site infections (SSI), 53e54, 129 Surgical sutures, 285e286 functions, 171 wound healing cascade, 171e172 Surgical wound closure, 171e172, 172f Suturable microfluidic systems, 140 Suture-based electrode approach, 140e141 Suture-based wound healing, 200 Suture materials absorbable sutures, 47, 200e202, 238e240 advanced surgical sutures, 249e250 biopolymers, 1e2 cellulose, 11e12 chitin and chitosan, 9 collagen, 4e5

polyhydroxyalkanoate (PHA), 10 polylactic acid (PLA), 5e6 silk, 6e8 animal-origin absorbable sutures, 20e21 antibacterial sutures, 243e244 biocompatibility, 255e259 biological properties, 255 braided suture, 238 characteristics, 47e50, 48t, 49f, 200, 253e255, 254f classification, 21t, 250e253, 250f material based, 251e252 origin based, 250e251 physical configuration, 252e253 size, 252 clinical evaluation, 118e119 commercial, 50e53 cytotoxicity, 259e261 deatraumatic sutures, 50 definition, 1 desirable and undesirable properties, 256t evolution, 48f function, 45e46 handling properties, 255 history, 46e47 knotless barbed sutures, 243 monofilament suture, 22, 49, 238 multifilament, 22, 49 nonabsorbable sutures, 202e207, 240e241 organism effect, 49e50 physical and mechanical properties, 254e255 preparation, 50 properties, 208f bio-based sutures, 217e218 biological properties, 208e209, 221e223 chemical properties, 208e209 chitosan-based sutures, 216e217 collagen-based sutures, 213e214 mechanical properties, 208e209 physical properties, 208e209 physico-mechanical properties, 218e221 polyamide suture, 212e213

Index

poly(e-caprolactone) based sutures, 211e212 polypropylene sutures, 216 polyurethane-based sutures, 214e215 silk-based sutures, 209e211 structural properties, 208e209 selection, 249e250 silk sutures, 20e21 smart sutures, 245e246 stem cell seeded suture, 244e245 sterilization, 12 structurally coated and uncoated sutures, 242 structural modification, 54e56 synthetic polymers, 1e2 synthetic sutures, 47 types, 21e22 wound closure, 249e250 Suturing technique, 77e78 Synthetic absorbable polymeric sutures, 200e201 Synthetic Body Fluid (SBF), 165 Synthetic nonabsorbable sutures, 251e252 Synthetic suture, 77e78 Synthetic suture polymers, 97e100

T Temperature sensors accuracy test, 135 localized skin temperature abnormalities, 133e134 temperature sensitive suture, 133e134 tensile stress tests, 135 thermal based smart suture, 134e135 thread-based diagnostics, 134 variations tests, 135 vivo tests, 135 TEMPO-mediated oxidation of regenerated cellulose (TORC), 11e12, 11f Tensile strength, 80e81 Tetracycline hydrochloride (TCH) treated sutures, 7 Therapeutic value, sutures

313

commercial sutures, 50e53 suture modification chemical modification, 56e60 drug delivery sutures, 60e65 stimuli responsive systems, 65e69 structural modification, 54e56 surgical site infections (SSI), 53e54 Thermal based smart suture, 134e135 Thermo-responsive polymers, 67e68 Thread-based diagnostics, 134 Tissue adhesive polymers advantages, 113e114 bioinspired tissue adhesives, 113e114 2-component injection procedure, 113e114 crosslinking, 113e114 cyanoacrylate glue, 114 sealant, 113e114 Tissue adhesives, 208 Traditional standard sutures, 179e180 Triclosan, 61, 88e89, 103 Triclosan-loaded L-lysine based nanogel grafted silk sutures (TLNGSs), 225

U Upper critical solution temperature (UCST), 68 Usage based suture categories, 243

V Vicryl (polyglactic acid), 188, 201e202

W Wicking process, 133 Wound breaking strength, 83e84 Wound closure devices, 1 Woven suture, 157

Y Yarn, 97

Z Zinc Bioglass, 165 Zwitterionic polymers, 57e58