Electrospun Polymeric Nanofibers: Insight into Fabrication Techniques and Biomedical Applications 3031314026, 9783031314025

Table of contents :
Preface
Contents
Recent Developments in Electrospinning Spinneret and Collector Assembly for Biomedical Applications
1 Introduction
2 Principle of Electrospinning Technique
3 Advances in Electrospinning Technique
4 Modification of Spinneret in Electrospinning Process
4.1 Multi-needle Electrospinning
4.2 Nozzle Free Electrospinning
5 Modification of Collector in Electrospinning Process
6 Conclusion
7 Future Trends
References
Fabrication of Multiscale Polymeric Fibres for Biomedical Applications
1 Introduction
2 Fabrication Techniques of Nano/Micro Size Electrospun Nanofibres
2.1 Coaxial, Emulsion and Co-electrospinning
2.2 Edge Electrospinning
2.3 Gap Electrospinning
2.4 3D Jet Writing
2.5 Caged Collector Electrospinning and Moving Spinneret
3 Applications of Multiscale Fibrous Scaffolds
3.1 Bone Tissue Engineering
3.2 Cartilage Tissue Engineering
3.3 Cardiovascular Tissue Engineering
3.4 Liver Tissue Engineering
3.5 Neural Tissue Engineering
3.6 Skin Tissue Engineering
3.7 Tendon Tissue Engineering
4 Conclusion
References
Techniques to Fabricate Electrospun Nanofibers for Controlled Release of Drugs and Biomolecules
1 Introduction
2 Electrospun Fiber Fabrication Techniques and Mechanism of Biomolecule Delivery
2.1 Electrospinning Process
2.1.1 Physical Adsorption
2.1.2 Covalent Immobilization
2.1.3 Blend Electrospinning
2.1.4 Coaxial Electrospinning
2.1.5 Emulsion Electrospinning
2.1.6 High Throughput Electrospinning
3 Application of Electrospun Nanofibers for Therapeutic Delivery
3.1 Transdermal and Wound Dressing
3.2 Drug Delivery Systems
3.3 Growth Factor Delivery System
3.4 Gene Therapy
References
New Prospects in Melt Electrospinning: From Fundamentals to Biomedical Applications
1 Introduction
2 Melt Electrospinning Configurations
2.1 Multi-Temperature Control Melt Electrospinning
2.2 Laser Melt Electrospinning
2.2.1 Spot Laser Beam Melt Electrospinning
2.2.2 Line Laser Beam Melt Electrospinning
2.3 Melt Coaxial Electrospinning
2.4 Needleless Melt Electrospinning
2.5 Gas-Assist Melt Electrospinning
2.6 Melt Electrospinning Writing
3 Biomedical Applications
3.1 Biosensors
3.2 Drug Delivery
3.3 Tissue Engineering
4 Conclusions
References
Centrifugal Spun Nanofibers and Its Biomedical Applications
1 Introduction
2 Fiber-Forming Systems
3 Principle of Centrifugal Spinning
4 Material and Machine Parameters Influencing the Fiber Formation
5 Application of Centrifugal Spun Nanofibers in Various Biomedical Applications
5.1 Tissue Engineering Applications of Centrifugal Spun Fibers
5.2 Drug Delivery Applications of Centrifugal Spun Fibers
5.3 Wound Dressing Applications of Centrifugal Spun Fibers
6 Advances in Centrifugal Spinning Process
7 Conclusion
8 Future Trends
References
Recent Advances in Electrospun Nanofibrous Polymeric Yarns
1 Introduction
2 Classification of Yarns Depending on Nanofiber Alignment
3 Electrospinning Strategies and Collector Designs for Nanofibrous Yarn Development
3.1 Disc-Shaped Collector
3.1.1 Electrospinning Towards the Center of the Disc
3.1.2 Electrospinning Towards the Edge of the Disc
3.2 Ring-Shaped Collector
3.3 Filament-Shaped Collector
3.4 Tube Collector
3.5 Cylindrical Static Collector
3.6 Glass Rod Collector
3.7 Friction Double Cylinder
3.8 Metal Frame
3.9 Liquid Bath as Collector System
3.10 Funnel-Shaped Collector
3.11 Hemispherical-Shaped Collector
4 Yarn Fabrication by Twisting Electrospun Membrane
5 Collector-Less Yarning Process
6 AC Electrospinning
7 Recent Advances in the Applications of Nanofibrous Yarns in Biomedicine
8 Conclusion
References
Fabrication of Textile-Based Scaffolds Using Electrospun Nanofibers for Biomedical Applications
1 Introduction
2 ELS NFs BMA
2.1 Wound Healing (WH) Properties
3 Musculoskeletal Complications
4 Cardiovascular Diseases (CVD)
5 Nephrology
6 Drug Delivery
7 Bone Regeneration (BR)
8 Gynaecology
9 Cancer Biology (CB)
10 Snakebite (SKB)
11 Neurology
12 Diabetes (DT)
13 Conclusion
References
Biomedical Applications of Electrospun Piezoelectric Nanofibrous Scaffolds
1 Introduction
2 Significance of Piezoelectric Polymers
2.1 Tissue Engineering Applications of Electrospun Piezoelectric Polymers
2.2 Piezoelectric Polymers-Based Self-Powered Implantable Biomaterials
3 Conclusion
References
Surface Modified Polymeric Nanofibers in Tissue Engineering and Regenerative Medicine
1 Introduction
2 Surface Modification of Polymeric Nanofibers
2.1 Hydrogel Coating
2.2 Chemical Treatment
2.3 Plasma Treatment
3 Conclusions and Future Outlook
References
Polymer/Ceramic Nanocomposite Fibers in Bone Tissue Engineering
1 Introduction
2 Polymer/Ceramic Composite Nanofibers in Bone Tissue Engineering
2.1 Silica-Based Ceramics
2.1.1 Bioactive Glass
2.1.2 Wollastonite (CaSiO3)
2.2 Calcium Phosphate (CaP)-Based Ceramics
2.2.1 Hydroxyapatite
2.2.2 Tricalcium Phosphate (TCP) (TCP; Ca3(PO4)2)
2.2.3 Tetracalcium Phosphate (Ca4(PO4)2O; TTCP)
2.2.4 Octacalcium Phosphate (OCP: Ca8(HPO4)2(PO4)45H2O)
2.3 Magnesium-Based Ceramics
2.3.1 Magnesium Silicate or Forsterite (Mg2SiO4)
2.3.2 Whitlockite (WH) (Ca18Mg2(HPO4)2(PO4)12)
2.3.3 Akermanite [AK] (Ca2MgSi2O7])
2.4 Carbon Nitride-Based Materials (C3N4)
2.5 Calcium Sulfate (CS)
2.6 Alumina (Al2O3)
3 Conclusion
References
Electrospun Fibrous Scaffolds for Cardiac Tissue Engineering
1 Introduction
2 Bioactive Nanoparticles Incorporated Nanofibers
3 Growth Factors/Cytokines Incorporated Nanofibers
4 Conductive Nanofibers
5 Conclusion
References
Electrospun Nanofibrous Scaffolds for Neural Tissue Engineering
1 Introduction
2 An Overview of Electrospinning
3 An Initiation to Tissue- Engineered Nerves
4 Electrospun Scaffolds for Neural Tissue Engineering
4.1 Natural Polymers
4.1.1 Collagen
4.1.2 Gelatin
4.1.3 Alginate
4.1.4 Chitosan
4.1.5 Silk
4.1.6 Miscellaneous
Keratin
Hyaluronic Acid
4.2 Synthetic Polymers
4.2.1 PVA
4.2.2 PLGA
4.2.3 PPy
4.2.4 PCL
4.2.5 PGS
4.2.6 Miscellaneous
PLA
PGA
PEDOT
4.2.7 Carbon-Based Polymers
Graphene
CNTs
5 Conclusion and Future Perspectives
References
External Stimuli Responsive Nanofibers in Biomedical Engineering
1 Introduction
2 External-Responsive Nanofibers
2.1 Thermo-Responsive Nanofibers
2.2 Magnetic-Responsive Nanofibers
2.3 pH-Responsive Nanofibers
2.4 Electrically-Responsive Nanofibers
2.5 Biomolecule-Responsive Nanofibers
2.6 Multi-responsive Nanofibers
3 Biomedical Applications of External-Responsive Nanofibers
3.1 Wound Dressings
3.2 Drug Delivery
3.3 Diagnosis
3.4 Scaffolds for Cell Culture and Delivery
4 Conclusions
References
Electrospun Antimicrobial Polymeric Nanofibers in Wound Dressings
1 Introduction
2 Antibacterial Constituent
2.1 Antibiotics
2.2 Nanofibers Loaded with Metal Nanoparticles
2.3 Plant Extracts
2.4 Biomacromolecules
3 Summary and Prospective
References
Application of Electrospun Polymeric Fibrous Membranes as Patches for Atopic Skin Treatments
1 Introduction
1.1 Challenges in Skin Treatment
1.2 Electrospinning
2 Electrospun Patches
2.1 Patches from Biodegradable Polymers
2.1.1 Poly(3-Hydroxybutyrate-Co-3-Hydroxyvalerate)
2.1.2 PHBV Fibers Blend with Evening Primrose Oil
2.1.3 Polycaprolactone
2.2 Patches from Non-biodegradable Polymers
2.2.1 PVB Membranes and Urea
2.2.2 PS and PA6 Composite Membranes
2.2.3 PI Membranes
2.2.4 PI Membranes with Chlorine
3 Summary
References
Nanofibrous Scaffolds for the Management of Periodontal Diseases
1 Introduction
2 Brief Overview of the Periodontium
3 Periodontal Diseases
4 Treatment Strategies
4.1 Traditional Treatment Strategies
4.2 Regenerative Strategies
5 Periodontal Regeneration and Nanofibrous Biomaterials
5.1 Nanofibrous-Occlusive Membranes
5.1.1 Drug Carrier Occlusive Membranes
5.1.2 Bioactive Nanoparticles-Containing Occlusive Membranes
5.1.3 Bilayered/Multilayered Occlusive Membranes
5.2 Nanofibrous Scaffolds (Grafting Nanofibrous Biomaterials)
5.2.1 Composite Nanofibrous Scaffold
5.2.2 Delivery Vehicle Nanofibrous Scaffolds
Drugs
Genetic Materials
Bioactive Factors
5.2.3 Injectable Nanofibrous Scaffolds
5.2.4 Multiphasic Nanofibrous Scaffolds
6 Implant-Related Nanofibrous Biomaterials
7 Conclusions and Future Perspectives
References
Recent Advances in Brain Tumour Therapy Using Electrospun Nanofibres
1 Introduction
2 Application of Electrospun Nanofibres in the Brain Tumour Research
3 Nanofibre as a Chemotherapeutic Delivery Platform
3.1 Carmustine or BCNU
3.2 Temozolomide (TMZ)
3.3 Paclitaxel (PTX)
3.4 Delivery of Other Therapeutics
4 Conclusion and Future Perspective
References
Layered Fibrous Scaffolds/Membranes in Wound Healing
1 Introduction
2 Skin Structure, Function and Wound Healing Process
2.1 Acute Wound Healing
2.2 Chronic Wounds (Longer Than 12 Weeks)
3 Polymers as Wound-Healing Materials
3.1 Natural Polymers
3.1.1 Alginate
3.1.2 Chitosan
3.1.3 Collagen
3.1.4 Gelatin
3.1.5 Hyaluronic Acid
3.1.6 Silk Fibroin
3.2 Synthetic Polymers
3.2.1 Poly(Lactide-Co-Glycolide) (PLGA)
3.2.2 Polyethylene Glycol (PEG)
3.2.3 Polyurethane (PU)
3.2.4 Polyvinylpyrrolidone (PVP)
3.2.5 Polycaprolactone (PCL)
4 Cutaneous Scaffolds with Added Therapeutic Agents for Wound Treatment
4.1 Growth Factors (GFs)
4.2 Antibiotics
4.3 Natural Substances
4.4 Antimicrobial Peptides (AMP)
4.5 Metal Nanoparticles (MNPs)
4.6 Metal-Organic Frameworks (MOFs)
5 The Conceptual of Design of Layered Scaffolds/Membranes for Wound Healing
6 Processing Techniques for Fibrous/Layered Wound Dressings
6.1 Multi-layered Electrospun Membranes
6.2 Combination of Three-Dimensional (3D) Porous Scaffolds with Nanofibers for Preparation of Layered/Fibrous Wound Dressings
6.3 Combination of Hydrogel Scaffolds with Nanofibers for Preparation of Layered/Fibrous Wound Dressings
6.4 Combination of Three-Dimensional (3D) Printed Scaffolds with Nanofibers for Preparation of Layered/Fibrous Wound Dressings
7 Summary
References

Citation preview

Advances in Polymer Science  291

R. Jayakumar   Editor

Electrospun Polymeric Nanofibers Insight into Fabrication Techniques and Biomedical Applications

Advances in Polymer Science Volume 291

Editorial Board Members Akihiro Abe, Tokyo Polytechnic University, Yokohama, Japan Ann-Christine Albertsson, KTH Royal Institute of Technology, Stockholm, Sweden Geoffrey W. Coates, Cornell University, Ithaca, NY, USA Jan Genzer, North Carolina State University, Raleigh, NC, USA Shiro Kobayashi, Kyoto Institute of Technology, Kyoto Sakyo-ku, Japan Kwang-Sup Lee, Hannam University, Daejeon, Korea (Republic of) Ludwik Leibler, Ecole Supe`rieure de Physique et Chimie Industrielles (ESPCI), Paris, France Timothy E. Long, Virginia Tech, Blacksburg, VA, USA Martin Mo¨ller, RWTH Aachen DWI, Aachen, Germany Oguz Okay, Istanbul Technical University, Istanbul, Tu¨rkiye Virgil Percec, University of Pennsylvania, Philadelphia, PA, USA Ben Zhong Tang, The Chinese University of Hong Kong, Shenzhen, Shenzhen, China Eugene M. Terentjev, University of Cambridge, Cambridge, UK Patrick Theato, Karlsruhe Institute of Technology (KIT), Karlsruhe, Germany Brigitte Voit, Leibniz Institute of Polymer Research Dresden (IPF), Dresden, Germany Ulrich Wiesner, Cornell University, Ithaca, NY, USA Xi Zhang, Tsinghua University, Beijing, China

Aims and Scope The series Advances in Polymer Science presents critical reviews of the present and future trends in polymer and biopolymer science. It covers all areas of research in polymer and biopolymer science including chemistry, physical chemistry, physics, and material science. The thematic volumes are addressed to scientists, whether at universities or in industry, who wish to keep abreast of the important advances in the covered topics. Advances in Polymer Science enjoys a longstanding tradition and good reputation in its community. Each volume is dedicated to a current topic, and each review critically surveys one aspect of that topic, to place it within the context of the volume. The volumes typically summarize the significant developments of the last 5 to 10 years and discuss them critically, presenting selected examples, explaining and illustrating the important principles, and bringing together many important references of primary literature. On that basis, future research directions in the area can be discussed. Advances in Polymer Science volumes thus are important references for every polymer scientist, as well as for other scientists interested in polymer science - as an introduction to a neighboring field, or as a compilation of detailed information for the specialist. Review articles for the individual volumes are invited by the volume editors. Single contributions can be specially commissioned. Readership: Polymer scientists, or scientists in related fields interested in polymer and biopolymer science, at universities or in industry, graduate students.

R. Jayakumar Editor

Electrospun Polymeric Nanofibers Insight into Fabrication Techniques and Biomedical Applications With contributions by K. Ashok
N. Ashok
M. Babu
R. Babu
L. Cai
A. Chandramouli
D. Chandrasekaran
K. Chatterjee
J. Chen
O. da Silva
T. T. Demirtas¸
I. M. El-Sherbiny
A. A. Elzatahry
A. Gu¨nyaktı
H. Hamedi
Y. S. Ibrahim
R. Jayakumar
R. Jeyanthi
F. Jiang
V. K. Kaliannagounder
A. Karakec¸ili
G. Kavitha
C. S. Kim
R. Ladchumananandasivam
E. Manikandan
A. M. Mansour
N. Mathivanan
D. Menon
S. Moradi
V. Muthuvijayan
S. V. Nair
C. H. Park
P. Pitchaimuthu
S. Pramanik
M. Rajput
G. D. V. Rengaswami
C. R. Reshmi
H. Ruan
D. Sankar
A. R. K. Sasikala
Y. Shi
S. Sowmya
U. Stachewicz
H. Thillaipandian
A. E. Tonelli
S. Tyeb
A. R. Unnithan
M. M. Zagho
C. Zhang
L. Zhou

Editor R. Jayakumar Polymeric Biomaterials Lab, School of Nanosciences and Molecular Medicine Amrita Vishwa Vidyapeetham (University) Kochi, India

ISSN 0065-3195 ISSN 1436-5030 (electronic) Advances in Polymer Science ISBN 978-3-031-31402-5 ISBN 978-3-031-31403-2 (eBook) https://doi.org/10.1007/978-3-031-31403-2 © The Editor(s) (if applicable) and The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 This work is subject to copyright. All rights are solely and exclusively licensed by the Publisher, whether the whole or part of the material is concerned, specifically the rights of translation, reprinting, reuse of illustrations, recitation, broadcasting, reproduction on microfilms or in any other physical way, and transmission or information storage and retrieval, electronic adaptation, computer software, or by similar or dissimilar methodology now known or hereafter developed. The use of general descriptive names, registered names, trademarks, service marks, etc. in this publication does not imply, even in the absence of a specific statement, that such names are exempt from the relevant protective laws and regulations and therefore free for general use. The publisher, the authors, and the editors are safe to assume that the advice and information in this book are believed to be true and accurate at the date of publication. Neither the publisher nor the authors or the editors give a warranty, expressed or implied, with respect to the material contained herein or for any errors or omissions that may have been made. The publisher remains neutral with regard to jurisdictional claims in published maps and institutional affiliations. This Springer imprint is published by the registered company Springer Nature Switzerland AG The registered company address is: Gewerbestrasse 11, 6330 Cham, Switzerland

Preface

This volume titled “Electrospun Polymeric Nanofibers: Insight into Fabrication Techniques and Biomedical Applications” highlights the different techniques employed in fabrication of polymeric fibers with distinct surface functionalization and vivid biomedical applications. The individual chapters in this volume explain the numerous fabrication techniques in use, namely, centrifugal, melt, co-axial, and yarning to acquire layered and tubular scaffolds that’ll find enormous purpose in the biomedical arena. There are also chapters that discuss in brief the chemical and surface functionalization of fibers, which enhances the biological properties of the electrospun fibrous scaffolds. The chapters presented in this volume talk about the recent advances and application of the multiscale fibrous and textile-based scaffolds in regenerative medicine and tissue engineering. The potential application of these electrospun polymeric fibrous scaffolds as biosensors, in other biomedical fields when developed as hybrids and external stimuli responsive scaffolds has also been highlighted. Some chapters of this volume throw light on the possible use of these fibrous scaffolds in cardiac, neuronal, and bone engineering. Possible use of the nanofibers as antimicrobial wound dressers and their use in atopic skin treatments and periodontal infection has also been mentioned. In addition to the aforementioned potential possibilities and applications, aspects of electrospun polymeric nanofibers in cancer theragnostic and brain tumor therapy have also been discussed in this volume. Overall, this volume provides the expertise knowledge about electrospun fibers fabrication at different scale with various techniques that will find abundant application as a scaffold in a biomedical field. This volume will also be extremely beneficial to material/biomaterials scientists, bioengineering and biotechnologists by providing a better insight on the subject concerning the innovative applications of fibrous scaffolds in biomedical and bioengineering arena. Kochi, India

R. Jayakumar

v

Contents

Recent Developments in Electrospinning Spinneret and Collector Assembly for Biomedical Applications . . . . . . . . . . . . . . . . . . . . . . . . . Hemamalini Thillaipandian, Pathalamuthu Pitchaimuthu, Dhandapani Chandrasekaran, and Giri Dev Venkateshwarapuram Rengaswami

1

Fabrication of Multiscale Polymeric Fibres for Biomedical Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Nivethitha Ashok, S. Sowmya, and R. Jayakumar

23

Techniques to Fabricate Electrospun Nanofibers for Controlled Release of Drugs and Biomolecules . . . . . . . . . . . . . . . . . . . . . . . . . . . . Monika Rajput, Suhela Tyeb, and Kaushik Chatterjee

37

New Prospects in Melt Electrospinning: From Fundamentals to Biomedical Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Moustafa M. Zagho, Yasseen S. Ibrahim, and Ahmed A. Elzatahry

69

Centrifugal Spun Nanofibers and Its Biomedical Applications . . . . . . . Hemamalini Thillaipandian and Giri Dev Venkateshwarapuram Rengaswami

81

Recent Advances in Electrospun Nanofibrous Polymeric Yarns . . . . . . C. R. Reshmi, Rosebin Babu, Shantikumar V. Nair, and Deepthy Menon

107

Fabrication of Textile-Based Scaffolds Using Electrospun Nanofibers for Biomedical Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . K. Ashok, M. Babu, G. Kavitha, R. Jeyanthi, R. Ladchumananandasivam, O. da Silva, and E. Manikandan Biomedical Applications of Electrospun Piezoelectric Nanofibrous Scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Afeesh Rajan Unnithan and Arathyram Ramachandra Kurup Sasikala

139

167

vii

viii

Contents

Surface Modified Polymeric Nanofibers in Tissue Engineering and Regenerative Medicine . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Nivethitha Ashok, Deepthi Sankar, and R. Jayakumar

177

Polymer/Ceramic Nanocomposite Fibers in Bone Tissue Engineering . . S. Sowmya, Nirmal Mathivanan, Arthi Chandramouli, and R. Jayakumar

191

Electrospun Fibrous Scaffolds for Cardiac Tissue Engineering . . . . . . . Nivethitha Ashok, Vignesh Krishnamoorthi Kaliannagounder, Cheol Sang Kim, Chan Hee Park, and R. Jayakumar

213

Electrospun Nanofibrous Scaffolds for Neural Tissue Engineering . . . . Sheersha Pramanik and Vignesh Muthuvijayan

229

External Stimuli Responsive Nanofibers in Biomedical Engineering . . . Hamid Hamedi, Sara Moradi, and Alan E. Tonelli

287

Electrospun Antimicrobial Polymeric Nanofibers in Wound Dressings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Yunfan Shi, Chenzi Zhang, Feng Jiang, Liuzhu Zhou, Ling Cai, Hongjie Ruan, and Jin Chen Application of Electrospun Polymeric Fibrous Membranes as Patches for Atopic Skin Treatments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Urszula Stachewicz Nanofibrous Scaffolds for the Management of Periodontal Diseases . . . Alaa M. Mansour and Ibrahim M. El-Sherbiny Recent Advances in Brain Tumour Therapy Using Electrospun Nanofibres . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Arathyram Ramachandra Kurup Sasikala Layered Fibrous Scaffolds/Membranes in Wound Healing . . . . . . . . . . Ays¸e Gu¨nyaktı, Tugrul Tolga Demirtas¸, and Ays¸e Karakec¸ili

313

335 361

409 425

Adv Polym Sci (2023) 291: 1–22 https://doi.org/10.1007/12_2022_134 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 28 August 2022

Recent Developments in Electrospinning Spinneret and Collector Assembly for Biomedical Applications Hemamalini Thillaipandian, Pathalamuthu Pitchaimuthu, Dhandapani Chandrasekaran, and Giri Dev Venkateshwarapuram Rengaswami

Contents 1 2 3 4

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Principle of Electrospinning Technique . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Advances in Electrospinning Technique . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Modification of Spinneret in Electrospinning Process . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1 Multi-needle Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.2 Nozzle Free Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 Modification of Collector in Electrospinning Process . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 Future Trends . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

2 3 5 8 9 10 13 18 18 18

Abstract Electrospinning is one of the versatile methods to produce fibers of submicron diameter and the heart of the spinning assembly is the spinneret, voltage assembly, and collector. The spinneret design helps in achieving uniform fibers or fibers with desired morphology. Porous, coaxial fibers, hollow fibers, and profile fibers have a significant impact on the growth and proliferation of cells. Moreover, the collector also plays a significant role in the assimilation of fibers onto the assembly. Different types of orientations of fibers and forms can be achieved by

H. Thillaipandian and G. D. V. Rengaswami (*) Department of Textile Technology, Anna University, Chennai, India e-mail: [email protected] P. Pitchaimuthu Department of Mechanical Engineering, Anand Institute of Higher Technology, Chennai, India D. Chandrasekaran Physics Instruments Co., Chennai, India

2

H. Thillaipandian et al.

the selection of proper collector assembly. The design of collectors is mostly carried out with potential end use and for particularly biomedical applications aligned fibers, tubular fibers and profile fibers aid in cell proliferation and functionality. The chapter discusses the recent developments in the spinneret and collector assembly. Keywords Biomedical application · Collector · Electrospinning · Spinneret

1 Introduction Nanofibers in the range of 1–100 nm are widely used in biomedical applications due to their larger surface area to volume ratio, higher chemical reactivity, and electrical and mechanical property. Nanofibers can be fabricated using various techniques such as drawing, template synthesis, phase separation, self-assembly, centrifugal spinning, and electrospinning process. The process and the limitation of each process are shown in Table 1. Among the other processes, the electrospinning technique is a versatile and straightforward technique used for the production of nanofibers in various forms such as porous, helical, multichannel, triaxial, Janus, hollow, bicomponent, and core-shell structures. The electrospinning process comprises four different regions such as polymeric solution loaded on the syringe and charged, linear jet region emerging from the tip of the spinneret/needle, stretching of jets and splitting them to form nanofiber, and deposition of formed nanofibers on the

Table 1 Nanofiber formation using various techniques Techniques Drawing

Template synthesis

Phase separation Selfassembly Melt blowing Microfluid spinning Centrifugal spinning

Process Production of discontinuous nanofibers from the drop of liquid using a sharp needle Synthesized by pouring the polymeric solution onto a template

Formation of nanofibrous matrix by removal of the solvent from two immiscible liquids Small molecules aggregated to form nanofibers using noncovalent bonds Uses air to draw polymeric solution to fibers Injection of different fluids in the microscale channel and followed by polymerization Uses centrifugal force for the production of nanofibers

Disadvantages Lower productivity due to the formation of discontinuous fibers Requires soluble porous template Production of discontinuous fiber Longer processing time Difficult in processing continuous and longer fibers Time-consuming process Formation of unstable fibers Suitable for thermoplastic polymers Selection of a suitable solidification process hinders the scalability Difficulty in the collection of nanofibers.

Reference [2]

[3]

[4]

[5] [6] [7]

[8]

Recent Developments in Electrospinning Spinneret and Collector. . .

3

collector. The jet formed by the polymeric solution undergoes major three changes due to the application of the electric field on the formation of fibers such as the formation of Taylor cone, the ejection of the straight jet, and whipping instability. Electrospun nanofibers are widely used for biomedical applications due to the utilization of natural and synthetic biopolymers as the polymers can be degraded based on intended applications. The fibrous nonwoven increases the interaction and bonding between the nanofibers and biomolecules. The electrospinning process allows the selection of a wide range of polymers and solvents as the functional group of the polymers offers flexibility, high mechanical properties, and controllable pore size suitable for various technical applications [1].

2 Principle of Electrospinning Technique Electrospinning is a versatile technique that allows the production of fibers in the range of nano to micrometers due to its unique characteristics utilizing electro-hydrodynamics which combines spraying and spinning process. The system comprises collecting device (collector), voltage supplier, and syringe pump with spinneret (needle). The role of the syringe pump ensures a constant flow rate of solution loaded in the syringe through the spinneret, a high-voltage power source is to electrically charge the polymeric solution in order to overcome the two forces acting on the solution, namely electrostatic repulsions between polymeric chain and Columbic force exerted by an application of the external electric field. The polymeric solution forms Taylor cone from the spherical shape as the polymeric solution leaves the tip of the needle due to these forces. Upon increasing the voltage beyond the threshold value, the cone stretches into a jet as the electrostatic force overcomes the surface tension of the polymeric solution followed by fiber formation. The formed fibers were deposited on the ground or negatively charged collecting device due to the difference in the electrical potential between the spinneret and collector facilitating evaporation of the solvent. The fibrous mat comprises randomly or aligned nanofibers by designing the selection of spinneret or collector. The materials used for the fabrication of the spinneret include metals or insulators whereas the collector includes aluminum foils and copper plates [9]. The types of spinnerets and collectors used in the electrospinning process are shown in Figs. 1 and 2.

Fig. 1 Types of spinneret used in electrospinning process

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Fig. 2 Types of collectors used in electrospinning process

The morphology of the fibers can be controlled by optimizing the process parameters such as environmental conditions such as humidity and temperature, material parameters such as polymer concentration, viscosity, molecular weight, selection of solvents, and conductivity of the polymeric solution, machine parameters such as needle tip to collector distance, applied voltage, and flow rate of the polymeric solution [10]. The humidity of the environment has an influence on the morphology of the fibers as it has effect on the solidification of fibers by evaporation of solvent. The thickness of the fibrous mat and evaporation rate of the solvent increase with an increase in relative humidity of the environment. Temperature has an influence on mass transfer between the interface of water vapors and solution jets. The surface of the nanofiber can be influenced by the temperature at which it is processed as low temperature results in a solid and smooth surface whereas higher temperature results in a rough and porous structure. Polymer concentration influences the spinnability of the fiber as a lower concentration of polymer results in rupture of polymeric chain into fine fragments resulting in the formation of beaded fibers but increasing the polymer concentration results in an increase in chain entanglement as the latter overcomes the surface tension of the solution resulting in bead free fiber formation. Viscosity of the polymeric solution has an effect on fiber formation as low viscosity of the polymeric solution results in droplet formation but on optimum viscosity uniform bead free fibers were formed. Diameter of the nanofibers is influenced by the molecular weight of the polymer as the higher molecular weight of the polymer results in larger diameter fibers and on the contrary lower diameter fibers can be formed at a low concentration of the polymer. Selection of solvent has a crucial role in the formation of bead free fibers as it influences two parameters such as solubilization of the polymer and possesses a lower boiling point.

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Highly volatile solvents are chosen for spinning the nanofibers due to the enhanced evaporation rate of solvents from the polymeric solution resulting in the solidification of fibers whereas nonvolatile solvents can be employed with modification in the electrospinning process with a coagulation bath due to poor evaporation rate of the solvents. Another parameter that influences the diameter of the nanofibers is the conductivity of the polymeric solution. Highly conductive solution uptakes the charge resulting in Taylor cone formation and forming uniform bead free fibers and a low conductive solution undergoes no changes in droplet and hinders the formation of Taylor cone resulting in beaded fibers. The needle to collector distance decides the diameter of the fibers as lesser distance results in a larger diameter of fibers due to poor evaporation of solvent and vice versa. Applied voltage has a crucial role in fiber formation as it is responsible for stretching the spherical droplets to Taylor cone to the jet of polymeric solution. Fiber formation of smaller diameters was facilitated when the applied voltage is higher than a critical value due to stretching and elongation of polymeric jet caused due to whipping instability. The diameter of the metallic needle is directly proportional to the size of fiber as larger diameter needle results in larger diameter fibers. Flow rate controls the diameter of the fiber as higher mass throughput results in a larger diameter due to poor evaporation [11]. Optimizing the process parameters results in the formation of uniform bead free fibers for various technical applications. Electrospun nanofibers were used in biomedical fields in the area of tissue engineering, drug delivery, cancer therapy, regenerative medicines, biosensors, biocatalysts, medical devices, and filtration applications [12].

3 Advances in Electrospinning Technique Electrospinning technique can be classified based upon the processing of polymeric solution into nanofibers other than the single needle electrospinning process. Electrospinning process offers various advantages of spinning a wide range of materials either in the form of dissolution, melt or dispersion in the solution. The structure can be oriented in a specific direction by controlling the various process parameters such as machine and material parameters and also controlling the electric field and type of collector [13]. Electrospinning technique allows the production of blended nanofibers by blending two polymers either natural or synthetic polymers. The blend electrospinning process allows the inclusion of drugs either in dissolved or dispersed state in the polymeric solution but the limitation of the process includes utilization of organic solvent resulting in denaturation of proteins, loss of activity of the protein, and loss in the sensitivity of the biological behavior of drug [14]. Coaxial electrospinning technique allows the production of core-shell nanofibers by preparing two different polymeric solutions. The solution was fed on two capillaries such as shell and core and the solutions were forced through respective capillaries and composite droplets were formed. On application of voltage, electrostatic repulsion between the

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polymeric chain results in elongation of shell liquid which creates viscous stress on the core resulting in rapid stretching of core solution thereby forming core-shell nanofibers [15]. The add-on advantage of coaxial spinning includes the production of core-sheath fibers by combining different materials within a single structure, hollow fibers by either removal of core or sheath material, protective casing of sensitive drugs within the polymer matrix and controlling the release kinetics, production of non-spinnable materials with an aid of electrospinnable polymeric solution [16]. Emulsion electrospinning technique allows the production of nanofibers from two immiscible liquids and rapidly stirring into emulsion followed by the electrospinning process. To stabilize the immiscible liquids, a surfactant or stabilizer is used. The technique is similar to the blend and coaxial electrospinning technique allows the encapsulation of small particles with sustained release and stable bioactivity rather than burst release of drugs. The technique is widely used for encapsulation of hydrophilic inorganic materials and proteins thus suitable for tissue engineering applications [17]. As the electrostatic force overcomes the surface tension of the polymeric solution resulting in the formation of Taylor cone, higher molecular weight and high viscosity polymers are difficult to electrospun using the traditional electrospinning process. Bubble electrospinning and air-jet assisted bubble electrospinning are the advances in the electrospinning process which is used for the polymers that require higher electrostatic force for attenuating into nanofibers. Bubble electrospinning process is an example of a needleless electrospinning process in which the free liquid is placed in the reservoir as the gas insufflation turns on bubbles are generated. On increasing the direct-current voltage beyond the threshold value, it results in the formation of instability polymeric jet resulting in the formation of thick nanofibers on the collector. The limitation of the technique includes the requirement of liquid with high surface tension suitable for bubble formation and formed nanofibers of straight, coiled, and helical fibers along with beaded structure [18]. Air-jet assisted bubble electrospinning (blown bubble Spinning or electrostatic-field-assisted blown bubble spinning) is similar to the bubble electrospinning process as it uses blowing air and an electric field to draw the fibers by increasing the efficiency of the bubble electrospinning process [19]. Gas assisted electrospinning technique is similar to the electro-blowing and meltblowing technique which uses poorly electrospinnable materials, namely the materials possessing low electrical conductivity or dielectric constant. Melt-blowing technique is similar to the melt electrospinning technique, in which the polymeric solution was prepared by melting the polymer and loading onto the syringe. As the polymeric solution emerges from the spinneret, high-velocity hot air was used to attenuate the jets into random nanofibers. The web offers high surface area per unit weight, high insulation value, and high barrier properties but with the limitation of thermal degradation of polymers [20]. Electro-blowing utilizes air for drawing the fibers similar to the melt-blowing technique. In this technique, the polymer was dissolved in a solvent and a homogenous solution was prepared. The solution was loaded onto the syringe pump and high voltage was applied to the needle in order to charge the polymeric solution and high compressed air is applied at the end of the

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nozzle to attenuate the fibers. Melt-blowing technique uses only compressed air whereas the electro-blowing technique relies on the electrostatic repulsion between the polymeric chain and compressed air to produce nanofibers [21]. Solvent free electrospinning technique offers the advantage of spinning the polymers into fibers without using conventional solvent thereby eliminating the fumes of solvent into the air or atmosphere thereby efficiently utilizing the polymer. The technique includes melt electrospinning, supercritical CO2-assisted electrospinning, anion-curing electrospinning, and UV-curing electrospinning. Melt spinning is another advancement in the electrospinning technique as it avoids the utilization of organic solvent rather than the polymer is melted into solution and spun similar to single needle electrospinning technique. The processing of the polymeric solution requires the constant rate of heating as the molten solution solidifies on cooling. The advantage of the technique includes a reduction in the needle to collector distance as the conventional electrospinning process uses the solvent for dissolution. The technique is environmental friendly and utilizes thermoplastic polymers but limitation includes processing temperature and viscosity of the polymeric solution hinders the reproducibility of nanofiber formation thereby requiring higher charge density [22]. Supercritical CO2-assisted electrospinning utilizes an electric field and supercritical CO2 solvent for the production of nanofibers. Supercritical liquid refers to processing the materials above a critical value of pressure and temperature with carbon dioxide as the solvent. The process is not widely used as a traditional electrospinning technique as it requires high pressure of 14 MPa in a dissolution of materials in the supercritical stage [23]. Anion-curing electrospinning technique utilizes polymer, cyanoacrylate monomer, and poly (methyl methacrylate) components. The role of poly(methyl methacrylate) is to increase the viscosity of the cyanoacrylate monomer as it solidifies into stronger fibers in the presence of anionic materials [24]. The difference between traditional and UV-curing electrospinning is the oxygen-free environment in order to solidify the nanofibers comprising UV curable material. The polymeric solution was spun into nanofibers under ultraviolent radiation in a nitrogen atmosphere with the advantage of solidification of fibers without evaporation of solvent [25]. Nanospider electrospinning or roller electrospinning comprises single needle electrospinning units along with a roller spinning electrodes, a tank, and a supporting material. The roller spinning electrode is placed in the tank containing the polymeric solution which is responsible for the production of nanofibers. The supporting material acts as a grounded electrode on which ultrafine nanofibers were electrospun continuously. The limitation of the technique includes the recovery of solvent used during dissolution and the production of nanofibers from low molecular weight is difficult [26]. Magnetic field assisted electrospinning uses the materials that respond to the magnetic field as two parallel permanent magnets are employed. The technique allows the uniform production of fibers without splitting compared to the traditional electrospinning technique. It was reported that the technique allows the production of fine fibers with better orientation/alignment due to the internal arrangement of polymeric chains of the electrospinning jet by the magnetic field [27]. Near field electrospinning is similar to the traditional electrospinning technique in which the

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distance between the needle tip to collector distance was greatly reduced. The technique utilizes a stable region of the jet from the needle tip to collector as the distance causes the fiber to deposit on the collector by an electrical charge. Compared to the traditional electrospinning process, low voltage can be employed for the production of nanofibers in the range of 100–600 V whereas the former utilizes higher voltage for the formation of fibers [28, 29]. Electrospun fibrous mats are widely used as protection membranes, battery separators, sensor manufacturing, biomedical applications such as tissue engineering, drug delivery, and wound dressing applications due to their high surface area to volume ratio, high reactivity, and tunable chemical, physical, and mechanical properties. Among the other nanofiber production techniques, the electrospinning technique is widely used for biomedical applications as the loosely deposited 3D porous structure which mimics the extracellular matrix (ECM) as the cells anchors to the structure promotes the cell growth making it suitable for tissue engineering applications, sustained and control release of drugs by the slow degradation of polymers produced through electrospinning process was higher due to higher surface area compared to cast film technique thus making the fibrous web suitable for drug delivery applications. The electrospun web is an ideal candidate for wound dressing applications due to its homogeneity structure which promotes the growth of fibroblasts and keratinocytes thereby enhancing the wound healing process compared to freeze-dried samples as the latter offers a heterogeneity structure [30]. The limitation of the electrospinning technique includes mass productivity hence, research works are being carried out to address the issue. The electrospinning technique was modified to address the productivity either by increasing the number of needles (spinneret), air-assisted needles, or needleless electrospinning process. The alignment of the nanofibers on the depositor can be modified based on the end applications by modifying the collector with suitable designs. Hence, the modification of electrospinning in terms of spinneret and collector is discussed in the chapter.

4 Modification of Spinneret in Electrospinning Process Modification of spinnerets was carried out to increase the quality and productivity of the nanofibers. The modification includes the utilization of multiple spinnerets and needleless nozzles in the traditional electrospinning process. Low-cost spinneret was designed by modifying McIntyre cannula which was used for simultaneous irrigation and aspiration in ophthalmic surgeries. The cannula allows the production of core-sheath fibers similar to the coaxial electrospinning technique as the latter offers low cost by five-folds compared to the spinneret used for the production of core-sheath fibers. The spinneret comprises outer 18G and inner 23G needles which allows the uniform flow of core and sheath fluids by means of the adaptor (T-joint) which acts as leakproof. The setup was modified into a coaxial spinning spinneret by grinding the bevel without distorting the annular spacing between the needles. The stable Taylor cone and encapsulation efficiency

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were achieved with the annular spacing of 0.2 mm for the passage of sheath fluid and 0.32 mm for core fluid. The needles were maintained at the uniform levels to transfer the electrostatic force between the two different liquids. The modified spinneret allowed the production of nanofibers by encapsulation of red blood cells (RBC) within the polycaprolactone (PCL) fibers, Bacillus subtilis in PCL nanofibers, polyvinyl alcohol (PVA)-lysozyme enclosed in a PCL sheath as the core fluids either release in burst manner or undergo thermal or enzymatic degradation which hinders the application in the biomedical area [31].

4.1

Multi-needle Electrospinning

The major limitation of the electrospinning technique is the mass production of nanofibers which can be achieved by multiple jetting of nanofibers from the spinneret. Uniform, steady, and rapid jets were formed with multi-spinneret assisted with airflow. It was reported that the airflow acting on the sheath of polymeric solution inside the spinneret resulted in the stretching of jets thereby causing thin nanofiber deposition without the assistance of voltage. The position of the spinneret played a major role as the spinneret arranged parallelly in the line caused electric field interference thereby restraining the jet formation and suppressing the electric field on the center of the spinneret by the peripheral spinneret thereby hindering the steady jetting. On the other hand, multi-spinneret arranged in a curved manner resulted in uniform jetting of fibers by reducing the electric field interference between the spinnerets. The polymeric solution flows through the outlet into the inlet port and fills uniformly on the 5  5 array of the spinneret and high voltage was applied onto each spinneret as it reaches the critical value, jetting of the polymeric solution was achieved by overcoming the surface tension and viscous force of the solution. The airflow acting on the sheath of the spinneret increased the stretching force and caused uniform and continuous ejection of multiple charged jets. It was reported that the addition of air pressure for multi-jetting reduced the critical voltage for both center and outer spinneret thereby aiding the mass production of nanofibers [32]. Profiled multipin electrospinning technique offers advantages such as prevention of needle clogging, particle settling, uncontrolled/uneven Taylor cone formation, utilization of low voltage, and uniform sized nanofibers compared to needle and needleless electrospinning process. Profiled multipin electrospinning comprises 21 half-sphere shaped profiled pins with a distance of 15 mm each within the circular disc. The spinneret assembly moves forward and backward during polymer loading cycle time carries the polymer solution and forms spherical liquid droplets. On increasing the voltage to 30 kV, uniform jets were formed from each profiled pin resulting in fiber formation. Nanofibers comprising cellulose acetate and polyvinyl alcohol of polymer concentrations of 15 and 10 wt%, respectively, were spun with an average diameter of 124 and 160 nm. The developed technique has the potential of embedding the zinc oxide onto the polymer matrix but resulted in increase in the diameter of fibers [33].

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Nozzle Free Electrospinning

The utilization of multi-needles offers limitations such as a larger operating space, designing of spinneret at the relative spacing between the needles which causes a strong charge repulsion between the polymer jets resulting in nonuniform fiber deposition and frequent cleaning of needles is necessary to prevent the clogging of needles. To overcome the limitation of the multi-jetting principle, needleless electrospinning was adopted for increasing the productivity of nanofibers. The principle involves the formation of multiple jets by immersing the spinneret in the solution bath and the application of a high electric field between the solution bath and collector [34]. Rotary cone was utilized in the place of the needle to produce super high mass throughput offered production throughput of 10 g min 1 compared to traditional technique as it offers 5 g h 1. The modified method comprises highvoltage power supply, a metallic cone, a direct-current (DC) electromotor, and an aluminum belt as the collector. The specification of hollow metal rotary cone bottom and top diameters were 30 and 7 mm, the height of the cone was 40 mm, and the thickness was 2 mm. The spinning solution was prepared by dissolving polyvinylpyrrolidone in ethyl alcohol. The formation of micro to nanofibers was noticed as the polymeric solution reached the metal cone as the latter was rotated above 50 rpm. Increasing the rotating speed of metal cone in the absence of high voltage resulted in fiber formation implying that rotational speed and electric field have a crucial role in fiber formation [35]. The spinnability of the PVA/water solution was attempted using cylinder and disc type nozzles in the replacement of the needle in the conventional electrospinning process. The specification of the cylinder nozzle includes 20 cm long and 8 cm in diameter whereas the disc nozzle includes the same length and thickness of 2 mm. The nozzles were beveled with a curve of 5 mm. The polymeric solution was half filled in the vessel and electrically charged by application of high voltage through copper wire. The thin fibers were formed by rotating the cylinder and disc at the speed between 40 and 50 rpm, increasing the speed resulted in the throwing of the polymeric solution, on other hand decreasing the speed below optimum speed resulted in uneven solution coverage causing poor jet formation. It was reported that increasing the voltage from 47 to 62 kV resulted in a decrease in the diameter of the fibers from 340 to 194 nm in disc type nozzle electrospinning but in cylinder type, a slight change in diameter was found thereby implying that disc type nozzle offered finer diameter of fibers and also occupied less space compared to cylinder type. The electric field plays a major role in charging the polymeric solution and initiating the jet formation. The electric field in the cylinder was higher at the edges compared to center of the cylinder resulting in poor jet formation whereas the disc type offered higher electric field at the center followed by a decaying field was reported toward the collector surface. The disc type nozzle offered a higher electric field, finer diameter, and higher production of fibers compared to cylinder type needleless electrospinning technique [36].

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Electric field intensity of the spinneret decides the nanofiber formation and their productivity. As the electric field plays a crucial role in jet initiation, modification of spinneret was carried out to study the profile of the electric field. Flat wheel was used as a spinneret for the production of nanofibers in the electrospinning technique. The specification of the flat wheel includes the diameter of 150 mm, the thickness of 2 mm, and width of 25 mm, respectively. The flat wheel was immersed in the polymeric solution and charged by immersing the electrode into the solution bath by the application of high voltage and a drum covered with aluminum foil was used as a collector and rotated at 150 rpm. The wheel was immersed in one fifth of the polymeric solution and charged from 45 to 75 kV high strength electric field. It was reported that multiple jets of polymeric solution arose at 60 kV and the rotation speed of the wheel was between 5 and 8 rpm. The nanofiber collection was higher at the edges compared to the middle due to the distribution of the electric field at the edges. Increasing the applied electric decreased the fiber diameter from 450 to 350 nm by increasing the voltage from 45 to 75 kV due to an increase in the jet extension and increasing the concentration of PVA resulted in the increase in fiber diameter from 431 to 557 nm on increasing the concentration from 7 to 10 wt% due to poor drawing of fibers. It was reported that the productivity from the flat wheel was higher than the cylinder but lower than the disk and ring type spinneret and 15to 40-folds higher than the traditional electrospinning technique [37]. In the case of traditional electrospinning, utilization of small diameter needleless causes corona discharge on increasing the voltage to a higher degree. The applied voltage employed for the production of nanofibers is less than 30 kV and decreasing the voltage results in the formation of coarse nanofibers. To overcome the limitation of humidity dependent needle electrospinning process, the spinneret was modified with a cone-shaped metal wire coil to increase the productivity and finer nanofibers. The spinning solution was prepared with PVA and the specification of the coil includes a copper wire diameter of 1 mm with a conical angle of about 120 and a gap between the adjacent wires of 1 mm and the height of the cone was 15 mm. The conical wire was utilized rather than a flat surface due to differences in geometry and fluidics. The former (conical wire) offers an open surface which allows the uptake of viscous fluid thereby creating slightly higher liquid pressure. This allows the uniform covering of liquid at the outer side of the cone and also application of the electric field was higher at the outer ring compared to the smaller diameter of the cone. The modified electrospinning technique allows the production of finer nanofibers by increasing the voltage to 70 kV. Stretching of fibers enhances the production of nanofibers which can be achieved by slowing down the evaporation rate of solvent from the polymeric solution which in turn changes the humidity around the collector created by the multiple jets from the copper conical wire. Thus, the technique did not induce the corona discharge as the solvent evaporation from multiple jets was slow and humidity was maintained at a low level due to the air exchange between the electrospinning zone and ambient environment [38]. Conical spinneret was used for the production of nanofibers in the place of a needle in the electrospinning technique. The spinning solution was prepared by dissolving PVA in distilled water and electrospun using a modified and conventional

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electrospinning process. The polymeric solution was prepared by dissolving 10 wt% PVA in distilled water and electrospun by varying machine parameters such as a voltage of 15–25 kV, needle top to collector distance between 10 and 20 cm, and feed rate between 1 and 5 ml/h. Two zones of electrospinning jet arise from spinneret such as axisymmetric jet arises from spinneret outlet and whipping instability as the jet spirals and collects on the collector. The latter zone was responsible for thinning of fibers and deposition of fine nanofibers onto the collector surface. It was reported that the fibers collected on the aluminum collector have no significant difference in the diameter but the deposition of fibers on the collector was smaller area due to focused and high electric field in conical spinneret compared to needle spinneret. The proposed modification has the potential for various technical applications by utilizing low voltage and reduced power consumption [39]. A copper terraced spinneret was used as fiber generator to simulate the uniform electric field in the electrospinning process. The bottom of the spinneret was connected to a high voltage of 55 kV and the distance between the needle tip and collector was 15 cm. The electric field at the curving point was found higher compared to other parts of the terraced spinneret as it efficiently generates more jets resulting in whipping instability followed by fiber formation. The modified spinneret offered another advantage of preventing the clog formation, unlike the conventional single needle electrospinning process. Bead free PAN nanofibers were formed with an average diameter of 449 nm due to the distance between every layer of the terraced spinneret to the collector. The modified spinneret allowed the mass production of nanofibers for industrial application [40]. Linear flume spinneret was utilized for the production of multiple jets by application of a uniform electric field. The polymeric solution comprising polyacrylonitrile (PAN) in dimethylformamide (DMF) was transferred to the flume spinneret. The rotating cylindrical collector was used to collect the nanofibers. The number of jets was found to be 2–3 jets/cm operating at 60 kV. It was reported that the highest electric field was noticed at the exit of linear flumes and decreased on reaching the collector as the distance between spinneret to the collector was maintained at 15 cm. It was reported that increasing the concentration of PAN resulted in an increase in the diameter of fibers due to more molecular entanglement. The production of nanofibers was found to increase with an increase in solution concentration and applied voltage but decreased with increasing the collector distance as the electric field on the polymeric solution reduces resulting in the drifting of nanofibers from the depositor. It was reported that utilizing the linear flume spinneret increased the production to 5 g/h which is 24-folds higher than the single needle electrospinning machine thereby offering the potential for producing high quality and productivity nanofibers [41]. The spinning solution was prepared by dissolving polyvinyl alcohol (PVA) in distilled water. The needleless electrospinning was carried out by a copper spiral coil which acts as fiber generator, the electrode was placed between the coil and aluminum foil as a collector. The specification of a spiral coil includes the coil length of 16 cm, a diameter of 8 cm, a spiral distance of 4 cm, and wire diameter of 2 mm. It was reported that jets from polymer solution were formed at a higher

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voltage of 40 kV and on further increase in voltage, the number of jets was found to increase. On increasing the voltage to 70 kV, corona discharge was carried out rather than the normal electrospinning process. It was reported that increasing the concentration of PVA resulted in an increase in fiber diameter and decreased the number of jets emerging from the spiral coils. The maximum concentration of PVA that can be spun into nanofibers was 12 wt%, on increasing the concentration resulted in the clogging of coils. It was reported that coil electrospinning allows the mass production of nanofibers by increasing the applied voltage and decreasing the collector distance. Thus, the technique allows the bulk production of nanofibers due to the strong electrostatic force created by the higher electric field compared to the conventional electrospinning technique thereby resulting in the formation of thinner fibers and higher production [42]. Helical nanofibers were fabricated by replacing the needle with hole as it supplies a uniform electric field resulting in uniform fibers. Co-electrospinning system with core-shell structure was fabricated with a flat spinneret (metal needle free spinneret). The electrode was inserted directly onto the hole rather than the needle as it reduces the complexity of co-electrospinning technique, unlike the metal needle spinning system. The spinning solution was prepared by dissolving thermoplastic polyurethane in a solvent mixture of dimethylformamide (DMF) and tetrahydrofuran (THF) in the volume ratio of 1:3 with a polymer concentration of 14 wt% and poly (m-phenylene isophthalamide) (Nomex) solution was prepared by dissolving the chopped fibers in N, N-dimethylacetamide (DMAc) and Lithium chloride anhydrous (LiCl) with a polymer concentration of 12 wt%. The flat co-electrospinning setup comprises aluminum electrode and an auxiliary aluminum plate in which the shell tube and the core tube were placed through the electrode. The auxiliary plate was placed on the plastic plate and the holes were drilled to force the polymeric solution in order to form core-shell nanofibers. It was reported that a gradual decrease in the electric field was noticed in flat spinneret compared to needle spinneret as it offered a sharp decrease in electric field with the increase in the distance from the spinneret. It was reported that core-shell nanofibers of 500 to 800 nm were produced uniformly with Nomex in the core and polyurethane in the shell using the flat spinneret co-electrospinning technique [43].

5 Modification of Collector in Electrospinning Process The collector used in the electrospinning process can be classified into two types based on the state of motion such as a stationary and moving collector. The collector allows the production of nanofibers with tailored structure by modifying the form of the collector as it is a complicated procedure. Electrospinning technique allows the production of randomly oriented fibers for various technical applications. Parallel aligned fibers were produced by modifying the electrospinning with two collectors by varying the distance between them. The spinning solution was prepared by dissolving various concentrations of polyvinyl

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alcohol in distilled water. The collector was modified by spacing two copper collectors at a particular distance and negatively charged as the polymeric solution was positively charged. The spinning process was carried out with a voltage of 15 kV, the distance between the needle tip and collector at 9 cm and a flow rate of 0.5 ml/h. It was reported that increasing the concentration of PVA increased the diameter of the fibers and increasing the distance between the collector plates resulted in parallel alignment of fibers with a decrease in the diameter of the fibers due to horizontal pulling force acting on the fibers. It was concluded that an average diameter of the fibers was formed of 377 nm with the distance between the parallel copper plate was 15 mm [44]. Rotating mandrel of conductive material was used as the collector for the production of aligned nanofibers. Different collectors were used for the study such as solid aluminum tubes, aluminum tube with Teflon coating. Polycaprolactone was dissolved in acetone solvent and spun using the electrospinning technique. The gap between the aluminum drum conducting collector with Teflon coating (nonconducting surface) aids in the stretching of fibers. It was reported that no notable difference in fiber diameter was seen on changing the collector but a difference in fiber alignment was noted. This was due to electrostatic forces acting on the charged fibers, namely the electric field acting on the fibers and Coulomb interactions between the positively charged nanofibers and the negatively charged rotating aluminum mandrel. The tensile strength of the nanofibers was found to increase with fiber alignment caused due to enhanced molecular orientation and crystallinity by the rotating mandrel with conducting and nonconducting segments. Thus, the developed web finds an application in tissue engineering for the treatment of acetabular labrum hip injury [45]. Nanomaterials for biomedical applications can be prepared by the modification of the collector to yield designed and patterned nanofibers such as meshes, nets, nanowalls, and aligned fibers. The movement of the collector either in rotating or raster motion results in the web formation in a similar manner. Hydroxypropyl methylcellulose (HPMC), polyvinyl alcohol (PVA), and poly(ethylene oxide) solution were prepared by dissolving in deionized water. The drug, namely amoxicillin and clavulanic acid was loaded in PVA solution to assess the drug release, bactericidal, and wound healing properties of the fibrous mat. PVA web was sandwiched between HPMC nanofibers on which PEO solution was near printed by moving the spinneret forward and backward on the copper wire. It was reported that gradual release of drugs was seen from the fibrous mat of 95% after 7 days after incubation against Gram-positive (S. aureus and E. faecalis) and Gram-negative bacteria (E. coli). The fibrous mat offered wound healing properties confirmed using MTT assay [46]. Starch, a polymer composed of two types of molecules: amylose and amylopectin. Amylose is a largely linear structure which readily associates side by side under favorable conditions but amylopectin is a highly branched structure that gives a globular bulky hydrodynamic shape, which is not suitable for electrospinning as it cannot be easily elongated and aligned in the extensional flow field of spinneret [47]. A modified method of electrospinning technique was proposed to produce

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pristine starch into nanofibers. Spinning solution was prepared by dissolving 15 wt% of Gelose 80 starch (80% amylose rich) in 95% dimethyl sulfoxide (DMSO). The solution was heated and cooled to room temperature. The modified electrospinning technique comprises immersing a grounded metal mesh in a coagulation bath and the bath was grounded. Ethanol was used as a coagulation medium. As DMSO was a relatively nonvolatile solvent, solid nanofibers were not formed on the collector due to poor evaporation of solvent from the polymeric solution under ambient conditions. Solvent miscible with DMSO but incompatible with starch was used as a solvent in a coagulation bath. As the solvent jet reaches the coagulation, DMSO is extracted and starch collapses in a fibrous form [48]. Production of starch fibers was attempted by modification of electrospinning by means of a collector bath in the place of aluminum collector which dehydrates the solvent used for dissolution of the polymer. The collector bath was grounded to neutralize the high voltage applied to the polymeric solution. Natural tapioca starch was dissolved in deionized water and the collector bath was filled with ethanol at 20 C. It was reported that upon electrospinning the hydrophilicity of the fibrous web increased due to the enhancement of the amorphous structure by decreasing the ordered polymeric chain. The uniform bead free fibers were formed at the concentration of 4.5 wt%, increasing the concentration resulted in blockage of the spinneret due to an increase in viscosity of the polymeric solution. Water swelling and disintegration study of the web play a major role in drug delivery applications. The water swelling ratio was found to increase with an increase in the concentration of the polymer, thereby representing the hydrophilicity of the electrospun mat. The disintegration study was carried out for starch in the form of powder and fibrous web using dibasic calcium phosphate dihydrate as disintegrating excipient. The powder tends to disintegrate 10 times higher than the fibrous web as the fibrous network resists the movement of water due to an increase in the surface area caused due to swelling of fibers [49]. Solid collector enhances the removal of charges in the polymeric solution created by the electric field by utilization of conducting surface and allows the uniform collection of fibrous webs. On the other hand, liquid collector possesses the disadvantages of transferring charges from the needle tip to liquid due to complicated interaction between the nanofibers and liquid and poor conductivity. Liquid collector offers low bulk density materials compared to solid collectors [50]. The morphology and diameter of fibers can be altered by changing the motion of the collector in the liquid medium. To study the interaction between the nanofibers and the liquid surface, the electrospinning setup was modified with a liquid collector which was given motion by a magnetic stirrer. Polystyrene (PS) solution was prepared by dissolving the 20 wt% polymer in tetrahydrofuran (THF) and N, N-dimethylformamide (DMF) at a weight ratio of 1:3. The modified electrospinning comprises water bath which was grounded by placing grounded electrical wire. To the study, the change in morphology of the nanofibers, the water flow velocity, and conductivity of the water collector were altered by a magnetic stirrer. It was reported that increasing the conductivity of the water by varying the concentration of sodium chloride (NaCl) resulted in the formation of finer fibers due to excellent electric

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conductivity and transfer of accumulated charges during electrospinning. The liquid collector allows the loose construction of nanofibers compared to the solid collectors due to the larger deposition areas resulting in low nanofiber density. The circumferentially aligned fibers were produced by depositing the electrospun web on the surface of the distilled water with liquid flow [51]. The collection and alignment of nanofibers in the electrospinning technique play a crucial role in cell growth and orientation for tissue engineering applications. To control the alignment of fibers, the collector in the electrospinning setup was modified by cutting the aluminum sheet into eight electrodes of the desired shape. The electrodes were spaced at a particular distance and secured between two upper and lower plastic plates by means of a connective screw to control the noise and remove the unused areas. Each electrode was operated individually by a separate power supply. The spinnability of polyacrylonitrile (PAN) was studied by dissolving in dimethylformamide (DMF) with two types of operation. The first type of operation includes operating each electrode separately which resulted in control of fiber alignment and deposition on each electrode whereas the second operation includes operating each electrode simultaneously which resulted in difficulty in controlling the fiber diameter and alignment but with less damage to the fibers. It was reported that operating a pair of electrodes in a clockwise direction resulted in regular and even fibers either preventing the slipping of fibers without causing damage to them as it was due to stretching of fibers between two electrodes [52]. The electrospinning technique utilizes high voltage for the production of submicron fibers by passing the electric field to charge the polymeric solution and create bending instability thereby causing stretching and fiber formation. The electrospinning technique was modified by placing the secondary electrodes at a right angle or parallel to each other in order to study the path of electric field current density to study the deposition of fibers on the collector. Polypropylene nanofibers were spun by dissolving in cyclohexane, acetone, and dimethylformamide (80/10/ 10) and spun using modified electrospinning setup. It was reported that the position of the electrodes has an influence on the potential field thereby altering the current density lines. The modification allows the uniform deposition of nanofibers at the desired location rather than the undesired location thus allowing the mass of nanofibers collected on the collector as it drifts away from the collector and collects on the other objects compared to the conventional electrospinning technique [53]. Cylindrical scaffolds find the application in the area of manufacturing vascular grafts for the treatment of renal diseases, surgical bypass, and dialysis. The commercially available vascular grafts have a diameter greater than 6 mm hence, production of vascular structure below the diameter has been carried out by modification of collector in the electrospinning technique. The alignment of nanofibers decides the growth of cells on the substrate as it mimics the extracellular matrix (ECM). Thus, alignment of nanofibers can be controlled by modification of collector utilizing an insulating layer between the conducting layer as it aids in a longitudinal orientation. But the limitation of the technique includes increasing the gap between the conducting collectors reduced the drawing force acting on the fibrous network resulting in higher diameter fibers. The work was carried out to overcome the

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limitation by producing cylindrical nanowebs with a better longitudinal orientation by modifying the collector in the electrospinning technique. The spinning solution was prepared by dissolving PAN in DMF in the modified electrospinning technique by designing the collector with conductive (metal) and conductive segments (coated with Teflon) as the fibers tend to deposit on the nonconductive surface with longitudinal orientation due to electric field between two adjacent conductive segments. It was seen that increasing the concentration of the polymer and the distance between the needle tip to collector distance resulted in an increase in diameter due to a decrease in electric field thereby reducing the stretching of the polymeric solution and causing higher diameter fibers. It was reported that the orientation of nanofibers in the web was decreased due to an increase in polymer concentration due to bending instabilities, increasing the needle tip to collector distance due to complete evaporation of solvents from the polymeric solution thus, making the fiber too dry for stretching and drawing into finer fibers with longitudinal orientation. The distance between the conductive segments has a major influence on the fiber orientation as the mean length of 10 mm distance resulted in fiber orientation of 87.37 by the spinning of the solution with a polymer concentration of 11 wt%, a voltage of 12 kV, a flow rate of 150 μl/h, conductive length of 5 mm, and a groove depth of 2 mm. The cylindrical nanofibrous web with longitudinal orientation finds the potential in tissue engineering and vascular graft applications [54]. Electrospinning technique is the facile way of deposition of nanofibers on the 3D electroconductive collector as the fibers are aligned onto the recess shape of the 3D nanofibrous macrostructures. But the limitation of the technique includes incomplete/irregular filling of recesses as it is difficult to replicate the complex geometries of the collector. The negative 3D ear cartilage was prepared by pouring polydimethylsiloxane (PDMS) monomer and silicone elastomer hardener in the ratio of 10:1 after degassing onto the template and cured. An electroconductive collector was prepared by pouring gelatin and alginate solution onto the negative PDMS ear cartilage mold. Polycaprolactone (PCL) solution was prepared by dissolving the polymer in chloroform and methanol solvent mixture and electrospun on 3D gelatin-alginate (75:25) which was placed on polymethyl methacrylate flat substrate. In order to ensure the conformability of nanofibers deposition, the outer edges of the hydrogel cartilage were flattened and electrospun. Similarly, another side of the ear cartilage, PCL nanofibers were electrospun. It was reported that hydrogel-based electroconductive collector enables the filling of recesses of ear cartilage mold due to the flexibility of the material causing deposition of uniform thickness of fibrous materials, unlike conventional electrospinning process. Thus conformal filling of nanofiber finds the potential in various biomedical applications such as tissue engineering and drug/cell delivery fields [55].

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6 Conclusion The major challenge in the production of submicron fibers is the dimensions, orientation, and morphology of fibers. All these can be concomitantly achieved by the selection of proper spinneret and collector assembly. The tortuosity the electrospun fibers offer provides excellent adhesion to cells and their proliferation. Moreover, with careful selection of spinneret assembly hollow fibers, porous fibers, and bicomponent fibers can be produced at ease for biomedical applications. The choice of collector assembly decides the orientation of fibers. Moreover, the collector assembly decides the micro properties such as crystallinity. Significant work is being carried out on attaining uniform thickness and coherent mat for applications.

7 Future Trends The art of electrospinning has matured itself and most of the research work has been carried out on the potential biomedical applications. The major challenge left before the researchers are the scalability and attainment of uniform thickness of the mat. More attempts have been made on the spinneret assembly to achieve a coherent mat and the design of the collector assembly also plays a major part in achieving the desired properties. Rotating collectors in a predetermined path like Spirograph can lead to fibers being collected uniformly with good tortuosity. Manufacturers are also attempting to combine electrospinning along with centrifugal spinning and solution blow techniques. The role of solvent selection also plays a major role in the fiber formation leading to fibers with consistent morphology. If the fibers can be produced at nano dimension consistently at a shorter duration, lot of potential applications exist apart from the biomedical applications.

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Adv Polym Sci (2023) 291: 23–36 https://doi.org/10.1007/12_2022_137 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 28 Septemper 2022

Fabrication of Multiscale Polymeric Fibres for Biomedical Applications Nivethitha Ashok, S. Sowmya, and R. Jayakumar

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Fabrication Techniques of Nano/Micro Size Electrospun Nanofibres . . . . . . . . . . . . . . . . . . . . . . . 2.1 Coaxial, Emulsion and Co-electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Edge Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3 Gap Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4 3D Jet Writing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.5 Caged Collector Electrospinning and Moving Spinneret . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Applications of Multiscale Fibrous Scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1 Bone Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 Cartilage Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3 Cardiovascular Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.4 Liver Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.5 Neural Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.6 Skin Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.7 Tendon Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract In the recent years, multiscale fibres have gained widespread attention for use in diverse biomedical applications. With the technological advancements in electrospinning technique, the dimensions of the multiscale fibres can be tailored to match the desired requirements of the target tissues and organs. These fibres combining the unique properties of nano- and microfibres can be fabricated using

N. Ashok and R. Jayakumar (*) Polymeric Biomaterials Lab, Centre for Nanosciences and Molecular Medicine, Amrita Vishwa Vidyapeetham, Kochi, India e-mail: [email protected] S. Sowmya Department of Periodontics, Amrita School of Dentistry, Amrita Vishwa Vidyapeetham, Kochi, Kerala, India

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several techniques. The engineered multiscale fibrous scaffolds are being utilised to deliver growth factors and cells of interest, thereby, aid in tissue regeneration. The elastic modulus of the multiscale fibrous scaffolds can be enhanced by introducing biomaterials like multilayer carbon nanotubes (MCNT), nanoclay, nanohydroxyapatite (n-HAp), and graphene oxide (GO). Moreover, studies on bone, cartilage, cardiovascular and liver tissue engineering have revealed that the multiscale fibrous scaffolds are capable of restoring the microarchitecture of the extracellular matrix (ECM). The current chapter focuses on some of the fabrication techniques which include coaxial, emulsion & co-electrospinning, edge electrospinning, gap electrospinning, 3D jet writing, and caged collector electrospinning & moving spinneret. Further, the application of multiscale fibrous scaffolds with respect to bone, cartilage, cardiovascular, liver, neural, skin and tendon tissue engineering has been discussed. Keywords Multiscale fibres · Microfibres · Cardiovascular tissue engineering · ECM mimicking · Fibrous scaffolds

1 Introduction One of the tactical approaches in tissue engineering field is developing functional scaffolds that would effectually help in repair and replacement of tissues and organs [1–4]. An ideal scaffold for transplantation should be consistent with the anatomical site of implantation, have flexible and foldable properties which allow for better attachment to the irregular surface of the tissues (e.g., smooth muscles, cartilages, tendons, ligament and bone). To be specific, along with the mechanical and structural architecture, the scaffold should assure active nutrient supply, cell adhesion, biological interaction and stable support. Hence, a potent approach in the accurate designing of scaffolds should be based on the biophysical characteristics of organs and tissues [5]. Electrospinning is one such versatile technique, employed in the fabrication of scaffold containing fibres, scaling from several nanometres to micrometres. The ease with which electrospinning technique can be reproduced has been a very big advantage gaining popularity in different fields including textiles, sensors, energy and biomedical applications [6, 7]. Engineered scaffolds with the right ratio of growth factors and cells should have the potency to regenerate when implanted inside the host tissue. In vivo, the cells are composed and exposed to mechanical and chemical factors, cell–cell interactions and structural factors comprising extracellular matrix (ECM). All these aid in the formation and function of tissues. The ECM is well defined with micro and nano scale topography of collagen fibres, which are crucial in the cellular proliferation, migration, differentiation and morphology in vivo [8–13]. There are several studies which have indicated that cells show improved function and behaviour in terms of regeneration, when the scaffold

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employed had been designed in such a way that it mimics the exact topographical organisation of the ECM [14–16]. Various techniques and approaches have been employed, such as 3D printing, lithography and electrospinning, to develop and fabricate scaffolds that resemble ECM [17–19]. In the past few decades, the technique of electrospinning has garnered a lot of attention as it resembles the ECM geometry. The micrometre scale helps in cell migration whereas the nano scale helps in protein interactions, thereby aiding in cell viability and adhesion [20, 21]. The study by Shalumon et al. evaluated the biological characteristics of electrospun porous polylactic acid (PLA) multiscale scaffolds. It was found that there was enhanced cell penetration and proliferation, which was owed to the combinatorial effect offered by nano and micro fibres, respectively [22]. Hence it can be concluded that an ideal scaffold should possess all the required characteristics that aid in cell growth and regeneration. This review specifically focuses on the different techniques employed in the fabrication of multiscale polymeric fibres and the various aspects of their biomedical applications [23].

2 Fabrication Techniques of Nano/Micro Size Electrospun Nanofibres In the concept of altering the internal structure and the size of a scaffold, there arrives a concept of altering its material to favour more cell adhesion and viability. Natural and artificial polymers are widely explored in tissue engineering applications. Natural polymers such as collagen, fibrin, chitosan, alginate, etc. are good in cell adhesion and proliferation and show good biodegradability as well. Synthetic polymers are favourable for fine-tuning biodegradation and immunogenicity. It can also be easily reproduced [24]. In the process of electrospinning, there are many parameters which influence the fibre surface morphology and other properties. It affects the stable formation of fibre, size, fibre position and so on. The stable fibre formation is based on three boundaries: (1) conductivity, which depends on the content of salt, (2) surface tension which is based on solvent and concentration of polymer and (3) viscosity of the solution which is based on polymer molecular weight and polymer concentration. The grounded collector also plays a vital role. Fibre arrangement can be modified with various collectors such as drums, flat plates, disk collectors and so on [25].

2.1

Coaxial, Emulsion and Co-electrospinning

Coaxial fibre consists of two layers (i.e. core and shell) wherein the core and shell are two separate entities. These fibres are formed with two different solutions. The surface modification can be done without affecting the core material [26]. The

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good part of using a coaxial method is the two phases of fibre swank with their properties and assert two unique properties of fibres. This method allows for electrospinning of the non-spinnable material using polymer–polymer, polymer– inorganic and inorganic–inorganic coaxial fibre [27]. The unique properties of oligomers, metal salts and enzymes at the nano-level maybe limited by various factors such as solvent conductivity, molecular weight, high surface tension and solubility. Thus this particular method helps to cover and accommodate the core and so on [28]. In biomedical research, enzymes, which is an unstable material is used as a core material [29]. The emulsion electrospinning technique uses a material with properties like hydrophilicity, ductility and low elastic moduli materials by introducing biomaterials [30–33] like nano-composites [34], multilayer carbon nanotubes (MCNT) [35], nanoclay [36], nano hydroxyapatite (n-Hap) [37], nano-biopolymers [38] and graphene oxide (GO) [39]. Co-electrospinning makes use of two different spinnerets with different natural and synthetic polymer solutions with different molecular weight. Two spinnerets eject the polymer solution when high voltage is applied and it is deposited on the collector. Depending on the properties and various factors, the diameter of fibre varies from nano to micron size [40–42].

2.2

Edge Electrospinning

This process utilises the narrow cylinder shape collector which has a hollow space inside. The spinnerets (bowl shape) are set inside the collector. Multiple orifices are made in the spinnerets. These spinnerets are subjected to high pressure and high voltage after feeding the polymer solution into the hollow shaft. The Taylor cone jet is formed at all spinnerets and the bowl starts to rotate with adjustable multi-speed, thus, multiple narrow fibres are formed and deposited at the inner wall of a collector. Kolos Molnar et al. fabricated poly-vinylpyrrolidone-based nanofibres [43]. N M Thoppey et al. fabricated polyethylene oxide (PEO) and PCL nanofibres, using bowl edge electrospinning method. The diameter of the obtained fibres was between 200 and 400 nm with a porosity of 67% (approx.) which was similar to the traditional single-needle electrospun fibre [44]. This method is efficient as it prevents the local spillage of the polymer solution.

2.3

Gap Electrospinning

It is a new, emerging and less investigated technique. Aligned fibres can be achieved via this method. Two negatively charged plates are arranged vertically and parallel with a constant distance, which acts as a collector. The polymer solution is ejected with a high voltage and on another side the collector collects the polymer solution forming a fibre. A slight drawback of this system is that fibres can be fabricated only

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for few tens of centimetres. On the other hand, aligned fibres can be fabricated which is suitable for soft tissue regeneration like tendon, etc. [45, 46].

2.4

3D Jet Writing

Precise stacking and aligned microfibres of biomaterials such as PLGA, PCL can be fabricated; less branching of fibre is achieved with the help of moving electron in 2d platform. It helps micron sized fibre to deposit on top of each fibre precisely. Jacob H. Jordahl et al. made a microfibre of poly(D, L-lactic-co-glycolic acid) for calvarial defect regeneration and to study cancer metastasis using small animals [47].

2.5

Caged Collector Electrospinning and Moving Spinneret

The specific architecture of fibre can be achieved by changing the shape and size of spinnerets. Using a hollow caged collector helps in aligning the fibre diameter. Jian Xie et al. fabricated Poly(L-lactic acid) (PLLA) nano and microfibres with diameter varying from 0.6 to 1.2-μm metre. Bone marrow stem cells (BMSC) were seeded onto the scaffold to investigate cell adhesion, viability and differentiation [48]. Bin Sun et al. developed poly(3,4-ethylenedioxythiophene): poly(styrene sulfonate)poly(vinyl pyrrolidone) (PEDOT:PSS-PVP) combination microfibrous array by using a moving spinneret and static collector, which has an advantage of making aligned fibres. This technique is suitable for fibres made of elastic material and also has application in semiconductor industries [49].

3 Applications of Multiscale Fibrous Scaffolds 3.1

Bone Tissue Engineering

The key objective of bone tissue engineering is to repair and regenerate the bone. Despite the inherent capacity to regenerate, bones are not potent enough to facilitate regeneration in the case of heavy injuries [50]. Bone tissue engineering has seen great evolution in the development of novel biocompatible materials from the first generation of inert biomaterial scaffolds to the second generation of bioactive materials facilitating biochemical reactions on them to the development of third generation biomaterials mimicking the ECM [51]. Tuzlakoglu et al. developed a new type of ECM mimicking scaffold by merging the nano- and micro-fibres for bone tissue engineering applications. He used the techniques of fibre bonding and electrospinning to fabricate this structure which is a blend of starch/polycaprolactone (SPCL, 30/70 wt%). The microfibre SPCL

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scaffolds were impregnated with electrospun nanofibres. The electrospinning of nanofibres randomly between the microfibres formed the nanobridges and hence restored the microarchitecture of the ECM. The human osteoblast-like osteosarcoma SaOs-2 cells and rat bone marrow cells exhibited well-spread morphology, cytoskeletal organisation, cell viability and alkaline phosphatase activity. The results suggested that these multiscale scaffolds can aid in 3D bone tissue engineering [52]. Another important concern of bone tissue engineering (BTE) is vascularisation post-implantation. Santos et al. investigated the effect of multiscale scaffold with a blend of starch/polycaprolactone on the adhesion, morphology, growth pattern, inflammatory expression profile and expression of structural proteins of the endothelial cells. The nano/microfibres showed the expression of structural proteins like vimentin and PECAM-1. These fibres also facilitated the migration and organisation of endothelial cells. These multiscale scaffolds aided in the 3D distribution of endothelial cells without compromising bone regeneration [53].

3.2

Cartilage Tissue Engineering

Cartilage tissue engineering plays a significant role in recapitulating the cartilages that have poor tendency to regenerate and repair. It becomes important to combine cells, scaffolds and other vital factors to promote cartilage regeneration [54]. Recent advances in cartilage tissue engineering introduce the electrospun nanofibres that are capable of mimicking the ECM components. But, it has a complication of reduced cellular infiltration. Leverson et al. tried a method to overcome this difficulty by studying the effect of electrospun multiscale scaffold comprising of two different materials regularly oriented but in two different scales. The multiscale electrospun scaffolds consist of fibrin nanofibres and polycaprolactone microfibres produced using dual electrospinning. The usage of natural materials like fibrin adds to better biocompatibility and biodegradability of the scaffold, thereby aiding improved cellular response. The nanoscale fibres mimic the ECM structures and further promote cellular attachment and spreading. The efficiency of the multiscale scaffold was measured through glycosaminoglycans (GAGs) production, which is an important component of cartilaginous ECM. The histological study results infer that there was also better deposition of proteoglycans when multiscale scaffolds were used. All these results prove that the performance of the multiscale scaffolds was better compared to the previously used microfibres alone. The result suggests that advances in electrospinning prove to help the field of tissue engineering by bringing both nanoscale and multiscale fibres together in orchestrating the cellular function and response [24].

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Cardiovascular Tissue Engineering

The myocardium or the cardiac muscle tissue plays a key role in the functioning of the heart as it aids in pumping the blood throughout the body. Cardiovascular tissue engineering is of great significance in restoring the heart tissue after heart attacks and injuries [55]. Sreerekha et al. developed a multiscale scaffold that promotes myocardial regeneration. The multiscale scaffold consists of nanofibres of size 50 to 300 nm made of fibrin and microfibres of size 2 to 4 μm made of poly(lactide-coglycolide) (PLGA) that structurally mimics the hierarchy of heart tissue. The scaffolds were fabricated using the simultaneous electrospinning of fibrin and PLGA with a rotating mandrel. The scaffolds exhibited degradation of fibrin and deposition of ECM. The survival and morphology of hMSCs grown on the fibres were maintained. The cells seeded on the multiscale scaffolds exhibited better cell attachment, spreading and survival compared to pure PLGA fibres. The sequential electrospinning significantly contributed to an enhanced cell infiltration. The cells on the composite scaffold were more viable and proliferative. Further, the multiscale scaffolds stimulated the differentiation of umbilical cord derived MSCs which was confirmed by the expression of cardiac specific proteins, namely troponin, tropomyosin, α-sarcomeric actinin, desmin and atrial natriuretic peptide. The synergistic effect of the microfibres and nanofibres promoted the differentiation of hMSCs into cardiac phenotype [56]. Kook et al. suggested that multiscale scaffolds comprising of electrospun PCL fibres and two layers of encapsulated fibrin and alginate hydrogels promote regeneration of cardiovascular tissue. The multiscale scaffold accomplished a triple cell culture system comprising of adipose-derived mesenchymal stem cells (ADSCs) with C2C12 myoblasts on PCL fibres and human umbilical vein endothelial cells (HUVECs) on fibrin hydrogel. The results suggested that the multiscale scaffolds stimulated the formation of new blood vessels in vitro which was evident from the expression of CD31 marker and that of cardiac phenotype. The stem cells in the implanted scaffolds expressed cardiac muscle specific genes and HUVECs in the scaffold produced blood vessels. Upon transplantation, the stem cells in the scaffolds secreted anti-inflammatory cytokines like IL-10 and TGF-β. Thus, multiscale scaffold proves to be an emerging strategy for cardiovascular regeneration [57].

3.4

Liver Tissue Engineering

Chronic liver disease is one of the leading causes of death worldwide, estimating to a mortality rate of two million deaths per year [58]. Currently liver transplantation is the only available option to overcome liver diseases to an extent. But, the success of the transplantation has been hindered by several factors like shortage of donors, limited number of hepatocytes and poor engraftment of the transplantation. But the huge demand for the donors creates the need to find alternative strategy like liver

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tissue engineering that can help in fabricating liver tissue for transplantation [59]. Liver tissue engineering employs the usage of hepatocytes or stem cell derived hepatocyte-like cells combined with biomaterial scaffolds which orchestrate together in restoring the hepatic system [60]. Kim et al. fabricated hepatic patch by stacking patient specific liver cell sheets formed on multiscale electrospun fibres as a mode of regenerative therapy for liver injuries. The liver cell sheets have alternating cell sheets of human chemically derived hepatic progenitor cells (hCdHs) and Human Umbilical Vein Endothelial Cells (HUVECs). This helped in mimicking the hepatic system by not only promoting the functional aspects of the patch but also enhancing the therapeutic effects in acute liver damage models. The electrospun sheets enhanced the functions of liver like secretion of albumin, cytochrome p450 activity and hepato differentiation of the chemically derived hepatic progenitor cells. The results suggest that the hepatic patch on transplantation has proved to show better hepatic repopulation and survival in mice [61]. The next level development in the field of liver tissue engineering is the model liver tissues that bring about the concept of functional bio-artificial liver support system. Verma et al. fabricated scaffolds by combining electrospun nanofibres and hollow fibre membranes. This was developed to check the performance of the scaffolds in overcoming the greatest challenge of liver tissue engineering, i.e. to support the function and maintenance of hepatocytes. The multiscale fibre matrix comprises the electrospun biocompatible nanofibres of polycaprolactone (PCL), gelatin and chitosan covering the hollow fibre membrane. This multiscale fibre matrix proved to be hemocompatible and also showed better cell adhesion and proliferation of HepG2 cells. The synergistic effect of the hollow membranes and the nanofibres deposited on them contributes to the enhanced mechanical strength as that of the ECM [62].

3.5

Neural Tissue Engineering

The vertebrate nervous system is a very complex system that plays a crucial role in the motor sensory pathway. Any damage or injury to the system can cause fatal consequences [63]. The limited regenerative capacity and the complicated physiology of the nervous system are the huge challenges in treating the nervous injuries. Neural tissue engineering is an emerging strategy which combines the fields of biology, neuroscience, material science and engineering in order to develop substitutes that bio-mimic the structural and functional nervous system in native state thereby aiding in regeneration of the damaged tissues [64]. Panseri et al. studied the effect of electrospun multiscale PLGA/PCL fibre tubes on neuronal regeneration of a 10 mm gap in the sciatic nerve. The size of the multiscale fibres ranged from 280 nm to 8 μm. The fibrous microstructure was biodegradable and was easily sutured in the ends of the nerve stumps. The graft hindered the infiltration of other cells into the conduit and the pores in the graft

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facilitated the movement of nutrients. The results after 4 months of surgery concluded that the multiscale scaffolds induced neuronal regeneration and reconnection of the 10 mm gap in the sciatic nerve, in comparison with the control groups. The nerve tissue regeneration was accompanied by myelination and collagen deposition. The conduit proved to be biocompatible since there was a formation of fibrous tissue capsule with less inflammatory response. It is quite evident from the migration of Schwann cells that these implants are novel and efficient tools for regeneration that overcome the challenges of neural tissue engineering, as Schwann cells aid nerve growth, innervation and nerve regeneration [65].

3.6

Skin Tissue Engineering

Tissue engineered skin constructs face huge demand as they mimic the structural and functional features of skin, further aiding the process of wound healing. Skin tissue engineering unravels novel techniques to fabricate combinations of biomaterial and scaffolds that can efficiently restore the ECM environment [66]. Leong et al. developed electrospun PCL multiscale fibres with interweaving micro (3.3 ± 0.6 μm) and nano (240 ± 50 nm) fibres. The multiscale fibres showed enhanced infiltration of human dermal fibroblasts (HDF). Further, significant expression of HDF secreted ECM proteins on the multiscale scaffold is suggestive of a promising potential in skin tissue regeneration [67]. Kim et al. developed 3D PLGA multiscale scaffolds that find application in dermal tissue engineering. These multiscale scaffolds were fabricated using hybrid electrospinning. The multiscale fibres constitute the electrospun PLGA nanofibres of size 530 nm and PLGA microfibres of size 28 μm. The normal human epidermal keratinocytes (NHEK) and normal human epidermal fibroblasts (NHEF) showed better growth and attachment on the multiscale PLGA fibres compared to the microfibres alone. The nanofibrous architecture complements the microscale fibres in terms of better mechanical strength. The results suggest that multiscale scaffolds can emerge as a novel strategy for skin regeneration [68].

3.7

Tendon Tissue Engineering

Tendons are the fibrous connective tissues that play integral role in the function of the musculoskeletal system. It helps in the movement of the bone by attaching the bone to a muscle. These tissues are prone to wear and tear due to their load bearing properties. Exposure to severe tension and trauma can also lead to tendon injures. Hence, tendon tissue engineering involves the fabrication of an ideal scaffold that mimics the mechanical properties of the tendon and thus aids in tendon regeneration [69].

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Jayasree et al. developed multiscale fibrous scaffolds of PCL micro/collagenbFGF nanofibres (mPCL-nCol-bFGF) by electrospinning. Multiscale fibres were fabricated by spinning PCL microfibres and collagen nanofibres simultaneously at room temperature. The introduction of collagen fibres helps in recapitulating the native tendon and aids in tenocyte migration and proliferation. The incorporation of the basic fibroblast growth factor (bFGF) into the nanofibres enhances fibroblast proliferation. The nanofibres also aid in the sustained release of the growth factor. The mechanical strength of the scaffold is very important in tendon tissue engineering. This work also saw the implementation of technical improvement technology like braiding that enhanced the mechanical properties of the electrospun scaffolds. The microfibres were able to bring back the mechanical features whereas the nanofibres facilitated cellular attachment. bFGF promoted tenocyte proliferation and favoured the expression of tenocyte markers. The loading of bFGF helped in the upregulation of non-ECM components like tenascin C and fibronectin that are involved in cell signalling and survival. The result suggests that micro/nanofibres are a better strategy for tendon regeneration [21]. Sensini et al. fabricated multiscale scaffold comprising of bundles of resorbable electrospun nanofibres of Poly-LLactic acid (PLLA). They demonstrated that the nanofibres favoured cell growth and migration in the interior of the scaffold thereby providing an ideal environment for cell proliferation [70].

4 Conclusion Numerous ways are being employed to alter topographic properties such as electrospinning, 3D printing, Plasma assist and Lithography. Fabricating multiscale electrospun fibres using various methods with structural changes is in great demand for mimicking the natural ECM and for better tissue regeneration. Micron diameter sized fibre contributes to sufficient pore size, while nanofibre favours cell to cell interactions and cell adhesion. Fabricating scaffolds using two different polymers has been the main method for producing such multiscale fibres due to the unique properties of the two different polymers. On the other side of the coin, engineering the fibre alignment is its drawback. Various 3D structured nano and microfibres have been developed for tissue engineering of nerve, cardiac, bone, neural, skin, tendon, liver, etc. The multiscale fibres are promising for providing many biological functions and to a great extent in vascularisation. The interaction of ECM with the fibrous system governs microtubule formation, its orientation and cell architecture. Future work should focus on mimicking the complex ECM of organs which would also enhance the biomechanics and chemical effects. Researchers should also focus on the development of multiscale aligned fibres which could help in tendon regeneration, and also favour the enhancement of other functionalities on the multiscale fibres.

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Adv Polym Sci (2023) 291: 37–68 https://doi.org/10.1007/12_2022_140 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 11 October 2022

Techniques to Fabricate Electrospun Nanofibers for Controlled Release of Drugs and Biomolecules Monika Rajput, Suhela Tyeb, and Kaushik Chatterjee

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Electrospun Fiber Fabrication Techniques and Mechanism of Biomolecule Delivery . . . . . 2.1 Electrospinning Process . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Application of Electrospun Nanofibers for Therapeutic Delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1 Transdermal and Wound Dressing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 Drug Delivery Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3 Growth Factor Delivery System . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.4 Gene Therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract Electrospinning is a nanotechnology-based platform with great potential in tissue engineering and related biomedical applications. The high surface-to-volume ratio of nanofibers makes them ideal candidates for the controlled delivery of drugs and biomolecules. The successful loading of multiple drugs and their controlled and sustained release at the targeted sites renders them suitable for guided tissue engineering applications. This chapter provides an overview of the electrospinning process and its advanced modification for successful encapsulation of drugs, biomolecules, and gene products to develop nanofiber-based therapeutic systems for biomedical applications, particularly focusing on the repair and regeneration of tissues. The gaps in the field and opportunities for future research are highlighted. Keywords Controlled release · Core-sheath nanofibers · Drug delivery · Electrospinning

M. Rajput, S. Tyeb, and K. Chatterjee (*) Department of Materials Engineering, Indian Institute of Science, Bangalore, Karnataka, India e-mail: [email protected]

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Abbreviations Ac-DEX BMP-2 BSA DDS EGF FDA HA MMP NaAlg PCE PCE PCL PDGF PdLA PDLLA PEG PEI PEO PGA PHBV PLA PLCL PLGA PLLA PU PVA PVP VEGF

Acetylated dextran Bone morphogenic protein-2 Bovine serum albumin Drug delivery system Epidermal growth factor Food and drug association Hydroxyapatite Matrix metalloprotein Sodium alginate Poly(ε-caprolactone)-co-poly(ethylene glycol) Polycarboxylate Poly (ε-caprolactone) Platelet-derived growth factors Poly(D, L-lactide) Poly(D, L-lactic acid) Poly(ethylene glycol) Poly(ethylenimine) Poly(ethylene oxide) Poly(glycolic acid) Poly(hydroxybutyrate-co-hydroxyvalerate) Poly(lactic acid) Poly(l-lactide-co- ε-caprolactone) Poly(lactic-co-glycolic acid) Poly (l-lactic acid) Polyurethane Poly(vinyl alcohol) Polyvinylpyrrolidone Vascular endothelial growth factors

1 Introduction Nanotechnology is the branch of engineering science involving nanoscale materials, which are finding increasingly greater use in biomedical applications [1]. Nanomaterials with excellent physicochemical properties and minimum toxicity can sense the local environment and stimulate biological responses to achieve the desired therapeutic outcomes. Zero-dimensional (carbon dots, quantum dots), one-dimensional (nanowires, nanotubes, nanofibers), and two-dimensional (graphene oxide, transition metal oxide, MXenes) nanomaterials have been widely investigated for diagnosis, imaging, and targeted therapeutics/drug therapy [1, 2]. Nanofibers, which are fibers with nanoscale diameters, have gained much

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attention as a promising platform for targeted delivery of drugs and therapeutic molecules delivery for various diseases [3]. Nanofibers mimic the structure of the native extracellular matrix of human tissues with several attractive attributes such as a high surface area-to-volume ratio, tunable porosity, and interconnectivity, controllable surface morphology, ability to functionalize with bioactive factors, high loading and entrapment efficiency, and ability to tailor for highly efficient delivery of a wide variety of biological therapeutics [4, 5]. A wide variety of materials have been studied for preparing nanofibers, such as natural and synthetic polymers, inorganic nanomaterials, composites, and biomolecules, which makes them a robust and promising candidate for healthcare and advanced biomedical applications [3, 5–7]. To date, there are several fabrication platforms that have been adopted for one-dimensional nanofiber synthesis, especially for targeted specific therapeutics delivery such as emulsion processes, phase separation, self-assembly, spray drying, and electrospinning [4, 8, 9]. However, these conventional drug/compound encapsulating techniques have inherent drawbacks. High-temperature control and intense labor are required in the emulsion process, accumulation of polymer or merging of droplets before hardening during phase separation, and bead formation during spray drying with loss of biological activity, rendering these methods unsuitable for mass manufacturing [6, 10]. Among all techniques, electrospinning has gained more attention for the large-scale production of nanofibers. The electrospinning method was first patented by Gooley in 1900 and 1902 [11, 12] and further revived by Reneker et al. for synthesizing one-dimensional nanostructures [13]. Electrospinning is a versatile, simple, cost-effective, efficient, and scalable approach to fabricate continuous non-woven polymeric nanofibers [5, 14]. The availability of a wide variety of electrospinnable polymeric material, fabrication of ultrafine fibers ranging from microstructures to nanostructures with a high surface-to-volume ratio, good porosity, interconnected structures, and excellent tensile strength make electrospinning a demanding method for various biomedical applications [15, 16]. Electrospinning utilizes the strong electrostatic repulsive force between the surface charges generated by the high voltage to synthesize continuous micro- to nanoscale fibers from the polymer solution [3, 8], as shown in Fig. 1. Table 1 compiles the key parameters that influence the electrospinning process. A wide variety of polymers, both natural and synthetic, ceramics, small molecules, and their blended combinations are used for nanofiber synthesis [8, 10, 17]. Aside from the more popular one-dimensional solid round nanofibers, two-dimensional nanofibers, including porous, hollow, and core-sheath structures, have been generated with the electrospinning method. The surface of the nanofibers can be functionalized during and after nanofiber synthesis while controlling the fiber structure, morphology, and spatial orientation by molecular moieties [18, 19]. Electrospun nanofibers have a wide variety of applications, including tissue engineering (wound dressing, drug/ compound delivery) [20], solar cells [21], supercapacitors [22], and sensors [23]. To date, various types of electrospinning techniques such blend electrospinning [24, 25], emulsion electrospinning [26–28], multiaxial electrospinning [29], and secondary carrier electrospinning have been used to load small-molecular drugs, various therapeutic agents, including proteins, DNA, and RNA for targeted delivery

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Fig. 1 Schematic of the conventional electrospinning process and the application of nanofibers in healthcare and biomedical applications. Figure reproduced from Shahriar et al. [1] as per CC-BY 3.0

Table 1 The influence of the electrospinning process and solution parameter on the properties of electrospun nanofibers Parameter Applied voltage Flow rate Tip-to-collector distance Type of the collector Concentration of solution Conductivity of solution Viscosity of solution

Effect High voltage generates thin fibers Increase in flow rate is associated with an increase in fiber diameter Influences the fiber diameter; shorter the distance, larger the diameter and longer the distance, smaller the diameter of fibers Random and aligned nanofibers could be fabricated by changing the type of collector. Random fibers formed over plate collector, whereas a drum and cylinder are yielded aligned fibers A higher concentration produces nanofiber with a large diameter. Low concentration leads to the sputtering of polymeric material High conductivity yields thinner fibers with fewer beads

Reference [39] [40, 41]

High viscosity yields large diameter continuous fibers, whereas low viscosity yields thin fibers

[47, 48]

[40, 42] [43, 44]

[45, 46] [47]

applications [30, 31]. However, the initial burst release of entrapped molecules is a common limitation with blended nanofibers, which leads to less effective recovery of the tissue on its implantation [6, 32]. Various chemical modifications or crosslinking methods for controlled and sustained release of molecules lead to poor biocompatibility [32]. So, to overcome the drawback of burst release from nanofibers, several researchers now focus on using coaxial electrospinning to fabricate core-sheath nanofibers. Coaxial nanofibers offer numerous advantages over one-dimensional round nanofibers [33]. Drugs or small molecules may be easily loaded and protected from a harsh environment. The tailored thickness of the core

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and sheath could provide the controlled and sustained release of molecules. They mimic the structure of the extracellular matrix and enable the modulation of the mechanical or physical properties through the choice of two different materials. For tissue scaffolds, a more cytocompatible material could be selected for the sheath layer and a different polymer solution for loading the drug in the core [34, 35]. The core-sheath double layer structure not only provides the controlled release of small molecules but multiple drugs and therapeutics can be loaded to enable synergistic or combinatorial treatments [36]. Unlike other electrospinning techniques, in the coaxial electrospinning method, fibers can be prepared with a material that is not spinnable alone along with other spinnable materials for generating multifunctional nanofibers [37, 38]. There is a wide range of natural and synthetic polymers available for the fabrication of electrospun scaffolds for tissue engineering applications [49, 50]. The widely used biomedical polymers are compiled in Table 2. The choice of material depends on the material’s physical and chemical properties for feasibility of electrospinning and the application it will be used for. Natural polymers are difficult to handle during electrospinning which required additional modification for better mechanical stability and processability during electrospinning [50]. However, natural polymers are more tend to change their biological characteristics during Table 2 Polymeric materials used in electrospinning to fabricate nanofibers and their applications Polymer Collagen I Collagen/PHBV Chitosan Chitosan/kefiran Chitosan/PCL Chitosan/PVA Silk fibroin Gelatin Gelatin/P.U. Gelatin/fibrinogen Gelatin/chitosan Alginate Alginate /lavender Alginate/chitosan Alginate/PCL PCL PCL/gelatin PCL/PEO PLA PLLA PLGA PVP/Zein

Application Skin tissue engineering; wound healing Tissue engineering To prepare non-woven fabrics and biomedical applications Biomedical applications Liver tissue engineering Improving cell proliferation Biomedical application, bone/skin tissue applications Tissue engineering Wound healing Myocardial regeneration Skin tissue engineering Biomedical applications Burn wound healing Bone tissue engineering Stem cell engineering Skin tissue engineering Bone tissue regeneration Drug delivery, biomedical application Drug delivery Cardiac tissue regeneration Drug delivery and biomedical application Drug delivery

Reference [56–58] [59] [60] [61] [62] [63] [64–66] [67] [68] [69] [70] [71, 72] [73] [74] [72] [75] [76] [77] [20] [78] [79] [80]

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electrospinning process whereas synthetic polymers due to their physical properties are easier to use for electrospinning. To date, PCL [51], PLA [52], PGA [53], and their copolymers PLGA [54] are among the synthetic polymers that have been used extensively in electrospinning fabrication technique for various tissue engineering and biomedical applications [55]. This chapter focuses on the application of coaxial electrospun fibers for tissue-specific drug/compound/gene delivery for different tissue-specific applications. It aims to present an overview of the nanofiber-based therapeutic delivery systems and their biomedical application and future perspectives, particularly highlighting the utility of coaxial fibers.

2 Electrospun Fiber Fabrication Techniques and Mechanism of Biomolecule Delivery 2.1

Electrospinning Process

Electrospinning is a promising, highly productive, and cost-effective technique to generate ultrafine fibers, which can easily be employed in the laboratory and also scale up to industrial process [4, 8]. In the electrospinning process, electrostatic repulsion forces are utilized to spin the polymer solutions to generate continuous fibers from the micrometer range to the nanometer range. It consists of three major components: (1) a spinneret, which controls the flow rate of polymer solution and goes to a high electric field, (2) a high voltage source, which stretches the polymer solution into the ultrathin fibers, and (3) a conductive collector, collects the electrospun fibers in static and dynamic conditions [1, 81]. During the electrospinning process, high voltage is applied between the spinneret and collector. The polymer solution droplet at the tip of the needle becomes charged and generates a Taylor cone due to the surface tension of the droplet. When the applied electric field overcomes the surface tension of the droplet, a finely charged jet is ejected from the Taylor cone and grows longer and thinner, which results in solidification of the polymer due to solvent evaporation, and the ultrafine polymer fibers are then collected on the collector [5, 14]. There are two orientations of electrospun nanofibers: (1) random and (2) aligned [10]. Random nanofibers are generated using a static plate collector, whereas aligned nanofibers are produced using a cylindrical collector (rotation at high speed) or magnetic collector [15]. The physical properties of the electrospun nanofibers, such as high surface-to-volume ratio, fiber diameter, porous structure, and surface morphologies, can be influenced by modulating the process, solution, and post-processing parameters [11, 18]. The process parameter includes applied voltage, flow rate, tip-to collector distance, and the use of coaxial/triaxial needles to generate hollow, core-sheath, multi-sheath structures; solution parameters are the molecular weight of polymers, conductivity, viscosity and dielectric constant of the polymer solution, temperature, and humidity. The post-

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processing parameters are heating rate, and temperature for inorganic materials, and environmental factors, such as temperature and humidity [1, 33, 82]. In general, drug or biomolecule delivery can be accomplished in two ways, either direct encapsulation of the biomolecules within nanofibrous scaffolds or loading of nanocarriers tagged or encapsulated with drugs/biomolecules (micelles, nanosphere, nanorods, etc.) in electrospun scaffolds [81, 83]. Electrospun nanofibrous scaffolds have attracted attention in the field of drug delivery or targeted gene delivery in biomedical applications [43]. However, to generate the drug/biomolecule-loaded nanofibers, the primary challenge is the selection of the fabrication method and modulation of other parameters (degradation, swelling, and release profile). Production of biomolecule-loaded nanofibers by electrospinning can be categorized into (1) Physical adsorption, (2) Blend electrospinning, (3) Melt electrospinning, (4) Covalent immobilization, (5) Emulsion electrospinning, and (6) Coaxial electrospinning, as compiled in Fig. 2.

2.1.1

Physical Adsorption

Physical adsorption is the simplest method to prepare electrospun scaffolds loaded with biomolecules or nanocarriers [85]. In this method, the biomolecule or drug is adsorbed on the nanofibers by dip coating. The fabricated electrospun nanofibers are directly dipped in the pure solution or emulsion of the molecule, and they associate with the fibers via electrostatic forces, hydrogen bonding, or van der Waals forces [19, 85]. However, the main drawback of this technique is the uncontrolled bulk release of adsorbed biomolecules before the desired time frame of release profile for maximal therapeutic benefit. It has been reported that the loading of BMP-2 adsorbed onto the PLGA/HA scaffolds with 100% loading efficiency reached 95% release by day 15 [86]. In case of gene delivery, plasmid DNA adsorbed onto the multilayer (3-aminopropyl)-4-methylpiperazine end-capped poly (1,4-butanediol diacrylate-co-4-amino-1-butanol) induced >90% transfection efficiency within initial 10–15 h indicating burst release of vector-DNA [85, 87].

2.1.2

Covalent Immobilization

In this approach, the drugs/biomolecules are immobilized onto the surface of the electrospun nanofibers via chemical modification [88]. In contrast to physical adsorption, this method yields a more stable interaction. Once the surface of the nanofibers is chemically modified, a stable covalent bond is formed between the nanofiber surface and drug/biomolecule. This approach is a two-step process, the first is to yield functional groups (-OH, -COOH, and -NH2) on the nanofiber surface, and the second is the covalent reaction between the biomolecule and functional groups [88]. There are two most common types of chemical functionalization: (1) aminolysis and (2) hydrolysis. Aminolysis incorporates free amine groups on the nanofibers by forming a covalent bond between a diamine and polymer surface.

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Fig. 2 Fabrication of functionalized nanofiber employing different strategies (a) physical adsorption, (b) blend electrospinning, (c) coaxial electrospinning, and (d) covalent immobilization. The figure reproduced from Feng et al. [84] with permission from Elsevier (License no. 5353761164086)

Hydrolysis involves acid or base treatment to generate functional groups on the polymer surface [85]. Typically, the covalent immobilization approach is used to modify the surface property of polymers. However, this approach is also utilized by

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researchers to immobilize the biomolecules for targeted delivery and to obtain a controlled and sustained release profile [89]. The studies reported that BSA immobilization on nanofibrous structures showed partially controlled release instead of burst release from the scaffold [90, 91]. Further, the other group was utilized to immobilize EGF on the nanofibers. Consequently, efficient delivery and release kinetics were obtained for the successful application of these scaffolds for regeneration in vivo [92]. Many groups developed MMP-responsive DNA, siRNA, and EGF immobilized nanofibers for the controlled release of biomolecules in response to high MMP concentration [93]. However, the covalent immobilization approach is complex, and during immobilization, it leads to changes in the physical properties of the fabricated nanofibrous scaffold such as compromised mechanical integrity due to loss of uniformity of the scaffolds. This approach represents an option for immobilizing multi-drugs or biomolecule onto the scaffold for controlled delivery to the targeted site.

2.1.3

Blend Electrospinning

Blend electrospinning is the most conventional approach to generate drug/biomolecule-loaded nanofibrous scaffolds [32]. In this approach, drug/biomolecules are mixed directly (dispersed or dissolved) in the polymer solution and used in the electrospinning process to generate hybrid scaffolds loaded with the molecules [24, 32]. With the aim to achieve a sustained release profile of encapsulated biomolecule, the hydrophilic polymers and their blends such as PEG, PVA, and gelatin has been used largely. The biomolecules are localized within the fabricated nanofibers, allowing a more sustained release profile. Various hydrophobic drugs, proteins, and genes such as BSA, EGF, VEGF, PDGF, antibiotics, and cytostatic agents were incorporated within the nanofibers using blend electrospinning [85, 94, 95]. This approach can meet the time frame for controlled and sustained release kinetics for wound regeneration, as a sustained release over several weeks can be obtained [96]. Although this approach is simple and highly reproducible, however, the activity of biomolecules incorporated within nanofibers can be highly compromised. The processes of homogenization and ultra-sonication to create emulsions can affect biomolecule activity. The organic solvents used for biomolecule dispersion in polymer solution can lead to changes in the conformation of the biomolecule and can lead to denaturation and loss of bioactivity. Especially the delivery of protein is highly compromised using blend electrospinning due to harsh environmental conditions used for encapsulation [97]. Most of the biocompatible polymers used in blend electrospinning are PCL, PLGA, PU, which dissolve in organic solvents, and incubation of proteins in polymer mix leads to the change in protein structures [32]. This technique is reported to be associated with low bioactivity of the delivered molecules for protein-based therapeutics [97]. Moreover, during electrospinning, the biomolecules distribute at the surface of nanofibers rather than within the nanofibers due to their inherent charges, which may lead to the rapid release of biomolecules [85]. To maintain the biomolecule stability, various

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strategies have been adopted, such as the addition of hydrophilic additives, which minimizes the interaction between the biomolecule and organic solvents, the use of salt complexation instead of emulsification, and the incorporation of hydrophilic polymers PEG in aqueous solution. The regulation of crosslinking between the hydrophilic linkers such as PEG-based linkers with polymer enables the regulation of time release of active molecules [85]. One study showed the time release profile of model active molecule from the PVA nanofibers, which were stabilized by PEG crosslinkers [98]. With the advancement in electrospinning technique, natural polymers such as silk fibroin have also been used for delivery of BMP-2 for bone tissue engineering [99]. Similarly, another group also showed the temporal programmed multi-agent release for combined therapy with drugs such as curcumin and doxorubicin within PEG-PCL and PVA, respectively [100, 101]. The typical release profile of drugs from blend electrospun scaffolds is at first an initial burst release followed by a sustained release. The release of drugs from nanofibers is controlled either by diffusion or erosion of the polymeric matrix [32]. For slowly degradable or non-degradable polymers, such as PCL, the protein release profile is governed by diffusion and is linear, whereas, for the degradable polymers such as PLGA, the drug release profile is controlled by erosion and shows sustained mode at first and then the release rate increases with polymer degradation [102]. Gene delivery via blend electrospinning is less complex in comparison to the protein release from the nanofibers. The first gene delivery application using plasmids in blend electrospinning has been reported by Luu et al. Plasmids can withstand the electrospinning process due to the protection from the complexation with vectors [103]. The pCMVβ plasmid encoding β-galactosidase was mixed with PLA–PEG– PLA copolymer and PLGA. The DNA maintained its structural integrity after being released from the PLGA nanofibers. It has been reported that gene expression showed two different release profiles based on the composition of nanofibers [87]. Luu et al. and Nie et al. reported burst release of DNA within 2 h followed by sustained release up to 20 days from PLA–PEG/PLGA polymer fibers and BMP2 release up to 4 weeks in vivo, respectively [104].

2.1.4

Coaxial Electrospinning

Coaxial electrospinning is an improvement of conventional blend electrospinning and a well-established technique for creating drug delivery nanofibers, in which two nozzles are connected to the high voltage source rather than one nozzle in a coaxial manner, as shown in Fig. 3a. It is essentially based on simultaneous co-electrospinning of two polymeric liquids [31, 32, 105]. Two different solutions are loaded within each nozzle, and the spinneret enables the generation of composite nanofibers with core-shell morphologies upon application of a high electric field. The inner core liquid is pumped via an internal needle, and shell material is pumped via an outer needle [82]. The classical electrospinning process has been modified to obtain the core-shell morphologies. The primary requirement is the use of electrospinnable shell polymeric materials with optimized molecular weight,

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Fig. 3 (a) Schematic of core-sheath electrospun nanofibers formation vial coaxial and emulsion electrospinning setup. (b) transmission electron microscopy microphotograph showing the coresheath morphology of nanofiber generated via coaxial electrospinning. (c–f) schematic showing encapsulation of gene vector in the core, the formulation of core and sheath protecting the vector from sheath organic solvent, gene delivery directed by creating porous heath layer, and modifying the sheath layer with polycationic polymers to enhance the transfection efficiency, respectively. The figure is reproduced from Lee et al. [30] as per CC-BY 4.0. (g) Schematic of tri-axial electrospinning setup using non-spinnable (core) and spinnable (sheath) solutions. Figure is reproduced from Luraghi et al. [31] with permission from Elsevier (License No: 5353760038786). (h) Scanning electron microscopy microphotograph showing the core (nanorod) and shell (microtube) structure fabricated using triaxial electrospinning process. The figure reproduced from [105] with permission from Elsevier (License No: 5353740257957)

concentration, and polymeric chain crosslinking in order to generate stable fiber jets. On the other hand, the core polymeric material should be non-spinnable, preferably polymers with low concentration. Both liquids used for generating fibers should not mix, and the interfacial tension between both liquid phases should be sufficient to draw the core liquid within the shell fiber jet [33, 34, 106]. The advantages of coaxial electrospinning are (1) the ability to produce nanofibers with core-sheath morphology from miscible as well as immiscible polymeric materials, (2) high loading and

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sustained release of drugs and biomolecules, and (3) less harsh processing conditions to deliver structurally stable biomolecules such as that it can protect and overcome the denaturation of drugs or biomolecules present in the biological system [34, 97]. Coaxial electrospinning relies on choosing the spinnable material for sheath structure and the non-spinnable material for core structure. In this technique, the drug or the biomolecule is loaded in the inner jet and co-spun along with the polymers of the outer jet. This arrangement provides protection to the loaded drug/ protein and facilitates its sustained release in biological environments [82]. It is reported that ketoprofen can be released from polyvinylpyrrolidone used as the sheath material and zein as the core material [80]. Similarly, non-spinnable dimethylformamide was used as sheath material and zein-ibuprofen as the core material [107]. The release profile from the core-sheath nanofibers is highly dependent on the type of drug and polymeric materials used for the core and shell. The core-sheath nanofibers also show an initial burst release followed by sustained release like in blend electrospinning. However, the burst release from coaxial fibers is significantly lower, with a more sustained release profile, due to the presence of a membrane barrier of a sheath that controls the diffusivity of the drug from the core [33]. To date, very few studies on protein/ gene delivery via coaxial electrospinning scaffolds for tissue engineering applications have been published. Biomolecules such as DNA, RNA, and protein are highly susceptible to unfavorable and harsh conditions [32]. The coaxial electrospinning process provides less harsh conditions for encapsulation and protection of the structure and conformation of biomolecules. Jia et al. reported the sustained release of VEGF growth factors for up to 28 days and improved cell adhesion and proliferation rate from the core-sheath nanofibers generated using PLGA as the sheath material and dextran as the core material for vascular tissue engineering [79]. The study reported that the diffusion of pDNA from the core-sheath nanofibers prepared using PCL as the organic sheath material which is loaded with PEI derivative hyaluronic acid as gene delivery vector to create shell and PEG to fabricate the core part. The results suggested the formation and diffusion of the gene-vector complex over a period of 60 days and transfected the cells present on the scaffold with 15% transfection efficiency [108]. Similarly, another study also reported the loading and release of green fluorescence proteintagged adenovirus from the core-sheath nanofibers. The results showed the cells with GFP expression for up to 30 days with a transfection efficiency of over 80%. The high transfection efficiency for the initial 2 weeks was due to the initial burst release of virus. It suggested that the different polymer concentrations have different poreforming abilities, which directly influence the release profile and transfection efficiency [109]. The sustained release of nerve growth factor to support nerve cell proliferation and differentiation using silk fibroin as the core material and poly-lactic acid as the sheath material [110]. The researchers developed a co-delivery system for BMP-2 and IGF-1 for improved osteoinductivity for bone tissue engineering applications from PLCL sheath material and hydrogel/emulsion of PLGA with heparin as the core material [111]. Similarly, the icariin-loaded in the silk fibroin/PLCL showed osteoinduction. Moreover, a liposome-loaded core-sheath system was developed,

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where the liposomes remained intact even under high shear stress and in the presence of chloroform in the shell polymer. The study suggested that the core-sheath technique preserves the bioactivity of the encapsulated enzyme, unlike in blended fibers [97]. Similarly, other systems encapsulating cells, yeast cells, and alpha granules have also been developed [32]. Multiaxial electrospinning can enables the fabrication of scaffolds with even more complex geometries. For example, in triaxial electrospinning, a spinneret consists of three nozzles for three different polymer solutions [106]. Triaxial electrospinning showed successful loading and release of the drug doxycycline from PCL/gelatin hybrid fibers [112], antimicrobial compound nisin, and other functional molecules [113]. A novel dual drug delivery system was reported and showed different release profiles of model drugs from the core and sheath of nanofibers. The system has three layers; the outer layer is the hygroscopic sheath layer, intermediate hydrophobic layer is of PCL to act as a membrane barrier between outer and core layers. Hygroscopic layer showed the initial burst release up to 80% within an hour, whereas the diffusion of drug from core showed 24× slower release than the coaxial fibers. This type of system could be used for both short-term and long-term treatment [106]. This technology allows multiphasic drug delivery profiles. Coaxial electrospinning has several advantages; however, the production rate is limited. The mesh layer obtained via coaxial electrospinning generally exhibits different properties of fibers across the entire layer. The optimization of liquids and their properties for various biomolecules may lead to the formation of defects and artifacts [32].

2.1.5

Emulsion Electrospinning

The emulsion electrospinning technique overcomes the disadvantages of blending and coaxial fabrication technique. This technique includes the emulsification approach in combination with both methods. Emulsion electrospinning uses only a single nozzle to generate core-sheath nanofibers, as shown in Fig. 3a [32, 114]. The emulsion is based using two immiscible polymer solutions during the electrospinning process, which is stabilized by using appropriate surfactants [114]. In this approach, the active biomolecule along with surfactant is mixed to create an emulsion by homogenization or ultra-sonication and then mixed with polymer solutions [27]. There are two different liquid phases, the continuous phase leads to the formation of the shell, and the droplet phase forms the core of the fiber. The processing conditions in this technique are slightly different from the blend and coaxial electrospinning. During the spinning process, the solvent evaporates from the continuous phase resulting in viscosity gradient and elliptical shape droplets in axial region [26]. The developed viscosity gradient allows the core material to deposit within polymer fiber [85]. Emulsion electrospinning generally uses two types of emulsions. (1) Water-in-oil emulsion, which is based on using a lipophilic material for the continuous phase and

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a hydrophilic material for the droplet phase [115]. This type of formulation is used to encapsulate polar and hydrophilic biomolecules and drugs. For stabilization of water-in-oil emulsions, Span 80 and Span 60 are used as surfactants with low hydrophilic–lipophilic balance [116]. The polymeric materials which are soluble in lipophilic solvents, and polyesters such as PCL, PLA, PLAGA, and PU, are used for the continuous phase to form sheath and water-soluble polymeric materials such as PVA, cellulose derivatives, alginate, chitosan, etc., are used for core formation. (2) Oil-in-water emulsion where continuous phase is formed by hydrophilic material and droplet phase is formed by lipophilic solution [117]. This type of emulsion is stabilized by surfactants such as Tween 20, Tween 80, and stearyl alcohol with high hydrophilic–lipophilic balance [118]. In addition to the above two types, a multiple emulsion system is also used where water-in-oil-in-water emulsion is produced by additional emulsification of oil-in-water in the water phase. This type of emulsion requires a careful selection of surfactants [119]. Along with emulsion formulation, the other material and process parameter also plays a crucial role. For forming a continuous phase, the electrospinnable material should have sufficient conductivity, high molecular weight, optimized polymer concentration, and low surface tension [32]. The droplet phase influences the internal organization of fibers. The core morphology plays an important role in deciding the release kinetics of drug/biomolecule from the core-sheath nanofibers. In emulsion electrospinning, release of drug takes place by both diffusion as well as dissolution of polymer matrix. In case of water-in-oil emulsions, the release rate depends on the core internal structure and the available contact points for core polymer dissolution. For the continuous core, the release occurs via capillary forces in core-sheath fibers, whereas for discontinuous core, the release rate depends on the interconnection of droplets [32]. From the drug delivery perspective, emulsion electrospinning has been used widely for the delivery of drugs and biomolecules [120]. The hydrophilic drug, Lidocaine hydrochloride, was successfully loaded within the PLA-based core-sheath nanofiber, where PVA was used in the aqueous phase for the core. The results suggested that PVA played a key role in the encapsulation of the drug, its distribution, and its release [121]. This technology is especially good for delivering protein molecules or susceptible molecules. The encapsulation of horse radish peroxidase and lysozyme showed the stability of enzymes and their bioactivity after release from the core-sheath nanofibers [122]. For long-term treatment, PLGA scaffolds were fabricated by applying negative voltage during electrospinning for long-term release of VEGF. The results showed the release profile up to 28 days [123]. Similarly, a co-delivery system for VEGF and PDGF has been developed using emulsion electrospinning [124]. The emulsion PCL scaffold loaded with bFGF showed sustained release of factor from the scaffold and enhanced bone regeneration in vivo [125]. Similarly, for induction of mesenchymal stem cell osteogenic differentiation, emulsion electrospun core-sheath nanofibers were loaded with PDGF [126]. The results from various studies demonstrated that emulsion electrospinning is an appropriate technique to generate core-sheath morphology nanofibers for encapsulating or loading various bioactive molecules. However, emulsion electrospinning also has one main disadvantage; this technique does not work with

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polymers of low interfacial tension. In addition, for emulsification, ultrasonication or mechanical mixing has to be carried out carefully in order to prevent damage to the loaded biomolecules or drugs.

2.1.6

High Throughput Electrospinning

The electrospinning technologies are now well recognized for their potential to generate drug-loaded nanofibers to deliver a variety of bioactive molecules. However, the technology is limited to small-scale production and cannot be easily scaled up. The throughput of this fiber-forming technique using the classic needle-based electrospinning process is limited and produces approximately 0.001–0.1 g/h [127]. This is due to the need for an optimized flow rate of the polymer solution, which is in the order of 10–100 μl/min [127]. Advancements in the electrospinning process to get high throughput were accomplished by using multiple needles simultaneously to increase the number of fiber jets. Various needle arrangements were adapted, such as triangular form, square, and hexagonal, to increase the jet distribution and spinning quality [128]. Figure 4a shows the various arrangement of a needle-based spinneret to produce multiple jets. In addition, to decrease the fiber deposition area and improve the fiber homogeneity, a metallic ring was installed around the electrodes [32, 129]. However, the multiple jet system has drawbacks with a non-uniform electric field. The appropriate arrangement for high throughput homogenous nanofiber fabrication requires large space along with high demand for regular cleaning of the spinneret. To overcome the problems in optimizing the needle arrangement for high throughput of the electrospinning process, the needleless electrospinning technique was proposed by Yarin and Zussman [130]. Figure 5a shows the schematic of high throughput production of nanofibers using a needless spinneret [131]. In 2004 the technology for mass production of nanofibers was commercialized and patented based on the work by Jirsak et al. under the name Nanospider™ [132]. The principle is based on the electrospinning of free liquid present on the polymer surface, which is destabilized under a high electric field. The fast-forming instability of surface waves, also known as Larmor–Tonks–Frenkel, led to the self-organization of jets to produce fibers [133]. Hence, needless electrospinning is the self-arrangement of jets on a liquid surface in response to surface tension and electrostatic forces. Self-organization of fluid is a consequence of instabilities created based on electrodynamics in electrospinning and thus considered as the key element for highly productive technology to generate nanoscale scaffolds [132, 134]. There are two types of needless electrospinning methods: (1) based on a range of electrodes, Shin et al. showed that a vertically charged threaded rod holding multiple drops generates multiple jets. They performed the experiment with a single rod and multiple rods in a linear array. A single rod of 50 cm in length produced polyvinyl propylene [PVP] fibers of 200–400 nm diameter at the rate of 4.5 g/h. (2) needleless electrode based on a rotating drum, disc, or coil in polymer bath [135]. Figure 4b showed the different spinnerets used in needless electrospinning. In this system, the polymer forms a thin layer over the rotating

Fig. 4 Different types of spinnerets used in the electrospinning process to create nanofibers. (a) Needle-based spinneret for low throughput production, and (b) needless spinneret for high throughput production of nanofibers. The figure was reproduced from Subrahmanya et al. [138] as per the CC BY-NC 3.0

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Fig. 5 (a) Schematic of multiple jet needless electrospinning set up for high production rate of nanofibers, (b) photograph showing the linear flume spinneret spinning process to generate multiple jets, and (c) scanning electron microscopy of fabricated nanofibers. Figure is reproduced from Wei et al. [131] as per the CC BY-4.0. (d) Schematic of needless coaxial electrospinning to produce core-sheath morphology nanofibers, and (e) photograph showing the generation of multiple jets from the needless spinneret during the electrospinning process. Figure is reproduced from Qin et al. [105] with permission from Elsevier (Ref. No: 5353740257957)

element and forms multiple jets. It has been reported that a concentrated electric field is crucial for needleless electrospinning, where the electric field profile along with the electric field strength on the spinneret directly influences the performance of the process. Thus, understanding the spinneret’s electric field would help to design high-performance needless electrospinning [133]. The software COMSOL Multiphysics 3.5a has been used to analyze the electric field profile and strength on ring spinneret for needless electrospinning using 3D finite element analysis. The result showed a strong electric field of 70 kV/cm intensity on top of the ring. The electric field is greatly influenced by the geometry of the ring and other process parameters. Wang et al. demonstrated the electric field distribution on different electrodes and found high strength at disc electrodes. Hence, using suitable geometry and process parameters, high-throughput and advanced electrospinning processes can be established [136]. Recently, the high throughput application of core-shell sheath nanofibers has also been reported. The coaxial needles for electrospinning to form the core-sheath nanofibers were first introduced by Forward et al., where a bilayer of two immiscible liquids was formed on the surface of a rotating wire electrode. Figure 5d, e show the schematic of needless coaxial electrospinning [137]. In this method, the electrode is immersed in the core polymer solution and then in the shell polymer solution, and under high electric field, a composite coaxial jet is formed from the surface of the bilayer. Another group introduced weir spinneret to generate core-sheath nanofiber morphology. In this method, the weir spinneret enabled the formation of a bi-liquid layer on the top of the linear spinneret and generated coaxial nanofibers [137]. They have also demonstrated the coresheath morphologies of fibers and their potential use in encapsulating drugs, and

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their time-dependent release profile. This method generates nanofibers ranging from 270 to 300 nm at a rate of 2.67 g/h, which is four times higher than the regular coaxial electrospinning process. Moreover, emulsion electrospinning from the needless wire-based electrode has been used for high throughput production of core-sheath nanofibers [114]. Recently, continuous phase of PCL with droplet phase of Pluronic F68 emulsions loaded with different active molecules (horseradish peroxidase, bFGF, TGFβ, IGF-1) was used to create core-sheath nanofibers by using static wire electrodes-based needless electrospinning. The results showed higher production rate along with preserved stability of loaded enzyme and molecules and their controlled release from the nanofibers. Along with that growth factor-loaded nanofibrous scaffolds showed a good stem cell proliferation rate. Thus, the use of needleless electrospinning for high throughput production techniques for core-sheath nanofibers with encapsulated bioactive molecules increases the potential use of nanofibrous scaffolds. The technology is more cost-efficient for developing drug-loaded core-sheath scaffolds and for scale-up quantities. Core-shell-based scaffold system has enormous potential to advance drug-loading studies and tissue regeneration. Further advancement is required, such as the use of medical-grade polymers, drugs, and bioactive molecules for rapid translation from laboratory research to clinical use.

3 Application of Electrospun Nanofibers for Therapeutic Delivery Nanotechnology enabled the understanding of molecular and cellular information with high selectivity and sensitivity. The electrospun nanofibers have been used extensively as a spatial template as it can mimic the extracellular matrix at the structural as well as functional level that helps to retain the cellular structure and function and delivery of target molecules at different cell types. It has a great potential to be used as a physicochemical guide and can be used in numerous applications, from biomedical applications to the textile industry. Here we describe the biomedical applications of nanofibers in therapeutic delivery, such as transdermal and wound dressing systems, drug delivery, growth factor delivery system, and gene therapy. There are two important considerations to achieve successful biomolecule delivery for any biomedical application: (1) retain the maximum bioactivity after loading and encapsulating within the carriers or scaffolds, and (2) tune the release profile to match the rate of regeneration of the target tissue [85]. These biomolecule delivery strategies using electrospinning technique are shown in Fig. 6. Table 3 compiles several strategies reported in the literature to prepare electrospun nanofibers for the delivery of biomolecules.

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Fig. 6 Biomolecule delivery strategies via different variations of the electrospinning process: (a) basic delivery system via blend electrospinning, (b) advanced delivery system via coaxial electrospinning, and (c) smart delivery system via response stimuli. Figure is reproduced from Shahriar et al. [1] as per CC-BY 3.0

3.1

Transdermal and Wound Dressing

Wound healing and tissue regeneration is one such thrust area where electrospun nanofibers act as drug reservoirs for sustained therapeutic benefits. In a reported study, electrospun scaffolds were fabricated from Ac-DEX loaded with an immunomodulatory receptor, Resiquimod, for the treatment of cutaneous disease [139]. The morphology and width of electrospun nanofibers revealed a direct correlation with their degradation rates. The faster degrading scaffolds showed a burst release of Resiquimod, whereas the slowly degrading fibers led to a sustained release profile. This tunable drug release characteristic of the electrospun scaffold shows its therapeutic potential for the treatment of various types of cutaneous

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Table 3 Polymeric nanofibers loaded with different drugs or bioactive molecules for controlled delivery for biomedical applications Polymer PVA/PLGA

Drug/bioactive molecule Gliclazide

PLGA/gelatin

Cefradine, 5-flurouracil; ciprofloxacin,quercetin Tetracycline Didofenac; Ibuprofen

Halloysite/PLGA PCL PVA/PCL PCL/collagen PLA PCE nHA/PLGA PLGA/graphene/H. A. PLLA Chitosan/PEO PVA/silicone Polydopamine coated poly l-lactide PCL/PEG PLGA/gelatin PCL PLLA Peptide amphiphile PDLLA/PLGA Chitosan/PCL PVA PCL PCL/PEG PLGA/gelatin PCL/collagen PEG/PEI

References [53]

Ag-chitosan nanoparticles with phenytoin Artemisinin Phosphorylcholine, curcumin, paclitaxel, ibuprofen Dexamethasone/BMP2 BMP2 bFGF/BMP2

Application Type 2 diabetes management Tissue regeneration and wound healing Antibacterial activity Inflammation and biomedical application Wound healing Malarial infection Cancer therapy, wound healing Bone regeneration Bone regeneration Bone regeneration

[159] [52, 160– 162] [163] [164] [165]

Heparin/VEGF VEGF/PDGF VEGF VEGF/BMP2

Aneurysm treatment Wound healing Islet transplantation Bone tissue engineering

[166] [167] [168] [169]

VEGF

Vascular tissue engineering Bone tissue engineering Neural regeneration Neural regeneration Neural regeneration Neural regeneration Wound healing Wound dressing Skin tissue engineering Skin tissue engineering Skin tissue engineering Skin tissue engineering Regenerative medicine

[170]

Blood vessel regeneration Cellular proliferation and differentiation Gene silencing Simulating gene delivery

[180]

Poly DL lactide/PEG

BMP2/VEGF NGF NGF NGF NGF/gNGF PDGF/EHGF/FGF/TGF-β EGF/FGF EGF EGF EGF EGF Non-viral vector-based DNA adsorption Plasmid-VEGF

PCL

microRNA

PCL/PEG PLA/PEG/PLGA

siRNA Plasmid [beta-galactosidase/ GFP]

[154, 155] [54] [51, 156, 157] [158]

[171] [172] [173] [174] [175] [176] [95] [177] [92] [178] [179] [152]

[181] [182] [153]

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diseases. The fabrication of antimicrobial nanofiber wound dressing from PEO and PdLA loaded with Nisapline (naturally occurring antimicrobial) [140]. The in vitro antimicrobial assays showed sustained diffusion of Nisapline for 4 days into the Staphylococcus aureus seeded plates. This resulted in a decrease of bacterial load by order of 105 compared to bacterial plates incubated with the control nanofibers. Another study described the fabrication of PCL electrospun nanofibers loaded with an anticancer drug, imiquimod, for the treatment of basal cell carcinoma [141]. The fibers showed sustained release of the drug for nearly 48 h. Basha et al. reported the fabrication of PVA blended Curdlan (β-1,3 glucan) electrospun water-soluble nanofibers crosslinked with glutaraldehyde [142]. The nanofiber mat showed an immunomodulatory effect and exhibited faster wound closure rate compared to the control PVA scaffold. Wistlich et al. fabricated peptide-functionalized electrospun nanofibers for imparting specific cell adhesion [143]. The developed nanofibers promoted the attachment of HT1080 cells in the presence of RGD peptides and can be used for immunomodulatory wound dressings [144, 145].

3.2

Drug Delivery Systems

Most of the available drugs are hydrophobic, and their therapeutic efficacy is limited by their low solubility, stability, and poor biodistribution in the human body [15]. The most common challenge with drug delivery is that the action of drugs is not targeted to the disease site, thereby compromising the desired therapeutic efficacy. The drugs often exhibit non-specific binding, induce systemic toxicity, and suffer from rapid elimination from the human system. This has motivated the development of drug delivery system (DDS) technologies, including the formulation, encapsulation, and delivery of drugs/biomolecules to a specified targeted site in vivo [35]. Moreover, the selected DDS technology also influences the absorption, distribution, release, and elimination of the drugs from the system with high loading and low toxicity. The release of drugs from the carrier depends on many physical properties of the carrier system, such as the fabrication method, degradation, and swelling for drug diffusion [146, 147]. Electrospun nanofibers as nanoscale DDSs are attracting the attention of researchers owing to their good biocompatibility, degradability, and high loading and entrapment efficiency, fulfilling the key prerequisites for an effective therapeutic delivery platform [14, 118].

3.3

Growth Factor Delivery System

The maintenance of the biological activity of growth factors is a key prerequisite for their successful delivery [94]. Growth factors are prone to rapid loss of their biological activity if subjected to chemical and physical modifications. Growth factors have short half-lives, and thus, bioactive scaffolds should maintain the

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temporospatial concentration of growth factors at the diseased tissue or condition for their release to facilitate the desired effect [85]. Thus, the scaffold fabrication, degradation behavior, and loading of growth factors are interdependent. For rapidly achieving the effective therapeutic concentration at the diseased tissue for regeneration, the scaffold can release growth factors initially as burst release and subsequently release factors in controlled, well-defined kinetics [148]. One such example is the controlled delivery of insulin loaded in electrospun nano-/micro-fiber patches. Sharma et al. prepared biodegradable PVA and NaAlg patches for transmucosal delivery of insulin [149]. The insulin release from the patch followed first-order kinetics, and the in vivo studies confirmed the delivery of active insulin resulting in better outcomes as compared to the commercial formulation. Another group prepared electrospun chitosan-blended PEO nanofibers for insulin delivery via buccal mucosa. The buccal permeability of the insulin released from the mat was studied ex vivo. Controlled release of insulin from the electrospun mat showed 16 times higher transbuccal permeability as compared to free insulin [150]. Similarly, electrospun nanofibers of PVA and PLGA loaded with gliclazide (an antidiabetic drug) were reported for the treatment of type 2 diabetes. These drug-loaded nanofibers were then encapsulated in a gelatin capsule for oral administration. The nanofibers showed a biphasic drug release profile with a rapid initial drug release followed by prolonged sustained release [53].

3.4

Gene Therapy

Gene therapy has been widely explored for the treatment of human diseases, including cancer. Gene therapy includes the transfection of cells using different genetic materials such as plasmid DNA, si-RNA, and micro-RNA to regulate the expression of the target protein at the targeted site. For successful gene delivery, it is important that the gene remains active after release from the scaffolds to interact with the host genome [30]. The gene products (DNA and sRNA) act intracellularly, bind to the host genome of the endogenous cells, and influence the molecular mechanism at the defect site, thereby converting the transfected cells to bio-activated actors to facilitate tissue regeneration [151]. For this purpose, before loading the genes into the scaffolds, the genes are packed within the vectors, which protect them from loss of biological activity due to degradation by extracellular enzymes and intracellular lysosomes during uptake by cells. However, a key consideration with gene delivery is the transfection efficiency which is influenced by the concentration of vector-gene loaded within scaffold, and its duration of release profile to achieve the desired effect [152]. For successful gene delivery, the release profile of the concentration of the vector-gene complex should be optimal to attain high transfection efficiency into the cell surrounding microenvironment along with the optimal time frame of its release and transfection [153]. The chapter summarizes the different types of electrospinning setups to fabricate different morphologies of nanofiber for the delivery of drugs, biomolecules, and

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gene products for guided tissue engineering applications. The coaxial and emulsion electrospinning processes are the choice of fabrication technology for bioactive molecule delivery in the field of drug delivery systems due to the success rate of delivery without compromising the stability and activity of molecules at the targeted sites. However, detailed studies are required on the pharmacodynamics and pharmacokinetics of molecules delivered by core-sheath technology. Acknowledgment The authors acknowledge support from the Department of Science and Technology (DST), Government of India, (DST/NM/NB/2018/119(G)).

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Adv Polym Sci (2023) 291: 69–80 https://doi.org/10.1007/12_2022_125 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 5 July 2022

New Prospects in Melt Electrospinning: From Fundamentals to Biomedical Applications Moustafa M. Zagho, Yasseen S. Ibrahim, and Ahmed A. Elzatahry

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Melt Electrospinning Configurations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Multi-Temperature Control Melt Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Laser Melt Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3 Melt Coaxial Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4 Needleless Melt Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.5 Gas-Assist Melt Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.6 Melt Electrospinning Writing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Biomedical Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1 Biosensors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 Drug Delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3 Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract Since melt electrospinning permits polymers with medical grade to be used without modification, this offers a remarkable benefit to gain regulatory support over biomedical systems prepared by solution electrospinning, which require expensive and challenging post-processing to remove toxic solvents. This chapter aims to identify the key components of melt electrospinning technique and addresses the

M. M. Zagho School of Polymer Science and Engineering, University of Southern Mississippi, Hattiesburg, MS, USA e-mail: [email protected] Y. S. Ibrahim and A. A. Elzatahry (*) Materials Science and Technology Program, College of Arts and Sciences, Qatar University, Doha, Qatar e-mail: [email protected]

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different types of melt electrospinning apparatus, including multi-temperature control, laser melt, coaxial, needleless, and gas-assist electrospinning as well as melt electrospinning writing. Different biomedical applications of melt electrospinning, involving biosensors, drug delivery, and tissue engineering, are also discussed. Keywords Biomedical · Biosensors · Drug delivery · Melt electrospinning · Tissue engineering

1 Introduction In principle, the melt electrospinning device consists of primary components such as a reservoir/syringe, needle, pump, voltage supplier, and collector. Typically, the melt electrospinning (solvent-free) process starts by heating the polymer with or without additives to form a polymer solution, hence ejected by the pump and forming a droplet at the tip of the needle [1, 2]. The polymer solution is then introduced to an electric charge force and the droplet starts to form when the electric force equals to that of the surface tension [1]. If the electric field is further increased, the electric force will break the surface tension of the polymer solution and create a liquid jet from the Taylor con. After that, the produced charged polymer jet passes by the most critical step in the process as it accelerates toward the direction of the collector. During this step, the jet develops in a nearly straight line [3]. After a short distance, the jet experiences an unstable bending/whipping form, increasing the jet’s traveled distance from the needle to the collector. The jet instability is proposed to be owned to the charge density and repulsive charge on the melt jet [4]. Nevertheless, the jet undergoes elongation and thinning due to the applied voltage and the traveled distance, controlling the fiber diameter before it solidifies and deposits on the collector.

2 Melt Electrospinning Configurations 2.1

Multi-Temperature Control Melt Electrospinning

The setup was proposed as an attempt to further decrease the diameter of the spun fiber. By controlling the temperature of several zones in the electrospinning, especially the needle-collector zone, where the melt jet experience thinning in diameter just like the solution jet [5–7]. By applying heat to the needle-collector zone, the fiber jet will maintain its met form and undergo a whipping motion. Hence, increasing the distance covered between the needle and collector before solidifying [8].

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2.2 2.2.1

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Laser Melt Electrospinning Spot Laser Beam Melt Electrospinning

In this design, the sample polymer is first melt pressed to form rod pallets and then subjected to a spot laser beam at the tip of the pallet from three angles by using four mirrors and an absorber from a high-power laser machine [9–11]. Moreover, in order to avoid burning the polymer pallet sample, N2 flow was used in conjunction with the laser beam. Although the laser output power causes thermal degradation and deteriorates the mechanical properties of the fiber, yet the fiber is still a functional material. The heat applied on the polymer is localized, unlike other melt electrospinning designs where the hole reservoir is subjected to heat [11].

2.2.2

Line Laser Beam Melt Electrospinning

To increase the fiber fabrication, the line laser technique was proposed [12]. Where a spot laser beam passed through an optical system, generating line laser beam which is then subjected on a sheet pallet creating multiple Taylor cones [13].

2.3

Melt Coaxial Electrospinning

This setup offers a facile strategy for coating solid materials (core-sheath) and creating hollow fibers [14]. The idea is based on merging the Taylor cone from the outer and inner needle creating fibers with various properties [15]. Just like the conventional electrospinning, the solvent of inner jet evaporates, and the fiber solidifies quickly to be encapsulated and create the core of the fiber. Additionally, the polymer in the inner needle, not its molten state, should be soluble in the solvent used in the outer needle and vice versa [14, 15].

2.4

Needleless Melt Electrospinning

Several different designs of melt electrospinning were reported using a needleless setup technique to increase the fabrication flow rate of fibers [16–19]. For example, using a hot reservoir to melt the polymer while a round disc is partially immersed in it and pulling out the melt polymer through disc surface [17]. Such a setup depends solely on the intensity of electric field to pull out the melt fibers from the disc surface. As no pump was utilized in the setup, increasing the applied voltage will create more Taylor cons on the disc surface [2]. Bubble melt electrospinning was also proposed as a needleless setup [20, 21]. The design is based on bubbling air from the bottom of

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the reservoir and creating a melt polymer bubble on the surface. Nevertheless, both setups reported inconsistent fiber diameter owned to varied distance covered between each jet, viscosity, and bubble size. Although, the difficulty in maintaining uniform fiber diameters, each setup is viable for mass production [2]. The needleless umbellate melt electrospinning, on the other hand, was equipped with a distributor to guarantee a consistent flow [16, 22].

2.5

Gas-Assist Melt Electrospinning

This design is a simplified approach for the multi-temperature setup, since controlling the temperature for several parts in the setup is quite difficult. The Gas-Assist setup is simply subjecting hot gas alongside the melt jet in a conventional melt electrospinning setup. Nevertheless, the process is based on the applied electric field not the temperature of the blown air or its speed [2, 23, 24]. In principle, the hot blown air supports the melt jet with heat causing a delay in its solidification before reaching the collector, hence reducing the fiber diameter [2].

2.6

Melt Electrospinning Writing

The technique supports a continuous melt electrospun fibers stacked layer-by-layer creating scaffolds [25]. The whole setup is connected to a computer-controlled device where all the desired parameters are set and in-situ controlled, like pump flow rate, collector speed, and nozzle-to-collector distance, making the device a 3D fiber printer [26–28].

3 Biomedical Applications Since melt electrospinning permits polymers with medical grade to be used without treatment, this offers a noteworthy benefit to receive regulatory support over other systems prepared using other techniques like solution electrospinning, which require post-processing expensive and challenging processes for toxic solvents removal [1]. This chapter discusses the different biomedical applications of melt electrospinning. Scheme 1 illustrates the different biomedical uses of melt electrospinning.

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Scheme 1 Schematic diagram of the different biomedical applications of melt electrospinning

3.1

Biosensors

Micro- and nanocapsules have great potential for functional devices. This technique with a coaxial spinneret is employed to encapsulate solids in a polymer network or composite to obtain phase change materials [14]. Owing to the fluidity of these thermochromic systems after melting, their industrial fabrication is limited [29]. These materials need to be stabilized in a solid matrix because they melt and crystallize with thermal cycling. For example, poly(methyl methacrylate) (PMMA) fibers were used to encapsulate these materials using coaxial melt electrospinning [29]. With a thermoresponsive core prepared from bisphenol A, 1-tetradecanol, and crystal violet lactone core and a PMMA shell with excellent optical transmission property, the resultant composites displayed excellent fluorescent and thermal energy management. The composite fibers had also potential in designing bodytemperature calefactive systems and temperature sensors. Furthermore, McCann et al. [14] described the preparation of phase change nanofibers using the same technique. The nanofibers consisted of composite sheath and long-chain hydrocarbon (octadecane) core. This process is used to electrospun paraffins and encapsulate them in one step. These shape-stabilized phase change exhibited huge heat of fusion of long-chain hydrocarbons, releasing huge quantities of thermal energy over a

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particular temperature range. The nanofibers also displayed innovative segmented structures for the core owing to the fast solidification of the long-chain hydrocarbon.

3.2

Drug Delivery

Controlled release systems with promoted and sustained drug release are highly required for wound dressings [30] and drug-loaded implants [31]. Melt electrospinning is widely used in pharmaceutics as it integrates the benefits of solvent-based electrospinning and melt-extrusion. Nagy et al. [32] designed drug delivery systems with improved dissolution using solvent-free melt electrospinning. This design was prepared using cationic methacrylate copolymer as a fiber-forming polymer matrix. Because of the soluble polymer in acidic media and their large surface area, these solvent-free melt electrospun fibers displayed an ultrafast drug release of thermally sensitive drugs such as carvedilol. This design represents a talented drug delivery system as it combines the large surface area properties of solvent-based electrospinning and effective amorphization and continuous process advantages of melt extrusion. Fast dissolving drug-loaded melt electrospun polymer mats were fabricated in the presence of plasticizers by Balogh et al. (Fig. 1, left) [33]. To evade undesired thermal degradation, plasticizers were used to lower the processing temperature. Carvedilol was introduced into amorphous methacrylate terpolymers and three different plasticizers (polyethylene glycol 1,500, Tween® 80, and triacetin) were used. The three plasticizers successfully lowered the processing temperature and enhanced the stability of carvedilol under conditions applied without plasticizers. Interestingly, the same group recently used melt

Fig. 1 Dissolution profiles of carvedilol. Fibers with carvedilol (20 wt%) and polyethylene glycol 1,500, Tween® 80, and triacetin (Left) [33]. A polycaprolactone scaffold coated with electrosprayed poly(lactic-co-glycolic acid) microparticles after 1 h of electrospraying. (a) Micro-computed tomography and (b–d) SEM images with different magnifications (Right) [36]

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blowing to obtain fast-dissolving high-quality micro- and nanofibers [34]. Due to their huge surface area, the melt-blown fibers dissolved within 2 min. Polymeric fibers can also be employed to introduce small particles, molecules, and proteins. In this regard, melt electrospinning can be used to fabricate scaffolds to enhance the stability of these species [14, 35]. For instance, biodegradable polycaprolactone microfibers were coated with high densities of poly(lactic-coglycolic acid) (PLGA) microparticles to encapsulate bovine serum albumin (BSA) for growth factor delivery (Fig. 1, right) [36]. This design was fabricated by using an optimized collector configuration. Integrating scaffolds with filled microparticles offered a reproducible scaffold coating. The immobilization of the microparticles on the surface of the scaffolds remarkably reduced the burst release of the protein. This approach provided a promising candidate for protein delivery and growth factor delivery treatments applied to bone, cartilage, and skin.

3.3

Tissue Engineering

Efforts are devoted for designing scaffolds to be combined with biological molecules or living cells to form a tissue engineering construct to promote the regeneration and repair of tissues [37]. The formed tissue engineering constructs should exhibit a structure that supports cell attachment and cell growth, excellent mechanical performance, and biochemical characteristics [38]. The fabrication process must be in a sterile environment and reproducible with low cost. Melt electrospinning provides additional benefits over solution electrospinning for tissue engineering applications. Additionally, the fibers can be collected onto cells or on water due to the use of water-insoluble polymers. A remarkable fraction of the melt electrospinning applications is concentrated on designing biostructures for vascular blood vessels, implant interfaces, and dermal substitutes [1]. Flexible tubular structures prepared by solution electrospinning are used in tissue engineering. However, the deposition of fibers is hard to be controlled owing to the chaotic nature of solution electrospinning, thus, melt electrospinning is favored in this regard. Brown et al. [39] melt electrospun polycaprolactone and prepared tubes using 20 μm diameter fibers with controllable micropatterns. The resultant tubes successfully supported the growth of primary human osteoblasts and mesothelial cells. This architecture is promising for cell spanning between adjacent fibers. Interestingly, the fiber winding angle controls the size, shape, and porosity of the scaffolds. In addition, the mechanical response to uniaxial tension and compression can be enhanced with lowering the winding angle. Designing layer-on-layer tissue made of polymers and cells in view of scaffolds fabrication was reported by Dalton et al. [40] They described a direct in vitro melt electrospinning, in which they melt electrospun a blend of poly(ε-caprolactone) (PCL) with poly(ethylene oxide)-b-poly (ε-caprolactone) with in vitro cultured fibroblasts. Small and homogenous fibers were produced with a high electric field. Fibroblasts were interacted to the melt electrospun fibers and detached from the substrate after 6 days from electrospinning.

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Fig. 2 Synthesis of osteoconductive biphasic scaffolds, cell seeding, and in vitro and in vivo implantation. (1 and 2) Harvesting and placement of cell sheets on the periodontal partition. (3 and 4) Positioning the design on the dentin block [42]

Surprisingly, the fibroblasts’ structure changed from flat and spread to spindleshaped and long when adhered onto the electrospun fibers. There is a shortage of appropriate models that imitate metastasis of human tumor cells to a human bone microenvironment. Thibaudeau et al. [41] described in vivo study using a human tissue-engineered bone structure to design a humanized xenograft model of breast cancer-induced bone metastasis in a murine host. They combined primary human osteoblastic cell-loaded fibers with recombinant human bone morphogenetic protein 7 and implanted subcutaneously in non-obese diabetic/ severe combined immunodeficient mice. Researchers paid their efforts to address the restrictions of current periodontal regeneration approaches. Cell sheet strategy can be combined with osteoconductive biphasic scaffolds. Costa et al. [42] studied the capability of osteoconductive biphasic scaffolds to regenerate cementum, periodontal ligament, and alveolar bone. A fused deposition modeled bone partition was adhered to melt electrospun periodontal partition coated with calcium phosphate layer, loaded with osteoblasts, and cultured in vitro for 6 weeks (Fig. 2). The calcium phosphate coating layer remarkably enhanced the alkaline phosphatase activity and mineralization. The coated scaffolds exhibited a significantly enhanced bone formation. Histological

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studies also revealed that the vascularization of the cell sheets was allowed by the large pore size of the periodontal partition. The inability to fully recapitulate the mechanical behavior and morphological organization of native cardiac tissues represents the main challenge in cardiac tissue engineering applications. Melt electrospinning can be employed to resolve these restrictions by designing well-organized fibrous materials with excellent architecture and biocompatibility compared with those fabricated by solution electrospinning. Castilho et al. [43] combined melt electrospinning writing with poly(hydroxymethylglycolideco-ε-caprolactone), and the resultant microfibrous (4–7 μm) scaffolds cellularly responded to the mechanical anisotropy. The resultant scaffolds exhibited a rectangular decoration that mimics the mechanical characteristics of the native myocardial tissues. Compared with electrospun PCL-based scaffolds, the melt electrospun poly (hydroxymethylglycolide-co-ε-caprolactone) fibrous scaffolds aligned more professionally with cardiac progenitor cells. This strategy offered a promising candidate with improved biological relevance and mechanical performance and can be considered as a potential framework for therapeutic viable in vitro cardiac engineered tissues. Bioactive electrospun structures built of polymer blends can release vascular endothelial growth factors in a sustained manner. Coaxial melt electrospinning can be employed in this subject to obtain high-performance engineered tissues. Seyednejad et al. [44] coaxially electrospun fibrous scaffolds with approximately 700 nm diameter. The composite scaffolds were based on PCL and poly (hydroxymethylglycolide-co-ε-caprolactone) and loaded with vascular endothelial growth factor (potent angiogenic factor) and BSA (protein stabilizer). The two polymers were miscible at the molecular level. The scaffolds prepared from poly (hydroxymethylglycolide-co-ε-caprolactone) (pHMGCL) displayed a remarkably higher surface hydrophilicity than those fabricated from PCL only. The scaffolds core exhibited a faster release of BSA than PCL scaffolds. Additionally, the incorporated protein resulted in initial higher numbers of adhered endothelial cells and preserved its biological behavior to support cell growth up to 7 days. The increased hydrophilicity of pHMGCL scaffolds resulted in a stronger interaction of human mesenchymal stem cells seeded onto the scaffolds than PCL scaffolds. Efforts were also devoted to designing 3D structures with tailored features using direct writing. Direct writing melt electrospinning can be employed to produce threedimensional scaffolds from micron-diameter fibers. However, disordered structures are obtained owing to the charge build-up on the deposited polymer, producing unwanted coulombic forces. Ristovski et al. [45] recently described an electrostatic control to reduce the destabilizing polymer charge effects, producing structurally suitable scaffolds for tissue engineering applications. They used dual voltage power supplies to enhance the control of fiber deposition and reduce the undesirable charge effects. Ordered scaffolds with up to 200 layers thick, diameter of 40 μm, and 1 mm fiber spacing, were obtained. The thickness of the resultant fibers was overall potential-dependent and specific tip/collector voltage-independent. Successful cell attachment was observed in vitro with little cell death after 7 days.

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4 Conclusions The melt electrospinning has made significant progress to keep up with the demands of polymeric materials for biomedical applications. As more research is done into current trends in biomedical applications, we can continue to utilize new information to generate informed materials that effectively perform, while remaining safe environmentally and health wise. The use of melt electrospun fibers shows much promise for the future of biosensors, drug delivery, and tissue engineering. As these areas expand, production costs will become more feasible on a large scale, and the safety of this technique will make it an attractive alternative to solution electrospinning. Acknowledgment The work was supported by NPRP grant NPRP12S-0309-190268 from the Qatar National Research Fund.

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Adv Polym Sci (2023) 291: 81–106 https://doi.org/10.1007/12_2022_131 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 21 July 2022

Centrifugal Spun Nanofibers and Its Biomedical Applications Hemamalini Thillaipandian and Giri Dev Venkateshwarapuram Rengaswami

Contents 1 2 3 4 5

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Fiber-Forming Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Principle of Centrifugal Spinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Material and Machine Parameters Influencing the Fiber Formation . . . . . . . . . . . . . . . . . . . . . . . Application of Centrifugal Spun Nanofibers in Various Biomedical Applications . . . . . . . . 5.1 Tissue Engineering Applications of Centrifugal Spun Fibers . . . . . . . . . . . . . . . . . . . . . . . 5.2 Drug Delivery Applications of Centrifugal Spun Fibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.3 Wound Dressing Applications of Centrifugal Spun Fibers . . . . . . . . . . . . . . . . . . . . . . . . . . 6 Advances in Centrifugal Spinning Process . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8 Future Trends . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract Fibers of micron and submicron diameters are finding potential applications in the medical field due to their high surface area and porosity. Electrospinning has been one of the potential methods to produce fibers in the submicron range by application of high voltage. But the major limitation is the speed and ease of production. Centrifugal spinning is an alternative technique for the production of submicron fibers compared to the electrospinning technique without the application of high voltage. The technique utilizes centrifugal force for the production of fibers and they can be produced at a shorter duration. The fibers are produced by overcoming the surface tension of the polymeric solution, followed by the ejection of polymer solution and stretching to produce the solidified fibers. The quality of fiber production depends upon the material and process parameters. Synthetic and natural polymers can be used to produce fibers that can find potential applications in the area

H. Thillaipandian and G. D. V. Rengaswami (*) Department of Textile Technology, Anna University, Chennai, India e-mail: [email protected]

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of biomedical, energy sensors, and filtration. The chapter summarizes the principle of centrifugal spinning, machine and material parameters and its potential application in the area of biomedical applications. Keywords Biomedical application · Centrifugal spinning · Machine and material parameters

1 Introduction Submicron fibers offer wide range of applications in the area of biomedical applications such as tissue engineering, drug delivery, dental applications, wound dressings, sanitary products, protective clothing, molecular filtration media, artificial blood vessel, biochip, and nanosensor due to higher surface area to volume ratio, excellent mechanical property, tunable structure, functionality, and high porosity compared to conventional fibers [1]. The fibrous structure enables the fabrication of 3D networks with an ability to incorporate the chemicals or tune the molecular alignment to increase the bioactivity between the substrate and cells, making it suitable for biomedical applications. The fabrication technique is chosen based on the type of polymer and solvent system used for the production. The polymeric materials can be reformed to nanofibrous structure either by means of dissolution in suitable solvents or melting at a higher temperature, followed by spinning the molten solution using the suitable processing technique, thereby ameliorating the orientation of the polymeric chains to yield the required functional properties based on end applications. Conventional spinning methods include extrusion of polymer solution, drawing the emerging filaments either by means of driving force such as electrical, gas pressurized, and centrifugal forces or mechanical drawing namely hot metallic rollers followed by solidification as the result of precipitation or drying to yield resultant filaments [2]. The conventional fabric manufacturing process can be adopted for the production of healthcare, nonimplantable and extracorporeal devices compared to implantable medical implants. Non-woven technique is the random arrangement of conventional fibers either by mechanical, chemical, or thermal means of bonding them and it is widely used for the manufacture of healthcare and hygiene products such as surgical gowns, caps, wipes and face masks and nonimplantable materials such as absorbent pads and wadding. Woven fabrics are produced by the interlacement of warp and weft threads and are used widely for the manufacture of bandages, gauze, artificial ligaments and tendons, surgical blankets, and covers. Knitting is the process of interlooping of course and wales used in the area of biomedical applications for the manufacture of cardiovascular implants, vascular grafts, heart valves, and surgical hosiery [3]. Implantable materials, namely tissue engineering, drug delivery, and wound dressing products, attract the utilization of nanofibers due to their functional and structural property.

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Centrifugal, force spinning, rotational jet spinning, or rotary spinning is a versatile technique that allows the production of submicron fibers utilizing a wide range of polymers. The advantages of centrifugal spinning compared to other nanofiber production techniques and the principle of fiber production and application of the centrifugal spun fibers are discussed in this chapter.

2 Fiber-Forming Systems The polymers used for biomedical applications should be biodegradable and biocompatible, thereby offering mechanical strength during their usage [4]. Nanofibers can be fabricated using various techniques utilizing a wide range of biopolymers such as natural and synthetic polymers as shown in Fig. 1. Drawing technique enables the production of discontinuous nanofibers by placing the drop of polymeric liquid on the surface, on which a sharp needle tip is made in contact and withdrawn. During withdrawal of the needle, polymeric solution on the needle tip is solidified by rapid evaporation of solvents resulting in single nanofibers. The technique is suitable for polymers that are viscoelastic and cohesive in nature. The technique depends on the size of redeposited droplets and the time of drawing as the size of droplets is reduced resulting in decrease in a the diameter of the fibers. Thus, the process offers a limitation of lower productivity due to the formation of discontinuous fibers [5]. Template synthesis is widely used for the production of inorganic nanofibers such as electronically conducting polymers, metals, semiconductors, and carbons. The fibers are produced in the form of nanotubes or nanofibers based on the type of molds or templates used. The monomers are filled into the molds and converted to polymeric nanofibers either by means of chemical or electrochemical methods. The nanofibers are separated from the template either by

Fig. 1 Methods of production of nanofibers

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dissolving or etching the template. The structure of the nanofibers depends on the pore dimension of the template [6]. The advantage of the technique includes altering the fiber diameter by changing the templates. The drawback of the process includes the requirement of a soluble porous template in order to remove the nanofibers in the case of conventional template synthesis and also continuous fibers cannot be produced. Phase separation is a process in which nanofibers are produced from the separation of two different phases caused due to physical inconsistency. The procedure includes dissolution of polymers and gelation, extraction of solvent, freezing, and freeze-drying, as these stages involve removal of the solvent phase leaving the other residual polymeric phase [7]. The polymer is dissolved into a homogenous solution either at room temperature or higher temperature allowing the solution to form a gel at gelation temperature. As the result of gel formation, the phase separation occurs, allowing the formation of a nanofibrous matrix upon removal of the solvent and drying process. The technique is suitable for the fabrication of porous fibrous membranes with a good mechanical property, but the process is suitable for selective polymers and requires a longer processing time and also the production of continuous and longer fibers with controlled fiber orientation is difficult. Self-assembly technique utilizes small molecules to form an aggregate using intermolecular interactions such as non-covalent bonds. The technology depends on the chemical nature of the small molecules, and the formation of nanofibers is time-consuming process, and the formed fibers are unstable in nature [8]. Phase separation and self-assembly techniques are not widely used in industrial applications due to time consuming and slow production process [9]. Melt blowing technique is similar to the melt spinning technique with a difference in submicron fiber production. The process sequence involves melting of the polymer chips and extruding the spinneret, followed by attenuating using hot high velocity air to produce ultrafine nanofibers, which are then collected on the collector in the form of random fibrous nonwoven. Melt blowing technique is suitable for thermoplastic polymers and not used for biopolymers such as protein or polysaccharides as the materials are processed at higher temperatures, which tends to denature or degrade thereby losing the functionality of the polymers [10]. Wet spinning is another technique widely used for conventional fibers as the polymers are dissolved in a suitable solvent to form a homogenous solution, then it is extruded through a spinneret into a coagulation bath. The latter contains a solvent that is nonsolvent for the polymer and upon solidification, the process yields micron fibers [11]. The fibers formed are highly porous due to the removal of solvents, thereby finding the application in tissue engineering application. The drawback of the technique includes excessive usage of chemicals during fiber manufacturing may toxicity to healthy cells upon application. Microfluid spinning enables the production of micro and nanofibers by injecting the two different fluids in the microscale channel through a separate channel. The core fluid is composed of polymeric solution, whereas the sheath act as a lubricant or crosslinking agent. The polymeric precursor is converted into solidified fiber using photopolymerization, ionic crosslinking reactions, phase separation, and solvent

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evaporation technique. The limitation of the process includes fabrication of microchannel, injecting the fluids into the channel, and selection of suitable solidification process, thereby hindering the scalability of the technique [12]. Solution blowing is a suitable method for processing materials with low dielectric constant and/or electrical conductivity. The technique involves forcing polymeric solution through a concentric nozzle as the droplet of a solution emerges from the inner nozzle, high pressure compressed air is passed through the exit of the outer nozzle. The change in the shape of a polymeric solution takes place from droplet to conical shape by the action of compressed air. As the critical value is attained, the aerodynamic forces overcome the surface tension of the solution. The jet emerging from the nozzle is stretched and collected on the collector by means of evaporation of the solvent. The limitation of the solution blowing includes the accumulation of evaporated solvents which creates environmental and health concerns [13]. Electrospinning is a versatile technique used for the production of nanofibers for various biomedical applications such as drug delivery, scaffold for tissue engineering, biosensors, wound dressing, and medical implants [14]. The fibrous structure offers advantages such as a higher surface area to volume ratio, tunable surface properties such as uniform pore size and porosity and mechanical property. The spinning system comprises the voltage supply, a syringe pump and a needle and collector. The polymeric solution is prepared by dissolving in a suitable solvent and loaded on the syringe pump. The metallic needle is connected to the negative terminal, whereas the collector is either ground or negatively charged. The application of voltage creates a charge between the polymeric chain, thereby creating instability. Upon increase in voltage creates repulsive force and on reaching a threshold value, the electrostatic force overcomes the surface tension of the polymeric solution resulting in the deformation of a circular droplet to a conical shape called as Taylor cone. Increasing the electric field causes the polymer chains to split into nanofibers allowing to deposit on the collector, which is placed at a known distance on evaporation of the solvents. The spinning of nanofibers utilizing the electrospinning technique can be influenced by three parameters such as machine, material, and environmental aspects. The material parameters such as viscosity, concentration, and conductivity of the polymeric solution, the molecular weight of the polymer, and selection of solvent, whereas the machine parameters, include application of voltage, the mass throughput rate, and the distance between the tip of the needle and collector. Temperature and relative humidity influence the spinning of nanofibers by controlling the fiber morphology as the environmental condition aids in the evaporation of the solvent [15]. The limitation of the process includes high electric field, usage of flammable solvents, low production rate and scalability and sensitivity to atmospheric condition. Hence, the centrifugal spinning technique allows the production of nano to microfibers by utilizing centrifugal force with the advantage of high production rate, fast and large-scale fabrication and usage of environmentally friendly solvents and the fibers web is widely used for technical applications as shown in Fig. 2.

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Fig. 2 Technical applications of centrifugal spun fibers

Fig. 3 Process sequence of centrifugal spinning

3 Principle of Centrifugal Spinning Centrifugal spinning enables the production of micron to submicron fibers comprising polymeric, inorganic, or composite materials. The process sequence of centrifugal spinning is shown in Fig. 3. The fibers were spun when the centrifugal force and hydrostatic pressure overcome the surface tension of the pendant drop formed by the polymeric solution at the orifice of the spinneret. The droplet is stretched into fine fluid and drafted into fibers by the frictional drag of air by means of evaporation of the solvent. The centrifugal force (F) acting on the polymeric solution can be calculated based on the mass of the material (m), the radius of the spinneret (r) and the speed of the spinneret (ω) using Eq. (1)

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Table 1 Parameters influencing fiber formation in centrifugal spinning process Material variables Viscosity of the polymer solution Surface tension of the polymer solution Polymer concentration and molecular weight

Machine variables Rotational speed of the spinneret Nozzle diameter Nozzle–Collector distance

F ¼ mrω2

ð1Þ

The centrifugal force decides the delivery rate and drawing force acting on the liquid to produce nanofibers, in comparison with other production techniques, the delivery rate is governed by a piston pump in electrospinning and extruder in melt spinning, whereas the drawing force is controlled by the applied voltage and electric field acting on the polymer chain in electrospinning process whereas hot air or heated rollers in melt spinning for drawing the molten liquid into nanofibers [16, 17]. Centrifugal spinning process can be categorized into two types such as melt centrifugal spinning and solution centrifugal spinning process. In the melt centrifugal spinning process, the polymer is melted by an external heating element, whereas in the solution centrifugal spinning, the polymer is dissolved in solvent and spinning was carried out. The advantages of the melt centrifugal spinning process include solventfree production of submicron fibers and an ecofriendly process compared to solution centrifugal spinning as the latter technique requires solvent recovery from the fibers but avoids the treatment of fibers at high temperature unlike the melt centrifugal process as it tends to degrade the polymer [18]. The morphology of nanofibers depends on the material and machine variables as shown in Table 1. Optimization of these parameters not only has control on fiber morphology but also on jet rupture and the formation of droplets due to PlateauRayleigh instability. Capillary number can be used to study the jet break-up, which can be defined as ratio of Weber number (We) to Reynold number (Re). The ratio gives the relationship between viscous force to the surface tension force. The capillary number, Weber number, and Reynold number can be calculated using Eqs. (2), (3) and (4) Capillary Number C a ¼

Weber Number ðW e Þ Reynold Number ðRe Þ

Weber Number W e ¼

ð2Þ

ρU 2 D γ

ð3Þ

ρUD η

ð4Þ

Reynold Number Re ¼

Where U, D, ρ, η, and γ are the speed of polymer exiting the spinneret, spinneret diameter, density, dynamic viscosity, and surface tension of polymer solution, respectively. The formation of shorter jet length and the earlier jet break up was seen in the polymeric solution with a low capillary number [19].

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4 Material and Machine Parameters Influencing the Fiber Formation The material variable has to be optimized based on the type of polymer and solvent used for the preparation of the spinning solution. In centrifugal spinning, solution viscosity plays a major role in determining the fiber morphology. The critical viscosity of the polymeric solution depends on the entanglement concentration of the polymeric chain (polymer chain entanglement density). If the polymer entanglement is low, the fiber formation on the polymer is not possible due to insufficient chain overlap, on the other hand, if the polymer chain density is equal to the critical value, the fiber formation is possible with beads. Uniform bead-free fibers are formed when the polymer entanglement concentration is greater than the critical concentration due to sufficient overlapping of polymer chains. Surface tension of the polymeric solution decides the fiber morphology as it is responsible for the jet stability. The centrifugal force when it overcomes the surface tension of the polymeric solution, results in the elongation of liquid droplets thereby resulting in stretching and elongation into fibers. Polymer concentration is another material variable in deciding the fiber morphology. If the polymer concentration is low, spinnability into nanofibers is limited due to insufficient chain entanglement and higher surface tension. If the polymeric solution finds an optimal value with suitable polymer concentration and polymer chain entanglement results in bead-free uniform fine fibers. Increase in the concentration of polymer results in an increase in larger and thicker diameter fibers greater than 1 μm can be produced and also results in clogging of nozzles in the spinneret. Further increase in the concentration of polymer, increases the viscosity of the polymer solution, thereby increasing the stress relaxation time by causing poor evaporation and resisting jet thinning into nanofibers [20]. The machine variables in centrifugal spinning include rotating speed, nozzle diameter and distance between nozzle and collector. The rotational speed has influenced the centrifugal force acting on the polymeric solution, which aids in solvent evaporation and elongation of the polymeric chain. Therefore, increasing the rotational speed increases the centrifugal force acting on the liquid jet emerging from the nozzle, which results in the production of finer fibers produced by extending and thinning the polymeric solution. The nozzle diameter in the spinneret has the influence on the mass throughput rate, which has a direct influence on the fiber diameter. Decreasing the nozzle diameter results in lesser mass throughput thereby decreasing the fiber diameter. On contrary, increasing the nozzle diameter, increases the diameter of the fiber. The centrifugal spinning machine utilizes a small diameter nozzle if fine fibers are desired. Nozzle–collector distance decides (NCD) decides the fiber morphology in relation to the evaporation of solvent from the solution. By altering the distance between the nozzle and collector, the change in flight time of the liquid can be achieved thereby enhancing the evaporation rate of the solvent from the polymeric solution. Increasing the distance, decreases the fiber diameter due to the longer travel of liquid from the nozzle tip to the collector [21].

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5 Application of Centrifugal Spun Nanofibers in Various Biomedical Applications Centrifugal spinning technique is widely used for biomedical applications such as wound dressing, tissue engineering, and drug delivery applications compared to other spinning techniques as the latter possess the disadvantage of material selection, safety and cost of processing. Centrifugal spinning offers a higher production rate and also offers processing of a variety of polymers in various forms such as melts, solutions, and emulsions [22]. This chapter summarizes the machine and material parameters and their application in the biomedical area.

5.1

Tissue Engineering Applications of Centrifugal Spun Fibers

The essential component of tissue engineering scaffold includes a porous polymer matrix, cells and regulators such as growth factors cytokines to restore, maintain, or improve the function of the cells or organs. The major requirement of the scaffold includes mechanical properties, pore size, microstructure, biocompatibility, biodegradability, and surface property that assists the cell growth [23]. Nanofibers are widely used for tissue engineering applications due to higher surface area to volume ratio, which offers the higher reactive sites for tissue growth. Polycaprolactone (PCL) scaffold was centrifugal spun into nanofibers by melt spinning and solution spinning principle to promote the growth and differentiation of cells in the area of biomedical applications. PCL melt spun fibers were prepared by heating the known weight of PCL polymer in the spinneret, and the temperature was varied from 120 to 250 C, then the setup was rotated at the speed of about 18,000 rpm for 30 s to produce the fibers on the collector placed at 14 in. distance. Solution-based PCL nanofibers were produced by dissolving the polymer at different weight percentages in methylene chloride. The solution was spun at room temperature with a nozzle to collector distance at 12.5 cm at a maximum speed of 9,000 rpm. The spun fibers were loaded with PC12 cells derived from the rat adrenal medulla to study the cell growth and differentiation after coating with collagen. Uniform and bead-free fibers were formed by heating the polymer to 200 C, upon increasing the heating temperature, the diameter of the fibers decreased due to a decrease in the viscosity of the spinning solution. In the case of solution spinning, the fibers were formed at the concentration of 15 wt % and no fibers were formed at 10 and 20 wt % due to low viscosity. The spectrum of PCL contains hydrocarbon, ether and carboxylic functional groups, upon treatment with protein amide bond was formed, indicating the attachment of proteins with collagen-treated PCL centrifugal spun fibers. PC12 neurons were attached and differentiated onto the melt and solution spun fibers with a diameter of 7 and 2 μm as the diameter of the fiber have the influence on cell migration and growth [24].

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Higher surface area to volume ratio of nanofibers offers physical adhesion and absorption of biomolecules making the substrate suitable for tissue engineering applications. A scaffold was prepared from PCL polymer utilizing needleless electrospinning (ES) and centrifugal spinning (CS) technology for bone tissue engineering application with the platelet as a drug delivery system. The concentration of PCL is varied for ES and CS as 24 and 40 wt % in chloroform/Ethanol solvent mixture (9:1). The spinning solution was loaded on the respective machine, and the spun fibers were sterilized at 70% ethanol solution. The sterilized fibrous webs were loaded with human mesenchymal stem cells and on which different concentration of platelet was introduced. ES-based PCL fibers were compact and uniform in size compared to CS fibers where the diameter varied from nano to micro-scale. Cell adhesion and proliferation were found to increase with an increase in the concentration of platelets in both technology-based fibers as it contains growth factors, cytokines, and chemokines as it plays a major role in cell communication and proliferation. Cell infiltration and proliferation were higher in CS-based fibers due to larger pore size and open structure compared to ES fibers. Cell proliferation was found to decrease in ES fibers due to difficulty in penetration of cells. Therefore, ES and CS fibers have potential in tissue engineering applications [25]. PCL was force spun into nanofibers by dissolving in chloroform and ethanol solvent mixture (9:1) as an artificial cartilage tissue engineering scaffold. The fibrous web was sterilized and washed with Phosphate Buffer Solution (PBS) and loaded with fibroblast cells to study cell migration and differentiation. It was reported that fibroblast cells were found to proliferate and promote growth which was confirmed using fluorescence microscopy [26]. Pore size of the fibrous substrate decides the migration of cells as the cell size ranges from 5 to 20 μm, whereas the pore size of the electrospun mat was 1 μm which hinders the penetration of the cells thereby hindering the tissue development on the substrate. Solvent free with large pores on the fibrous substrate was attempted using the centrifugal spinning process to promote cell infiltration and cellular growth for tissue engineering applications. Melt centrifugal spinning comprises a rotating disk, heating unit, and electromotor for controlling the speed of rotation from 350 to 2,000 rpm. The temperature of the heating unit can be varied from 20 to 300 C. Poly Latic acid (PLA) polymers were introduced on the heated rotating disk through the filler tube and melted, the resultant solution was spun into fibers. It was found that the centrifugal spun fibers offered wide distribution of fiber diameter ranging from nano to micro-scale. Cell proliferation was higher in centrifugal spun fibers compared to electrospinning nanofibrous web when cultured with MC3T3-E1 cells. Solvent-free melt processing technique offered less toxicity and a more ecofriendly process compared to electrospinning which offered lower cytotoxicity and greater proliferation, making the centrifugal spinning process suitable for tissue engineering applications [27]. Centrifugal spun fibers comprising PLA and PLA/BaTiO3 (barium titanate) were prepared for tissue regeneration application. The fibers were aged UV/O3 treatment and cytotoxicity of the fibrous web was analyzed by seeding human dermal fibroblast cells. It was reported that the addition of fillers has no influence on the wettability profile of the PLA fibers. Increasing the concentration of

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fillers (BaTiO3) resulted in increase in the diameter of the fibers. The cell viability on the aged samples was noticed when treated with UV/O3 for 5 min and a filler of 5 wt %. The antibacterial efficiency of the developed samples was tested against Staphylococcus epidermidis and it was found that increasing the concentration of BaTiO3 reduced the growth of bacteria by reacting with active oxygen species (ROS) generation, thereby causing bacterial damage [28]. Poly (L-lactic acid) (PLLA), synthetic polymers are used widely for tissue engineering applications as it degrades lactic acid and are removed from the human body through metabolic pathways. The scaffold was prepared by dissolving PLLA and polyvinyl pyrrolidone (PVP) at different ratios in dichloromethane (DCM) and spun using the centrifugal jet spinning technique. The attachment and proliferation of fibroblast cells on the substrate were confirmed using fluorescence microscopy after treatment with Cell Titer Blue dye. It was reported that increasing the concentration of PLLA resulted in an increase in the diameter of nanofibers but on an increment of rotational speed, the diameter of the fibers was reduced. The surface roughness of the web was increased from hydrophilic to super hydrophilic by modifying the concentration of PLLA by decreasing the content of PLLA and the contact angle of the web decreased. It was reported that the developed fibrous mat offered higher cell attachment and proliferation for the application of dermal tissue scaffolds [29]. Centrifugal melt spinning and solvent-assisted solution spinning were carried out using a commercial cotton candy machine with heat and speed control. Poly (DL-lactide-co-glycolide) (PLGA) was added directly to the spinneret as the drum was preheated to 150 C. The melt was spun into fibers, and the cell culture was filled overnight. Polystyrene (PS) was spun into nanofibers using a solvent-assisted spinning technique by means of dissolving tetrahydrofuran solvent. The solvent was transferred to the spinneret without heating and spun into fibers to which the cells were seeded. Fibroblast cells were seeded and the cell viability was studied under a fluorescent microscope with the staining agent namely Calcein AM (green) and Proprium Iodide (red) to quantify the living and dead cells. It was reported that the cell viability was found higher in PLGA compared to PS, which may be attributed to the difference in scaffold fabrication. Thus, it was reported that the glass transition temperature (Tg) of the polymer decides the suitable fabrication technique as higher Tg can be spun using a solvent-assisted spinning, whereas lower Tg polymers can be spun using the centrifugal melt spinning process. Thus, the random 3D network can be used for tissue engineering applications [30]. Porous structure comprising Ethyl Cellulose (EC) and Polyvinyl pyrrolidone (PVP) by means of centrifugal and electrospinning technique. The solution was prepared ideally for both spinning techniques by means of dissolving EC and PVP at a different mass ratio in aqueous ethanol with a total polymer concentration of 15%. The effect of EC and PVP were found to influence the surface property of the produced fibers. The centrifugal spinning was carried out at 3,500 rpm and the distance between orifice and collector at 12 cm, whereas electrospinning was carried out at 7 kV voltage and needle tip to collector distance of 12 cm. It was reported that rough morphology was formed on increasing the concentration of water and EC

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content. Increasing the water in ethanol/water mixture resulted in cracks on the surface of the fibers, whereas EC content on the total polymer ratio resulted in groove formation. It was concluded that the hydrophobicity of centrifugal spun fibers (148.59 ) was higher compared to electrospun fiber (102.65 ), which was confirmed by measuring the contact angle on the produced substrates. Thus, the centrifugal spinning process provides potential applications in area such as tissue engineering, drug release, filtration and electrode applications due to its porous structure [31]. Alignment of fibers has a crucial role in cell adhesion and proliferation suitable for tissue engineering applications. The alignment of micro and nanofibers mimics the extracellular matrix (ECM), which can be altered using a centrifugal spinning process. The spinning solution comprising of PCL and gelatin was prepared by dissolving the polymers at the different ratios in Trifluoroethanol (TFE) and spun at a laboratory-based centrifugal setup at 5,000 rpm. It was reported that the diameter of the fibers decreased with an increase in the concentration of gelation from 824 to 265 nm. The porosity of the developed nanofibers was found to increase with gelatin concentration from 86 to 93%, whereas the porosity requirement for ideal tissue engineering application ranges from 60 to 90%. The proliferation of the cells was studied by seeding NIH 3 T3 and HaCaT cells on the centrifugal spun fibers. It was found that cell proliferation increased to 65% in PCL/gelatin (70/30) concentration. Further increase in gelatin increased the swellability, thereby reducing the pore size and also cell viability. The centrifugal spinning enables the production of proteinbased polymers with a higher degree of alignment suitable for tissue engineering applications [32]. Poly (3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV), a thermoplastic polymer used widely for tissue engineering applications as it degrades D-L-β-hydroxybutyrate, which is a constituent found in the human blood. The polymer is produced by a reaction between Polyhydroxyalkanoates (PHA) and a copolymer, hydroxyvalerate (HV), as the PHBV offers chain flexibility and processibility into fibers. PHA, a polymer produced by microorganisms such as bacteria and haloarchaea is used for tissue engineering applications due to biocompatibility and biodegradability [33]. PHBV polymeric solution was spun into submicron fibers by dissolving the different concentrations of polymer in chloroform and spun at different rotating speeds. It was reported that the viscosity. Concentration and rotating speed of the spinneret decide the morphology of the fiber, as the viscosity of the polymeric solution tends to increase with the concentration of the polymer. Uniform bead-free centrifugal spun fibers were achieved at 25 wt % concentration of the polymer with a viscosity of 386 cP (centipoise) spun at 9,000 rpm [34]. The spinnability of pristine chitosan was difficult due to the polyelectrolyte nature of the polymer and high viscosity. Carboxylated chitosan and polyethylene oxide (PEO) solutions were prepared by dissolving in deionized water at different concentrations and spun using the rotary jet spinning technique with machine parameters of rotating speed at 4,500 rpm, distance from spinneret to the collector of 13 mm and nozzle diameter of 0.1 mm. The formed fibers were uniform and bead free with the diameter varying from 50 to 500 nm, thus making the fibrous web suitable for a

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tissue engineering and antibacterial wound dressing applications [35]. The prerequisite for tissue engineering scaffold is to mimic the extracellular matrix by maintaining the hierarchical three-dimensional architecture. Rotary jet spinning (RJS) technique was utilized to produce a fibrous web as it allows cellular microenvironments to produce functional and structural tissues. PCL, PCL/Collagen and PCL/gelatin fibrous web were produced by dissolving the polymers with total solid content of 6 wt% in HFIP, and the solution was spun utilizing the RJS and ES system in order to compare the tissue growth on the substrate by seeding ventricular cardiomyocytes, fibroblasts and cortical neurons of the neonatal rat. It was reported that the modified spinning system allows the production of nanofibers with fiber alignment comprising protein, collagen and gelatin, and polymer PCL. The technique allows the production of fibers comprising high protein content with a higher production rate compared to the electrospinning technique. The developed substrate offered higher cell alignment, maturation, and organization compared to the electrospinning technique [36]. Janus type of polymeric nanofibers was constructed with two different materials within a material as the structure offers different properties on either side of the material. The fabrication of biphasic Janus fabric was carried out using the centrifugal jet spinning process as it depends on rotational speed and solution characteristics for the production of scaffold for tissue engineering application. Centrifugal jet spinning process (CJS) is similar to the coaxial-spinning process, but the Janus type of fabric maintains a different phase with the retention of maximum contact between the phases. CJS system consists of two reservoirs with a single sidewall orifice at a distance of 1.025 mm apart. PCL and PCL/GE solutions were prepared by dissolving in 1,1,3,3,3-hexafluoro-2-propanol (HFIP), and the solutions were fed onto two reservoirs and the spinning was carried out at three different rotating speeds. The enzyme study using lysozyme and cardiac valve interstitial cells were seeded on both sides of the substrate to study the tissue growth. It was reported that increasing the rotational speed decreased the fiber diameter due to increased centrifugal and hydrostatic force acting on the polymeric solution. The enzyme solubility was found to increase with an increase in gelatin concentration in PCL/GE blend. The Janus nanofibers offered high amount of cell growth with increasing gelatin concentration in the blend compared to the hydrophobic PCL layer. The developed web with biphasic properties can be used for tissue engineering applications by loading different drugs on both sides of the spun nanofibers [37].

5.2

Drug Delivery Applications of Centrifugal Spun Fibers

The major requirement of drug delivery fibrous scaffold is the controlled and sustained release of drugs rather than burst release as it avoids the side effects such as infections, vomiting, fatigue, loss of taste, anemia and destruction of the healthy cells and immune system. The controlled drug release is attained by the slow degradation of the polymer, thereby promoting wound healing applications [38].

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Release of biomolecules from the electrospun nanofibers was limited due to reduced pore size and thickness. Emulsion centrifugal spinning process was carried out to produce core-shell fibrous structure in order to release the biomolecules drug release application. Pluronic F-68 was dissolved in ethanol and which biomolecules were included, the solution was blended to prepare emulsion by dissolving in PCL in chloroform/ethanol mixture (9:1). The shell structure is responsible for cell adhesion as PCL forms the continuous phase, and the core is responsible for the release of biochemicals as Pluronic F-68 forms the droplet phase. The release of biomolecules lasted for 7 days and regeneration of tissues was enhanced upon the incorporation of platelet lyophilizates as it contains tissue growth factor. The developed substrate offered cell proliferation which was confirmed upon seeding fibroblast and osteoblast cells, thus making the emulsion centrifugal spinning suitable for drug delivery applications to release susceptible bioactive molecules [39]. Rotary spinning of PCL/PVP was attempted with drug tetracycline for drug delivery application. PCL/PVP solution was prepared by dissolving in chloroform/ methanol solvent mixture (9:1) with a total polymer concentration of 12 wt%. To the prepared solution, 0.2 wt % of tetracycline was added, and the resulting solution was centrifugally spun. Increase in the diameter of the fibers was noticed from 300 to 927 nm due to the addition of PVP, which is a hydrophilic polymer that tends to absorb moisture from the atmosphere. The weight loss and degree of swelling play a major role in drug delivery application as the developed web possessed a 60–130% swelling percentage, and the weight loss was seen in pristine PVP followed by the blend (PCL/PVP) compared to hydrophobic PCL nanofibers. The antibacterial efficacy of the developed substrate was tested against gram-positive (S. epidermidis, B. megaterium) and gram-negative bacteria (E. coli, P. aeruginosa) as skin pathogens. The drug offered higher antibacterial activity against gram-positive bacteria with the maximum zone of inhibition of 42 mm compared to gram-negative bacteria of 33 mm as tetracycline binds the ribosomes of the bacteria, thus preventing attachment of aminoacyl tRNA to tRNA – ribosome complex, thereby inhibiting protein biosynthesis of the bacteria. The drug release was slow due to the hydrophobicity of PCL and structural stability of CS fibers. Upon incorporation of PVP, the drug release was increased from 12 to 74% within 24 h as the rapid release of the drug offers inhibition against the dermal infection and controlled release prevents the growth of secondary infection. It was reported that the addition of PVP in hydrophobic PCL increased cell adhesion and proliferation thus, the substrate acts as a potential wound dressing with inhibition of microorganisms [40]. Carvedilol is an adrenoreceptor antagonist drug used for the treatment of hypertension and coronary artery disease, but the drug release through oral administration was limited due to poor solubility of the drug in aqueous solvents. The drug stock solution was prepared by dispersing carvedilol in ethanol followed by the addition of citric acid monohydrate, and the dispersion was stirred till the formation of a clear solution. The total polymeric concentration was fixed to 50 wt % and the spinning gel was prepared by blending carvedilol solution with hydroxypropyl cellulose and spun through the rotary spinning technique. It was reported that the average diameter

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of the fibers was 135 μm. To assess the oral dispersibility of the drug-loaded microfibers, the fibers were milled along with citric acid anhydrous and sodium bicarbonate as excipients. The drug release profile was independent of the pH of the applied medium. It was reported that rotary spinning of carvedilol from crystalline to amorphous transition, thus making the milled fibrous web as orodispersible tablets [41]. Drug-loaded nanofibers offer a higher surface area to volume ratio and also drug encapsulation ability, high drug loading, and stability. Sustained drug release of anticancer drugs inhibits the growth of cancer cells by reacting with DNA and RNA synthesis leading to cell death. Polycaprolactone was spun into nanofibers using force spinning containing mercaptophenyl methacrylate functionalized carbon nanoonions (PCL/f-CNOs) composite for a sustained drug release profile. The polymeric solution was prepared by dissolving PCL, doxorubicin and functionalized carbon nano-onions in trifluoroacetic acid and spun. It was reported that increasing the concentration of f-CNOs resulted in the decrease in fiber diameter to 353 nm as the pristine PCL offered fiber diameter of 596 nm, upon the incorporation of doxorubicin resulted in a decrease in the average diameter of fiber around 470 nm due to amalgamation and electrostatic interaction with PCL. Wettability of the fibrous substrate decides the cell adhesion and proliferation as the wetting behavior of the fibrous substrate was decreased either due to electrostatic interactions between PCL and f-CNOs or the hydrophobic nature of f-CNOs. The release of the drug from the substrate was decided by the porosity, which was proportional to polymer concentration as the average pore diameter of (PCL/ doxorubicin /f-CNOs) was 6.03 μm. The cell viability of the substrate was studied by loading with fibroblast cells and it was reported that significant cell viability was seen due to higher hydrophobicity and lower degradation rate of (PCL/doxorubicin/f-CNOs) nanofibers. The drug release from the developed nanofibers was pH responsive as the drug doxorubicin released 87% at pH 6.5 whereas at pH 5, complete release of the drug around 99% was noticed caused due to π-π stacking interactions between DOX and f-CNOs [42]. Centrifugal spinning enables the production of microfibers containing low aqueous solubility drugs along with fiber-forming polymer. Olanzapine (OLZ) and piroxicam (PRX) are poorly water-soluble drugs spun into fibers as a solid dispersion in sucrose microfibers for the treatment of psychoses and arthrosis. The modified centrifugal spinning comprises of rotating spinneret, collector and heating unit. The spinneret is formed by placing the upper aluminum plate over the concave bottom plate with a gap of 0.8 mm as it is responsible for fiber production. The mixture containing (90 wt%) sucrose and (10 wt %) was mixed using a mortar and a known amount of mixture was transferred to preheated spinneret and the melt solution was spun at 2,400 rpm at room temperature. The drug loading efficiency of the microfibers was studied by dissolving fibers in dimethyl sulfoxide, as the drug and fibers were soluble in the solvent and the resultant was analyzed using UV (Ultraviolet Visible) spectroscopy. Saturation solubility was carried out by dissolving the microfiber-containing drug and pristine drug at PBS at a pH of 6.8. Undissolved material corresponds to saturation. Invitro dissolution study was carried out using USP type II paddle apparatus, which contains phosphate buffer at pH of

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6.8 as it represents the healthy oral saliva. The microfibers were disintegrated into powder and loaded into gelatin capsules with an onset dissolution time of 3 min. The liquid was taken to study the drug dissolution time utilizing UV detection, and the graph was plotted against the drug release profile against time. It was reported that incorporation of drug molecules within the polymer matrix resulted in an increase in fiber diameter due to poor drug–carrier interactions representing entrapped drugs within the amorphous sucrose structure. Highly stable solid dispersions can be prepared either by hydrogen bonding or hydrophobic interactions between the drug and polymer. It was reported that hydrogen bond interaction was found between the proton donors of sucrose molecules as it contains free hydroxyl group and proton acceptor in the tertiary amines from azepine and piperazinyl rings in OLZ and SO2, amide groups and tertiary amines of PRX. It was reported that dissolution of drugs was improved by increasing the concentration of sucrose due to hydrogen bonding interaction between water-soluble polymer and the drug. The dissolution at in vitro condition was higher compared to the pristine drug due to high surface area and low density of the microfibers resulting in systemic absorption leading to improved oral bioavailability compared to conventional solid formulations [43]. The alignment and stacking of the fibers with different stack angles can be attained by combining centrifugal spinning with the electrospinning (CES) process with multiple orifices for fiber production. Polyvinyl pyrrolidone (20 w/v %) was dissolved in ethanol followed by the addition of Tetracycline Hydrochloride (TE-HCl) (5 w/w%) antibiotic, and the homogenous solution was spun using the CES system. The modified CE system comprises the spinneret, DC (direct current) motor, syringe pump and high voltage supplier. The four nozzles in the spinneret were placed at 90 angle at a distance of 5 mm to avoid the accumulation of electric field. The circular collector was placed at a distance of 9–14 cm, which was covered with conductive iron wire and grounded to fringe the applied electric field. UV spectrophotometry was used to study the antibiotic release from the PVP nanofibers. The spinning system deposited the fibers on the collector based on the direction of rotation of the spinneret as the alignment influences cell proliferation and orientation. The uniform bead-free fibers were formed at 20 w/v% of PVP, 70 rpm of spinneret speed and 9 cm of a spinneret to collector distance. It was reported that drug release from the random arrangement of fibers was higher compared to stack and aligned fibers due to reduced order and enhanced aqueous interaction. It was concluded in the literature that controlled drug release can be obtained by controlling the orientation and stacking of fibers using the CES system [44].

5.3

Wound Dressing Applications of Centrifugal Spun Fibers

Wound can be defined as a disruption in the continuity of the epithelial lining of the skin caused due to chemical, thermal and physical damage. Wounds can be classified into two categories based on the nature of the repairing process such as chronic or acute wounds. Acute wounds can heal completely within 8–12 weeks, whereas

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chronic wounds heal slowly for a period of more than 12 weeks. Wounds can be classified into three categories based on the number of layers of damage such as superficial (wounds involving only on the epidermis), partial thickness (wounds involving the dermis and epidermis), and subcutaneous wounds (wounds on the subcutaneous or deeper tissue). The stages of wound healing involve hemostasis and inflammation, migration, proliferation, and remodelling phases to re-establish the integrity of the damaged or lost tissue [45, 46]. Micro-fibrous textile materials are widely used for wound dressing applications as it mimics the Extra Cellular Matrix (ECM) found on the skin due to high surface area to volume ratio, porous and conformable structure to the wound surface. Electrophoretic deposition (EPD) is a versatile and simple technique used for biomedical applications as the biopolymers are charged into a stable colloidal suspension and moved through the liquid by application of the electric field and the polymers are deposited on an oppositely charged conductive substrate thereby forming the intended material for wound healing applications [47]. In the EPD technique, the electric field is applied between cathode and anode electrodes, alkaline and acidic pH gradient occur between the charged electrodes. PCL nanofibers were spun using the centrifugal spinning technique with 15% polymer dissolved in chloroform. Chitosan was deposited on the surface of PCL fibrous membrane by the EPD technique. The deposition of chitosan using the EPD technique was varied by changing the parameters such as pH, voltage and molecular weight and concentration of the polymer as it influences the deposition and weight add percentage (%). It was reported that an increase in molecular weight of the chitosan resulted in an increase in weight add % due to uniform coating of the polymer at interstices of the fibrous web. Increasing the voltage resulted in an increase in add on up to 10 V and upon an increase in voltage resulted in a decrease due to the evolution of hydrogen gases from the electrode resulting in a rough surface with larger pores. It was reported that lower pH resulted in poor deposition due to a decrease in ionization and charge density. Uniform deposition of chitosan with pore structure was seen at pH 5.5, which is ideal for wound dressing application. It was concluded in the literature that the EPD of chitosan resulted in uniform deposition with a medium molecular weight with a pH of 5.5, voltage of 5 V and duration of 10 min, and deposition of chitosan can be tailored for various biomedical applications [48]. Electrophoretic deposition and centrifugal spinning technique allow the deposition of chitosan along with other bioactive molecules such as protein, drug, growth factors, and active agents suitable for tissue engineering application [49]. Wound dressing was prepared by the centrifugal spinning of Polycaprolactone (PCL) on which electrophoretic deposition of chitosan/Polyethylene Glycol (PEG) along with silver nitrate (AgNO3) was carried out. The spinning solution was prepared by dissolving PCL in chloroform and spun at the speed of 5,000 rpm. Chitosan, pH-sensitive polymer as it turns to a gel state above pKa of 6.3. The cathode alkaline region can be utilized for the deposition of the chitosan on metallic implants. For EPD, chitosan was dissolved in 2% acetic acid and pH was adjusted to 5 by the addition of sodium hydroxide (NaOH). Silver nitrate solution by dissolving AgNO3 in acetic acid, and the solution was added dropwise onto the

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chitosan solution to prepare 1 mM concentration. Graphite sheet was used as an anode and cathode electrode, the centrifugal spun PCL was placed on the cathode electrode. To the final solution, 0.5% of PEG was added and EPD was carried out at a distance of 5 cm between the electrodes at 5 V voltage for 10 min. The EPD-treated samples were dried at room temperature to vaporize the moisture content from the web. It was reported that silver nanoparticles of size up to 15 nm were formed by the reduction of PEG confirmed through Transmission Electron Microscopy (TEM). The developed substrate offered a slight increase in hemolysis ratio as it refers to the destruction of red blood cells on contact with the material after the introduction of PEG, which can be related to the release of silver ions. The developed substrate offered antibacterial activity against P. aeruginosa and S. aureus, which enhances wound healing and offers infection control. The toxicity of the developed substrate was studied using MTT (3-(4, 5-Dimethylthiazol-2-yl)-2, 5-diphenyltetrazolium bromide) assay and it was found that the introduction of PEG resulted in a slight decrease in cell viability due to burst release of silver ions from the coated textile substrate [50]. Controlled release of drugs prevents the burst release of the drugs thus reducing the frequent drug administration, which harms the adjacent living and healthy cells. Poly Latic acid (PLA) and gelatin (GE) blend force spun fibers undermine the cell growth by enhancing the adhesion, proliferation, and migration in the presence of antibacterial agents Ciprofloxacin hydrochloride (CFH), thus the resultant substrate enhances the wound healing process. PLA and GE solution (85:15) was prepared by dissolving in hexafluoro-2-propanol to which CFH was added and the solution with dispersed antibacterial agents was force spun with a syringe to collector distance of 15 cm. It was reported that increasing the concentration of CFH from 0 to 12% resulted in an increase in the diameter of the fibers from 513 to 622 nm, which can be contributed to the increase in solution viscosity and chain entanglement and also improved the surface smoothness of the fibers. The latter prevents the cause of secondary injury to the wound thus making it ideal for wound dressing application. The contact angle of PLA/GE nanofibers was 81 which was due to the blending of hydrophilic GE and on increasing CFH, the contact angle of the substate decreased due to the hydrophilic antibacterial agent, thus developing a moist environment for promoting the wound healing process. The substrate offered antibacterial activity against Staphylococcus aureus and Escherichia coli with the zone of inhibition of 44 mm and 36 mm on increasing the concentration of CFH. It was reported that the burst release of drug from the fibrous mat was initially seen due to surface adsorbed agents, followed by sustained release due to encapsulation of drug within the polymer matrix confirmed by an in vitro drug release study [51]. Fibrous scaffold for wound healing application was prepared by blending bacterial cellulose (BC) with polylactic acid (PLA) and polycaprolactone (PCL) by rotary spinning technique. The blend of BC with PCL and PLA was prepared by dissolving the polymers in chloroform and force spun at 36,000 rpm to produce a bandage-like fibrous structure. It was reported that the incorporation of BC in PLA and PCL resulted in bead-like structures along with fibers. It was reported that the average diameter of the fibers was 19 μm and 6 μm for BC blend with PLA and PCL,

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respectively. Higher surface area to volume ratio of the fibrous scaffolds offers cell adhesion and proliferation, thereby aiding the wound healing process and enhancing the mechanical property of the fibrous web [52]. Carboxymethyl chitosan (CMCS) with a degree of carboxylation is 85% and polyethylene oxide (PEO) was dissolved in deionized water and force spun using self-designed centrifugal spinning machine. It was reported that increasing the concentration of CMCS resulted in an increase in the shear viscosity of the solution. The ideal water vapor transport rate for wound healing application is 2,000–2,500 g2day1, whereas the developed substrate offered 2 2,748–2,983 g day1 due to increased porosity of the fibrous web. Hydrophilicity and absorption of liquids allow the penetration of biomolecules, cells and nutrients to aid the wound healing process. The fibrous web offered a liquid absorbing ability of 29 g/g on testing with 2.5 mmol L1 calcium chloride dihydrate and 142 mmol L1 sodium chloride as it represents the liquid oozing from the wound site. It was reported that the antibacterial activity of the developed substrate was higher grampositive bacteria (S. aureus) compared to gram-negative bacteria (E. coli) due to the difference in an antibacterial mechanism. In the case of gram-positive bacteria, electrostatic interaction between the substrate and teichoic acids on the cell surface causes the cell disruption of bacteria whereas, in gram-negative bacteria, chitosan creates a chelation effect by reacting with cations when pH > pKa value and also reacts with anionic parts of the lipopolysaccharide found on the outer cell wall causing cell death, thus the fibrous web has the potential in wound dressing applications [53]. Centrifugal melt spinning was adopted to spin the polymers without solvent and offers a high production rate. Polyethylene glycol was used as polymer and drugs used for incorporation within the polymer matrix include Ibuprofen (IBU), Tinidazole (TNZ), and nifedipine (NF) for wound healing applications. The release of drug from the substrate was studied using a release medium based on the drug used as PBS for IBU, SLS for NF and deionized water for TNZ, respectively, and the drain was taken for UV spectrophotometer. The drug release from the fibrous substrate was initially high compared to the pure drug release profile due to the change in crystallinity of the drug caused by the changes in solid state of the drugs. It was reported that the fibrous web offers modulable drug release for wound healing applications [54]. Co-axial centrifugal spinning is adopted for spinning core-shell fibers comprising Carboxylated chitosan (CCS) and PEO with two types of drugs such as ibuprofen (anti-inflammatory drug) and human epidermal growth factors (hEGF) for drug delivery and wound healing applications. The spinning solution was prepared by dissolving CCS and PEO in deionized water to which two types of drugs were added and centrifugal spun. The anti-inflammatory drug incorporated solution was loaded onto the spinneret nozzle of 22G acts as a sheath layer, whereas hEGF incorporated was loaded onto the spinneret nozzle of 30G, which acts as a core layer. The developed core-shealth fibers were tested for drug release, cell viability and antibacterial evaluation against the selected bacteria Escherichia coli (E. coli), Pseudomonas aeruginosa (P. aeruginosa), and Staphylococcus aureus (S. aureus).

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The release of bioactive drug, hEGF was moderate due to encapsulation of the drug within the sheath structure. The average diameter of the core-sheath fibers was 1,154 nm and monoaxial fibers was 481 nm thus causing a slow release of drugs as the erosion rate of the sheath is lower than that of the core. The cell viability was seen higher in hEGF incorporated core-sheath fibers compared to pristine CCS-PEO nanofibers. It was reported that drugs incorporated fibrous mat offered Zone of Inhibition of 12.5 mm, 16 mm, and 10.7 mm activity against E. coli, S. aureus, and P. aeruginosa. The substrate offered wound healing and drug delivery application with a higher production rate compared to the traditional co-axial electrospinning process [55]. Multicomponent fibers are produced by modifying the centrifugal spinning with multiple spinnerets, thereby controlling the physical and chemical properties for drug delivery, tissue engineering, and functional textile applications. The multicomponent fibers were produced by sectioning spinneret into three subdisk which are integrated in a single axis in a vertical direction. The three different polymeric solutions were prepared by dissolving polystyrene, polymethyl methacrylate, and polyvinyl pyrrolidone in chloroform and loaded onto each subdisk of the spinneret and the solution was spun. It was reported that the multispinning system increased the production time of fibers by 300 folds compared to the electrospinning process. The developed fibers were used for mask filter application as it offers 97% droplet capturing efficiency, thereby having the potential to filter fine dust and infectious viruses [56].

6 Advances in Centrifugal Spinning Process Nozzle-free centrifugal spinning was carried out to produce the nanofibers comprising poly- (methyl methacrylate) (PMMA) by means of dissolving in chlorobenzene with the polymer concentration of 5 wt %. The solution was placed into a chuck of spin coater and rotated at 3,000 rpm. Thus, the fibers were produced along with the cup-shaped beads and the produced fibers were hydrophilic in nature. Then, it was reported that the centrifugal spinning is suitable for the polymers which are not amenable to the electrospinning process [57]. Nanofibers were produced by modifying the centrifugal spinning by utilizing the solution blowing technique for biomedical applications. The modified apparatus provides the advantages compared to the electrospinning by offering higher production rate, environmentally friendly and toxic-free chemicals and avoids the utilization of a higher electrostatic field for the production of fibers. Solution blown centrifugal spinning utilizes an air source in addition to the centrifugal force as the solution is fed into the nozzle in the spinneret, the droplet deforms allowing the solvent to evaporate leaving the fibers along with beads on the collector. It was reported that the fiber morphology was influenced by airflow rate, solvent evaporation rate, viscosity, and surface tension of solution [58]. Centrifugal spinning combined with UV-initiated polymerization was carried out to produce fibers comprising of ene functional monomer (Dipentaerythritol

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pentaacrylate), thiol functional monomer (pentaerythritol tetrakis) along with polyethylene oxide (PEO), photoinitiator (Irgacure) and ethyl acetate. PEO polymer was responsible for fiber formation and ethyl acetate for controlling the viscosity of the photopolymerization gel. The premixed monomer was fed onto the centrifugal spinning system which is equipped with a UV light source, an igniter, a mercury short arc lamp and UV transparent reflective mirror. The fibers were photopolymerized with a constant source of light with an intensity of 1,160 mW/ cm2. It was found that the monomer mixes viscosity and spinneret hole decided the fiber diameter by means of controlling the delivery rate. The elasticity of the monomer mixes has no significant influence on the fiber diameter. The centrifugal spinning has the potential in fabricating high-performance fibers due to its scalability and tunable machine parameters [59]. Ultrafine nanofibers were produced by combining centrifugal field and electrostatic field as the Centrifugal Electrospinning (CE) technique utilizes low voltage and slower speed compared to the conventional electrospinning process. The syringe was placed on the rotating disk which is supported on the insulated motor responsible for revolution of the disk and the collector was placed below the support disk. Polymeric solution comprising Polystyrene (PS) and polymethyl methacrylate (PMMA) were prepared by dissolving in tetrahydrofuran (THF) and polyvinylpyrrolidone (PVP) was dissolved into ethanol. The spinning solutions were loaded and spinnability into fibers was studied. It was reported that combing the centrifugal force and electrospinning process resulted in the reduction in applied voltage, rotating speed, and collector distance by 2.8–6.0 kV, 360–540 rpm, and 2.0–4.0 cm, respectively [60]. Electrostatic-assisted centrifugal spinning setup was made by modifying the centrifugal spinning process that includes spinneret, motor, collector, and insulator. The vertical electrostatic field was applied to the spinneret by means of a carbon brush that charges the polymeric solution. The collector was grounded by placing the conveyor belt on the metallic plate. The insulating coupler was used between the motor shaft and spinneret shaft to prevent the damage to the motor from high voltage. PS was dissolved in distilled water and the solution was spun using the modified centrifugal spinning process. It was reported that in conventional CS process, the fibers tend to migrate towards the center of the collector due to unstable airflow by the spinneret whereas in modified CS, uniform deposition of fibers was seen due to the action of electrostatic and centrifugal force. It was also seen that increasing the voltage resulted in a decrease in the diameter of the fibers due to the vertical arrangement of electrostatic force and higher drawing of polymeric jets and uniform bead-free fibers were formed at 8 kV. It was also reported that reducing the nozzle size from 23G to 30G decreased the fiber diameter from 1,848 to 500 nm and an increase in the rotating speed of the spinneret resulted in a reduction in the diameter of the fibers due to increased airflow and better evaporation of solvents [61].

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7 Conclusion One of the major challenges in the production of submicron fibers is the ease of production and consistency in terms of diameter along with uniformity of the web. Conventional fiber formation systems yield fibers in the range of 700 to 1,000 nm. With the advent of electrospinning, fibers in the range of 100-300 nm can be produced with the limitation that the production rate of fibers is low. Centrifugal spinning overcomes the above limitation and the fibers can be produced with ease, and the production rate is also significantly higher. Both polymers of natural and synthetic origin can be produced easily with the centrifugal spinning system. Hybrid techniques coupling with centrifugal spinning with electrospinning is also on the rise and it can still pave the way for ultrafine fibres. The potential of centrifugal spinning is to be yet explored by the research community and it has wide potential in the biomedical applications.

8 Future Trends Centrifugal spinning has got the attention of the global research community for its ease of production and versatility, But the major challenge in embracing the technique is in the development of a collection system and a lot of work has gone in this direction to develop the suitable design. Moreover, efforts are being taken by various groups to control the fiber length so that a nonwoven mat with defined pore size and distribution is achieved. Spinning of natural polymers like gelatin and chitosan along with synthetic polymers is likely to open a plethora of applications in the biomedical domain. The major challenge now before the research community is to come up with a design that encompasses the advantages of melting and having a solution as starting dope for the production of fibers. Moreover, it can be adopted easily provided all the natural polymers or its blends can be produced with ease using this technique, and a lot of work is being carried out in this direction.

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Adv Polym Sci (2023) 291: 107–138 https://doi.org/10.1007/12_2022_142 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 Published online: 3 January 2023

Recent Advances in Electrospun Nanofibrous Polymeric Yarns C. R. Reshmi, Rosebin Babu, Shantikumar V. Nair, and Deepthy Menon

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Classification of Yarns Depending on Nanofiber Alignment . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Electrospinning Strategies and Collector Designs for Nanofibrous Yarn Development . . 3.1 Disc-Shaped Collector . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 Ring-Shaped Collector . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3 Filament-Shaped Collector . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.4 Tube Collector . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.5 Cylindrical Static Collector . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.6 Glass Rod Collector . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.7 Friction Double Cylinder . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.8 Metal Frame . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.9 Liquid Bath as Collector System . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.10 Funnel-Shaped Collector . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.11 Hemispherical-Shaped Collector . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Yarn Fabrication by Twisting Electrospun Membrane . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 Collector-Less Yarning Process . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6 AC Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 Recent Advances in the Applications of Nanofibrous Yarns in Biomedicine . . . . . . . . . . . . . 8 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

108 109 110 110 112 113 115 117 118 118 120 121 123 126 127 127 128 130 132 132

Abstract Electrospinning technology has advanced significantly over the past two decades, and fibrous materials in multitudes of geometries, ranging from the conventional 2D membranes to 1D fiber bundles or yarns and recently 3D textile constructs have been developed. Of these, innovations in the engineering of electrospun fibers in the form of yarns, which circumvent the limitations of

C. R. Reshmi, R. Babu, S. V. Nair, and D. Menon (✉) Amrita School of Nanosciences and Molecular Medicine, Amrita Vishwa Vidyapeetham, Kochi, Kerala, India e-mail: [email protected]

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conventional electrospun fibers in terms of their mechanical characteristics, have emerged important. Nanoyarns which consist of thousands of nanofibers bundled together are advantageous because the fibers in yarns retain their nanoscale diameters required for functional benefit, but concurrently exhibit much improved mechanical properties that is key to the manufacturing of scaled-up products. This book chapter elucidates the innovations in nanofibrous yarn fabrication using DC and AC fields, diverse collector designs and modifications in the electrospinning assembly. The advantages and limitations of each of the methods utilized are discussed by highlighting the process yield, mechanical strength, scalability, and ease of fabrication. These nanofibrous yarns by retaining the functionality of fibers also meet the mechanical requirements of various textile processing techniques for developing scalable constructs. The chapter concludes with an overview on the diverse applications of electrospun nanoyarns, with emphasis in the field of biomedicine. Keywords Biomedical scaffolds · Collector design · Electrospinning methods · Nanofibrous yarns

1 Introduction Electrospun nanofiber membranes cut across multidisciplinary fields of biomedical research owing to their remarkable features such as high surface area to volume ratio, interconnected ultrafine fibrous structure, porosity, tortuosity, permeability, and mechanical properties [1, 2]. Few significant applications include its use as drug delivery systems [3], biosensors [4], tissue engineered scaffolds [5], wound dressing materials [6], filters and advanced textiles [7]. New innovative electrospinning methodologies have been explored to develop nanofibrous constructs with diverse 3D morphologies and patterns. Amongst them, nanofibrous bundling techniques to form yarns have grabbed immense attention because of its high processability [8]. Advances in nanofiber yarning technology have gained impetus in developing multidimensional functional materials for diverse biomedical needs. Over time, various electrospinning strategies have been adopted that directly convert nanofibers to mechanically strong yarns of uniform diameter. Such nanofibrous yarns can then be processed into multidimensional constructs by adopting the traditional textile technology techniques such as weaving, braiding, knitting, or embroidery, thus opening new avenues in biomaterials fabrication, as demonstrated by our group in the fabrication of vascular grafts, vascular patches, drug eluting implants, bone grafts, etc. [7, 9–11]. Literature reports several electrospinning approaches to fabricate nanofibrous yarns. Alteration in the collector design/geometry is the key strategy adopted to generate nanofibrous continuous and twisted yarns. Researchers have utilized collectors in the form of disc, funnel, ring, filament, etc. for this purpose. In recent years,

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our group has successfully demonstrated the feasibility of hemispherical-shaped collector for fabricating mechanically strong, uniaxially aligned/twisted and continuous nanofibrous yarns [8]. This book chapter gives a comprehensive overview of the recent advancements on electrospinning collector design and methodologies for obtaining mechanically stable, high-throughput nanofibrous yarns with uniform dimensions. A detailed elucidation of the methods of electrospinning and its setup, collector design and the optimization of various process parameters are also provided. This article concludes with an overview of the applications of nanoyarns in various realms of biomedical technology and the challenges in yarn fabrication.

2 Classification of Yarns Depending on Nanofiber Alignment Yarns with different nanofibrous arrangements are developed by various methods of electrospinning. It is the alignment of nanofibers within these yarns which dictates its diverse characteristics. Depending on the nanofiber alignment, electrospun yarns are mainly categorized as (1) randomly bundled and (2) unidirectionally twisted as schematically represented in Fig. 1. Uniaxially aligned and twisted nanofibers in yarns impart superior mechanical strength than that of randomly oriented ones, which thereby confer good processability using different textile technology techniques. This aside, unidirectionally aligned yarns show improved crystallinity and solvent wicking nature. Hence, fabricating continuous uniaxially aligned yarns having uniform twist should be the major focus during the electrospinning process

Fig. 1 Schematic representation and SEM images of nanoyarns with various alignment (A, A1) highly aligned yarns (B, B1) twisted yarns (C, C1) core-sheath yarns. Reprinted with permission from Refs. [8, 9, 12]

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design. Diverse collector designs have been extensively explored in literature to obtain mechanically stable, aligned nanofibrous yarns using different polymeric materials. These designs are detailed in the following section.

3 Electrospinning Strategies and Collector Designs for Nanofibrous Yarn Development 3.1 3.1.1

Disc-Shaped Collector Electrospinning Towards the Center of the Disc

Recently, several studies have focused on developing electrospun yarns using rotating metallic discs as the collector for obtaining uniaxially twisted continuous yarns. Well-aligned polyacrylonitrile (PAN) nanofibrous continuous yarns were fabricated using a setup that comprises two oppositely charged metal needles focused towards a neutral metal disc, a hollow tapered and insulated metal rod which is placed in front of the disc between the needles as shown in Fig. 2A. Nanofibers were hooked from the oppositely charged fiber jets deposited on the center part of the metal disc and drawn through the hollow metal rod using a metal wire that acts as a guide, and finally gathered on the rotating take-up roll. Uniaxially aligned nanofibrous yarns were obtained in this process as depicted in Fig. 2A1. The rotating speed of the metallic disc is a critical parameter which influenced the stability of the spinning triangle cone formed at the center of the metal disc, rate of bundling, twisting and alignment of the nanofibers in the yarn. Nanofiber alignment is critically influenced by parameters such as distance between the neutral metal disc and hollow metal rod, that between the needles, and the disc rotation speed. Increase in the distance between the metal disc and hollow rod and its rotation speed led to a decrease in the yarn diameter [13–15]. Such disc-shaped collectors are considered to be one of the methods that yield continuous, uniaxially aligned nanofibrous yarns with improved tensile strength. However, the major limitations of this process include the difficulty in hooking the nanofibers through the hollow metal rod under high electric fields and its scalability.

3.1.2

Electrospinning Towards the Edge of the Disc

In some studies, the above metallic disc collector is used to deposit nanofibers at the edge/rim of the disc rather than its center. Here, a positively charged single spinneret focused the jet towards the grounded rim of the disc having a thickness of 1–2 cm, set at a tip-target distance of 10 cm, as shown in Fig. 2B. Randomly oriented nanofibers were deposited on the disc rim which is subjected to a translational motion. The deposited ribbon-shaped membrane was then peeled off from the other end of the disc, twisted and then spooled to obtain fibrous yarns whose

+ve high voltage

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Fig. 2 Schematic of the yarning setup using disc-shaped collector. (A) Electrospinning towards the center of the disc and (A1) Electron micrographs of yarns obtained by this method. (B) Electrospinning towards the edge of the disc and (B1) Electron micrographs of yarns obtained by this method. SEM images reprinted with permission from Refs. [14, 16]

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morphology is given in Fig. 2B1. This method yielded yarns with a very low mechanical strength of typically 0.26 cN/dtex, which makes it difficult to be processed via the textile techniques. This suggests that yarning towards the edge of a disc is an inefficient strategy owing to the difficulty in forming continuous yarns with high throughput and strength [16, 17].

3.2

Ring-Shaped Collector

Metallic rings have also been utilized for producing continuous nanofibrous yarns. Here, dual spinnerets connected to an equal and oppositely charged high voltage focused the polymer jet to the center of a grounded aluminium ring collector, at an inclination of 60° with respect to the plane of the collector. This ring was fitted to a plastic plate containing a ball bearing for its free rotation. The ring collector had typical dimensions of 100 mm outer diameter and 88 mm inner diameter, respectively, and set at a rotation speed of 600–1,300 rpm as shown in Fig. 3A. This rotation of the ring collector ensured alignment of the deposited nanofibers as shown in Fig. 3A1 [18]. Electrospinning onto the rotating ring collector resulted in the formation of a membrane, which was then hooked using a plastic rod to yield a fibrous cone. From this, a nanofibrous yarn was drawn and collected to obtain a yarn

Fig. 3 (A) Schematic of the yarning process using a ring-shaped collector and (A1) the SEM image of the nanofibrous yarns obtained using this setup. SEM image reprinted with permission from Ref. [18]

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spool. Uniformly twisted yarns were obtained owing to the synchronous rotation of the fibrous cone with the ring collector. The nanofiber yarns produced by this technique have reported a twist angle of 54.4° between the fibers and a tensile strength of 93.6 MPa (elongation 242.6%). One of the major limitations of this technique is its scalability.

3.3

Filament-Shaped Collector

Nanofibers deposited and wrapped helically on a filament-shaped collector have also been used to produce yarns. In this strategy, a double spinneret electrospinning method was used, wherein two spinnerets were connected to equal and opposite high voltages. The spinnerets were focused towards the center of a rotating metallic disc through which a single core filament was fed under tension. The nanofibers deposited on the filament collector were wound to obtain a fibrous yarn spool as shown in Fig. 4A. Electrospinning parameters such as spinneret angle, twist rate of the filament, and uptake speed were optimized. Continuous nanofibrous yarns were obtained at a spinneret angle of 90° with respect to the filament plane, disc rotation speed of 500–750 rpm and uptake speed of 1.5 cm/s, respectively [19]. In another study, nanofibers were deposited on a filament placed at the rim of a neutral metal disc-shaped collector. During electrospinning, the nanofibers were focused towards the edge of the disc-shaped collector containing the single core filament, resulting in the filament being wrapped with nanofibers as depicted in Fig. 4B. The thickness of the disc is a critical parameter in this method. Nanofibrous yarns with fiber diameter ranging from 220 nm to 260 nm and strength of 3.25 (cN/dtex) were obtained with the rotating disc having a thickness of 5 mm [17]. In another study, a needleless electrospinning strategy was adopted to develop nanofibrous yarns using a filament-shaped collector. Herein, a polyester yarn was used as the filament-shaped collector, which was translated through a roller system, for a scaled-up electrospinning process as shown in Fig. 4C. This setup could produce continuous nanofiber covered polyester yarns at very high applied voltages of ~60 kV, distance between electrodes at 170 mm and thickness of the core filament at 470 μm. This nanofibrous yarn displays a combination of the high mechanical strength of the core filament and the high surface area to volume ratio of the nanofibrous shell [20]. In another nanofibrous yarning assembly, the collector is comprised of monofilaments of unidirectionally aligned polyamide (Diameter: 50 μm) mounted on a cardboard frame at 10 mm spacing from each other. This collector device is placed 12 cm below the spinneret which is connected to a high potential for nanofiber coating. After coating, the monofilaments were twisted into hybrid yarns. Figure 4D gives the schematic of this assembly. The electrospinning parameters such as flow rate (0.5 mL/h), tip-to-target distance (18 cm) and applied voltage (15 kV) were optimized. The electrospinning parameters, twist ratio and coating time helped to tune the nanostructured morphology of these hybrid yarns [21].

Fig. 4 Schematic representations of the yarning process using different types of filament-shaped collectors. (A) Spinning to the center of a rotating metallic disc through which a single core filament is passed, (B) Electrospinning onto a filament placed at the rim of a neutral metal disc-shaped collector, (C) A needleless scalable yarning process on polyester yarns, (D) Spinning onto a collector comprised of unidirectionally aligned polyamide monofilaments, (E) Electrospinning onto stainless steel monofilament, (F) Electrospinning onto a cotton thread drawn through the center of a rod-shaped collector. (B, D) Reprinted with permission from Refs. [20, 21]

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A robust and automated method for manufacturing continuous nanofibrous yarns over a stainless steel (SS) filament-shaped collector is also reported. Here, the grounded SS filament was placed underneath the spinneret on which the polymeric jet was focused to form a dense submicron fibrous cover. This fibrous mesh was detached to yield a long thread/yarn which was post-possessed by stretching to 300%, creating continuous yarns, as can be seen from Fig. 4E. The entire filamentshaped collector unit was composed of a feeding unit, wiping unit and a winding unit. An SS wire (diameter 100 μm) served as the feeding unit which was placed below the spinneret, by maintaining a tension during operation. A wiping unit in the electrospinning jet area prevented electrospun fibers from bridging between the wire and the spinneret. A cutter wheel served as the winding unit that separated the nanofibrous yarns from the SS wire. Continuous multifilament yarns were also prepared by assembling multiple one meter length yarns manually and twisting them into a ply yarn at 400 twists/m [22]. In another study, the core filament collector was a micron thick cotton thread combined with a metal disc and a rod-shaped collector (explained in Sect. 3.1). The cotton thread was drawn through the center hole of the disc-shaped and rod-shaped collector. The electrospinning setup consists of two oppositely placed syringes arranged symmetrically on both sides of the metal disc and the hollow metal rod and connected to positive and negative potentials, respectively. The oppositely charged nanofibrous jet of PAN got deposited between the disc and the rod through which the cotton yarns were fed. During this process, the cotton yarn was coated with the nanofibers via the rotation of the disc, as schematically depicted in Fig. 2A. The morphology of the spun yarn was precisely engineered by controlling the nanofiber diameter and its layer thickness [23]. In another electrospinning design, two spinnerets were placed face to face in opposite direction between which a core filament was passed at a particular speed as shown in Fig. 4F. Continuous fibrous yarns obtained in this way were twisted and collected using the core filament and wound as spools [24]. The advantages of the electrospinning process using filament-shaped collector are that it yields nanoyarns that are well guided by a core filament and hence continuity of the yarns is well assured. Also, its mechanical strength is high compared to that of the yarns produced by other collector geometries due to the presence of the core monofilament. Nevertheless, here the nanoyarns are wrapped over a core filament and hence cannot be categorized as completely nanofibrous. Also, the adhesion strength of the electrospun nanofibers onto the core filament material is an important aspect that decides the quality of yarns for diverse applications.

3.4

Tube Collector

A rotary metal tube collector is another collector design studied for the development of nanofibrous yarns. In a sole study, the collector was a metal tube of dimensions (outer diameter 110 mm, length 200 mm) rotating at 1000–3000 rpm. A membrane

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Fig. 5 Graphical representation of the yarning assembly using (A) rotating tube collector and (B) triple tube array. (B2) Represents the morphology of the nanofibrous yarns produced by these methods, respectively. (A1, B1) Reprinted with permission from Refs. [25, 26]

was formed due to the deposition of nanofibers on the mouth of the rotating tube collector from two spinnerets connected to two equal and oppositely charged potentials, respectively. This membrane was converted into a hollow fibrous cone. Nanofibers from this cone were hooked using an insulator rod, resulting in a continuous fiber bundle that was wound to obtain yarns. Figure 5A shows the graphical representation of this setup. Twisting of nanofibrous yarns was imparted by the rotation of the tube collector. Yarn post-processing strategies such as stretching and twisting were also integrated with the electrospinning system as shown in Fig. 5A. The yarn uptake rate (0.01–10 m/min) and twist rate (10,000 turns/min) were optimized [25]. Figure 5A1 shows the SEM image of the nanofibrous yarn obtained after stretching and twisting. In another study, a triple tube array consisting of a pair of grounded rotating metal tubes and a rotating plastic tube placed in between them were used to obtain aligned electrospun nanofibers as yarns. For this, a single spinneret connected to a high potential was placed in front of the triple tube array. The two metallic tubes were placed face to face such that the electric vectors produced near the gap points to their edges, resulting in the pulling of nanofibers across the gap due to the transverse electrostatic force. Tube rotation in opposite direction resulted in aligned electrospun

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nanofibers getting twisted as cone shapes in the gaps between the rotating tubes. Insertion of a rotating plastic winding tube into the middle of the gap helped to obtain continuous twisted yarns onto it. Adjustments in the rotation speed of the metallic and plastic tubes aided in controlling the twist of the nanofiber yarns. Figure 5B depicts the setup used for this yarning process and 5B1 represents the morphology of the resultant nanofibrous yarns. Typically, a rotation speed tuned from 1,000 rpm to 3,000 rpm for the metallic tube, and between 1 rpm and 5 rpm for the winding tube yielded. PAN yarns with diameter of 10 μm and ultimate strength of 100–180 MPa for different twist angles [26].

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Cylindrical Static Collector

Static cylinder is another collector design studied for the development of nanofibrous yarns, which utilizes a two-spinneret assembly placed in opposite directions, connected to equal and opposite potentials. The nanofibrous jet was focused towards the center of a static grounded metallic cylinder placed in between the two spinnerets. A yarn take-up twister unit was integrated with this assembly to obtain continuous yarns. The nanofibrous yarn formation mechanism is hypothesized as contributed by the displacement of electrons on the surface of the cylinder by which half of the cylinder gets positive charge and the other half a negative charge. This results in the movement of the nanofibers from the charged spinneret to the part of the cylinder with opposite charge, and the formation of a symmetric triangle zone. The deposition of nanofibers onto a piece yarn placed at the convergence point of the triangle zone aided in taking up the fibrous yarn using a rotating plate with a winding setup, as schematically shown in Fig. 6A. Alterations in the dimensions and geometry of this triangle zone influenced the fiber diameter, tension and ultimately the yarn strength. Figure 6A1 depicts the morphology of the yarns thus developed. The polymers used in this process should adhere to the metallic cylinder for obtaining continuous yarns. In typical experiments using this geometry, continuous yarns of Nylon 66, polyacrylonitrile, PVA, PVP, etc. have been fabricated. Composite nanofibrous yarns, especially that of TiO2 loaded ones, were also developed using this method. Typically, in all these experiments, the optimized parameters are: distance between spinnerets: 18 cm, feed rate: 0.2 ml/h, applied voltage: 13 kV, cylinder dimensions: 6 cm diameter and 30 cm length. The yarn diameter as well as its continuity was found to be influenced by the take-up and twisting speed used [27–30]. One of the major limitations of this assembly is the process scalability owing to the use of only two spinnerets. Uniform alignment and twist of the yarns obtained here are majorly controlled by the “take up-twister unit,” while in other cases, the collector setup itself dictates the same.

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Fig. 6 (A) Graphical depiction of the yarning assembly using a cylindrical static metallic collector. (A1) Morphology of the yarns developed using this assembly as seen from the electron micrograph. (A1) Reprinted with permission from Ref. [28]

3.6

Glass Rod Collector

In this technique, a glass rod introduced between a spinneret and collector enabled nanofibers to be slowly drawn out in the form of yarns, by the continuous movement of the glass rod away from the spinning zone. During spinning, the collector and spinneret were maintained at negative and positive potentials respectively, as shown in Fig. 7A [31]. The SEM image of nanofibrous yarns produced using this technique is displayed in Fig. 7A1. In another study, instead of the glass rod, a grounded needle was used to guide electrospun yarns. Here, the grounded needle induced the bundling of nanofibers which were then pulled away from the spinning zone yielding yarns, which were wound as spools. The nanofibers thus produced were well aligned with slight twisting [32]. Mechanical properties of these yarns were improved by the postprocessing techniques such as stretching and annealing [33]. Although yarns were obtained in this method, continuity of the yarns is questionable.

3.7

Friction Double Cylinder

Frictional double cylinder collector setup is a scalable strategy for the fabrication of nanofibrous yarns with high alignment. In this method, a pair of friction rollers was arranged parallelly below a needleless electrospinning unit. The nanofibers were

Fig. 7 (A) Nanofibrous yarning setup using a glass rod collector and (A1) SEM image of a twisted and aligned nanofibrous yarn generated using this setup. (A1) Reprinted with permission from Ref. [32]

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Fig. 8 Schematic representation of electrospinning setup for nanofibrous yarn collection using friction rollers

focused towards the middle of the two friction rollers at a slit, which was on the same central axis as the needleless spinning unit. While the needleless spinning unit was connected to a positive potential, a negative potential was applied to a metal bar placed behind the slit between the two friction rollers. A pressure gradient generated by an airflow helped to focus the nanofibers towards the collector setup. This airflow augmented the nanofiber yield and yarn properties. Thus, a synergism was generated by the negative pressure gradient in the spinning field and the electrostatic attraction between the spinneret and collector. This aided in developing continuously twisted nanofibrous yarns that were wound into a spool. The optimized electrospinning parameters are: applied voltage (34 kV), flow rate (48 ml/h); working distance (40 cm), pressure gradient (0.2–0.8 MPa) and winding speed (40 rpm) [34]. Using this setup (Fig. 8), continuously twisted yarns of PAN, PVDF, and PU were prepared, demonstrating its suitability for electrospinning many polymer solutions into yarns.

3.8

Metal Frame

In this method, a rotating metal frame consisting of an insulating U-shaped frame within it was used to produce discontinuous nanofibrous yarns. During electrospinning, the inner non-conducting frame was held stationary, while the outer one was rotated. In the process, nanofibers got attracted to the inner frame

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Fig. 9 (A) Schematic representation of the electrospinning method using metal frame collector for nanofibrous yarn formation, (A1) SEM image of nanofibrous yarns produced using this setup. (A1) Reprinted with permission from Ref. [35]

and were stretched across a pair of parallel plastic tubes, resulting in an aligned fibrous array oriented perpendicular to the tubes, as depicted in Fig. 9A. These aligned fibers in the inner frame were twisted, yielding a uniaxially aligned discontinuous nanofibrous yarn as shown in Fig. 9A1 [35]. The major limitation of this method is the discontinuity in fiber formation, which limits its further applications in textile technologies.

3.9

Liquid Bath as Collector System

One of the most studied systems for fabricating electrospun yarns is the liquid-bath collector system. A liquid collector offers several advantages over a solid substrate, wherein the nanofibers deposited on a liquid substrate are easily maneuvered, making it advantageous for the bundling of nanofibers into yarns. In this method, which is depicted in Fig. 10A, the spinnerets were maintained at a positive potential, while the liquid bath placed at the bottom of the spinnerets was kept at a negative potential. Nanofibers deposited on the surface of the bath were collected using a glass rod, and this fibrous bundle was wound into a spool. The yield of the nanofibrous yarn was mainly influenced by the applied electric field as well as the bundling and drawing processes. The bundling process includes a combination of wet, wet-dry, and dry processes. The transformation from nanofibers into yarn filaments mainly occurs during the wet process. The wet-dry and dry processes decide the crystallinity, alignment degree of nanofibers as well as its diameter. These yarns can be further post-processed by heating and plying [33]. Here, multiple needles can be used to obtain nanofibrous yarns with better throughput. This method can also be adopted for the fabrication of composite yarns. One such study reported the fabrication of polymer/CNT composite nanofibrous yarns using CNT liquid bath as the collector, as shown in Fig. 10B. Herein, the polymer solution

Fig. 10 (A–D) Schematic representation of different liquid bath-based collector systems used for fabricating nanofibrous yarns. (A–C) Reprinted with permission from Refs. [35–38]

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was loaded in a syringe and the collector bath containing CNT dispersion was placed below the spinneret at a working distance of typically 40 mm. A grounded aluminium foil was placed at the bottom of the liquid bath of height 5 mm. A glass rod introduced into the liquid bath helped to extract the yarns out of the bath, which was wound on an uptake roller with a velocity of 0.6 m/s. The composite yarns thus obtained showed a uniform distribution of CNT throughout the yarn [36]. Dynamic liquid coagulation baths have also been used as a collector setup in fabricating nanofibrous yarns. Here, the dynamic movement of the liquid is utilized for the bundling of nanofibers into yarns. This is designed such that a vortex created from the outflow of water from a basin (with a hole of diameter 5 mm) aids in bundling the nanofibers into continuous yarns, with a collection speed of >60 m/ min, as depicted in Fig. 10C. Herein, the optimized parameters for obtaining PVDF yarns include: applied voltage 12 kV, tip-target distance 12 cm and a constant water level height, which was maintained by recirculation. A metallic wire inserted into the basin helped to remove any residual charge on the water surface. The resultant yarn was collected onto a yarn winding setup placed below the reservoir. Liquid properties including viscosity, surface tension, as well as the fluid interface and hydrodynamic interactions control the nanofibrous yarn morphology. Besides these, other parameters such as nature of the polymer, applied voltage, uptake speed, velocity of the water in the vortex, etc. influence the yarn formation, yield and its properties [37]. In another research, the same principle of vortex water bath was used as the collector setup for nanofibers, with the variation in the yarn drawing and winding setup. These were placed on the top of the dynamic water bath and controlled by a rotating cylinder, unlike the bottom configuration discussed earlier [38]. Figure 10D gives a schematic of this setup. This method is limited by the water bath used as the collecting medium, which can cause short circuit when operated at high potentials. Also, only hydrophobic polymers can be processed as yarns using this setup. Likewise, this system cannot be utilized for loading water-soluble drugs or molecules within the yarn.

3.10

Funnel-Shaped Collector

By utilizing two pairs of oppositely charged spinnerets and a funnel-shaped rotating metal collector, continuously twisted nanofiber yarns have been developed. Nanofibers ejected from the oppositely charged spinnerets were deposited on the mouth of the rotating funnel-shaped collector, thus forming a nanofiber web. This web was drawn out using an insulating rod resulting in a fiber bundle and wound to form continuous yarns. The rotation speed of the collector dictated the fiber diameter and twist angle of the nanoyarns. The graphical representation of this assembly is shown in Fig. 11A. In one such study, PAN nanofiber yarns were fabricated under the following optimized conditions of electrospinning: applied voltage 20 kV, flow rate 6.4 mL/h and tip-target distance 18.5 cm. An improvement in the yarn mechanical properties

Fig. 11 Graphical depiction of (A) funnel-shaped collector assembly used for fabricating nanoyarns. A scaled-up assembly comprised of (B) a pressure gradient and (C) pressure infusion pump

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was observed with increase in twist angle. Typically, an angle of 65° yielded yarns with diameters of 70–216 μm and a tensile strength of 50.71 MPa, with an elongation at break of 43.56% [39]. Twisted assemblies of polyacrylonitrile (PAN), polyvinylidene fluoride (PVDF), poly(L-lactide) (PLLA) and polycaprolactone (PCL) nanofibers have also been fabricated likewise. Silk fibroin (SF) and poly(L-lactide-co-caprolactone) blend and PCL/silk/carbon dot composite yarns are also reported by the same method [40–44]. Funnel collectors have been demonstrated to yield continuous yarns in a scalable manner. In one such study, a multiple conjugate electrospinning apparatus consisting of several infusion tubes connected to an array of spinnerets of opposite polarities was used. These were arranged symmetrically on both sides of an unearthed funnel collector yielding continuous PAN yarns with high throughput. Here, the presence of the rotating funnel-shaped collector altered the electric field distribution, resulting in an electrostatic induction which bestows the funnel edge with charges that are opposite to the nearby charged spinnerets. This enabled the charged jets to get attracted towards the funnel edges, yielding a hollow inverted cone-shaped nanofiber web. An oriented twisted fiber bundle was formed by drawing the fibers from the cone apex. Typically, continuous PAN nanofiber yarns were produced at optimized parameters: Applied voltage (18 kV), distance between positive and negative spinnerets (17.5 cm), overall flow rate (3.2 ml/h). The optimized nanofibrous yarns with twist angle of 41.8° have relatively good values of tensile strength (55.70 MPa) and elongation at break (41.31%). This method can also be extended to fabricate multifunctional composite yarns [45]. In another study, a negatively charged funnel-shaped metal collector (Outer diameter: 110 mm, Inner diameter: 40 mm, Length: 85 mm) coupled with dual spinnerets connected to positive potential was used to fabricate nanoyarns. In this study, polysulfone amide yarns were prepared under the optimized electrospinning parameters of flow rate (0.5 ml/h) and voltage applied (30 kV) [46]. Another group utilized a vacuum pump with a power of 360 L/min to produce a negative pressure towards the center of the funnel collector, thus driving the nanofibrous jet towards the collector, thereby increasing the yarn yield. The electrospun jet entered the funnel instantly after it got ejected from the spinnerets and formed aligned fibers along the air flowing direction. PAN/LiCl composite nanoyarns were fabricated by this method [47]. In another scaled-up setup using funnel-shaped collector, an innovative spinneret consisting of an air chamber was used to increase the throughput of nanofibers as seen in Fig. 11C. This system also utilized two pairs of oppositely charged spinneret units. The optimized parameters are: applied voltage (34 kV), air flow rate (1,200 mL/min) and overall solution flow rate (32 mL/h). Twist angle is a critical parameter that influences the yarn tensile strength and elongation at break, with an increased trend noticed for increasing twist angle. At an optimized twist angle of 49.7°, these nanoyarns exhibited a strength and elongation of 0.592 cN/dtex and 65.7%, respectively [48]. Amongst all the geometries adopted for developing yarns based on DC electrospinning, this technique stands out owing to its potential to generate

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continuous, uniaxially aligned, nanofibrous yarns with relatively good mechanical strength and scalability.

3.11

Hemispherical-Shaped Collector

The collector design adopted here is a rotating metallic hemisphere coupled with dual spinnerets connected to equal and opposite high voltages. An array of equidistant point electrodes (or conducting tines) was placed at the periphery of the hemispherical rotating collector, having a variable rotation speed from 100 rpm to 1,100 rpm. Dual spinnerets placed at 45° with respect to the center of the hemisphere and charged oppositely were utilized here as depicted in Fig. 12A. During electrospinning a fluffy mass was formed at the center of the hemisphere due to the convergence of the electric field at the tines. This fluffy mass was hooked with an insulating rod and the yarns thus formed were collected as a spool. Electron micrographs of the resultant yarns are displayed in Fig. 12A1. The optimized parameters include: winding speed (0.6 m/s), collector rotation speed (~900 rpm). Mechanically strong, aligned, and twisted yarns were produced in a continuous manner with an overall efficiency of ~94% [7, 8].

Fig. 12 (A) Schematic depiction of the hemispherical collector used for fabricating continuous nanofibrous yarns. (A1) SEM images of the resultant yarns showing the twist and alignment of nanofibers in the yarn. (A1) Reprinted with permission from Ref. [8]

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4 Yarn Fabrication by Twisting Electrospun Membrane Twisting of electrospun membranes is another simple strategy to produce discontinuous nanofibrous yarns, wherein the nanofibers are randomly oriented. In one such study, PVDF nanofibrous membrane was transformed as a yarn after simple electrospinning at optimized parameters [flow rate: 0.5 mL/h, working distance: 15 cm and voltage: 15 kV]. This membrane was then converted into a nanofibrous yarn using the post-processing techniques of drawing and then twisting, as evident from Fig. 13A. This rotation under tension allowed the nanofibers to be pulled from the web and get aligned as can be seen from the electron micrograph of the yarns in Fig. 13A1. Twist parameter is an important factor that dictates fiber alignment in the yarns and its mechanical strength. Process parameters including rotating and winding speeds and the diameter of the roller control the twist parameter [49]. Compared to non-woven PVDF nanofibers without twisting, the mechanical properties of the yarns were dramatically improved with increased width of the non-woven nanofiber membrane and the number of twists [50].

5 Collector-Less Yarning Process Collector-less approach has also been adopted for the fabrication of electrospun yarns. Two spinnerets connected to equal and opposite potentials, respectively, and placed face to face were used for the production of nanofibers (Fig. 14). These nanofibers travelling through air helped to form continuous yarns in this substrateless approach. In a specific study, PAN nanofibrous yarns were obtained using the above process. The nanofibers charged with a voltage of the same value but opposite polarity got attracted and discharged at a midpoint, from where a continuous strand of nanofibers was taken up using a guide wire. However, this process

Fig. 13 (A) Schematic representation of electrospun membranes being transformed into nanofibrous yarns and (A1) SEM image of the yarn. Reprinted with permission from Ref. [43]

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Fig. 14 (A) Schematic depiction of the collector-less nanofibrous yarning assembly. (A1) SEM image of the corresponding yarn obtained. (A1) Reprinted with permission from Ref. [52]

yielded a bulk strand of nanofibrous yarn of low orientation, strength, and uniformity [51]. Nevertheless, another group reported the fabrication of continuous, wellaligned nanofibrous yarns of PVA and PVP using the same method. Herein, the two spinnerets placed face to face enabled the nanofibers to get bundled via electrostatic attraction between fibers, yielding a neutral yarn which could be wound continuously as a spool [52].

6 AC Electrospinning Direct current (DC) electrospinning in its numerous forms, which include needle and needleless variants are widely investigated as detailed thus far. However, fabrication of yarns using DC electrospinning has limitations with its industrial scalability. Another disadvantage is that all its variant setups discussed above require different geometries of collector. This makes the process difficult to integrate it with other processing methods, due to the presence of high electric field within the spinning zone. Contrary to the detailed investigations on DC electrospinning, there are very few reports on electrospinning using AC fields for nanofibrous yarn fabrication [53– 56]. In AC electrospinning, needleless spinning-electrodes are utilized for generating a highly productive smoke-like aerogel. This nanofibrous aerogel resembles a thin plume of smoke which emanates from the spinning electrode, without the need for a separate collector. The entanglement of nanofibers within this plume aids in making a compact bundle, thus helping its manipulation into either continuous wound ligaments or deposition on a flat surface or twisting as yarns. This collector-less process is efficient due to the generation of a continually charge-changing virtual counter-electrode which is composed of nanofibers (Fig. 15).

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Fig. 15 (A) Schematic representation of AC electrospinning assembly. The components include: (1) electrode, (2) nanofibrous plume, (3) core yarn, (4, 5) twirling devices, (6) core yarn feeding unit and (7) final winding setup. (B) Optical image showing the nanofibrous plume emitted from the spinning electrode. (A1) is the SEM image of the nanofibrous yarn produced by this method. (A, B, and A1) Reprinted with permission from Ref. [58]

In a typical study on AC needleless electrospinning, a high-voltage transformer [Conversion ratio: 36,000/230 V], a residual-current device, infusion pump, and a metal rod electrode were utilized. The output voltage (0–250 V) was controlled using a variable transformer for a 230 V AC input, with a maximum output current of 4 A. The metal rod which is used as the spinning electrode was fed continuously with a polymeric solution. At a high AC field, nanofibers were formed between the top of the spinning electrode and a virtual counter-electrode (composed of emitted nanofibers). At each half-wave of the AC field, the virtual counter-electrode is recreated periodically, thus yielding a continuous nanofibrous plume. PVB and PAN nanofibers were produced using this method at an operating voltage of 30 kV [57]. In another study, the group investigated a collector-less and needleless AC electrospinning approach to fabricate composite yarns whereby nanofibers were wound around a classic thread. Integrating a twirling device with the setup allows for additional twisting of the nanofibers around the core yarn. Typically, the core yarn was fed horizontally with respect to the spinning space [150 mm above the spinning electrode]. The nanofibrous plume generated from the spinning electrode was perpendicular to the yarn core axis. During rotation of the twirling devices, the core yarn gets tightly enveloped by nanofibers. The key parameters that determined the quality of the nanofibrous yarns include linear density and diameter of the core yarn, yarn tension and winding speed, angular speed of both twirling devices and yarn mechanical properties. The resulting yarn was composed of nanofibers up to 80% of its weight, with a relatively high throughput of 10 m/min, which could be increased to 60 m/min [58].

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7 Recent Advances in the Applications of Nanofibrous Yarns in Biomedicine Electrospun nanofibrous membranes are already established in the field of biomedicine due to its porosity, surface area to volume ratio and its resemblance to the structural features of native ECM. This in turn is known to promote cellular adhesion, spreading, migration, proliferation, and differentiation [59]. The recent translation of 2D electrospun membranes to 1D yarns has also gained wide acclaim in biomaterials research on account of its additional characteristics such as improved mechanical strength and processability. Processing of 1D nanoyarns into 3D textiles by the textile engineering techniques of weaving, knitting, braiding, etc. has opened up new avenues in fabricating fibrous scaffolds as well as implants for diverse applications in tissue engineering, drug delivery and biosensors. Such scaffolds can resemble the hierarchical and anisotropic structures as well as the strainstiffening properties of native tissues. Owing to these features, such 1D and 3D fibrous constructs find immense potential for applications as surgical sutures and acellular or cellular scaffolds that can promote the repair and regeneration of nerve, blood vessels, bone, cartilage, etc. In addition, these yarns can also be loaded with various drugs/biologics, offering sustained release platforms. Nanofibrous yarns have been explored for use as sutures in several applications. In one such study, multi-plied yarns have been developed into a thread to increase its mechanical properties and utilized as sutures in tendon repair [60]. These sutures showed negligible immunogenicity and significantly higher neovascularization. Varieties of bioactive materials/drugs/biologics including curcumin [24], vascular endothelial growth factor (VEGF) [61], aceclofenac [10], insulin [10], cefazolin [62], and silver nanoparticles [63] were incorporated into nanoyarns to impart a predetermined biological functionality. Drug loaded nanoyarns exhibited sustained release characteristics, without comprising its mechanical behavior. Nanoyarns have also been investigated for its utility as nerve guidance conduits. In a recent study, PLGA-based nanoyarns coated on PLLA microfiber yarns were studied as nerve guidance conduit due to its resemblance with the fascicle structures of native peripheral nerves, and longitudinal alignment of axons in fascicles. The resemblance of a bundle of electrospun nanoyarns with natural nerve trunk has been capitalized in this study. These nanoyarns demonstrated a notable promotion in cellular proliferation as well as phenotypic maintenance [64]. In another study, electrospun PCL nanoyarns coated with polypyrrole (PPy) were developed and the growth and proliferation of Schwann cells (SCs) on them were studied [65]. Nanoyarns have also been maneuvered into 3D constructs by utilizing weaving technology for its application as nerve conduits. Herein, PLLA yarns were co-woven with polypyrrole coated PLLA yarns and ultrathin copper/platinum wires and were developed as conduits that can be electrically stimulable. Enhanced neurite outgrowth and length of rat dorsal root ganglion sensory neurons were observed on these conduits in an in vitro study [66].

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Nanotextiles made by interweaving nanoyarns have also been explored for their potential in the engineering of tendons. Depending upon the yarn weaving density, properties such as porosity, pore size, mechanical and biological properties were modulated [67, 68]. In one such research, thymosin beta-4 (Tβ4) loaded nanoyarns exhibited promotion of the tenogenesis of stem cells [69]. Mechanical properties of the yarns were enhanced by plying together or braiding multiple yarns to get a more stable and suitable platform for tendon regeneration [70]. This enabled significant tenogenic differentiation of stem cells in comparison with aligned electrospun membranes [70]. In another study, basic fibroblast growth factor (bFGF) was incorporated into PCL-micro/collagen-nano hybrid yarns and braided into textile patterns for Achilles’ tendon reconstruction [71]. Nanofibrous yarns also represent potential materials in the development of bone scaffolds. Chopped PCL copolymer nanoyarn encapsulated type I collagen hydrogels have been investigated for bone tissue engineering [72]. This injectable hydrogel promoted osteogenic differentiation of stem cells and was found to be suitable as a bone scaffold. Silk Fibroin (SF)/PLLA nanoyarn-based 3D constructs significantly promoted the regeneration of new bone tissues in an in vivo study [73]. Likewise, hydroxyapatite (HA) particle decorated SF/PLCL nanofibrous yarnbased scaffolds exhibited improved cellular response [74]. A post-soaking method was employed to deposit HA nanoparticles on SF/PLLA nanoyarn-based 3D scaffold, which also demonstrated notable increase in the osteogenic differentiation of stem cells [75]. Nanoyarns have also been used as mechanical reinforcements in biodegradable polymeric scaffolds for bone tissue engineering applications. Various drugs/biologics have also been loaded into such matrices for enhanced biological functionality. Gelatin/hydroxyapatite nanocomposite scaffolds reinforced with PLLA or PLLA/gelatin nanoyarns have proved to be an excellent biodegradable matrix that promoted osseointegration in critical sized mandibular defects in rabbit and pig models [76–80]. The mechanical characteristics of the nanoyarn reinforced matrix were such that it also aided in dental implant placement for prosthetic rehabilitation. This nanocomposite matrix also promoted the sustained release of antibiotics and biologics/growth factors, which in turn was found to be effective in osteomyelitis treatment as well as in promoting osseointegration, respectively, in animal studies [79, 81, 82]. Nanotextiles made from electrospun yarns in the form of vascular grafts and vascular patches have also found potential use in the repair and regeneration of cardiovascular defects. PLLA nanoyarns woven into flexible conduits tested as small diameter vascular grafts were found to be robust, suturable, kink proof, and non-thrombogenic when tested in vivo in rabbit model [9]. These small diameter grafts demonstrated surgical feasibility, safety as well as increased transmural endothelial ingrowth when implanted in pig carotid artery. It also exhibited 100% patency at 2 and 4 weeks, with minimal changes in lumen size, flow velocities and inflammation response, as compared to ePTFE [83]. The same nanofibrous yarns of PLLA and PCL/collagen were also developed using a different weaving strategy as nanotextile vascular patch. These nanotextiles displayed good mechanical

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characteristics along longitudinal and transverse directions and exhibited good endothelial cell proliferation in its preliminary studies, suggesting its potential used in vascular patch material [7]. Drug loaded nanoyarns have found immense value in cancer therapy as implantable sustained release depot. In a specific research, paclitaxel was loaded within polydioxanone nanofibrous yarns and processed as 2D nanotextiles for implantation in the peritoneal cavity of mice. Modulation in the drug release kinetics was established through alterations in the packing density of the nanofibrous yarns in the nanotextile. Sustained drug release for a duration of 3 months was established in vivo [11]. This biodegradable nanotextile further demonstrated enhanced antitumor efficacy and safety through metronomic intraperitoneal chemotherapy in a metastatic ovarian cancer model. No signs of systemic or organ toxicity were observed for mice implanted with the nanotextile, indicating its utility as a drug delivery depot for various applications [84].

8 Conclusion Last two decades have seen remarkable progress in the processing of nanofibers as nanoyarns and its implementation in various applications. Significant strides have been made in developing scalable strategies for yarn fabrication by electrospinning. Various types of collector designs were studied to produce continuous, uniaxially aligned yarns with good mechanical strength and throughput. However, the nanofibrous yarns produced till date have not attained mechanical characteristics superior to microyarns. Despite this, biomedical textiles such as 2D woven, knitted and braided patterns, and also 3D textiles have been constructed using these nanoyarns. Nevertheless, to optimize the morphology, structure, biological properties and scalability of both nanoyarns and nanotextiles still remains challenging. Future efforts should be devoted to the large-scale manufacturing of electrospun nanoyarns with good reproducibility, high yield and processability through facile methodologies. Acknowledgments The authors acknowledge the financial support from Department of Science and Technology (DST), Government of India, through the “Thematic Projects in Frontiers of Nanoscience & Technology” (SR/NM/TP-15/2016G) for the development of nanoyarn technology.

References 1. Suja PS, Reshmi CR, Sagitha P, Sujith A (2017) Electrospun nanofibrous membranes for water purification. Polym Rev 57:467–504 2. Sagitha P, Reshmi CR, Sundaran SP, Sujith A (2018) Recent advances in post-modification strategies of polymeric electrospun membranes. Eur Polym J 105:227–249

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Adv Polym Sci (2023) 291: 139–166 https://doi.org/10.1007/12_2022_135 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 30 September 2022

Fabrication of Textile-Based Scaffolds Using Electrospun Nanofibers for Biomedical Applications K. Ashok, M. Babu, G. Kavitha, R. Jeyanthi, R. Ladchumananandasivam, O. da Silva, and E. Manikandan

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 ELS NFs BMA . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Wound Healing (WH) Properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Musculoskeletal Complications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Cardiovascular Diseases (CVD) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 Nephrology . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6 Drug Delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 Bone Regeneration (BR) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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K. Ashok and M. Babu Department of Microbiology and Biotechnology, Faculty of Arts and Science, Bharath Institute of Higher Education and Research (BIHER), Chennai, Tamil Nadu, India G. Kavitha P.G. and Research Department of Physics, A. M. Jain College, (Affiliated: University of Madras), Chennai, India R. Jeyanthi Department of Zoology, Presidency College, Chennai, Tamil Nadu, India R. Ladchumananandasivam (*) Post-graduate Programme in Mechanical Engineering – PPGEM, Federal University of Rio Grande do Norte (UFRN), Natal, Brazil O. da Silva Department of Textile Engineering – DET, Federal University of Rio Grande do Norte (UFRN), Natal, Brazil E. Manikandan (*) Department of Physics, Thiruvalluvar University College of Arts and Science (TUCAS) Campus, Thiruvalluvar University, Thennangur, Tamil Nadu, India UNESCO-UNISA Africa Chair in Nanosciences/Nanotechnology Laboratories (U2AC2N), College of Graduate Studies, University of South Africa (UNISA), Pretoria, South Africa

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8 Gynaecology . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 9 Cancer Biology (CB) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10 Snakebite (SKB) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11 Neurology . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12 Diabetes (DT) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 13 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract Even with its simplicity, ease of use, and wide variety of applications, the ELS technique has gained in popularity. The characteristics of EFs fibres (FBs) can be affected by changing process variables or polymeric solution (PLs) conditions. Many of the elements that impact ELS, however, are linked. An ideal ELS method keeps these parameters constant and consistently creates NFs with consistent physicochemical properties. The PLs might be aqueous, polymeric melt, or emulsion, resulting in NF production in various forms. Inverting the polarity and altering the collector design can also change the NF characteristics. Blending, surface functionalization, and emulsion generation are all methods for incorporating the active moiety into polymeric FBs. The multilayer polymer covering permits the incorporated active moiety to be released constantly, and the NFs may be modified to carry a variety of medications. Polymer (PLY)-derived EFs and NFs are utilized for DD, ANC, WH, BS, SiRNA delivery, stem cell treatment, and growth factors. This review compiles papers concerning the utilization of EFs and NFs in biomedical applications (BMA). Keywords Drug delivery · Electrospinning process · Female reproductive health · Nano FBs · Wound healing and cancer biology

1 Introduction Electrospinning/electrostatic spinning (ELS) is a simple and adaptable production technology that produces polymeric solutions (PLs) or melts into continuous nanoNFs and nonwoven (NWs)-based textiles (TS) using strong electrical currents (Fig. 1) [1–25]. Electrospun (EFs) NFs and associated NW textiles provide structural benefits such as mechanical resistance (MR), nanoscale (NS) interstitial space (ISS), and adjustable porosity (AP). Because of their features, EFs TS can be used as scaffolds (SS) for a variety of applications. EFs materials increasingly attracted interest in recent years, not simply in typical textile industries like clothing such as poly(allylamine hydrochloride) (PAH) [26], cotton (CO) [27], TPU/TBAC [28], polyacrylonitrile NF yarns [29], hyaluronic acid (HA) NFs [30], chitosan (CHS) [31], CHS/poly(ethylene oxide) [19], HA-based silk

Fabrication of Textile-Based Scaffolds Using Electrospun Nanofibers. . .

Fig. 1 ELS types and parameters

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fibroin (SF)/zinc oxide (ZnO) [17], ZnO [14], PVA and CHS mats [18], CHS/poly (vinyl alcohol)/graphene oxide (GO) [32], CHS-polyethylene oxide NF mats [33], thermoplastic polyurethane (TPU) [34], gelatin/PCL composite [35], poly(l-lactic acid) (PLLA) [36], carbon nanotube (CNT)-filled NF yarns [37], polyacrylonitrile NFs [38], graphene nanoribbons (GNR) [39], poly(L-lactide) (PLLA) yarn [40], aligning [41], yarn/hydrogel (Hs) composite [42], polyvinylpyrrolidone mat-FBs [43], cellulose [44], superhydrophobic (SH) NFs [45], zein and zein/poly-L-lactide NFs [46], poly(vinylidene fluoride-co-hexafluoropropylene) NFs [47], PHB [48], polyurethane NFs [49], ZnO/Nylon 6 NFs mats [50], deoxybenzoin [51], CO [52], zinc oxide nanoparticles (ZnO NPs) [53], polycarboxylic acids [54], nylon-6 Spidernet NFs [55], and lignin/poly (vinyl alcohol) [56] but also in cutting-edge sectors including fundamental and applied biomedical research (BMR) [57] for wound healing (WH), medication delivery technique development, biosensors (BS), and tissue engineering (TE) among other things. Furthermore, recent Coronavirus disease outbreak has as a consequence of this, need for disposable NW face masks (FM), demonstrating the use of EFs textiles like FM filters in filtering NS pollutants and protecting against airborne viruses such as SARS-CoV-2 and COVID-19 [57]. COVID-19 FM NFs that are appropriate include nylon 6-polyacrylonitrile [58], poly(l-lactide)/zein NFs [59], CNT [60], polyethylene terephthalate (PET) [61], carboxymethyl cellulose (CMC) [62], cellulose (CE) [63], activated carbon (AC) and carbon nanofiber (CNFs) [64], polybenzimidazole NFs [65], CHS [66], CHS NPs [67], keratin (K) [68], gelatin/β-cyclodextrin NFs [69], silk [70], polycaprolactone (PCL) [71], poly(ethylene oxide) NFs [72], carboxymethyl CE NFs [73], chitin (CHT) [74], starch [75], pullulan [75], bioactive glass (BS) [76], latex [77], polyimide (PI) and polyamide (PA) [78], and polyhydroxyalkanoates [79].EFs fabrics are made up of both natural and synthetic PLYs such as Col [80, 81], polysaccharide [82], gelatin/PCL [83], CHS-collagen (Col) hydrogels (Hs) [84], gelatin- hydroxyapatite [85], silk [86], poly (ε-caprolactone) [87], polyurethane [88], CHS [89], polystyrene (PS) [90], poly-Llactide NFs [91], co-polyether sulfone [92], Hs [93], silk, keratin, elastin, and resilin proteins [94], glycosaminoglycan [95], BMP-silk [96], tussah silk [97], HAHs [98], polylactic acid/polyglycolic acid [99], MBG/PEGylated poly(glycerol sebacate) [100], PLGA/TiO2 CNT [101], PLGA [102], PLGA/HA NFs [103], PLGA [104], Fe3O4NPs [105], hydroxyapatite/PLGA [106], PCL/HA [107], ZnO CNT [108], acrylic FBs [109], BSA [110], Nylon-4,6 (PA 4,6) [111], silk fibroin (SF)//CHS/ magnetite [112], PAN [113], carboxymethyl cellulose-pullulan Hs [114], carboxymethyl CHT Hs [115], SF-BS [116], reinforced Hs/silk [117], CdS nanocrystallites on SF [118], polybenzimidazole NF [119], styrene–butadiene–styrene [120], poly(ethylene terephthalate) [121], and polyaniline [122] are promising materials for BMA (Fig. 2).

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Fig. 2 BMA of EFsNFs [123–165]

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2 ELS NFs BMA 2.1

Wound Healing (WH) Properties

WH is a treatment that aids in the repair and healing of wounded tissue [166– 173]. The four continuous and partially overlapping phases of this WH process are as shown in Fig. 3 [174]. Beeswax, milk, soil, herbs, and animal fats are just a few of the ingredients used to cure wounds in Sumerian and Egyptian times. In recent years, WH techniques have changed significantly. Over time, WH products have evolved from topical ointments and lotions to standard CO and wool gauze dressings. Although most of these products aid in the healing of acute wounds (AWs), they will not assist chronic or complicated wounds (CWs). As a result, gauze and CO treatments have largely given way to a new generation of dressings. The main goal recently has been to concentrate on preventing, appearing, or eradicating infections, as well as accelerating the WH process by renewing the skin's structure and function [175–179]. EFsNFs mats are one of the possible choices, since they offer a wide variety of prospective wound dressing (WD) applications. ELS are a low-cost, scalable, adaptable, and extremely easy approach for producing NFs from a variety of synthetic and biological components [129, 180–184]. MO penetration can be reduced by changing the porous size of the EFs mats, but oxygenation may still flow freely through WD and reach the injured area. Interestingly, water vapour (WV) transfer can be changed to provide the right amount of moisture for WH. Because of their enormous surface area, NFs are good for medication loading and dispersion [185]. The NFs matrix can contain adsorbent drugs, natural substances, or bioactive chemicals [186]. The selfstanding EFsNFs dressings make wound care a breeze [187–189]. Natural biomaterials (NBs) have lately attracted attention due to their remarkable biocompatibility and regenerating potential [21, 190–194]. A single biomaterial, on the other hand, has drawbacks such as low mechanical characteristics and a single function. As a result, innovative medical biological dressings might include drug loading, altering NBs, and merging with other materials.

Fig. 3 Process of WH

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ELS's NFs have a high specific surface area, porosity, liquid absorption, and semi-permeability [195]. It might have the same shape and molecular qualities as the organic ECM, that aids cell adhesion, migration, and proliferation. Because of their strong biodegradability and biocompatibility, as well as their mechanical properties and physicochemical parameters, NBs are often used in the area of skin implants and WDs [196, 197]. Additionally, the bioactivity and clinical effects of the ELS dressing are yet inadequate. To help with this problem, bioactive ingredients like growth factors (GFs), vitamins, antimicrobial agents (AMAs), and other similar compounds can be added to the EFsNFs, and the gradual release of active substances in the WD can help not only limit infection but also promote WH and tissue regeneration (TR). In the future, the WD field's growth trajectory will most likely be EFs composite NFs with bioactive chemicals.

3 Musculoskeletal Complications Tendon and ligament (T/L) tissues have similar compositions, structures, and mechanics. Both tissues have a uniaxial aligned (UNA) ECM (mainly Col 1A) and are heavily loaded in one direction, resulting in extremely anisotropic mechanical characteristics [198]. Due to its high ECM density, Col organization, and low vascularity, the T/L has a limited potential to regenerate [199–202]. Because of its scar-mediated healing response and incapacity to regenerate, T/L has been investigated for TE techniques to replace damaged or diseased tissue [203–207]. T/L Col FBs are tightly packed and stacked in parallel arrays. UNA NFs mats are therefore commonly utilized for T/L regeneration. UNA NFs can be captured using a rotating collector or extra parallel electrodes. On uniaxial NFs, both adipose-derived mesenchymal stem cells (MSCs) and induced pluripotent stem cells show tenogenic differentiation. T/L tissue, on the other hand, has a significant degree of anisotropy. 3D SS with braided, woven, or knitted yarn networks is favoured over 2D mats. By incorporating Hs into the NFs matrix, biomolecule and cell encapsulation can be enhanced. Prior to T/L regeneration, unidirectional PCL NFs were coated with CHS/HA Hs [207].

4 Cardiovascular Diseases (CVD) Rheumatic heart disease (RHD), BV diseases, atherosclerosis (hardened arteries), peripheral vascular disease (PVD), myocardial infarction (MI), and coronary artery disease (CAD) are all examples of CVD difficulties (CAD) [208]. Nicorandil (NCD) is a drug that has agonistic effects for both the ATP-sensitive potassium channel and the polyatomic ion nitrate channel. It is commonly used to treat angina (AN) or angina pectoris [209].

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For its poor absorption and late beginning of action, as well as significant side effects like increased turnover rate and mucosal ulcers, NCD's usage as an anti-AN medication has been restricted (MUs). To overcome these shortcomings, NCD was EFs combined with polymeric NFs containing riboflavin (Vitamin B2), HA, and PVA to deliver a sublingual dosage for the treatment of angina pectoris. Riboflavin integration in the NFs SS was thought to cure mucosal ulceration (MU), whereas HA was thought to enable rapid inflammation healing in injured tissue by lowering the amount of pro-inflammatory cytokines (CKs). This nano-sized drug-loaded fibre mat, on the other hand, was able to sustain regulated release of NCD over an extended period of time. Pharmacokinetic (PK) studies revealed that the newly developed formulation kept a therapeutic level four times longer and had a fourfold longer biological half-life (t1/2) than commercially available NCD. Furthermore, no MUs were found after histological examination of the administration site for the indicated formulation [210]. Carvedilol (CAR) is an antihypertensive medication that inhibits both alpha and beta adrenergic receptors by binding to them. It is used for the treatment of congestive heart failure (CHF) [212]. According to Potrc et al. (2015), EFs PCL NF SS were explored as drug carriers (DCs) for weakly water-soluble (WS) CARs in the oral cavity. The CAR's crystallinity reduced after absorption into the PCL NFs, and the average size of a drug-laden PCL NFs is proportional to the amount of loaded drug. The drug was molecularly entangled in the PCL NFs and, to a lesser extent, in the scattered nanocrystal formation. CAR was shown to be released from PCL EFsNFs in as little as four hours, suggesting a considerable improvement in the dissolving rate of the weakly WS drug [211]. As a result, ELS is a groundbreaking nanotechnology (NT)-based technique for increasing the rate at which waterinsoluble medications dissolve.

5 Nephrology ELS is a flexible, long-lasting, and expense NF production technique. This method is used to create NF mesh with silicon (Si) and aluminium (Al) incorporated. The Si/Al ratio [213] was used to compute the creatinine (Cr) adsorption level. Despite the fact that their technology is only a prototype, it looks to be a feasible dialysis solution. A novel NF mesh has also been developed by researchers to remove toxins from the bloodstream (BS) [214]. Because of their inherent ability to eliminate particles, the kidneys (KIs) are great targets for nanoparticles (NPs) (12 nm). Renal excretion of particles smaller than two nanometers is caused by the KIs function, with decreasing clearance at six nanometers and no renal excretion of particles bigger than eleven nanometers [213, 215]. When NT, TE, and material sciences (MS) are combined, they provide an enthralling BMA. The exponential growth of these new sectors has considerably enhanced human life quality. A thorough examination of the characteristics of several biomaterials is necessary for the development of a new substrate for KI regeneration. As a result, this chapter focuses on the use of advanced

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nanomaterials (CNTs and NFs) in the therapy and repair of KI failure, as well as nanomaterial-based adsorbents and membranes in wearable blood purification devices and synthetic KIs [216].

6 Drug Delivery In the formation of EFs NF, PLY characteristics, solution viscosity and conductivity, solvent type, applied voltage, tip-to-collector distance, and relative humidity all play a part [217–219]. For EFsNFs to have good characteristics, these criteria must be addressed. In the realm of DD, several ELS methods, such as direct or coaxial ELS, have been successfully used to construct NFs with a variety of drug release (DR) characteristics, including immediate, biphasic, postponed, or sustained release. The impact of different factors and PLYs on the production of NFs using the EFs approach [219] was addressed in this chapter. It also highlighted the uses of EFsNFs in DD, particularly in oral dose forms with various release mechanisms, as illustrated in Figs. 4 and 5.

Fig. 4 Steps involved in the DR

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Fig. 5 Applications of EFsNFs used in DD

7 Bone Regeneration (BR) Cells and signalling molecules, in addition to SS, are significant components in regenerative medicine (RM). The same may be stated for BR, which can benefit from their use. Synthetic calcium phosphate (TCP, HA) bone replacements are the most often employed SS in the treatment of bone diseases. Despite having a high degree of osteoconductivity for tiny bone defects, the extent of the defect that can be repaired is limited [220]. As a result, therapies that combine the usage of SS with the other cells and signalling molecules are being investigated in order to improve the efficacy of bone regeneration therapy. Because of its potential applicability in SS, IPS cells have generated a lot of research in recent years [221, 222]. Yamanaka factors (YFs) are injected into somatic cells, such as skin cells, to create multipotent cells capable of converting into a range of body cells. Because of their potential uses in RM and the treatment of incurable illnesses, these cells have aroused a lot of curiosity. There has also been research on the therapeutic use of IPS cells in bone repair through seeding on NFs SS [223, 224].

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8 Gynaecology Since no EFsNFs-based devices have yet been authorized, this promising technology might be used in a variety of medical applications, many of which are related to women's reproductive health (WRH). EFsNFs offer a lot of promise for generating new product designs and management systems because of their flexibility in adjusting DD properties and 3D form. EFsNFs have been employed as TE implant (IMP) coatings to prevent biofilm development and minimize inflammation and calcification on vascular grafts (VGs) [225]. IMPs utilized in WRH, such as IUDs and contraceptive (CP) IMPs, may benefit from similar coating methods. STI prevention, IUDs, and therapies are the three core areas of WRH identified as most eligible for EFsNF improvement [225]. As a DD approach, EFsNFs have several benefits, including configurable DR and the potential to encapsulate a wide range of therapeutic medicines, and they are well positioned to address the present WRH issues. This adaptability may pave the path for future multipurpose preventive technology (MPTs). These devices are intended to prevent many STI indications in a single dose. MPTs are now being studied in the form of intravaginal rings, gels, the SILCS diaphragm, polyoxymethylene (POM), dapivirine (DPV), levonorgestrel (LNG), and rilpivirine [226–230] and have the potential to provide a revolutionary means of contraception. As a customized DD system, EFsNFs have the potential to allow the creation of products that enhance the health of women all over the globe, with MPTs providing a particularly promising platform for improvement in women’s reproductive health. EFsNFs have yet to be approved as a commercially viable dosage form, despite its enormous potential as a drug delivery platform. To reduce the WRH gap, current ELS technology must be improved to match the production capacities needed for a commercial product, and these materials must be tested for safety and effectiveness in clinical trials.

9 Cancer Biology (CB) Cancer (CA) therapy is one of the most essential components of therapeutic management. CA cells have several distinct features, including uncontrolled proliferation, longevity, and the capacity to spread [231]. This chapter will look at the current uses of EFsNFs in CA therapy, with an emphasis on DD and CA cell detection/ sensing, as well as the obstacles and future perspectives for EFsNFs applications in CA research. The most often used CA treatment techniques include conventional surgery, radiation, and chemotherapy [232]. These alternatives, however, are frequently restricted and insufficient [233]. In fact, at therapeutic doses, the most widely used chemotherapeutics against a variety of solid and haematopoietic malignancies, such as breast CA, osteosarcomas (OS), aggressive lymphomas (LY), and leukaemia (CLL), are generally restricted by severe and devastating side effects

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[234]. Furthermore, many patients are required to take excessive doses of medicines, either orally or through systematic injection, in order to optimize therapeutic results, which may cause adverse effects in healthy tissues [235]. Researchers are presently concentrating their efforts on postoperative chemotherapy applications in order to produce novel and sophisticated therapies that selectively target CA cells in order to reduce frequent systemic side effects and toxicity. Micro and NPs, liposomes, and NFs materials are examples of targeted micro-carrier technologies that are meant to transport drugs to a specific location in a regulated manner. The controlled release of medicines is a difficult task in which systems deliver an ideal quantity of medication to a specified target at a predefined pace and for a certain period of time [236]. Controlled-release systems provide benefits over traditional pharmacological treatments. To begin, with a single dosage, they keep the medicine in the proper therapeutic range in the blood. Avoiding the reciprocated inclination of injectable drugs into the blood, hence the results in alternating periods of ineffectiveness and toxicity. Another advantage is that drugs may be delivered to a specific physiological target. With a smaller dose of the required medicine and a lower frequency of drug administration, the pharmacological characteristics of free drugs can be enhanced, and patient compliance can be raised [237–239]. In addition, unlike sustained formulations like suspensions and emulsions, which are influenced by environmental conditions and hence subject to patient variability, controlled-release formulations and procedures allow the drug to be retained in a polymeric state. Two key processes govern the polymeric system network: (1) The most frequent release is drug diffusion, in which the drug flows in a concentration gradient from the system's inner side to the system's outer side, eventually reaching the body; and (2) PLY degradation or drug cleavage from a PLY bond to enable DR are examples of chemical processes [240]. EFsNFs are a favourable substrate for DD because of their large surface area to volume ratio [241, 242]. They could allow for a lot of DR antibiotics, anti-cancer medicines, proteins, and nucleic acids have all been given in various ways. EFsNFs have lately gained appeal as a local medicine delivery strategy after surgery to remove solid tumours [243]. The active molecules might be chemically or physically linked to the fibre surface, or they're just mixed with PLYs. Various loading processes will be covered in the following sections.

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Snakebite (SKB)

Several serious illnesses, including as CVD, DT, and CA, are on the rise as a result of today's lifestyle. Early discovery and treatment can reduce rates of death in many cases. Clinical concerns such as SKB envenomation, in addition to lifestyle-related disorders, are a substantial health concern across the world. Over 5.4 million SKB cases are reported globally each year, with over 100,000 people dying as a result of SKBs, including over 50,000 deaths in India alone [244]. The great majority of deaths happen in rural and agricultural areas. SKB has also been identified as a

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“neglected tropical condition” by the WHO, making it a global health concern [245– 247]. Although antiserum (AS) is the sole treatment option for SKB patients, it may not always work to prevent venom-induced bleeding, necrosis, nephrotoxicity, and hypersensitivity [248]. In animals, the development of AS is delayed, necessitating the deployment of a cold chain. As a consequence, a successful therapy for SKB has been developed. The employment of non-physician practitioners (NPs) to dispense medications is becoming more widespread [249]. Titanium dioxide (TiO2) is one such promising chemical that has been utilized in a variety of biological applications and has qualities that vary due to the production processes used [249].

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Neurology

Using EFsNFs meshes with TE techniques for neural tissue reconstruction has a lot of potential. The composition and design of EFs mats have been reported to alter brain cell function in several investigations; however, the effect varies depending on the kind of neural cell employed. According to recent clinical research, placing EFs mats in three-dimensional (3D) patterns can encourage tissue regeneration and restore full function. SS design will be prioritized in the future due to the rising difficulty of producing functioning 3D brain tissues. Due to the dense packing of FBs, traditional ELS are restricted in their ability to form transmembrane or plate matrices, limiting cells release [161]. Different techniques are being investigated, including the use of sacrificial agents during spinning and the mixing of micro and NS fibre shapes to increase porous size [250], but the use of 3D readers instead of horizontal collectors [251]. In the brain and spinal cord, the difficulty is not only enhancing neuronal survival, regenerating axons across the injury site, and/or mending connections with the target of innervation, but also managing the inflammatory response, which causes further injury [252]. In the future, EFs SS might be used to deliver neuroprotective drugs and bioactive agents (BAs)/growth factors (GFs) to improve brain tissue growth. To get the required response, different materials, GFs, and drugs may need to be combined and placed in the EFs mats. To achieve functional recovery, different cell types may need to be mixed with EFs mats. Damage to the nervous system is a difficult issue that requires innovative solutions to help restore function [253].

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Diabetes (DT)

NFs are capable of transporting a large spectrum of small substances to lesion sites. As stated previously, various proteins and nucleic acids were loaded into NFs or functionally linked to the surface of FBs. The release duration and loading capacity of NPs with unique GF incorporation by NFs are increased. GFs formed by gelation NPs (GNPs) were used to deliver VEGF and PDGF for DT issues, with Col and HA

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functioning as NFs mesh. In the dual source dual power method, VEGF-loaded GNPs and bFGF HA solution, as well as PDGF/VEGF-loaded GN and EGF Col solution, were employed as EFs. The NPs-decorated NFs simultaneously released four separate GFs, revealing NFs' capacity to transport bioactive compounds [253]. NFs mediated GFs used for DT complications include heparin mimetic peptide amphiphile (HMPA) [126], HBPA, peptide/heparin hybrid, PCL, PCL-PEG block copolymers, PCL-PEG diblock copolymer, PCL and PEG, polycaprolactone (PCL)/PEG/PCL triblock copolymer, PLGA, PLA-PVA, PELA, PLGA/CNC, Col/HA, and PHBV [126, 243, 250, 254–262].

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Conclusion

With its ease of use, scalability, and flexibility in changing fibre diameter from micrometre to NS range, the ELS technique has aroused the interest of a number of BMAs. Despite the fact that ELS has been around for some decades, the methodologies and equipment utilized in the process are constantly changing. Through coaxial and emulsion spinning arrangements, ELS developed from an SN design to multi-nozzle variants. More research is being done to enhance fibre characteristics and simplify the production process by altering the nozzle architecture and collector design. This study looks at EFsNFs that have been impregnated with NPs and covers a number of essential features of ELs in the usage of EFsNFs in drug administration. There is also discussion of studies on the biological uses of EFsNFs. EFsNFs may now be used to provide a variety of antibacterial (ANB) and anti-cancer (ANC) therapies by carefully selecting PLYs. Techniques for massive, consistent production of NFs with acceptable architectural and material characteristics should be established in order to make future breakthroughs, notably in the realm of drug administration. Despite the efforts of researchers and academic specialists, the majority of EFsNFs research is done in vitro. Future advances in the field of EFsNFs will necessitate in vivo testing. To meet current and future pharmaceutical delivery demands, researchers in this field must devise ways for using NFs in BMA.

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Adv Polym Sci (2023) 291: 167–176 https://doi.org/10.1007/12_2023_144 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 Published online: 11 February 2023

Biomedical Applications of Electrospun Piezoelectric Nanofibrous Scaffolds Afeesh Rajan Unnithan and Arathyram Ramachandra Kurup Sasikala

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Significance of Piezoelectric Polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Tissue Engineering Applications of Electrospun Piezoelectric Polymers . . . . . . . . . . . 2.2 Piezoelectric Polymers-Based Self-Powered Implantable Biomaterials . . . . . . . . . . . . . 3 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract The recent developments in the field of smart biomaterials and the research related to their properties have provided ground-breaking approaches to the field of biomedical research. Piezoelectric materials are one of the most promising smart materials for biomedical applications which can generate electric signals when mechanically stimulated. Electrospun nanofibers are an exciting class of materials for biomedical applications due to their topological and mechanical properties which can directly relate to the characteristics of biological materials and extracellular matrices. Electrospinning technology can be evoked to create micronano-size fibres and scaffolds with various in situ functionalities such as piezoelectricity which can be directly accessible to the cellular level and hence makes it a versatile tool for bioengineering applications. Electrospinning itself can be used to evoke piezoelectricity in certain polymers and hence the combination of electrospinning and piezoelectricity can be a better option for the development of the next generation of smart biomaterials. The chapter gives a technical overview of the properties and the applications of piezoelectric nanofibers and their potential applications in tissue engineering to implantable self-powered biomedical devices.

A. R. Unnithan (✉) and A. R. K. Sasikala Faculty of Life Sciences, Centre for Pharmaceutical Engineering Science, School of Pharmacy and Medical Sciences, University of Bradford, Bradford, UK e-mail: [email protected]

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Keywords Bioengineering · Biomaterials · Electrospinning · Nanogenerators · Piezoelectricity · Tissue engineering

1 Introduction Tissue engineering is focused on the development of rebuilding and repairing injured/unhealthy cells or tissues via creating biological alternatives utilising bioinspired materials or biomaterials [1, 2]. The most critical part of this bioengineering process is the biomaterial scaffold which provides a guided protective platform to cells to control and direct the cellular matrix synthesis, cell growth and hence accelerate the development of new tissue. The niche and the environment created through the biomaterial scaffold must remain high in generating enough physical and biochemical signals capable of obtaining long-lasting cellular activity between cells or tissues and that can be very crucial for the regeneration of damaged tissue [3]. Due to the ability to mimic the composition and structures of the extracellular matrix (ECM), the electrospinning-based nanofibrous scaffolds are extensively utilised for tissue regenerative applications. The content and structural morphology of each fibre and the developed scaffolds can be modified to control the cell behaviours like cell addition, proliferation, cell migration and cellular differentiation [4–7]. Along with the topological signals and exterior physiochemical signals, the developed nanofibrous scaffolds can also be used to promote the regeneration of defective tissues. Electric fields are omnipresent throughout the body and play a crucial role in various biological processes such as cell signalling, neural regeneration, osteogenic regeneration, cardiac tissue function, etc. [8]. The presence of electric fields is inevitable in many biological functions, e.g., neural cells can experience up to 140 mV/mm of the electric field during neuronal wound healing [9, 10]. Hence neuronal wound healing can be enhanced using electroactive scaffolds to direct the neural cells to grow on precise patterns. Similarly, different kinds of cells rely on mechano-electric stimulations for enhanced regeneration and reconstruction. But the safe and effective way of giving electric stimulation to cells remains a challenge. Most of the existing electrical stimulation procedures require invasive electrodes. Among those, the piezoelectric material-induced electric stimulations are considered effective and safe for the development and regeneration of electroactive tissues such as tendons, bone, etc. Recently there has been a wide interest in developing bioactive piezoelectric biomaterials for tissue regeneration [11, 12]. Piezoelectric biomaterials hold tremendous potential as they can develop bioelectric signals in response to mechanical deformations. Piezoelectricity is a material property and the basic requirement for this is the non-centrosymmetric orientation of ions which will ultimately generate a permanent dipole. An applied mechanical stimulation will cause a separation of charges in piezoelectric materials which leads to the

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development of a local electric field. In the case of piezoelectric biomaterials, these stimulations can be provided in various ways such as ultrasound, physiological stress, etc.

2 Significance of Piezoelectric Polymers Piezoelectric materials can generate electric potential when strained through applied stress, which is called the direct piezoelectric effect. There are materials whose polarisation can be switched by an externally applied electric field, those are called ferroelectric materials. In order to initiate this response, the material has to be poled within a high electric field to align these dipoles. So, the ferroelectric material-based biomaterial scaffolds should be poled before application. This can be accomplished by introducing the scaffolds to an electric field higher than the scaffold material’s coercive field. This is usually done at raised temperatures to enable the alignment of dipoles in the applied electric field’s direction. Such corona poling is preferred for biomaterials compared to other poling methods such as thermal poling due to the possibility of contamination [13, 14] from thermal poling. Both corona poling and thermal poling were utilised to pole the complex biomaterial scaffolds for introducing the piezoelectric behaviours. The extent of voltage and temperature required for poling depends on the material’s piezoelectric characteristics and is modified to obtain the most effective piezoelectric poling. Interestingly the piezoelectric electrospun nanofibers possess tremendous application in bioengineering applications owing its ability to provide electrical stimulations which are analogous to the bioelectric signals to enhance tissue regeneration. Piezoelectric nanofibrous scaffolds provide exceptional properties due to its topological properties and dimensional features and hence can be applied for numerous biomedical applications. But the development of piezoelectric ceramics-based electrospun nanofibers is more difficult due to the physical properties such as hydrolysis, gelation and condensation when compared to polymers. This issue has been solved by introducing the synergic approach of combing electrospinning and sol-gel. This hybrid process allows the development of different compositions, sizes and morphologies of ceramic nanofibers. Similarly, the single-step fabrication of piezoelectric composite nanofibers can also be done using this method by utilising nanoceramics as fillers. All of these advantages positively contributed to the development of non-toxic lead-free piezoelectric nanofibers and scaffolds for biomedical applications. Among these, the development of self-powered implantable devices, nanogenerators and electroactive scaffolds for tissue engineering is gaining much interest due to their innovation in approach [15, 16]. Electrospun piezoelectric nanofibers have triggered the research interest in developing electroactive tissue engineering scaffolds due to their ability to provide electric stimulations for tissue regeneration. Among these electrospun nanofibers, the biomedical application of fluoropolymers such as the PVDF (polyvinylidene fluoride) is widely explored due to their chemical inertia, stability and processability. In addition

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to that the piezoelectric properties and excellent biocompatibility made PVDF an excellent candidate for developing stimuli-responsive smart scaffolds for tissue engineering.

2.1

Tissue Engineering Applications of Electrospun Piezoelectric Polymers

As explained earlier, the piezoelectric behaviour of PVDF contributes to the creation of electrical potential upon mechanical deformation. During electrospinning, polymer stretching generates the formation of the polar β phase from the non-polar α phase, which contributes to the piezoelectric nature [17, 18]. Such nanofibrous piezoelectric scaffolds that can create electrical stimulations are widely considered for bone and cartilage tissue engineering applications. In a recent study [19], oxygen plasma-treated electrospun PVDF scaffolds were developed to obtain high piezoelectricity and were compared with the drop-cast membranes for their efficacy as potential osteogenic scaffolds. The results showed that the PVDF has more β phase in electrospun-scaffolds compared to drop-cast scaffolds. The oxygen plasma-treated PVDF scaffolds showed higher wettability and improved surface functionalities. The piezoelectric electrospun-scaffolds showed excellent biocompatibility with Saos-2 cells upon seeding. The electrospun scaffold exhibited excellent support to cells by inducing osteoblast phenotype and cell spreading. The electrospun piezoelectric nanofibers enhanced intracellular Ca2+ currents as measured through the activated cell numbers. The drop-cast based non-piezoelectric scaffold didn’t activate many cells as compared to the piezoelectric scaffold. So this study showed that the combination of the electrospun piezoelectric PVDF nanofibers with oxygen plasma treatment can generate a smart scaffold that is capable of stimulating electrically excitable cells, without the help of external electrical stimulation. In another study [20], flexible nanofibrous (PVDF-TrFE) scaffold was developed and assessed for in vitro tissue regeneration and tissue growth. The electrospun PVDF-TrFE nanofibrous scaffolds were electrospun and then β phase was enhanced through heat treatment for improving its piezoelectric properties. The mechanical stimulations were provided to the scaffold using a dynamic bioreactor where physiological frequency-based cyclic compressions were provided and the differentiation of MSCs towards osteogenic and chondrogenic lineages was evaluated. The results revealed that the differentiation of MSCs was affected by the level of piezoelectric effect of the scaffold. Moreover, it was also found that the higher piezoelectric effect promoted osteogenesis and the lower level of piezoelectric effect supported chondrogenesis. The piezoelectric nature of the PVDF-TrFE scaffold supported greater MSC differentiation as evidenced by the higher gene expression and the increased matrix synthesis under mechanical stimulation demonstrating the effect of piezoelectricity on cellular differentiation (Fig. 1).

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Fig. 1 Graphic representation of the bio-mechano stimulation mediated calcium signal activation pathway [21]

In a recent study [11], the electrospinning collector design was modified to generate various combinations of piezoelectric nanofiber patterns such as single pattern aligning, randomly oriented pattern aligning, radially aligned patterns, repetitive patterns of gradient alignment and random patterns of electrospun PVDFBaTiO3-Multiwalled carbon nanotubes based piezoelectric nanofibrous scaffold. These topographically modified piezoelectric scaffolds can regulate and manage electrical cues and biomechanical signals to control neuronal cell responses. The capability of such topology-modified piezoelectric scaffolds to enhance neuronal cell differentiation and regeneration can be successfully used for long-term non-invasive neural regenerative applications making them the best candidates for nerve repair. Also, the study on Hippo/YAP signalling pathway provided clues on how the cellular behaviour is controlled on the piezoelectric scaffold. The topographical cues along with piezoelectric properties will enable the stress fibre development in neuronal cells and hence stimulate nuclear YAP localisation. So this study effectively fabricated electrospun-scaffold incorporating various topographical characteristics and structural alignments in the fibres to study the impact on neural

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regeneration. Hence such developed scaffolds will be a better option to deliver various mechanotransduction and electrical stimulations which can have an impact on the cell morphology and cell proliferation through their mechanosensitive and voltage-gated cellular channels. The regeneration of electroactive tissues such as muscle tissue regeneration can be significantly enhanced by the application of piezoelectric scaffolds as the muscle cells react to electro-mechanical stimulations and tend to maintain the fibrillar structure. The studies have reported the impact of topography and polarisation of PVDF piezoelectric electrospun nanofibers on the enhancement of the tissue response on muscle tissues [22]. The studies revealed that the aligned PVDF nanofibers with a negative surface can provide better stimulations for improved myoblast regeneration [23] and hence the vital application of electrospinning-based nanofibrous biomaterials cardiac bioengineering applications has been widely reported. The studies showed that the required adhesion of cells created the instantaneous contraction within post-24 to 48 h seeding. The cell viability assay also revealed higher biocompatibility on day 3 and day 6 without any notable cytotoxicity [24]. In another study, better cell attachment and piezoelectric effect have been promoted through a ZnO nanoparticle (NPs) incorporated P(VDF-TrFE) composite piezoelectric nanofibres [25]. The studies revealed the attachment and proliferation of hMSCs and the improved angiogenesis by utilising the bioelectric properties of the composite nanofibers and ROS (reactive oxygen species) generated by the ZnO NPs. According to the study, a 2 wt% of ZnO NPs composite piezoelectric scaffolds were biocompatible and enhanced cell attachment and proliferation. The in vivo studies also confirmed the cytocompatibility and the angiogenesis of the composite piezoelectric ZnO/PVDF-TrFE nanofibrous scaffolds. Studies have been conducted to enhance the cellular response and the bioelectrical properties of the piezoelectric scaffolds. In one of the works [26], a magnetic nanofilm of PCL was covered with microfibers of PVDF-TrFE to conserve the contractility of cardiomyocytes and to enhance its biocompatibility, cell attachment and proliferation. The scaffold showed higher piezoelectric constant and enhanced the human and rat cardiac cell attachment and the presence of PCL-based layer provided the necessary mechanical stability and contractability for the myocytes cultured on the developed scaffold.

2.2

Piezoelectric Polymers-Based Self-Powered Implantable Biomaterials

The development of implantable biomaterial devices (IBD) can enhance the efficiency of diagnostic tools and which in turn can improve the human life status by early detecting the presence of pathogens affecting the body along with supporting healthy tissue regeneration. Numerous studies have been conducted to develop lightweight and flexible IBDs to reduce interference with daily actions or activities. Energy harvesting from body movements is continuously studied for developing

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Fig. 2 Schematic representation of the various application of piezoelectric biomaterials-based nanogenerators [27–32]. (Reproduced with permission [12])

self-powered IBDs which will translate biomechanical energy to bioelectrical energy. In this case, piezoelectric materials will possess a vital role as they will translate the physical energy into electrical signals (Fig. 2). So piezoelectric polymer-based nanogenerators are broadly used to generate biophysical energy-dependent IBDs. The development of flexible electrospun piezoelectric nanofibers has significantly fixed many issues related to the development of IBDs coupled with irregular human organs [33]. To comply with the necessities of wearable IBDs electrospun PVDF-TrFE and PVDF polymer-based nanogenerators were developed. These electrospun nanofibers-based flexible electronics can be interlaced into flexible textiles and incorporated with fabrics to harvest biomechanical energy through physical movements [34]. Recently, these PVDF and its copolymer-based piezoelectric nanofibers are the core of attention in developing digital health monitoring systems due to their excellent features such as biocompatibility, chemical inertness, easy preparation methods and cost-effective mode of development compared to the existing systems [35]. The self-powered batteryless piezoelectric IBDs won’t depend on an external power supply and are hence considered the most reliable and efficient systems for developing the next generation of implantable biomaterial devices when considering the safety issues associated with batteries [36]. The long-time demand for developing the healthcare monitoring device and endless power sources can also be solved by the electrospun piezoelectric

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nanofiber-based nanogenerators [37]. They can be used to detect numerous human gestures such as breathing, throat movements, heart beating, arm and shoulder movements, etc. [35].

3 Conclusion The electrospun piezoelectric nanofibrous scaffolds are currently attracting significant scientific attention in the field of healthcare technologies due to their excellent biocompatibility nature and inherent mechano-bioelectrical properties. These outstanding properties can be utilised in novel and interesting ways to repair and restore tissue and body functions. The fascinating technology of electrospinning helps the development and production of piezoelectric nanofibrous scaffolds in combination with bioceramics, nanoparticles and biomolecules. Electrospinning helps us to modify the parameters, which directly affect the piezoelectric properties of each nanofiber and scaffold in general. From a future perspective, intense research efforts on controlling the piezoelectric properties of piezo-scaffolds including the impact of different piezopolymers, various fabrication techniques including random, aligned and combination nanofibers and their effect on various post-treatment methods like plasma treatments, etc. In future, the need for personalised health technologies will be on the rise and hence the development of high-performance piezoelectric devices will be a fascinating research area for many decades.

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Adv Polym Sci (2023) 291: 177–190 https://doi.org/10.1007/12_2022_143 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 Published online: 3 January 2023

Surface Modified Polymeric Nanofibers in Tissue Engineering and Regenerative Medicine Nivethitha Ashok, Deepthi Sankar, and R. Jayakumar

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Surface Modification of Polymeric Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Hydrogel Coating . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Chemical Treatment . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3 Plasma Treatment . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Conclusions and Future Outlook . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract Fibrous scaffolds are commonly used as scaffolds in tissue engineering as they have remarkably high surface-area-to-volume ratio. Surface modification of fibrous constructs by employing innovative post-processing techniques such as chemical modifications, using coating gradients or active cues, avenues for tuning cellular respones. Most of the commercial polymers/unmodified polymeric scaffolds contribute little to generate biological activity. The below described methods offer plethora of chances by altering the complexity of fibers for efficient scale-up that would be relevant for industrial applications. It is also to be taken into consideration that while modifying the surface of these fibers, the mechanical properties of the materials are to be retained as such. In this chapter, we have discussed about the various modification chemistry and techniques to improve biological activity of the fibrous scaffolds.

N. Ashok, D. Sankar, and R. Jayakumar (✉) Polymeric Biomaterials Lab, School of Nanosciences and Molecular Medicine, Amrita Vishwa Vidyapeetham, Kochi, Kerala, India e-mail: [email protected]

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Keywords Chemical treatment · Fibers · Hydrogel coating · Plasma treatment · Surface modification

1 Introduction The multi-fibril ECM structure that promotes cell adhesion, proliferation, and differentiation is mimicked in three dimensions by electrospun fibers [1]. An ideal fibrous scaffold should assist in the synthesis of ECM, offer mechanical support for physiological processes, and gradually be replaced by tissues that have undergone regeneration. The design takes into account appropriate materials with specific properties including biocompatibility, bioresorbability, biodegradation kinetics and mechanical strength comparable to ECM [3]. For the production of fibers, synthetic and biodegradable polymers like poly(lactic acid) (PLA) and poly(lactic-co-glycolic acid) (PLGA) are widely utilized [2]. Cell adhesion is greatly influenced by the surface. Hydrophilic surfaces are best for tissue engineering since their contact angle should be less (100°) of aliphatic polyester present in polymer fiber prevents cell attachment [4, 5]. An ideal fibrous scaffold should have surface biological moiety and bulk characteristics. However, the synthetic polymer fiber often achieves bulk qualities but falls short of biological functions at the surface [6]. Signal transmission and cell anchoring on ECM are two ways that cells and ECMs interact with integrins, also known as cell adhesion molecules (CAM), signalling complexes, which are responsible for identifying the fate of cells. The majority of the time, complex signalling molecules including growth and differentiation factors determine how cells migrate, differentiate, proliferate, arrange their ECM, and remodel. Actin filament complexes formed by the coordination of ECM and the cytoskeleton of a cell, initates cell attachment. The cell membrane’s integrins aid in mediating adhesion [7]. Thus, a key component for better cell–material interaction in nanofiber tissue engineering is the modification of surface characteristics [4, 8]. When biomaterials are implanted within the body, a number of interactions take place, including blood fluid contact, protein adhesion, and subsequently cell adhesion on the implant’s surface. Surface hardness, chemical ratios, energy, and roughness are the only factors that affect the events [9]. Numerous techniques, including co-axial electrospinning, physical and chemical modification, plasma assisted method, LBL method, mixing of two or more polymers, and others [10], are used to convert the surface to hydrophilic and fix biological molecules onto it. In this review (1) The permanent alteration of the surface by covalently joining the biomolecules is thoroughly examined. Chemical stimuli that expose the amino, hydroxyl, and carboxylic acid groups on biomaterials aid in the biomolecules attachment, substantially immobilizing the connected biomolecules [11, 12]. (2) Using plasma treatment to modify the surface free energy of polymeric

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substrates and their polarity to transform fiber into a hydrophilic surface. As a consequence, only the fiber surface is changed; the bulk characteristics very much stay the same. In addition, modifying the surface’s wettability by covering hydrogel has been efficient in numerous ways.

2 Surface Modification of Polymeric Nanofibers The electrospun fibers hold enormous surface area and association of different techniques such as micro patterning [13], photolithography [14], layer-by-layer deposition [15–18], and hydrogel coating [19, 20], plasma and chemical treatment [21] are ample to modify the surface area of these fibers [22]. In brief, hydrogel coating, plasma and chemical treatment will be the subject of this section.

2.1

Hydrogel Coating

Hydrogels consist of cross-linked polymers which swell but are insoluble when in contact with water. Hydrogels are of high importance as they help in cell growth, wherein they aid in diffusion of nutrients, providing an environment for the cells to grow through the hydrogel network. Owing to the unique properties, namely biodegradability, biocompatibility, and composition close to native ECM, hydrogels composed of natural polymers and/or polysaccharides, viz. chitin, chitosan, gelatin, hyaluronic acid, etc., found widespread application in the field of tissue engineering [23–26]. In a study by Deepthi et al. (Fig. 1), chitosan-hyaluronic acid gel was coated on PCL random and aligned multiscale fiber, first on one layer of the electrospun scaffold. Followed by this is the process of arranging another hydrogel coated layer to form a bi-layer scaffold, which is left overnight for drying. The scaffolds are cross-linked using EDC and then lyophilized completely [27]. The coating using hydrogel would hold the cells, and in addition, act as a reservoir, holding the growth factor that will aid in the growth of ligament cells. The same group had developed a scaffold [28]: electrospun aligned poly(L-lactic acid) (PLLA) fibers with chitosan-collagen hydrogel coating (Fig. 2). The coated membrane is rolled and is further coated with alginate as a preventive measure for peritendinous adhesion. The aligned fibers mimic the collagen fiber bundles whereas the hydrogel layer mimics the glycosaminoglycans of tendon sheath, all of which aid in tendon regeneration. A biodegradable scaffold was constructed by Hayami et al., using electrospun poly(E-caprolactone-co-D,L-lactide) fibers, which were embedded in primary ligament fibroblast loaded photocrosslinked N-methacrylated glycol chitosan hydrogel, which forms a better substitute for existing constructs in ligament tissue engineering [29]. One of the drawbacks of electrospinning process is that it produces meshes which are thin with large lateral dimensions, which is well suited for vascular and skin

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Fig. 1 Flow chart depicting the steps involved in the fabrication of chitosan-hyaluronic acid gel coated PCL multiscale fiber [27]. Reproduced with permission, Copyright Elsevier

Fig. 2 Schematic representation of chitosan-collagen/PLLA/alginate scaffold [28]. Reproduced with permission, Copyright Elsevier

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Fig. 3 (a) Representation of simultaneous electrospinning which produces multiscale fibrous scaffolds (b) Pictorial diagram showing the steps involved in the alginate coating of the braided scaffold [38]. Reproduced with permission, Copyright American Chemical Society

applications [30, 31], but limits applicability in terms of three-dimensional tissues such as ligament, tendon, muscle, etc. To address this issue, Thayer and group had developed thick meshes by rolling/stacking [32–34]. These layers are then bound together with the help of polyethylene glycol diacrylate (PEGDA) or fibrinogen which polymerizes in situ to arrange itself as an interpenetrating hydrogel network [35, 36]. In an attempt to improve viscoelastic property, Freeman and group fabricated a scaffold comprising of PLLA fibers in combination with PEGDA hydrogel [37]. Anjana et al. had developed a braided fibrous construct (Fig. 3) comprising micro PCL-nano collagen-bFGF, which would be suitable for tendon tissue engineering and coated with alginate as the final step to prevent peritendinous adhesion. One of the major components of tendon ECM is collagen, which along with the aligned fibers offers the functional and structural properties of tendon [38].

2.2

Chemical Treatment

Modifying the fiber’s surface can be carried out by presenting functional moieties by means of alkaline hydrolysis wherein the ester linkages of the polyester backbone are broken forming free carboxylic acid and hydroxyl groups [39]. The process of aminolysis, where one of the di-functional groups reacts with the fiber surface and the second group is left for attaching to topographic cues or proteins [40–42]. A study by Yoo et al. showed the use of aminolysis followed by EDC coupling, in an attempt to attach thiol and DHPA [3-(3,4-dihydroxyphenyl)] on to the surface of

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PCL fibers. The study showcased improved cell adhesion and viability in comparison with unmodified PCL fibers [40]. It has been observed that aminolysis and alkaline hydrolysis lead to polymer degradation and chain scission, and to address these concerns, chemical cues have been introduced on the surface in the form of immobilized peptides. These peptides which are present on the surface of the fibers help in promoting integration with the host tissue, by providing signals to the cells [43, 44]. Viswanathan and colleagues had developed amphiphilic diblock copolymers based on PLA, by means of a thiol-ene reaction. To the PEOGMA block an RGD peptide was conjugated. This modified fiber’s surface, by presenting them with cell adhesive and cell inert sites, which mimics a similar condition found in the ECM. Results showed enhanced proliferation and adhesion of human mesenchymal progenitor cell [45]. Another powerful tool that aids in surface modification by the formation of hetero-atom linkages between two different functional groups has emerged, named as the “Click chemistry”. Over the recent years, several click chemistries have attached bio-functional moieties to the surface of PCL fibers. This presents the surface with a hydrophilic base for cell attachment and during cell seeding it inhibits molecular weight degradation [46–48]. Lancusˇki and group developed PCL scaffolds functionalized with azide groups. Further alkyne group containing bioactive molecules were conjugated to the scaffold surface, enabling them for CuAAC reaction. But, it was observed that the residual copper from the scaffold yielded aldehydes which eventually led to biocompatibility issues and side reactions [49, 50]. In order to eradicate copper from the surface, during this modification process, researchers had established “click” reactions, namely, thiol-ene reactions, thiol-Michael addition, strain-promoted alkyne-azide cycloadditions (SPAAC) and oxime ligation [51–56]. A well-controlled reaction technique termed “multi-click” chemistry has been efficiently in use for attaching multiple functional groups to the fiber surface [54, 56, 57], which has exhibited potential synergistic cellular interactions [58–60]. This technique has provided the opportunity to attach multiple functional moieties to fiber’s surface, which paves the way for promising result as it closely mimics the complex native environments which presents multiple factors.

2.3

Plasma Treatment

It is a process of surface modification of polymer fibers that alters the surface for a few nanometers without changing the bulk characteristics of the fibers [61– 63]. When microwave or radio frequency waves are introduced into an ionizing gas chamber at low pressure or in a vacuum, plasma is created. When this plasma reacts with the substrate’s surface, it changes the functional group and also physically alters the surface pattern and roughness [63, 64]. The surface of the material is functionalized with hydroxyl, carboxyl, and amine groups by using non-polymerizable gases. Since plasma is a charged particle, it possesses high energy. Thus, it modifies the surface properties via etching, grafting, functionalizing,

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etc. [63, 65, 66]. Air, ammonia, and oxygen are some gases that are used for plasma treatment, which gives rise to the carboxyl, hydroxyl, or amine functional group on the surface [10, 67–70]. Hence, choosing a perfect source for plasma is considered, for creating various functional groups onto the surfaces, which directly alters or enhances the surface’s properties like wettability, cell adhesion, biocompatibility [71, 72]. This surface modification also covalently immobilizes various extracellular proteins such as fibronectin, collagen, laminin, and gelatin onto the surface, which enhance cell adhesion and its proliferation [73–75]. Furthermore, plasma treatment eliminates surface contamination and commonly results in a hydrophilic surface when inert or less polymerizing gases are employed. In general, for biocompatible fibers, argon, and air gases are used for treatment for increasing wettability [76, 77]. The extent of surface modification depends on other factors like temperature, pressure, and gas mixture ratio which are to be critically controlled. In general, plasmas are generated in a PVD setup, wherein the air inside the chamber is evacuated, low pressure or vacuum is introduced, so it can also be known as cold plasma treatment, hence, samples are not affected by temperature. Microwave or RF wave energy source is used to generate plasma (Fig. 4). By using this technique hydrophobic polymeric surfaces are turned into hydrophilic surfaces by fixing various chemical moieties such as hydroxyl, amine, and carboxyl, which augments cell adhesion and subsequent growth [78–81]. Plasma-treated polymer nanofiber surface displayed enhanced cell adhesion [82]. A detailed study about the effect of surface plasma treatment toward cell–material interaction using argon and nitrogen on PCL nanofibers, microfibers and multiscale fibers was carried out. The treatment rendered the surface with oxygen and nitrogen containing chemical

Fig. 4 Pictorial representation of method employed in plasma treatment

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groups, which imparted greater hydrophilicity, protein adsorption thereby improving cell attachment and proliferation, by regulating the interaction of adhesion molecules [83]. Another study determined a correlation between plasma treatment and fiber morphology in inducing tenogenesis. Argon plasma treatment rendered the surface of aligned PCL fibers interspersed with sacrificial collagen nanofibers, with more oxygen containing functionalities and random nanoroughness. This enhanced vitronectin adsorption and mesenchymal stem cell attachment, proliferation, and elongation. Tenogenesis was induced on these fibers triggered by vitronectin mediated cell attachment, cell elongation and cell shape change, rhoA activation and tendon-related marker expression [84]. Hence, a positive correlation between surface protein adsorption and surface free energy was proved [83]. Griffin et al. fabricated polyurethane composite and plasma treated with different gases like oxygen, argon, and nitrogen. Differentiation studies using adipose tissuederived stem cells on these polyurethane composites indicated that Ar plasma treatment had bought in sufficient chondrogenic and osteogenic differentiation [85]. Protein adsorption on the surface plasma-treated surface is also an important basis for cell adhesion. The air plasma-treated, core-shell PVA-PLLA fiber showed enhanced protein adsorption onto the surface [86]. It is a fact that plasma treatment alters the surface properties and not the bulk. Another fact that should be taken into consideration is the plasma setup, where it should be cost efficient, thus bulk quantities of electrospun fibers surface can be modified economically.

3 Conclusions and Future Outlook Owing to the ECM mimicking properties, possessed by the polymeric fibers, it has drawn immense attention from researchers all over the world, belonging to the biomedical community. In this chapter, we have discussed about the strategies undertaken for surface modification, focusing on functionalization approaches and also treatments which preserve the fiber integrity. It is highly anticipated that the fibrous architectures that greatly mimics the native ECM architecture will favor cellbased outcomes, since these fibrous constructs are better able to provide the resources as in the biochemical environment. The surface modification strategies that have been discussed in this review may facilitate simplistic and efficient coupling mechanisms to introduce various chemical functionalities, retaining the scaffold integrity and, thereby, enriching the chemical and biochemical factors. Hence, this review is concluded by understanding the different functionality aspects that are presented by these scaffolds, which opens avenue for approaches to design spatially patterned factors which hold the potential to closely connect the synthetic materials with the biological environment. The future research on designing nextgeneration fibrous scaffolds must consider the link between manufacturing, mechanics, topology, and surface chemistry to overcome the drawbacks of biomaterials pertaining to scalability, controlled and tailored cellular response, and scaffold integrity.

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Adv Polym Sci (2023) 291: 191–212 https://doi.org/10.1007/12_2023_145 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 Published online: 11 February 2023

Polymer/Ceramic Nanocomposite Fibers in Bone Tissue Engineering S. Sowmya, Nirmal Mathivanan, Arthi Chandramouli, and R. Jayakumar

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Polymer/Ceramic Composite Nanofibers in Bone Tissue Engineering . . . . . . . . . . . . . . . . . . . . 2.1 Silica-Based Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Calcium Phosphate (CaP)-Based Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3 Magnesium-Based Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4 Carbon Nitride-Based Materials (C3N4) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.5 Calcium Sulfate (CS) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.6 Alumina (Al2O3) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract In the past three decades, many types of electrospun fibers have been developed and utilized for innumerable tissue engineering applications. Precisely, in the area of bone tissue engineering, composite bone grafts are being employed to overcome the limitations associated with autografts, allografts, and xenografts. A composite bone graft comprises a polymeric osteoconductive matrix and an osteoinductive or osteogenic material. Moreover, polymeric and ceramic-based engineered nanocomposite fibers developed through electrospinning are relatively similar to composite bone grafts. The polymers offer biocompatibility, resorbability, and flexibility to the nanocomposite fibers. In addition, polymeric matrices provide structural support for cell adhesion and subsequent cellular processes leading to

S. Sowmya Department of Periodontics, Amrita School of Dentistry, Amrita Vishwa Vidyapeetham, Kochi, Kerala, India N. Mathivanan, A. Chandramouli, and R. Jayakumar (✉) Polymeric Biomaterials Lab, School of Nanosciences and Molecular Medicine, Amrita Vishwa Vidyapeetham, Kochi, Kerala, India e-mail: [email protected]

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tissue formation, whereas bioactive ceramics provide strength and osteoconductivity to the nanocomposite fibers. An array of ceramic particles such as bioactive glasses, wollastonite, hydroxyapatite, tricalcium phosphate, tetracalcium phosphate, octacalcium phosphate, magnesium silicate, whitlockite, akermanite, carbon nitride, calcium sulfate, and alumina have been explored in bone-specific applications to gain successful outcomes. The current chapter emphasizes the properties, cytocompatibility, and in vivo performance of ceramic-incorporated nanocomposite fibers and fibrous scaffolds in bone tissue engineering applications. Keywords Bone tissue engineering · Ceramic · Electrospinning · Fibers · Polymer

1 Introduction Bone tissues are highly complex, vascularized, and have a predominant role in mechanical support and movement of the body [1]. Bone remodeling is an essential physiological process that involves resorption and reformation. At the time of an injury, bone defects smaller than the critical size can remodel and heal spontaneously [1, 2]. But in certain conditions, this may be followed by trauma, and tissue loss which may require surgical intervention. The large bone defects of tibia and femur, and critical-sized bone defects in weight-bearing bones cannot heal spontaneously and may require surgical intervention, therefore being some of the crucial medical conditions which prolong and cause inconvenience to the patients [3]. At present, majority of the large bone defects may require the Gold standard autologous or allogenic bone grafting [4]. These grafts contact, integrate, and deposit cells onto the surface of the native bone and provide the desired matrix for new bone tissue growth [5, 6]. The growth factors in the autograft bone and its osteoinductive properties also play an important role in new bone formation by providing signals to the osteoprogenitor cells [7]. Thus bone healing via autogenous grafting provides satisfactory achievement but the limitations include surgical complications, lack of donor harvesting sites, and donor site morbidity [8, 9]. On the other side, bone allografts may satisfy the availability of bone grafts but their efficacy and bioactivity are limited as compared to that of autografts [10]. To overcome the limitations associated with autografts and allografts, composite bone grafts are being developed and employed in clinical applications. An ideal bone graft or substitute should meet basic requirements such as porosity, biocompatibility, biodegradability, ability to support cell growth, etc. For the past four decades, researchers have been closely working with clinicians for developing newer bone regenerative materials and strategies to address patient needs; however, they have not been able to achieve significant outcomes. Electrospinning is a technique used to prepare nano to micron diameter fibers via a syringe and needle system which act as an electron where potential is applied and a contact is made with the collector, the counter electrode. As a result of the applied

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Fig. 1 The diagram shows the Electrospinning set-up

voltage, the charged polymer solution deforms into a Taylor cone (cone shape structure) at the tip of the needle and facilitates the flow of the solution toward the collector (Fig. 1) [11]. This method is popular for making fibers for tissue engineering applications since it is highly flexible, efficient, and cost-effective [12]. The surface roughness, morphology, and other properties of the fibers change according to the parameters. Such parameters control the fiber diameter, spinnability, and alignment of the fibers. The diameter is greatly swayed by the molecular weight and concentration of the polymeric solution, ambience humidity, and instrument parameters, especially the distance between the collector and the syringe, applied voltage, and flow rate of the solution. Spinnability predominately relies on solution viscosity, conductivity, and surface tension. The formation of fibers can be modified by using different collectors such as drums, flat plates, disks, and so on [13]. In the past three decades, electrospun fibers are being developed using various artificial and natural polymer solutions, by incorporating many osteoinductive and osteoconductive materials. Fibers made using these combinations act as an artificial bone extracellular matrix (ECM) and provide a favorable ambience for cells to adhere and proliferate without any external factors. This chapter emphasizes the work done in the past 5 years exclusively on ceramic-incorporated electrospun fibers for bone tissue engineering (Fig. 2).

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Fig. 2 Different ceramic particles incorporated electrospun fibers for bone tissue engineering (BTE)

2 Polymer/Ceramic Composite Nanofibers in Bone Tissue Engineering 2.1 2.1.1

Silica-Based Ceramics Bioactive Glass

The first bioactive glass (BG), termed 45S5, was invented at the University of Florida by Larry Hench with a composition of 24.5 wt% of Na2O, 24.5 wt% of CaO, 45.0 wt% of SiO2, and 6 wt% of P2O5 [14, 15]. It is a widely accepted material that not only assists in the formation of new bone cells but also bridges hard and soft tissues. After its implantation in vivo, it undergoes a morphological change into amorphous or crystalline CaP, on the surface of the implant which helps in integrating the implant with the native tissue. It also releases ions that can stimulate osteogenesis and angiogenesis. The rate of degradation of BG is controlled by changing its chemical constituents. Its structural properties can be modified by altering the chemical composition or its ambient temperature. The porous architecture of trabecular bone can be formed by BG. Researchers are focusing on overcoming the limitations of BG such as its brittleness and mechanical strength [16]. Keyvan Shiran et al. prepared nanofibres using poly (ε-caprolactone) PCL, different wt% of PCL/gelatin, PCL/gelatin/BG (5%), and PCL/BG (5%). The results showed that the mean fiber diameter was decreased to 167 nm with the addition of

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gelatin and BG into PCL. The BG-containing scaffolds showed apatite layer deposition after immersion in simulated body fluid (SBF) for 28 days. The tensile strength and Young’s modulus of BG-containing scaffolds were enhanced by about twofolds as compared to the non-BG scaffolds [17]. Aysen Akturk et al. made gelatin/silver nanoparticles (AgNps)/BG fibers with a mean diameter of 472 nm approximately [18]. Yaping Ding et al. prepared fibers of Poly(hydroxybutyrate) (PHB)/PCL/58S BG (60SiO2-36CaO-4P2O5). PHB, PCL, and BG were mixed in different weight ratios such as 1:0, 20:1, 10:1, and 5:1. The corresponding scaffolds showed a tensile strength of 2.8, 2.2, 2.9, and 1.9 Mpa, respectively. The scaffolds supported the adhesion and proliferation of MG63 osteoblast cells. The cytotoxicity assay showed that cell proliferation was higher in 1:0 and 5:1 ratio groups than control on day 3. Thus, they could be a suitable material for bone tissue engineering [19].

2.1.2

Wollastonite (CaSiO3)

It consists of 48% CaO and 52% SiO2 and has two structural forms, α- and β-wollastonite. Alpha wollastonite (pseudo wollastonite) has a pseudo hexagonal structure and beta wollastonite has a triclinic structure [20]. Wollastonite is active in body fluid and stimulates the formation of mineral deposition at the surface; has a good biocompatibility and biodegradation property. Its linear degradation helps in new bone formation and might assist in bone healing if it is used as a graft material. Fast degradation in physiological conditions and low mechanical strength are some of the limitations of wollastonite materials [21]. Beta wollastonite can be identified when wollastonite undergoes calcination between 800 and 1,130°C, whereas alpha wollastonite is identified when it is calcined above 1,130°C [20]. Abudhahir et al. fabricated PCL with copper-doped wollastonite electrospun fibrous scaffolds. Scaffolds immersed in fetal bovine serum (FBS) for 28 days showed higher protein adsorption whereas biomineralization analysis showed mineral formation at the surface of the scaffolds. When mouse MSCs were seeded onto the scaffold, osteoblast differentiation was confirmed by the expression of ALP, type I collagen, and Runx2. The scaffolds also showed an antibacterial effect against E. coli and S. aureus [22].

2.2

Calcium Phosphate (CaP)-Based Ceramics

The presence of calcium and phosphate in bone was discovered in 1769. Calcium phosphate (CaP) cement consists of calcium and phosphate ions which are cationic and anionic [23]. Many CaP-based porous scaffolds [24], coatings [25], and cements [26] have been developed till date. CaP is the major content in bone and is very well known and preferred for its biocompatibility and its ability to dissolve in body fluids. Biodegradation and ion release are also characteristic features of CaP-based ceramics. It controls cell adhesion, proliferation and favors new bone formation

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[27]. Cellular and non-cellular biomineralization processes increase the concentration of phosphate and calcium ions, thereby inducing the formation of bone minerals on CaP surfaces. They also affect the expression of osteoblast differentiation markers such as COL 1, BMP-2, OP [28], ALP [29], and Runx2 [30]. CaP plays a major role in cell adhesion and tissue formation through selective protein adsorption [31, 32], and mineralized tissue formation in bone regeneration [33]. Ca ions aid in bone maturation via calcification, whereas bone formation is through cell signaling. Ca induces mature cells via the formation of nitric oxide and promotes precursor cells for bone growth [34, 35]. By initiating the ERK 1/2 pathway, Ca induces the osteoblast to synthesize and mineralize the bone [36]. It also regulates the PI3K-Akt pathway in stretched osteoblasts [37]. Additionally, it modulates osteoclast emergence and resorptive functions [38]. Phosphorous exists in the form of PO43- in several parts of the human body which include protein, nucleic acid, ATP [39], and in physiological processes [40]. Phosphate manages the growth and differentiation of osteoblasts and osteoblast lineage through ERK 1/2 and IGF-1 pathways and increases BMP expression [41, 42]. Moreover, phosphate has a negative feedback interaction between the RANK-ligand and its receptor signaling. To facilitate the inhibition of osteoclast differentiation and bone resorption, phosphate regulates the ratio of RANK-ligand: OPG [43, 44]. CaP is reported to possess both osteoinductive and osteoconductive properties, wherein osteoinduction is a process to induce the progenitor cells to differentiate into osteoblast lineage, whereas osteoconduction is the formation of bone onto the material surface. Both these properties support cell adhesion and proliferation [45, 46].

2.2.1

Hydroxyapatite

Hydroxyapatite (HAp) (Ca10(PO4)(OH)2) is a naturally occurring CaP in human bones that consists of a large amount of inorganic compounds. With its CaP ratio of 1.67, it is the most commonly preferred bioceramic for bone regeneration applications [39, 47]. It is stable in human physiological conditions and body fluids [48]. HAp is utilized for a wide range of bone tissue engineering namely maxilla, mandible, alveolar bone, calvarium, and long bones [49]. It is combined with various biodegradable polymers namely PLA, PCL, PVA, PGA, gelatin, chitosan, silk fibroin, PLGA, PEG, and so on for making electrospun fibers and scaffolds [50]. Salifu et al. fabricated gelatin-HAp electrospun fibrous scaffold with varying concentrations of HAp from 0 to 25 wt%. 25 wt% HAp-gelatin scaffold at 20 kV showed the highest cell adhesion, cell proliferation, and ECM production. Fiber orientation improved the mechanical properties, with Young’s modulus and tensile strength in the range of 0.5–0.9 GPa and 4–10 MPa, respectively [51]. Furthermore, 40 wt% nano-HAp/gelatin showed a higher Young’s modulus of around 10.2 ± 0.8 GPa, as demonstrated by Catledge et al. [52]. Gautam et al. fabricated HAp/gelatin/PCL nanocomposite scaffold wherein gelatin-PCL nanofibers were

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electrospun followed by surface treatment with 1 wt% nHAp for different time periods. The fiber diameter and pore size for 10, 20, and 30 min of surface treatment were in the range of 505 ± 146 nm, 615 ± 269 nm, and 477 ± 108 nm, and 4.9 ± 1.6 μm, 4.7 ± 1.04 μm, and 1.9 ± 0.47 μm, respectively. Fibers treated with HAp for 20 min showed sufficient nHAp deposition and were stable for upto 2 weeks in PBS following which the fibers started degrading. Moreover, at the end of 4 weeks, nHAp was totally absent on the scaffold surface and the fibers were found to be merged. The scaffolds were cytocompatible with well-enhanced cellular proliferation [53]. Sani et al. developed an electrospun scaffold with modified chitosan (mCS) incorporated with PCL via ring-opening polymerization method to form a copolymer, and nano-HAp was added into it. The fibers were fabricated in various ratios of mCS-PCL and PCL (0:100, 5:95, 10:90, 20:80, and 30:70). The average diameters of the corresponding ratios were 862 nm, 325 nm, 314 nm, 280 nm, and 265 nm, respectively. Ninety percent PCL was chosen for further analysis due to the smooth, uniform, and narrow fibers with the absence of beaded structure. Further, to fabricate nanocomposite scaffolds, 1, 3, 5, and 10% (w/w) nano-HAp were added to 90% PCL. The average fiber diameter was 320 nm, 332 nm, 341 nm, and 385 nm and porosity was 89%, 86%, 85%, and 85%, respectively. Regarding degradation of the scaffolds, though it was enhanced with the addition of nano-HAp, this enhancement in weight loss was less than 10% at the end of 8 weeks. The bioactivity of 90% PCL with different ratios of nano-HAp was investigated after 1, 2, 4, and 6 weeks. The apatite layer formation was more pronounced in scaffolds containing nano-HAp in comparison to non-Hap-containing scaffolds. Further, cytocompatibility was assessed using NIH-3T3 mouse embryonic fibroblast cells. The scaffold with 3% nano-HAp showed higher metabolic activity with more than twofold increase, which indicates higher osteoblast differentiation in the presence of nano-HAp. Gene expression analysis of 3% nano-Hap-containing scaffold also showed higher OCN, ALP, and COL-1 expression in comparison to the non-HAp 90% PCL scaffold (Fig. 3), making HAp a suitable material of choice for bone tissue engineering [54].

2.2.2

Tricalcium Phosphate (TCP) (TCP; Ca3(PO4)2)

TCP, also known as bone ash, ample in Ca and P, is a widely studied material after HAp. It has a Ca:P ratio of 1.5, partitioned into two phases namely alpha and beta with a crystal structure of monoclinic and rhombohedral space group [55– 57]. α-TCP is formed at 1125°C or even higher and β-TCP is formed between 900 and 1,100°C [58, 59]. β-TCP is suitable and used for bone regeneration since its structure is stable and more biodegradable than α-TCP, but is less stable and degrades faster than HAp [60]. On the whole, it is used in the form of bone substitutes and bone cement [61, 62]. It is also biocompatible due to its good resorption properties [60]. β-TCP helps in promoting precursor cells like osteoblast and bone marrow stromal cells [23, 63]. These properties are due to its nanoporous structure which aids in cellular adhesion and biomineralization [64].

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Fig. 3 MG-63 cells OCN, ALP, COL-1 mRNA expression with no scaffold (negative control) and 90%PCL/mCS-PCL with 0 and 3% nano-HAp PCL90 and PCL90;n3) [54]. Reproduced with permission, Copyright Elsevier

In the PLLA/CMS/β-TCP electrospun scaffold, increasing β-TCP concentration leads to a decrease in mechanical strength and increases the hydrophilicity [65]. Topsail et al. made nanocomposite fibers of 10 wt.% Polyurathene (PU)/ 3 wt.% Chitosan/3 wt.% β-TCP, which had a good tensile strength and cell adhesion when cultured using L929 fibroblast cells. Amoxicillin drug was loaded and release was determined in vitro. It showed that TCP-incorporated nanocomposite fibers are suitable for bone regeneration [66]. Baykan et al. fabricated PCL/β-TCP with β-TCP concentrations of 10, 20, 30, and 40% individually. Thirty percent β-TCP scaffold showed promising mechanical strength making it suitable for in vivo subcutaneous implantation in the epigastric groin fascia of rats. In vivo analysis showed ectopic new bone formation [67]. Zhang et al. made a biomimetic nanofibrous composite membrane of gelatin/β-TCP. In vitro analysis using rat bone marrow stem cells showed osteogenic differentiation, which may be attributed to the activation of calcium-sensing receptor signaling. On in vivo analysis, implantation of the scaffolds into rat calvarial critical-sized defects for 12 weeks showed a good osteogenesis and higher calcium-sensing receptor (CaSR) expression than pure gelatin nanofibers [68]. Castro et al. made poly(lactic-co-glycolic acid)/β-TCP electrospun membranes with varying concentrations of β-TCP (5, 10, and 20 wt%). Fibers showed good mechanical strength for a higher amount of β-TCP upto 10%. Cell proliferation was analyzed upto 7 days, wherein 10 and 20% β-TCP scaffold showed agglomeration of β-TCP nanoparticles making it less suitable for the proliferation of osteoblast cells. Whereas, 5% β-TCP showed good osteoblast proliferation making it suitable for guided bone regeneration [69]. In a similar study, Ezati et al. fabricated a PCL/gelatin/chitosan/β-TCP electrospun composite membrane with different concentrations of β-TCP. Although 5% β-TCP fiber showed good mechanical strength, 3% β-TCP showed enhanced cell growth and type 1 collagen expression in comparison to the other variants, thus making 3% β-TCP scaffold a good material for guided bone tissue regeneration [70].

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199

Tetracalcium Phosphate (Ca4(PO4)2O; TTCP)

Tetracalcium Phosphate (TTCP) is a type of calcium phosphate with a Ca/P ratio of 2. It is formed at the temperature range >1,300°C in the system (CaO-P2O5). It shows reasonable reactivity and simultaneous solubility when combined with acidic CaP such as dicalcium phosphate anhydrous (DCPA, monetite) or dicalcium phosphate dihydrate (DCPD, brushite) [71]. It is used as an automatic setting cement for bone tissue engineering, which slowly forms HA under human body conditions [72– 74]. There are two approaches to synthesize TTCP. The first is direct solid-state synthesis wherein calcium carbonate (CaCO3) is mixed along with dicalcium phosphate anhydrate (CaHPO4) and heated at high temperatures between 1,450 and 1,500°C for 6–12 h. However, TTCP obtained from this method is not pure. Contrastingly, the wet process method provides high-purity TTCP which uses co-precipitation technique, followed by calcination at 1,500°C [75].

2.2.4

Octacalcium Phosphate (OCP: Ca8(HPO4)2(PO4)4
5H2O)

OCP has a Ca:P ratio of 1.33 and a triclinic crystal system with space group P1 [76]. Its properties include good biodegradability, osteoinductivity, and osteoconductivity. During hydrolysis, OCP quickly changes into HA, which is an irreversible process. A few animal and clinical studies on OCP have shown enhanced new bone formation as compared to other CaP ceramics [75–77]. Heydari et al. made fibers with PCL and varying concentrations of OCP (0, 5, and 10 wt%). The average diameter obtained was 1.55 μm ± 1.39 μm, 0.1–1.7 μm, and 0.52 μm ± 0.3 μm and the tensile strength was 3.19 ± 0.68, 4.34 ± 0.44, and 4.33 ± 0.32 (MPa) which showed that increasing OCP concentration leads to an increase in tensile strength. The cell viability and proliferation test for 28 days using HG-292 osteoblast cells showed that 10% OCP/PCL showed higher cell viability and proliferation at day 9, whereas no cell proliferation was observed for pure PCL scaffold [78]. Miyatake et al. investigated the effects of partial hydrolysis-induced compositional and structural changes of OCP on its osteoconductive properties when implanted in rat tibia for 56 days. For this, non-stoichiometric OCP was compared with slightly hydrolyzed OCP (low crystalline LC-OCP), and fully hydrolyzed apatitic product of OCP or biodegradable β-TCP ceramic. Results confirmed that the partially hydrolyzed OCP (LC-OCP) with a Ca:P ratio of 1.37 showed maximum enhancement in bone formation, with suppression of osteoclastic activity and reduced inflammation [79]. Wang et al. made a nanofibrous electrospun membrane of poly(3-hydroxybutyrate-co-4-hydroxybutyrate) P(3HB-co-4HB)/OCP, with varying (0, 5, and 10) wt% of OCP. The composite membrane with 10% OCP showed reduced fiber diameter with an enhanced average tensile strength of 2.73 MPa in comparison to the other two membranes. The nanofibrous membranes (5 and 10%) showed superior biocompatibility, better MSC proliferation and adhesion as compared to P(3HB-co-4HB) membrane. OCP-containing membranes also showed enhanced osteogenic differentiation of MSCs. The osteogenesis-related

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genes were upregulated and the gene expression was more in OCP membranes than in pure P(3HB-co-4HB) membranes. ALP activity and alizarin red (ARS) staining confirmed the biomineralization of MSCs on the nanofibrous membranes. Further on in vivo implantation in rat calvarial defects, P(3HB-co-4HB) with 10% OCP showed enhanced neobone formation with higher bone volume/total volume (BV/TV) ratio and bone mineral density (BMD) [80].

2.3

Magnesium-Based Ceramics

Magnesium (Mg) is an important trace element present in bones. It plays an important role in DNA stabilization, bone metabolism, and development. It aids in osteoblastic cells’ proliferation and adhesion, immunomodulation and angiogenesis [81]. Mg also helps in bone regeneration by increasing the production of Collagen X and VEGF in the bone cells. Mg-containing scaffolds have shown good cytocompatibility, enhanced osteogenic differentiation, and ALP activity. On the other hand, Mg deficiency leads to lower bone density and may result in osteoporosis [81].

2.3.1

Magnesium Silicate or Forsterite (Mg2SiO4)

Kharaziha et al. developed nanofibrous poly(ε-caprolactone) (PCL)/forsterite scaffolds using electrospinning. Forsterite (Mg2SiO4) nanoparticles were surface modified by esterification using dodecyl alcohol to provide an interfacial adhesion with a uniform dispersion of the nanoparticles in the polymer matrix. These modified forsterite nanoparticles were combined with PCL for the fabrication of nanofibrous scaffold. Furthermore, to improve the hydrophilicity, surface-modified forsterite nanoparticles and the nanofibrous composite scaffolds were treated in boiled water for the removal of dodecyl chains. Surface modification significantly improved the tensile strength and toughness of the scaffolds in comparison to the unmodified samples. Hydrophilicity enhanced the degradation rate, bioactivity, cellular adhesion, and proliferation on the nanofibrous scaffolds. Thus, surface modification with hydrolytic treatment proved to be an effective method for the fabrication of nanofibrous scaffolds with improved properties [82].

2.3.2

Whitlockite (WH) (Ca18Mg2(HPO4)2(PO4)12)

Whitlockite (WH) is high in Mg, Ca, and P and is the second most abundant mineral in the hard tissues. Based on the Mg2+ content, WH occupies 20–35 wt% of the bone mineral [83]. The high concentrations of PO43- and Mg2+ aid in bone regeneration through osteoblast activity and inhibit osteoclast pursuit without disturbing the natural HAp formation [84]. Yegappan et al. developed a WH-based injectable

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nanocomposite hydrogel with an angiogenic drug, dimethyloxallylglycine. The endothelial cells on the nanocomposite hydrogel showed a significant cellular migration with a capillary tube-like formation. The nanocomposite hydrogel also showed enhanced expression of RUNX-2, COL, and OPN proteins which was attributed to the presence of Ca2+, Mg2+, and PO43- ions in WH NPs, thereby favoring osteoblast differentiation and mineralization. Overall, the study confirmed the effects of WH and dimethyloxallylglycine on in vitro osteogenesis and angiogenesis [85]. Zhang et al. fabricated composite electrospun membrane composed of PCL and WH (with WH as 0, 5, 10, and 15%). The ionic release was determined which was found to increase with increasing WH content. Mg2+ in WH reached a release equilibrium within 8 days whereas Ca2+ and PO43- were released steadily for up to 12 days, providing an appropriate ambience for osteogenic differentiation and angiogenesis. In comparison to pure PCL, the membrane with WH showed higher mineral deposition on the seventh day and the Ca:P ratio was found to be 1.6. The concentration of WH significantly improved the mineralization ability. The membranes supported the osteogenic differentiation of BMSCs wherein gene expression levels were upregulated with an increase in WH content. In addition, the membranes assisted in angiogenesis by directly stimulating endothelial cell migration and upregulating VEGF expression in vitro. Subcutaneous implantation of the nanofiber membranes in SD rats confirmed its in vivo biocompatibility with more neovascularization in the PCL/WH membrane than pure [86].

2.3.3

Akermanite [AK] (Ca2MgSi2O7])

Akermanite (AK) is a calcium magnesium silicate-based bioceramic with favorable mechanical properties, good bioactivity, and biocompatibility [87]. Bafandeh et al. fabricated fibers of poly(vinyl alcohol) (PVA)/chitosan (CS)/AK composite containing 0, 0.5, 1, and 2 wt% of AK. In composite fibers with 2% AK, bead formation was observed which could be attributed to the high viscosity of the solution. The average fiber dimeters and tensile strength of composites containing 0.5, 1, and 2% AK were 89 nm and 10.6 MPa; 86 nm and 12.5 MPa; and 91 nm and 2 MPa, respectively. This result suggested that AK concentration >1% decreases the mechanical strength, thereby making 1% AK-containing fibers appropriate for bone tissue engineering applications. Glutaraldehyde crosslinking of composite nanofibres with 1% AK increased its mechanical strength to 12.5 MPa as compared to the fibers without crosslinking [88].

2.4

Carbon Nitride-Based Materials (C3N4)

Graphitic carbon nitride (g-C3N4) is a polymeric substance. It has a tris-triazinebased pattern which consists of carbon, nitrogen, and some impurities of hydrogen

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atom. It has electron-rich properties as compared to other carbon-based materials [89]. Due to the presence of nitrogen and hydrogen atoms, it has basic surface functional properties. Due to its light weight, simple fabrication, low cost, adjustable and adaptable properties, it can be used in hydrogels to enhance mechanical properties [90]. Owing to its remarkable electrical, optical, and thermal properties, it is used in electronics, sensors, and storage devices. Furthermore, its biocompatible and non-toxic nature makes it suitable for applications in the area of biomedical research. However, poor dissolution ability and large particle size of bulk g-C3N4 are some of its limitations [91]. Awasthi et al. fabricated PCL/g-C3N4 nanofibrous scaffolds with different concentrations of g-C3N4 (0.5, 1, and 2 wt%). The average fiber diameter was 0.8 μm for pure PCL, and 0.38, 0.2, and 0.1 μm for 0.5, 1, and 2 wt% of PCL/g-C3N4 nanofibrous scaffolds, respectively (Fig. 4). Though the addition of g-C3N4 nanosheets decreased the fiber diameter of the nanofibrous scaffolds, mechanical properties, biodegradability, and biocompatibility were enhanced [91]. The scaffolds were cytocompatible and supported cell adhesion and proliferation when seeded with MC3T3-E1 cells. The cell viability was higher on PCL/gC3N4 nanofibrous scaffolds as compared to pure PCL. This increase in cellular viability was consistent with the increase in culture duration as confirmed by CCK-8 assay. The nanofibrous scaffolds also showed enhanced ALP activity and mineralization which could be attributed to higher surface wettability and increased surface area-tovolume ratio of the nanofibrous scaffolds [91].

2.5

Calcium Sulfate (CS)

Calcium sulfate (CS) exists in three forms based on the levels of hydration which include CS dihydrate (CaSO4
2H2O) or Gypsum, CS hemihydrate (CaSO4
0.5H2O) or Plaster of Paris, also called bassanite and CS anhydrous (CaSO4). When gypsum is heated to above 100°C, it loses water through calcination resulting in the formation of Plaster of Paris or CS hemihydrate. The hemihydrate form exists in two

Fig. 4 SEM images (from left) (a) PCL only fiber (b)–(d) PCL/Graphite-C3N4 composite (0.5, 1, 2 wt% of Graphite-C3N4) nanofibrous scaffolds, (e)–(h) higher magnification images and at right (i) histogram of average fiber diameters of different composite fibers [91]. Reproduced with permission, Copyright Elsevier

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forms, namely α and β [92]. CS is biodegradable, osteoconductive, and biocompatible. However, its rapid rate of resorption, pH, and ion release behavior are unfavorable for bone regeneration. To overcome these negative effects, CS is often combined with other bioceramics or polymers for bone tissue engineering applications [93]. Shams et al. prepared Bioglass (BG) nanofibers (3, 5, 10, 15, 25, and 35 wt%) (BGF) which were added to CS cement forming a nanocomposite. When 15 wt% of BGF was added to 85% CS cement, the initial and final setting times were reduced from 12 and 27 min to 6 and 21 min, respectively. CS cement 85%/BGF 15% also possessed good mechanical strength of 29.54 MPa; however, thereafter with increasing BGF concentration, the mechanical strength was found to decrease. Apatite formation was seen on the surface of CS cement 85%/15% BGF nanocomposite after 14 days of soaking in SBF. hBMSCs viability was higher on CS cement/BGF than on BGF and CS cement. In osteogenic differentiation, CS cement 85%/15% BGF nanocomposite showed higher ALP activity and calcium content as compared to the other groups. Thus, CS cement 85%/15% BGF nanocomposite proves to be an ideal candidate for further in vivo investigations [94]. Zhou et al. fabricated electrospun fibers of PCL/CS hemihydrate (CSH) in different weight ratios of 90/10, 80/20, and 70/30, respectively. Figure 5 shows the SEM images of PCL/CSH fibers in different weight ratios [95]. Among all the four groups, i.e., PCL, PCL/10% CSH, PCL/20% CSH, and PCL/30% CSH, PCL/20% CSH composite fibers possessed the highest fiber diameter of 2.41 μm with the highest tensile strength and elastic modulus of 3.68 MPa and 7.4 MPa, respectively. Further on increasing the CSH content, tends the fiber diameter and mechanical parameters were found to decrease. PCL/20% CSH also showed good bioactivity when immersed in SBF with an agglomeration of clustered apatite formation at the surface of the composite fibers. EDX analysis confirmed the presence of Ca, P, and O, with a Ca:P ratio of 1.7, whereas, pure PCL fibers did not show any apatite formation, thus indicating poor bioactivity [95].

2.6

Alumina (Al2O3)

Alumina (Al2O3) is a bio-inert material with no bone-bonding ability and is used for the adsorption of toxic heavy metal ions such as arsenic or arsenate. It is being applied in orthopedic joint prostheses [96] and dental implants [97] due to its excellent mechanical strength, good corrosion resistance, and biocompatibility more specifically the bio-inert property [98]. However, owing to its non-bioactive nature, dental implant detachment from the surrounding bone, and loosening in joint replacements have been reported [99, 100]. Kurtycz et al. prepared PLA/Al2O3 nanocomposite fiber mats wherein Al2O3 nanopowder with an average particle size of 21.3 nm was incorporated into PLA. Increasing concentration of Al2O3 nanoparticles, ranging from 0 to 10 wt%, resulted in thicker fiber formation. Indirect cytotoxicity evaluation confirmed the non-toxic

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Fig. 5 SEM images of pure PCL and PCL/SC: (a) PCL; (b) PCL/10%CS; (c) and (d) PCL/20% CS (low and high magnification); (e) PCL/30% CS [95]. Reproduced with permission, Copyright Elsevier

nature of the nanocomposite fiber mats [101]. Toloue et al. prepared poly (hydroxybutyrate) (PHB)/Chitosan/Alumina (Al2O3) nanowires fibrous composite scaffolds with varying concentrations of Al2O3 nanowires (0–5%), 9% PHB, and 0–20% chitosan, respectively. The addition of chitosan and Al2O3 into PHB significantly enhanced the fiber diameters and tensile strength of the composite scaffolds. With increasing concentrations of Al2O3 especially with respect to 5% Al2O3, agglomerations were observed on SEM analysis which could lower the tensile strength of the composite scaffold. PHB-Chitosan/3%Al2O3 showed a tensile

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strength of 11.18 ± 1.24 MPa which was the highest among all other scaffolds. Further, adhesion and proliferation of MG-63 cells was also highest on the PHB-Chitosan/3%Al2O3 composite scaffolds as compared to PHB and PHB-Chitosan. This enhancement in cell number could be due to the high hydrophilicity and surface roughness of the composite scaffolds [102]. Esfahani et al. prepared polyamide 6 (PA6)/hydroxyapatite (HA) coated zirconia-toughened alumina (ZTA) nanocomposites wherein PA6/HA was coated through electrospinning. The coated ZTA nanocomposites showed homogenous, bone-like apatite formation after 21 days of immersion in SBF. SEM images confirmed the smooth homogenous mineral surface formation on the surface whereas the uncoated ZTA showed very minimal mineral formation. Both the coated and uncoated samples were cytocompatible and did not exhibit any toxicity toward MG63 cells [103]. These studies, thus, confirmed the biocompatible nature of Al2O3.

3 Conclusion A lot of research work is ongoing in the field of bone tissue engineering. Herein, we have reviewed the major work done exclusively on ceramic-incorporated nanocomposite fibers for bone tissue engineering. These ceramics include bioactive glasses, wollastonite, hydroxyapatite, tricalcium phosphate, tetracalcium phosphate, octacalcium phosphate, magnesium silicate, whitlockite, akermanite, carbon nitride, calcium sulfate, and alumina. Irrespective of the type and form of ceramic material, the polymeric/ceramic nanocomposite fibers showed enhanced mechanical properties with improved bioresorbability and flexibility. Moreover, cytocompatibility, cellular proliferation, and osteogenic differentiation were well enhanced in comparison to ceramic-free polymeric controls. Some nanocomposite fibers also demonstrated higher bone volume/total volume ratio and bone mineral density on in vivo implantation with an overall enhancement in bone regeneration and angiogenesis. Despite the significant accomplishments, ready-to-use nanocomposite fiber-based products are still lacking. Hence, future research should focus on product development and clinical trials to understand and address the challenges in a real clinical scenario.

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Adv Polym Sci (2023) 291: 213–228 https://doi.org/10.1007/12_2023_147 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 Published online: 22 March 2023

Electrospun Fibrous Scaffolds for Cardiac Tissue Engineering Nivethitha Ashok, Vignesh Krishnamoorthi Kaliannagounder, Cheol Sang Kim, Chan Hee Park, and R. Jayakumar

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Bioactive Nanoparticles Incorporated Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Growth Factors/Cytokines Incorporated Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Conductive Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract Worldwide, cardiovascular diseases and associated morbidity have been on a constant rise for the past few years. However, heart transplantation remains a distant option for many owing to demand-supply issues. A completely bio-engineered heart is still a far vision for the hoping patients. The evolving area of cardiac tissue engineering provides a chance to fabricate bioactive scaffolds which would aid in functioning and supporting the cardiac tissues by mimicking the mechanical, chemical, and biological properties of the native tissue. Numerous fibrous scaffolds have been studied and are currently under study in cardiac tissue

Nivethitha Ashok and Vignesh Krishnamoorthi Kaliannagounder contributed equally to this work. N. Ashok and R. Jayakumar (✉) Polymeric Biomaterials Lab, School of Nanosciences and Molecular Medicine, Amrita Vishwa Vidyapeetham, Kochi, India e-mail: [email protected] V. K. Kaliannagounder, C. S. Kim, and C. H. Park Department of Bionanosystem Engineering, Graduate School, Jeonbuk National University, Jeonju, South Korea Department of Bionanotechnology and Bioconvergence Engineering, Graduate School, Jeonbuk National University, Jeonju, South Korea Division of Mechanical Design Engineering, Jeonbuk National University, Jeonju, South Korea

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engineering. This review overviews the latest developments of the electrospun fibrous scaffold decorated with growth factors/cytokines, bioactive nanoparticles, and conductive fibers in cardiac tissue engineering that will help in the beneficial functioning of the cardiac system. Keywords Cardiac tissue engineering · Conductive fibers · Electrospinning · Electrospun fibers · Growth factors · Nanoparticles · Scaffolds

1 Introduction Cases of cardiovascular diseases have been witnessing a steep rise in recent years, giving a toss at the quality of life. This has garnered an immense focus on developing the best scaffold using tissue engineering techniques. Though researchers have developed numerous scaffolds which are both biocompatible and biodegradable, special consideration must be taken when fabricating scaffolds for cardiac tissues. The scaffold should withstand the pressure and retain its elasticity as it will be placed close to arteries and veins that continuously pump blood [1–5]. The human heart is one of the most complex organ with various cell compositions, such as resident cardiac cells, fibroblasts, endothelial cells, etc. and it has been reported that the heart lacks potential embryonic regeneration potential and is devoid of regenerative capacity. Presently, mitigation of cardiovascular diseases by drugs and other clinical practices is in use and has effectively aided in improving patients’ quality of life and survival rate [6, 7]. However, this is seen only as a temporary solution that’ll last only for a short duration. Post-graft surgical complications, a critical shortage of donor organs, and partial interference of pharmaceutical treatment courses have placed prominence on the scaffold-based therapeutic methods and concerning research approach. To regain functional significance and regeneration, delivering these engineered scaffolds directly at the site required is what will lead to the hasty recovery of the tissue. Recent advanced delivery approaches, including fibrous patches and injectable hydrogels, have garnered significant attention owing to their flexibility, compliance, and versatility. Fibrous scaffolds having multiscale dimensions and conductivity, incorporated with bioactive nanoparticles and growth factors, have provided many opportunities to expand research and understanding [8– 10]. Electrospinning has become an effective and emerging strategy in fabricating fibrous scaffolds through different techniques and combining various other materials [11–14]. Fibrous scaffolds are highly favored out of the many existing, owing to their unique property of mimicking the natural ECM and its components [15, 16]. The electrospinning technique employed in the fabrication of fibrous scaffolds presents a wide array of advantages: increased porosity and poreinterconnectivity, larger surface areas, and a site for cell adhesion to grow and proliferate. In addition, the electrospinning scaffold could be easily tailored to the

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tissue-specific engineered scaffold by selective combinations of natural and synthetic polymers with targeted biomolecules like small bioactive molecules, localized delivery of growth factors, drugs, smart electro-active biomaterials, etc. Thus, with the development of cardiac tissue engineering, there has been a firm hope in restoring the regeneration potential of cardiac muscles, thus paving the way for an opportunity to achieve a permanent cure. In this review, we have elucidated different types of evolving electrospun fibrous scaffolds engineered with bioactive nanoparticles, growth factors, and electrically conductive biomaterials, which would find prospective applications in cardiac tissue engineering.

2 Bioactive Nanoparticles Incorporated Nanofibers Materials in the nanoscale that can induce a biological response upon interaction with cells, tissue, or proteins are termed “bioactive nanomaterials.” This has garnered immense attention over conventional materials owing to their bioactivities, such as mimicking bio-matrix, which aids in tissue regeneration, stimulating cell adhesion, differentiation, and bonding with soft and hard tissues [17–20]. Furthermore, these versatile properties of bioactive materials have unveiled the scope of altering cellular interactions and functions, thereby prompting a specific response from tissues, making it valuable in regenerative medicine, therapeutics, and diagnostics [21–26]. A scaffold developed by Venugopal et al., composed of PCL and collagen (type I and III), had a tensile strength of 7.79 MPa and a tensile modulus of 18 MPa, which was perfectly apt for mimicking a blood vessel conduit. Upon seeding of coronary artery smooth muscle cells, it was observed that these cells had a higher rate of proliferation on the scaffold [27]. According to Heydarkhan Hagvall, scaffolds with the combination of PCL with elastin, collagen, gelatin, or other natural polymers exhibited enhanced tensile strength in comparison with the other existing hybrid scaffolds [28]. In a study conducted by Kai and the group, it was observed that the PCL/gelatin composite fibrous scaffold had exhibited mechanical properties and anisotropic wetting characteristics, which were similar to the native cardiac anisotropy. Upon seeding cardiac myocytes from rabbit on the scaffold, the biological components and structured topography prompted greater cell attachment and alignment [29]. Protecting the cells from oxidative damage by effectively scavenging them from cellular ROS (reactive oxygen species) has been extensively achieved using cerium oxide nanoparticles (nCe). Recent studies have depicted the link between reducing the levels of ROS, which has helped suppressing cardiac hypertrophy. A study by Jain et al. has explained the fabrication of PCL-gelatin blend and nCe decorated PCL nanofibers by employing an electrospinning technique. This scaffold was found to be cytocompatible and indicated a sharp reduction in the levels of ROS when the primary cardiomyocytes seeded on the scaffold were subjected to induced H2O2 oxidative stress. The results of this study looked promising and suggestive of the potential application of this scaffold as a cardiac patch containing

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anti-hypertrophic and anti-oxidant properties [30]. One of the most interesting and emerging areas of research is marine biopolymers as a bioactive functional ingredient. Abundantly present in marine animals and seaweed, these hold a wide variety of functionality and beneficial biological activities. These marine biopolymers as cell proliferation scaffolds, release modifiers, and bioadhesives have been attracting a constantly increasing interest in the applications of the cardiovascular tissue engineering field, owing to their impeccable biodegradability, biocompatibility, and unique physicochemical properties [31].

3 Growth Factors/Cytokines Incorporated Nanofibers Growth factors (GFs) are multifunctional soluble molecules, usually a secreted protein or a steroid hormone, which have been extensively used in tissue engineering applications nowadays [32]. They extend a therapeutic assurance with their capability of regulating different cellular activities, mainly cell proliferation, migration, differentiation, and tissue healing [33]. Moreover, they act as signaling molecules between cells. However, their function varies; vascular endothelial growth factors (VEGF) promote blood vessel differentiation, whereas nerve growth factors (NGF) promote neuron differentiation [34]. Widely, research has been carried out to understand the significance of growth factors in cardiac tissue regeneration applications. These biomolecules enhanced angiogenic and antiapoptotic properties, specifically cardiomyocyte proliferation, extracellular matrix (ECM) remodeling, stem cell recruitment, and immunomodulatory functions [35]. The major contributors to cardiac regeneration include, but are not limited to, VEGF, fibroblast GF (FGF), platelet-derived GF (PDGF), insulin-like GF (IGF), epidermal GF (EGF) families, and so on [36, 37]. However, the systemic administration of these GFs appeared ineffective because of their short in vivo half-life, reduced bioactivity and stability, and poor availability at the target sites. Owing to this, repeated injections are required in higher dosages, which further results in adverse side effects and higher treatment costs [38]. Thus, using biomaterials offers a promising approach by delivering growth factors directly at the target site in a controlled and sustained manner, which overcomes the aforementioned challenges [39]. Some commonly used biomaterials are cardiac patches, injectable hydrogels, vascular grafts, and nanocarriers. Among these, injectable hydrogels have been extensively utilized for cardiac regeneration, but the intramyocardial needle can further damage the affected tissues, and high burst release of GFs has been observed, which are the drawbacks [40, 41]. An alternative approach is the electrospun nanofibrous patch, which is porous with an ECM-mimicking surface and supports the GFs to release in a controlled way (Fig. 1). This also enhances cellular adhesion and proliferation on the scaffolds. Another advantage of nanofibrous scaffold is the ability to easily incorporate the GFs during electrospinning [42]. Kerignard et al. developed a group of hybrid patches with diverse copolymers composed of degradable nanofibers incorporated with IGF and hepatocyte GF (HGF). This is further

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Fig. 1 (a) TGF-β3 release from PCL-PLGA nanofibers was regulated by varying the percentage of co-incorporated BSA, with 10% BSA meshes releasing considerably more bioactive TGF-β3 (n = 5; ^p < 0.05). (b) The electrospinning technique was used to fabricate aligned PCL-PLGA nanofibers containing 10% BSA, with and without TGF-β3. (c) TGF-β3 content and scaffold morphology of as-fabricated meshes revealed that the high-dose group had around twice as much TGF-3 as the low-dose group (n = 5; ^p < 0.05), but there was no significant difference in fiber diameter or growth factor loading efficiency between groups. (d) The bust release profile of the growth factor within the first 6 h of incubation, and (e) the sustained release profile of TGF-β3 from both low-dose mesh and high-dose mesh during 14 days (n = 5; *,^p < 0.05). Note: *p < 0.05, different from the previous time point; ^p < 0.05, the difference between groups. Reproduced with permission [42]

associated with a collagen membrane in different methods to prolong the GF release up to 3 weeks. The results show that the patches are cytocompatible, and enhanced proliferation of myoblast cells was noticed due to the presence of IGF and collagen. This indicates that these patches could be used for cardiac muscle regeneration [43]. In another study, Tambrchi et al. attempted to differentiate cardiomyocytes from human adipose-derived mesenchymal stem cells (Ad-MSCs) on a polycaprolactone (PCL)/polyaniline (PANI) nanofibrous scaffold integrated with either 5-azacytidine alone or combined with TGF-β. Ultimately, cardiomyocyte-specific gene expression levels revealed that the developed scaffold has the potential ability to differentiate into cardiomyocytes with the presence of both drug and GF [44]. The same group of

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authors developed an identical scaffold with a similar property with PCL/polylactic acid (PLA) as nanofibrous electrospun material. However, PLA has poor mechanical properties and gets degraded easily. Thus, it is blended with PCL to avoid each polymer’s drawbacks and attain an improved property [45]. An alternative finding by Sajjad et al. employed a combination of polyvinyl alcohol (PVA) and alginate sulfate (ALG-S) electrospun scaffold to conjugate TGF-β1, which can be delivered in a sustained manner for the application of various tissue engineering. Since TGF-β1 was utilized, the seeded MSCs can differentiate into a cardiomyocyte, which can be advantageous for cardiac regeneration [46]. Interestingly, Hariharan et al. studied the controlled release of VEGF from the PCL/gelatin/silk fibroin (SF)/ VEGF core-shell nanofibers to initiate the differentiation of MSCs into vascular smooth muscle cells (VSMCs). The results prove that the scaffold exhibits a morphological difference and enhanced cell proliferation due to the presence of VEGF. The functionalized nanofibers mimic the native ECM, which promotes the attachment and differentiation of MSCs, for vascular tissue engineering applications [47]. Cytokines are similar to GFs, which are small, secreted protein-based cell signaling molecules produced by almost every cell that control the growth, cellular activity, and movement of the immune cells toward the site of inflammation or infection [48]. They are described as immunomodulatory agents because of their capability to modulate or alter the immune response system. They are mostly used in cancer treatments to prevent or manage side effects during chemotherapy and as biomarkers for detecting diseases [49]. The major examples of cytokines are interleukins (IL), interferons (IF), and tumor necrosis factor-alpha (TNF-α). Others include granulocyte-macrophage colony-stimulating factor (GM-CSF), thymic stromal lymphopoietin (TSLP), leukemia inhibitory factor (LIF), and so on, which are used for different applications. Depending on their functions, cytokines can be pro-inflammatory, which accelerate immune responses (e.g., IL-1β, IL-6, TNF-α), and anti-inflammatory, which prevent inflammation by suppressing immune cells (e.g., IL-4, IL-6, IL-10) [50, 51]. During tissue damage, the immune system plays a significant role. The injury initiates coagulation and acute inflammatory response, and the macrophages in tissues react to these injuries, further providing control over tissue homeostasis. But there are specific macrophages that should be functioned at each stage of regeneration and, in addition, to ease the rate of inflammation, which cytokines can control. Cytokines can induce M2 differentiation and promote epithelialization by the proliferation or differentiation of progenitor cells, which are essential in tissue regeneration [52, 53]. Specifically, in cardiac diseases, myocardial infarction (MI) generates continuous over-reactive inflammatory responses, resulting in heart failure. In detail, pro-inflammatory cytokines such as TNF-α and IL-8 trigger subsequent immune responses and cytotoxic injuries. However, antiinflammatory cytokines, for instance, IL-10, can assist in cardiac repair by initiating ECM remodeling and stabilizing the matrix caused by the inflammatory responses [54]. Though the cytokines have a very short half-life in vivo as GFs and continuous high-dose administration of these systemically can make the tissues cytotoxic. Thus, they can be easily administrated and delivered to the target site using biomaterials as

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a delivery tool, lowering the side effects and improving efficacy. Biomaterial-incorporated cytokines can also maintain their biological activity and bioavailability in in-vivo circumstances [36].

4 Conductive Nanofibers In a normal heart, the cardiac conduction system (CCS) comprises electrical nodes, conductive bundles, and fibers that start, propagate, and conduct electrical currents for atrial and ventricular contractions [55]. Effective heart function necessitates rhythmic atrial and ventricular contractions, which rely on the spontaneous stimulation of pacemaker cells, followed by the signaling and production of action potentials (AP) by individual cardiac myocytes (CMs). From the cellular point of view, the electrophysiological properties of CMs arise due to the distribution and concentration of charged ions (sodium, calcium, potassium, and chloride) between the cell membrane. The unequal distribution of charged ions between the cell membrane of CMs causes membrane potential, and the change in membrane potential due to ion transfer is the AP. The AP consists of two phases: depolarization (i.e., membrane potential rising phase) and repolarization (i.e., membrane potential falling phase). These electrical polarization cycles are regulated based on the concentration gradient of ions and the cell membrane’s voltage and receptor-gated ion channels. Briefly, these ion channels play a vital role in electrical signal coupling through cell– cell communication via transferring ions and small molecules. Thus, any functional abnormalities in the regulation of ion channels affect the normal AP generation and conduction and lead to myocardial infarction [56]. Understanding the brief cardiac electric conduction mechanisms and scaffolding, a tailored conductivity using natural or synthetic materials with biomimetic microenvironments could be a promising strategy for cardiac regenerative therapy. Looking into conductive biomaterials, there are three major categories: carbon nanomaterials (graphene, carbon nanotubes (CNT), carbon nanofibers (CNF)), conductive polymers (polyaniline (PANI), polypyrrole (PPY) and poly (3,4-ethylenedioxythiophene) (PEDOT)), and metal nanomaterials (gold (Au), silver (Ag), molybdenum (MoS2), and selenium (Se)) [57]. Furthermore, the application of conductive materials in biomedical implants (flexible, intelligent sensors, and tissue regenerative smart patches) has started to widen due to the need for conductive microenvironments for the healthcare monitoring system and to achieve the functional reconstruction of bone, neurons, and cardiac tissues. Recently, S. Mombini et al. developed PVA/CS/CNT conductive composite nanofibrous scaffold. In this work, different concentrations of the CNT (1, 3, and 5 wt%) reinforced PVA/CS/CNT with an average of 292 ± 5.5 nm fiber diameter with a conductivity of 0.041 mS/cm, which is similar to the native myocardium conductivity range from 0.016 mS/cm (longitudinal) to 0.05 mS/cm. Among different concentrations of CNT 1 wt% (PVA/CS/CNT1), the enhanced mechanical strength of around 100% more compared to pristine PVA/CS scaffold and in vitro

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studies revealed cell viability and proliferation of mesenchymal stem cell (MSC). Differentiation gene expression studies also evidenced the >4-fold increased relative gene expression of cardiac markers (Nkx2.5, Troponin I, and β–MHC) for the PVA/CS/CNT1 scaffold with electrical stimulation (ES) compared to control [58]. Recently, the same group conducted a similar study with a multiwall carbon nanotube (MWCNT) reinforced into CS/PVA NFs [59]. In this work, the addition of 2 wt% of MWCNT to CS/PVA showed the conductivity of 1.17 mS/cm and in vitro differentiation studies with unrestricted somatic stem cells (USSCs) in the presence of 5-azacytidine (small molecules aids in cardiac differentiation) and with pulsed ES applied at 1.25 Hz of frequency, and 10 V of voltage showed an enhanced expression of troponin I, CX43, and β-MHC genes to 172, 5.3, and 64-times to undifferentiated cells. Another study by Shokrael et al. also evidenced that 2 and 3 wt % MWCNT loaded PU NFs showed the conductivity of 0.054 and 0.47 mS/cm, respectively, and the cell viability and morphology of H9C2 and HUVECs (umbilical vein endothelial cells) [60]. M. Tashakori-Miyanroudi et al. developed a cardiac scaffold with collagen-containing carbon nanofiber (Col-CNF) and investigated the regeneration of myocardial injury in animal models [61]. Herein, in male rats left anterior descending blockage-induced ischemia region was developed by surgical procedure. After 4 weeks, followed by the implantation of the Col scaffold and Col-CNF scaffold, the Col-CNF group evidenced the significant upregulation of α-actinin (indicates a formation of new cardiac cells), CD31 (indicator from endothelial cells involved in angiogenesis), PDGFRα (indicated a formation of epicardial-derived germ cells), and downregulation of caspase 3 (indicator of cell death via apoptosis) expression. Recently, CNF-reinforced printable alginate (Alg)/Ge conductive scaffold was reported by A. Serafin et al. [62]. Their results showed that CNF incorporation improves the shear-thinning behavior for printing, mechanical strength, and conductivity up to 534.7 ± 2.7 kPa and 0.41 ± 0.02 mS/cm. The other study with 0.05 wt% GO incorporated polyethylene terephthalate (PET), conductive core-shell NFs with a conductivity of 0.0013 mS/cm, and in vitro studies HUVECs and H9C2 cells proliferation and morphological elongation were noted [63]. Similarly, the other study reported the conductive MoS2/nylon NFs and the in vitro studies with embryonic cardiac cells evidenced the scaffolds cytocompatibility and the gene expression. Immunostaining studies demonstrated the maturation and upregulation of cardiac functional genes such as GATA-4, c-TnT, Nkx2.5, and α-MHC (cardiac marker) in the MoS2/nylon NFs compared to bare nylon NFs [64]. K. Huang et al. recently reported a flexible, intelligent array patch based on polyurethane (PU) and nanofiber/PPy composite for cardiac strain sensing monitor and drug delivery [65]. In this work, the waterborne polyurethane (PU) with polycaprolactone-gelatin (PCLG) nanofiber (NFs) composite films were fabricated, on which 300 nm gold layered was sputtered. Subsequently, porous dexamethasone (DEX) drug-PPY electrode was surface coated using an electroplating technique. It showed cardiac monitoring with micro-stain as low as 0.1% with the response within 0.15 s and controlled drug delivery of 54.9 μg cm-2 DEX within 1 min for applied 0.6 voltage (V). Recently, an interesting finding was noted by Y. Liang et al. with PPY-incorporated silk fibroin (SF) fibers for cardiac tissue regeneration [66].

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This study evidenced that among a different concentration of PPY-SF NFs, a 7% SF scaffold with 30 wt.% of PPY nanoparticles to SF (7%PPY30) showed robust mechanical properties with the highest conductivity of 0.52 mS/cm. Further, this study showed enhanced electromyography (EMG) signal detection from the single muscle mode with 7%PPY15 compared to 7%PPY0 (control). Also, in vitro studies with CMs revealed that the PPY hybrid mat supports the CMs contraction with elongated organized sarcomeric striations (Fig. 2). Similarly, M. Khorram et al. developed a natural polymer-derived chitosan (CS)/collagen (COL)/polyethylene oxide (PEO) composite blend with different weight % of PPY. Herein, COL was employed, as it is one of the prime proteins of the extracellular matrix (ECM); meanwhile, PEO is known for its resistance to protein absorption and platelet adhesion, a much needed characteristic for cardiac implants. In this work, PPY/CS/COL/PEO blend solution was electrospun and crosslinked with glutaraldehyde to control the degradation of natural polymers and achieved conductivity up to 1.64 mS/cm with better cell adhesion, growth, and proliferation properties [67]. Also, Fakhrali et al. reported the PPY-coated electroconductive poly(glycerol sebacate) (PGS)/polycaprolactone (PCL) nanofibrous structure that could be applicable as the cardiac patch. In this work, a different molar concentration of PPY coated on the surface of the fiber was evaluated and reported a 0.05 M PPY-coated PGS/PCL NFs exhibits 1 mS/cm of conductivity with in vitro biocompatibility of fibroblast cells [68]. Similarly, the mechanically robust conductive nanofibrous patches were developed using the composite (PU) and (PLA) solution and coated by interfacial polymerization of polyaniline (PANI) and were investigated electroconductivity of the nanofibrous patch and the proliferation of hy926 endothelial cells with electrical stimulation by Mousa et al [69]. The other study reported PCL gelatin (Ge)/PANI conductive scaffold for cardiac tissue regeneration by O. Gil-Castell et al. [70]. Herein, blending of Ge with PCL aids in the tailoring of degradation rate and provides arginyl-glycyl-aspartic amino acid sequences (RGD), which act as a biochemical signal for cell adhesion, migration, and proliferation. A different weight % of PANI nanoparticles was incorporated with PCL/Ge scaffold for conductivity. The scaffold exhibited low cytotoxicity and enhanced proliferation of HL-1 cells, an immortalized line from murine atrial cardiomyocytes. The analysis of macrophage M1 and M2 profiles with PCL/Ge/PANI scaffold evidenced lower expression levels of the pro-inflammatory cytokine TNF-α for the M1 profile and CD206 for the M2 profile than control cells. The novel multiscale conductive user design oriented layer-specific fibrous scaffold was developed by Lei et al. using melt-based electrohydrodynamic (EHD) printing [71]. The multiscale conductive scaffold consists of 9.5 ± 0.8 μm PCL microfibers mimicking collagenous fiber structure and 470 ± 76 nm PEDOT-PEO submicron conductive fibers mimicking conductive Purkinje fibrous system of the native cardiac ECM. The design and thickness of the scaffold could be controlled as needed for implant condition is the advantage of this EHD printing technique. Further, a pattern with varying angles was developed, and H9C2 MCs cells, proliferation, cellular orientation (Fig. 3), and quantification of cardiomyocyte-specific protein expressions and

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Fig. 2 Cytoskeleton and cardiac-specific protein expression of neonatal rat cardiomyocytes (NRCMs) on ES(PPy, SF) Mats. (a–d) F-actin staining of NRCMs cultured on pure SF and PPY-SF mats on Day 3. The arrows represent stress fibers. (e–g) Immunostaining of sarcomeric α-actinin (red) and nuclei (blue) of NRCMs on 7%PPy15, 7%PPy0, and the culture plate on Day 10. The corresponding FFT image is presented at the bottom left. (h–j) Enlarged view showing the

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beating behaviors of primary cardiomyocytes on the printed scaffolds were evaluated.

5 Conclusion In recent years, significant advances in cardiac tissue engineering have paved the way for substantial contributions in developing and fabricating novel electrospun fibrous scaffold materials. Though we have been able to address minor problems by developing cardiac patches and vascular grafts, etc., the bigger problems are yet to be addressed. Consequently, to overcome the intricate interplay between various factors influencing cardiovascular tissue engineering, electrospun fibrous scaffolds, and various additives like bioactive nanoparticles, growth factors have become necessary in aiding the functionality and supporting the regenerative capacity of the heart. These factors underpin the need to understand the features that can be incorporated in fabricating and designing efficient scaffolds.

 ⁄ Fig. 2 (continued) expression of CX43 (green) and sarcomeric α-actinin (red). Double-headed arrows showed the direction of the sarcomere. (k–m) α-actinin organized z-band shown by arrows. (n–r) Quantification of α-actinin coverage and CX43 coverage, sarcomere length, Z-band width, and cell aspect ratio (n = 3). (s) Sarcomere structural illustration. *p < 0.05, **p < 0.01. Reproduced with permission [66]

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Fig. 3 Hybrid EHD printing technique to construct micro/nanoscale fibers and their impact on H9C2 cell adhesion and proliferation. (a) Pictorial representation of the fibrous cardiac extracellular matrix (ECM) and the gradual transformation of cellular orientations in heart tissue. (b) The construction of microscale PCL fibers and sub-microscale PEDOT: PSS-PEO fibers utilizing melt-based EHD printing and solution-based EHD printing, respectively. (c) The PCL and PEDOT: PSS-PEO fiber diameter. (d) The microscale PCL fibers fabricated by EHD printing and (e) their impact on the H9C2 cell attachment. (f) The sub-microscale PEDOT: PSS-PEO fibers fabricated by EHD printing and (g) their impact on the H9C2 cell attachment. (h) The microscale PCL fibers and sub-microscale PEDOT: PSS-PEO fibers that were perpendicularly oriented during EHD printing, and (i) their impacts on cell attachment. Reproduced with permission [71]

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Adv Polym Sci (2023) 291: 229–286 https://doi.org/10.1007/12_2022_130 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 5 July 2022

Electrospun Nanofibrous Scaffolds for Neural Tissue Engineering Sheersha Pramanik and Vignesh Muthuvijayan

Contents 1 2 3 4

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . An Overview of Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . An Initiation to Tissue- Engineered Nerves . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Electrospun Scaffolds for Neural Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1 Natural Polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.2 Synthetic Polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 Conclusion and Future Perspectives . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract The repairing procedure in the nervous system is intricate and brings significant difficulties to investigators. The complication of the structure and function of the nervous system, and its slow rate of regeneration, make it further challenging to treat in comparison to other human tissues when damage takes place. Furthermore, the existing therapeutic modalities comprising the utilization of conventional grafts and pharmacological actives have numerous shortcomings and cannot completely rehabilitate injuries to the nervous system. Though the peripheral nerves regenerate to some extent, the consequent findings are not satisfactory, especially for severe injuries. The continuing functional loss owing to inadequate regeneration of the nerve is a significant problem around the world. Therefore, a successful therapeutic approach to bring functional rehabilitation is immediately required. Lately, tissue engineering methods have enticed many scientists to lead tissue regeneration efficiently. Majorly, the electrospinning method has come into the limelight for the fabrication of the scaffolds as they can develop fibrous meshes with fiber diameter in nanoscale dimensions. The electrospun

S. Pramanik and V. Muthuvijayan (*) Department of Biotechnology, Bhupat and Jyoti Mehta School of Biosciences, Indian Institute of Technology Madras, Chennai, Tamil Nadu, India e-mail: [email protected]

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substrates have a high prospective in mimicking the structure of the natural extracellular matrix. These produced fibers can be random or oriented to assist the extension of neurite via contact guidance. In this book chapter, we have demonstrated the principal parameters necessary for suitable electrospinning. Further, we have discussed the recent advances of electrospun polymeric scaffolds in neural tissue engineering. Finally, the challenges and future potentialities have been addressed. Keywords Electrospinning · Nerve injury · Nerve tissue regeneration · Polymer · Scaffolds

1 Introduction Traumatic damage to the nervous system impacts a huge population globally every single year. Serious injuries that induce significant nerve defects can guide to lifetime impairment in sufferers, decreased life quality, and financial and social problems [1, 2]. Despite the inherent regenerative capability of the adult peripheral nervous system (PNS), natural nerve reconstruction in PNS has been constantly linked to inadequate functioning results if no medicinal intervention is endorsed [3]. Contrasting PNS neurons, neurons of the central nervous system (CNS) fail to regenerate by themselves after damage to the nerve due to the inhibitor’s presence inside myelin and the glial scar’s development [4, 5]. Hence, clinical administration of CNS injuries, like, spinal cord injury, has less potential, while clinical remedy of peripheral nerve damages has obtained substantial advancement through combined efforts for a long time [6], primarily based on signs of progress in tissue engineering recently. A novel neural tissue engineering area has firmly become evident, promptly matured, and captivated the focus of investigators and practitioners. The advancement of tissue-engineered nerve grafts enables the substitution of conventional neurorrhaphy via the technique of neural grafting to address nerve damages. As already discussed, the regenerative capacity of the mammalian nervous system is usually restricted. Thus, patients having injuries/traumas in the nervous system commonly endure damage to sensory/motor function and neuropathic pains. For the purpose of facilitating neural regeneration, numerous therapeutic advances have been endeavored. Direct end-to-end surgical rejoining is a standard procedure of neural transection injury therapy when the gap in the injury is insignificant in the PNS. Usually, nerve autograft is believed as the ‘gold standard’ for connecting bigger gaps in nerve defects. Nevertheless, the lack of grafts from the donor, the prospective function loss at donator sites, and the necessity of several surgeries are some of the causes that restrict the autograft’s utilization. Other grafts, like, allografts and xenografts [7–9], are furthermore taken into account to replace autologous

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neural grafts. Nevertheless, these therapies are hindered due to rejection by the immune system and the probability of ailment transmission. Presently, traditional medication does not have efficient and prosperous therapies for these damages/injuries, and symptom treatment is usually the most satisfactory solution. For the purpose of reversing it and achieving the neuron’s functional reconnection, tissue engineering presently opts for the utilization of cells, biomaterials, and biomolecules to mimic the primary anatomy as far as possible. Both natural and synthetic biomaterials have displayed constantly promising results in neural tissue engineering, involving the outgrowth of neurites, human nerve stem cell differentiation, and bridging the nerve gap [10, 11]. For all of the aforementioned reasons, tissue-engineered scaffolds might assist as a substitute preference for implantation to promote nerve damage repair. The physicochemical characteristics of artificial grafts can be customized depending on applications. To decrease the incidence of immune responses, materials possessing biodegradable and biocompatible properties are often utilized. Appropriate structural and biochemical cues may be furnished to advance tissue reformation by governing various factors, like morphology, framework, and components of the scaffold. Such artificial grafts might be even altered to offer axons, a lenient substrate, to permeate the distressed region in tissue regeneration of CNS [12]. Imitating the structure of the native extracellular matrix (ECM) has been considered a general technique in tissue engineering. The ECM performs a significant role in monitoring cellular activities by affecting cells through various topographical cues and biochemical signals [13, 14]. ECM has two primary constituents, namely, polysaccharides and fibrillar proteins. Intrinsically, nanofiber frameworks have been utilized widely as prospective tissue engineering scaffolds. It is usually assumed that an ECM’s close mimic will deliver more contributory circumstances for activities of cells like adhesion, migration, proliferation, and differentiation [15]. In the meantime, nanofibers display an exceptionally high ratio of surfacearea-to-volume. The aforementioned characteristic promotes the biochemical’s (proteins, drugs, and nucleic acids) released from the fibers, acting as a carrier [16– 18]. Furthermore, the enormous surface raises the contact region among cells and the fibers; as a result, improving the cell’s chemical absorption. Lately, electrospinning, one of the limelight approaches to fabricating nanofibrous constructs, has fascinated enhanced concern for regenerating tissue because of the aforementioned properties. In this chapter, we have endeavored to throw some light on the fabrication of nanofibrous scaffolds via the electrospinning method. Further, we have addressed the recent investigations on electrospun nanofibrous constructs for neural tissue engineering. Finally, the challenges and future perspectives for the progress of electrospun nanofibrous constructs for engineering functional nerve tissues are also discussed.

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2 An Overview of Electrospinning The technique of electrospinning is a straightforward procedure of manufacturing constant fibers having micron to sub-micron regime diameters [19, 20]. It is an appealing procedure for the polymeric biomaterial’s processing into nanofibers. The resilience of the approach is apparent from the comfort of the formation of fibers appertaining to an extensive scope of materials comprising polymers (natural and synthetic), composites, and ceramics [21]. Similar to the traditional fiber spinning procedure, factors like concentration, viscosity, and flow rate of the polymer solution might be controlled during fabrication to modify the fiber’s dimension. Moreover, due to the electrospinning’s nature, the electrical field’s strength and field pattern might be varied to tailor the resulting construct’s framework and morphology. At present, there are two conventional electrospinning arrangements, horizontal and vertical. Electrospinning is conducted at room temperature with atmospheric conditions. The standard establishment of the electrospinning device is displayed in Fig. 1a, b. Fundamentally, an electrospinning apparatus comprises three significant elements: an elevated supply of voltage power, a spinneret, and a grounded collector (usually a plate or rotating mandrel). The electrospinning principle is to employ an electric field to attract a solution of polymer or melt from an aperture to a collector. The electrospun nanofibers varying from 50 to 1,000 nm can be generated by implementing an electrical potential to a solution of the polymer [22, 23]. The polymeric solution is retained at the capillary tube’s tip because of its surface tension and electrical voltage employed to deliver a charge to the solution. A force is generated in the polymeric solution due to mutual repulsion of charge, which is directly inverse to the surface tension of the polymeric solution. An enhancement in the electrical potential primarily guides to stretching the hemispherical surface of the polymeric solution at the capillary tube’s tip to create a conical form referred to as Taylor’s cone [24]. An additional enhancement makes electric potential achieve a critical value to overpower the forces of surface tension to induce the jet formation discharged from the Taylor cone’s tip. The charged polymeric solution experiences unsteadiness and slowly thinned in air principally due to extension and evaporation of solvent [25]. Eventually, the generated randomly orientated nanofibers are gathered on a metallic stationary/rotating collector [26].

3 An Initiation to Tissue- Engineered Nerves Regeneration and functional restoration of nerves is an exceptionally complicated procedure that includes multivariable mechanisms that are needed to be realized at the levels of molecule, cells, and tissue. Depending on the assembled understanding of neural regeneration, which comprises long-term investigations of the researchers, an advanced theory on “biodegradable tissue-engineered nerves” has been

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a) Syringe Polymer solution

Spinneret

High Voltage

Fibers Collector

b) Collector

Syringe Polymer solution Spinneret

Fibers High Voltage

Fig. 1 Schematic representation of the arrangement of electrospinning device (a) standard vertical set up and (b) horizontal arrangement of electrospinning device, reproduced with permission from [27], copyright 2010, Elsevier

progressively accepted. A tissue-engineered nerve is implanted inside the damaged nerve to connect the defects of the nerve. An ideal neural graft must have the subsequent characteristics: (a) Controlled rate of biodegradability in the human body; (b) biocompatibility with neural cells and tissue; (c) decreased immunogenicity; (d) promotes angiogenesis and exchange of metabolism; (e) guides to reduces the formation of the scar; (f) huge biomaterials accessibility for graft construction. The aforementioned fundamental conditions are generally recognized as guidelines for tissue-engineered neural graft development [28]. Xiaosong Gu postulated an innovative theory concerning nerve regeneration which is illustrated and discussed in Fig. 2.

Fig. 2 An innovative theory about the construction of tissue-engineered nerves. (a) Diagrammatic representation of a nerve scaffold consisting of a porous, biomaterial-associated nerve guidance conduit (NGC) and intraluminal filaments. (b) Diagrammatic representation of the regeneration process of damaged neurons with the assistance of a nerve construct, which permits vascularization from the stump of the nerve and infiltration of trans-wall vessels of blood to furnish nourishment for migration of Schwann cell and extension of axons and to promote target muscle’s reinnervation. (c) Diagram depicting that a nerve

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Like every other tissue-engineered scaffold, tissue-engineered neural grafts consist of a nerve construct (or template) with additional molecular and cellular cues. During the last few decades, great efforts have been made to select the biomaterial and fabricate the scaffold to optimize the nerve scaffold generation. Various biomaterials (synthetic and natural) have been widely used [29], and also new inorganic compounds with unusual surface characteristics have begun to draw the researcher’s attention besides conventional biomaterials [30–32].

4 Electrospun Scaffolds for Neural Tissue Engineering 4.1 4.1.1

Natural Polymers Collagen

Collagen proteins are made up of three alpha strands that come together to form a polymer. Almost 28 types of collagen have been identified. The amino acid structure is Gly-X-Y-, with glycine being required throughout every third location to permit collagen’s tight packing arrangement. Collagen is determined by the existence of 4-hydroxyproline, which is produced after post-translational modifications [33]. Collagen’s adaptability is due to its extensive dispersion throughout the body. Low antigenic and allergic characteristics, cytocompatibility, decent water uptake, ease of isolation from various sources, and the flexibility to alter biomechanically through cross-linking are some of the advantages of collagen [34]. Electrospinning, alongside other techniques, is used to create collagen-based structures. When collagen absorbs water, it loses its architectural and tensile integrity. Collagen, including other biopolymers and synthetically-derived polymers, can be utilized to adjust features such as structural rigidity, permeability rate, compression modulus, cell number, and cellular metabolic function to avoid this [35]. Numerous tissues, notably blood vessels, skin, bones, tendons, and cardiovascular tissues, contain collagen, especially type I. Due to this, collagen type I frameworks have been frequently employed as a substitute extracellular matrix (ECM) in the growth and repair of tissues. Electrospinning produces a matrix of nanosized or microsized fibers that mimics the fibrous character of the original ECM. These fibrous matrices are often used as substrates for tissue regeneration. Adjustments in electrospinning variables (solvent, polymer, solution concentration, voltage,  ⁄ Fig. 2 (continued) construct safeguards and leads to regrowth of axons and migration of Schwann cells. (d) Scheme displaying favorable basal lamina tubes reinnervation in the distal nerve stump via corresponding motor, sensory, and sympathetic axons. (e) Diagram depicting that chitooligosaccharides (COS) improve neural regeneration and promote proliferation of Schwann cell by the axis of miRNA-27a/ FOXO1, reproduced with permission from [28], copyright 2015, Springer

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needle-ground length) can modify scaffold characteristics such as fiber diameter, pore-volume, interfiber length, and fiber organization, resulting in a wide range of scaffolding structures that can imitate the natural ECM arrangement. Such matrices have already been extensively used for neural tissue engineering because electrospinning can yield substrates with elevated amounts of alignment, mainly axially oriented and unidirectionally oriented scaffolds [36]. Collagen is extensively explored as a biocompatible material for neural biomedical applications, and as a consequence, many collagen-based neuronal for directed peripheral nerve restoration are commercially accessible. Collagen is presently the only biomaterial that has been licensed for clinical use in neural tissue creation. In 43% of patients, NeuraGen® was found to be highly efficient in peripheral nerve repair [37]. In its initial clinical trial, Neuromaix®, a potentially commonly produced collagen nerve guidance, demonstrated exceptional success in bridging extensive nerve gaps. Collagen-based nerve conduits were clearly the strongest compatible nerve scaffolds now accessible in healthcare situations, and their efficacy is typically similar to autologous nerve transplantation, the clinically standard measure [38, 39]. Three distinct conduits, comprising polyurethane (PU), polyurethane/collagen (PU/C), and novel conduits built on polyurethane/collagen/nanobioglass (PU/C/ NBG), have been created in an investigation by Nasab et al., utilizing the electrospinning method. The MTT and DAPI analyses confirmed that the conduits were safe and that the cells were well attached. On the whole, the results of different experiments revealed that the innovative PU/C/NBG conduits had more desirable properties such as porosity (60–90%), wettability (88.6
), and biocompatibility than PU and PU/C-based conduits. These conduits may represent a good contender for peripheral-based nerve rejuvenation and axon growth because they offer repair possibilities. The incorporation of nanobioglass produced 0.23 mm-thick conduits, whereas the addition of collagen decreased the thickness of the conduits from 0.26 to 0.19 mm. The rate of degradation was less in PU/C/NBG as compared to PU/C because NBG acts as a bridge in the scaffolds and prevents the hydrolytic cleavage of collagen. Collagen-based scaffolds usually support cell adhesion because of the RGD sequence in the collagenous structure. Hence, these biocompatible electrospun scaffolds could potentially be used in neural tissue engineering [40]. The goal of the research by Gawrońska et al. was to see if the topology of the surfaces and physiological stimulation with nerve growth factor (NGF) had a synergistic influence on the behavior of PC12 cells. Using either blended or coaxial electrospinning, three kinds of substrates built on collagen type I, collagen type III, and poly(L-lactide-co-caprolactone) were created with synchronized fiber alignment. Encapsulating the growth factors inside the core-shell nanomaterials was discovered to preserve the growth factor’s functional properties. Nanofibers produced via co-axial electrospinning enabled the controlled and sustained release, and they followed the Korsmeyer–Peppas and Higuchi equations. This caused PC12 to engage with the materials and discovered that the cells developed more effectively (as shown in Fig. 3 (i)) on core-shell scaffolds, exhibiting bi- and tri-polar alignments while maintaining their usual phenotypic, relative to the other scaffolds used in the study. Neurite outgrowth reached 39.74  13.84 μm for core-shell PCB.

Fig. 3 (i) (a) Bio-performance of developed electrospun constructs and TCP, (b) medium neurite length; (c) % of cells bearing neurite, reproduced with permission from [41], copyright 2019, Royal Society of Chemistry; (ii) Schematic representation of the PLGA and collagen-associated conduit with gelatin perfusion (Gel@PLGA/Col), modified with permission from [43], copyright 2022, Elsevier; (iii) Diagrammatic representation of optimization of vapor treatment via glutaraldehyde for Collagen/Silk Tissue Engineering constructs prepared via electrospinning, reproduced with permission from [46], copyright 2017, American Chemical Society

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Collagen type I, collagen type III, and PCL conduits with oriented core-shell fibers have been providing new opportunities that can be used as platforms for peripheral nerve repair [41]. Tubulation restoration of sciatic nerve transection by nerve guidance conduits (NGCs) lowered the likelihood of axonal misalignments considerably, although NGCs’ therapy impact is limited by their single-hollow construction. Bioinspired multifunctional scaffolds can be utilized to create a micro-sized architecture necessary for directing cellular migration and encouraging tissue maturity. This approach drives the advancement of neuronal engineering by controlling nerve formation and repair. NGCs have shown tremendous promise in imitating the complicated PNS circuitry in the therapy of PNS injuries [42]. To imitate the complicated PNS architecture for tissue regeneration affecting the sciatic nerve, Zhao et al. fabricated a gelatin-based poly (lactic-co-glycolic acid) (PLGA) and collagenous tissue-derived bioinspired electrospinning-based conduits (Gel@PLGA/Col (as displayed in Fig. 3 (ii)). The Gel@PLGA/Col conduit demonstrated excellent nerve regenerating ability in a sciatic neural defect implant model having a massive gap. Histopathological examination of the recently created nerve tissue also revealed that the bioinspired electrospun conduit played a significant role in provoking axon growth and nerve consistency. The results also showed an increment in myelin tissue positivity in the recently created nerve tissue of 14.43% and 13.81% at 6 and 12 weeks, respectively, which are equivalent to autologous nerve transplantation. Furthermore, functional examinations of nerve and muscle tissue revealed that such scaffolds increased functional capacity after injury by restoring synaptic communication. Hence, the biomimetic fluffy sponge-like electrospun scaffolds offer exceptional nerve reconstruction and show a lot of promise for peripheral nerve restoration [43]. In another investigation, Mohamadi et al. prepared electrospinning-based scaffolds using poly (ε-caprolactone)/collagen/nanobioglass, which promoted human endometrial stem cells (hEnSCs) to proliferate and adhere to the scaffolds. Hence, these conduits could be suitable for peripheral neural engineering [44]. Similarly, Ebrahimi et al. fabricated polycaprolactone/collagen electrospun scaffolds based on electrospinning. They were found to enhance the attachment and differentiation of hEnSCs into motor neurons [45]. Also, Zhu et al. used glutaraldehyde vapor treatment to improve the conformational transitional ability of silk (as shown in Fig. 3 (iii)). However, overly treated scaffolds (Collagen/Silk) showed reduced cell-matrix adhesion [46].

4.1.2

Gelatin

Because of its bioactivity and biosafety, gelatin, among the first biological tissue engineering substrates, has been extensively used in regenerating a variety of tissues. Its limited mechanical characteristics and quick degradability, on the other hand, severely limit its use in repairing 3-D tissue [47]. Because it possesses the RGD sequencing of collagen, gelatin, which is formed from the hydrolytic cleavage of collagen, is generally considered a good choice for tissue regeneration because it breeds an efficient milieu for cell adhesion. Furthermore, gelatin has a low

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hypersensitivity, high degradability, cytocompatibility, and low market availability, all of which contribute to its widespread use in healthcare and biopharmaceutical industries. Poor tensile performance in physiological conditions and quick degradation rate has been significant disadvantages [48, 49]. Gabapentin (GBP), an antiepileptic medicine used as a painkiller to manage neuropathic symptoms, was combined with gelatin (Gel) and cellulose acetate (CA) in a study by Farzamfar et al. to create a possible neural tissue-engineered scaffold. The drug-loaded 3D scaffolds were made employing the wet-electrospinning method with CA/Gel [1:1 (w/w)] solutions in water/ethanol (3:7) (v/v) coagulating pools consisting of 3%, 6%, and 12% (w/v) of GBP. The shape, contact angle, porosity, mechanical characteristics, and cellular responsiveness of the scaffolds were assessed. The scaffold made from a 6% (w/v) GBP solution was selected to be the best structure for other in vivo testing in a Wistar rat sciatic nerve lesion model. The outcomes of the sciatic function score, hot plate latent scores, the weight-loss proportion of the wet gastrocnemius muscle, and histopathology using hematoxylin-eosin staining showed the GBP-loaded scaffold significantly improved the rejuvenation of the conceived injury, demonstrating its potential for neural bioengineering [50]. Recent research suggests that structured and chemically modified scaffolds can affect the direction of brain cell growth. In a study by Vashisth et al., aligned nanofibers (ANFs) were manufactured by the electrospinning and then treated to create a customized hybrid scaffold (HANF). The concentric rings of ANFs in the patterned scaffold were strengthened in a gellan–gelatin hydrogel matrix having excellent biocompatibility to offer enough mechanical properties and contact assistance for neural cell attachment and proliferation in in vitro tests. Quercetin was placed into the electrospun nanofibrous substrate as a therapeutic moiety to repair brain tissue. The strengthened ANFs improved the scaffold’s mechanical properties and offered a tubular nerve conduit architecture for neuronal cell development. In vitro cell culture settings were used to test the scaffold architectures’ influence on cellular behavior. It was discovered that functionalized structured scaffolds promoted guided neural cell development with favorable cellular cultures appearance while also displaying minimal cytotoxicity against nerve cell lines after days 3 and 7. The drug release showed a rapid burst release, with 60% of the drug released within 24 h, followed by sustained release up to day 5. The findings showed that the produced scaffolds had the ability to guide nervous tissue repair and could be employed as neuronal regenerative scaffolds [51]. PGS (poly(glycerol sebacate) is a biocompatible and biodegradable synthetic polymer that is becoming more popular in biomedical applications. A unique ternary PCL/gelatin/PGS mix nanofiber with a variety of biochemical constituents, mechanical characteristics, and controlled biodegradation levels was designed and manufactured in a study by Behtouei et al. PGS-gelatin blends are frequently employed to improve electrospinability. However, their poor mechanical characteristics, absence of structural support in an aqueous solution, and unamplified degrading behavior have restricted their utility. In a ternary composition, combining PGS and gelatin with PCL may increase their qualities. While examining the ternary

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mixes of PCL/gelatin/PGS, the inclusion of gelatin was expected to increase hydrophilic nature, resulting in improved biocompatibility and regulated degradability. Electrospun nanofibers with no beads were created by raising the polymer content, voltage, and needle separation from the collector. When measured against gelatin/ PGS binary nanofibers, ternary blend nanofibers with an equivalent weight proportion of polymers, T33 (containing 33 wt% each of PGS, gelatin, and PCL), have a 4-times improvement in compressive strength (7 MPa) and an 89-times boost in the break at elongation (1760.6%). In vitro research on glioma cells revealed that C6 glioma cells adhered well and proliferated effectively. The results showed that these scaffolds have the potential to be used in biomedical nerve applications [52]. In another investigation, Niu et al. used electrospinning to prepare gelatin/ poly (l-lactic acid) scaffolds for neural tissue regeneration (as displayed in Fig. 4 (i)). These scaffolds enhanced the secretion, proliferation, and elongation of glial-derived neurotrophic factor (GDNF). The developed construct was useful for SCs myelination and the epineurium remolding epineurium in the damaged site, which could efficiently restore the damaged nerve’s motor and sensory functions and hinder the target muscle tissue’s atrophy. Overall, the study displayed that the collaborative influence of nano topographical and biochemical cues on scheming biomimetic constructs could effectively advance neural tissue regeneration [53]. Similarly, Karbalaei Mahdi et al. fabricated electrospun polycaprolactone/ gelatin (PCL/GEL) nanofibrous scaffolds (as depicted in Fig. 4 (ii)). These scaffolds promoted the growth and proliferation of human pluripotent stem cells after an incubation time of 14 days. Their study demonstrated that bi-electrospun nanofibers of PCL/GEL had the ability to improve and ameliorate differentiation of hiPSCs to nerve cells. In general, the developed scaffolds appeared to be a practicable, dependable, and readily accessed composite for advanced neural tissue engineering experimentation [54]. Lee et al., on the other hand, used PCL/gelatin electrospun fibers to prepare scaffolds for neural tissue regeneration [55]. Gelatin was also used in a composite scaffold comprising gelatin/cerium oxide nanoparticles (nanoceria), which were found to enhance topographical stimuli and antioxidant effect on the cultured neural cells [56]. In another investigation, Wang et al. developed a threedimensional conductive PEDOT/Chitosan/Gelatin (PEDOT/Cs/Gel) construct by in situ interfacial polymerizations (as displayed in Fig. 4 (iii)). The results demonstrated that PEDOT/Cs/Gel scaffold substantially improved the differentiation of neural stem cells toward the neurons and astrocytes with enhanced expression of genes and proteins. The overall results depicted the prepared scaffold as a promising substrate for investigations in nerve tissue engineering [57].

4.1.3

Alginate

Alginate is a carbohydrate composed of two families of amino acids: D-mannuronic acid and L-guluronic acid. This polymer comes from brown seaweed and contains carboxyl groups between the sugars. Alginate is a biological polysaccharide that has been utilized in the biomedical field, such as wound healing and cosmetics since it

Fig. 4 (i) A biomimetic tubular nanofiber consisting of PLLA/gelatin was prepared to govern the Schwann cell’s proliferation, migration, and function in the regeneration of nerve, reproduced with permission from [53], copyright 2021, Elsevier; (ii) Diagrammatic representation of neural differentiation of mammalian induced pluripotent stem cells on nanofibers of bi-electrospun PCL/GEL, reproduced with permission from [54], copyright 2017, Elsevier; (iii) Schematic representation of neural stem cell’s 3D culture within conductive PEDOT layer-combined chitosan/gelatin constructs for nerve tissue engineering, reproduced with permission from copyright [57], 2018, Elsevier

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was licensed by the FDA [58]. Sulfated alginate has a structure comparable to heparin and is known for its good blood compatibility [59]. The growth and maturation of stem cells can be aided by alginate platforms [60]. Alginate has been utilized to assist in the maturation of embryonic stem cells into neural cells. Alginate is widely employed for tissue regeneration in the areas of skin, bone [61], cartilage [62], and cardiovascular constructions [63]. This is because of its biocompatibility, non-immunogenicity, cheap rate, and resemblance to ECM glycosaminoglycan [64]. Electrospinning of alginate stiff polymeric chains is still challenging due to strong intermolecular and intramolecular hydrogen bonding. It must be noted that the number of hydrogen bonds reduced when the -OH groups were replaced with sulfate groups. This makes the electrospinning procedure more practical. Additionally, alginate’s electrospinnability might be improved by utilizing a second polar polymer such as polyvinyl alcohol [65, 66]. In a study by Hazeri et al., electrospinning was used to create polyvinyl alcohol/ sulfated alginate (PVA/SA) nanofibers with a typical fiber diameter of 169–488 nm and varied sulfated alginate ratios (10, 20, and 30 wt%). The MTT assay and scanning electron microscopy data revealed that a PVA/sulfated alginate nanofibers with 30% SA offered good surface adhesion for C6, Schwann cells, as well as human bone marrow mesenchymal stem cells (hBMSCs). The neuronal proliferation of hBMSCs was confirmed using RT-PCR and immunocytochemistry for the MAP-2 biomarker. For up to 14 days, the production of MAP-2 confirmed neuronal differentiation. Thus, a PVA/SA nanofibrous conduit containing 30% SA was considered a desirable material for developing mesenchymal stem cells and can induce neurogenesis [59]. Electrospun fibers, which distinguish stem cells and neurons while also directing neurite outgrowth, are a potential element of biomaterial structures for neural tissue creation. Nevertheless, a method has yet to be established for safeguarding neurons, glia, and stem cells sown on the electrospinning-derived fibers between the lab environment and the operating conditions [67]. Miller et al. reported on an effort to employ cell-encapsulating hydrogel fibers generated via interfacial polyelectrolyte complexation (IPC). Interfacing acid-soluble chitosan (AsC) and cell-laden alginate and electrospinning them onto spindles of oriented electrospun fibers resulted in IPC-hydrogel fibers. Before IPC-fiber spinning, cortical neurons, L929 fibroblasts, and primary spinal astrocytes were combined into alginate hydrogels. After being encapsulated in IPC hydrogels for 30 min, 4 h, 1 day, and 7 days, the survival of each cell type was tested. IPC-hydrogel fibers composed of water-soluble chitosan (WsC) were used to encase specific neurons (WsC). Tuj1 was used to label neurons and check for neurite development. AsC-fiber neuron lifespan was lower than that of AsC-fiber astrocytes ( p < 0.05) and WsC-fiber neurons ( p < 0.05). As forecasted, neuronal and glial cell survivability was lower than L929 fibroblasts ( p < 0.05). Neuronal cells in WsC-fabricated IPC-hydrogel fibers extended neurites consistently, unlike AsC fibers. While some neurons remained encapsulated inside IPC-hydrogel fibers, others de-encapsulated and expanded neurons on electrospun fibers that could not completely merge with IPC-hydrogel fibers [68].

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Alginate was also used in a composite scaffold containing graphene nanosheetssodium alginate/polyvinyl alcohol electrospun fibers for neural tissue engineering, which showed a fourfold increase in tensile strength on the addition of 1% w/w graphene [69]. Alginate/PLA nanofibrous scaffolds were prepared by electrospinning by Xu et al.. After treating SA/PLA nanofibrous membranes, a calcium ion substitution step was employed to attach the alginate ion in the condition of gel-like calcium alginate (CA) [70]. Alginate magnetic short nanofibers/gelatin/ superparamagnetic iron oxide nanoparticles were prepared by electrospinning into a 3D hydrogel, which maintained the cell viability for up to 7 days, and the storage moduli of these composites were close in value to that of neural tissues [71]. In another study, graphene/carbon nanotubes/polyvinyl alcohol/sodium alginate conduits were made by the dynamic needleless linear electrospinning method to prepare neural-mimicking scaffolds on a large scale. These scaffolds had good thermal conductivity and hydrophilic characteristics [72].

4.1.4

Chitosan

Chitin can be present in the cell walls of crustaceans, including insects, lobsters, and shrimps, and also in bacilli and fungus. Chitin is then alkaline deacetylated to produce chitosan. The extent of chitin deacetylation is determined by the procedure used [73]. Solubility, tensile stability, and decomposition are all affected by the molecular weight and the degree of deacetylation [74]. Chitosan was found to be a suitable choice for scaffold assembly based on mechanical qualities, growth factor delivery, and other factors [75]. With the advancement of peripheral nerve tissue regeneration, biodegradable nanomaterials research has wide-ranging applicability. Due to its molecular resemblance to glycosaminoglycan, chitosan (Cs) is a good choice for brain tissue regeneration. On the other hand, Cs are frequently utilized in conjunction with other substances due to the absence of elasticity and flexibility. The freeze-drying process was used to make three composite scaffolds: Cs/hyaluronic acid (HA)/Gel, Cs/polyethylene glycol (PEG), and Cs/collagen (Col). These polymers’ tensile characteristics, swellability behavior, porosity, and conductance have been studied. The inclusion of other elements lowered the mean pore size while improving the tensile capabilities of the composite scaffold when compared to pure Cs. The findings showed that each of such Cs-containing scaffolds was biocompatible and cytotoxic, whereas Cs/PEG scaffolds had greater cell survivability and were able to stimulate PC12 cell attachment, propagation, and maturation. The Cs-based hybrid scaffolds thus created could be beneficial in the regeneration of brain tissue [76]. In a study by Wu et al., a multipurpose gene delivering nanovector with a CS core and polyethyleneimine (PEI) side-arms that were linked via PEG to arginineglycine-aspartate (RGD)/twin-arginine translocation (TAT) was created. RGD/TAT sequences were coupled for better targetability and cellular absorption, while stemmed PEI with a molecular mass of 2,000 Da was employed to strike a compromise between bioactivity and transfection effectiveness. The nanovector was

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compatible with tissues and demonstrated outstanding DNA condensing ability; the resultant DNA complexes were well-generated, had tiny particle sizes, and had a modest positive charge. In tumor (HeLa cells) and healthy cells (NIH 3T3 cells and 293T cells), a greater gene transfection rate was confirmed when contrasted to PEI (25 kDa). More crucially, cells transfected with the chitosan-graft-PEI-PEG/pCMVEGFP-Ntf3 combination generated persistent neurotrophin-3 (NT-3)with a gradual, cumulative rise in content, which encouraged neurite propagation and stimulated neurogenesis of neural stem cells. The results indicated that these multipurpose copolymers could be helpful to nano-based vectors for genetically modifying cells and could be used in tumor therapy and tissue regeneration [77]. In the research by Sadeghi et al., Nanofibers comprising PCL, chitosan, and PPy were electrospun to integrate the benefits of electrospun nanofibrous topology with the versatility of chitosan and PPy. Electrospinning was used to create various ratios of PCL/CS/PPy polymeric scaffolds, which were then examined for surface features, bioactivity, and water solubility. The findings demonstrated that chitosan in the scaffold increased the hydrophilic nature of the substrate by up to 66% (123  2.3 for PCL to 41.37  3.51 for PCL/chitosan), as seen from a reduction in contact angle. The fibers had a mean diameter of 30–180 nm, which had been impacted by the chitosan content, as an increment of up to 30% in CS concentration reduced fiber diameters from 124 to 36 nm. In vitro tests with PC12 cells demonstrated that the PCL/CS/PPy nanofibers promote cell adhesion, dispersion, and growth, with a 356% growth in proliferation compared to pure PCL and PC12 neuronal elongation. The results showed that the PCL/CS/PPy nanocomposites scaffolds promote PC12 cell attachment, expansion, and multiplication. As a result, this scaffold may be helpful as a neural tissue replacement [78]. Karimi et al. developed electrospun fibrous scaffolds comprised of poly (3-hydroxybutyrate)/chitosan, which demonstrated aligned and random fibers in rat neural cells (B65 cell lines) [79]. In another investigation, chitosan/gelatin electrospun fibers were prepared by Gnavi et al., wherein the random fibers promoted cell adhesion and proliferation while the aligned fibers axon and SC-directed growth [80]. PCL/chitosan [81], chitosan/polyurethane [82], shell/core chitosan/ poly (3,4-ethylene dioxythiophene) (PEDOT) [83] are some of the other investigations which displayed promising results.

4.1.5

Silk

Silk, a fibrous protein that forms made by silkworms and spiders, has unique qualities that make it appropriate for use as a biopolymer. Silk has good tensile properties, good biocompatibility, low immunogenicity, low bacterial adherence, and manageable biodegradability [84]. Silk is also a versatile product that has been utilized to make biomimetic constructions such as hydrogels, scaffolds, nanofibers, nanoparticles, and films. Silk hydrogels are flexible and long-lasting biomaterials that are frequently utilized in neural tissue regeneration because they may preserve structural stability longer than some other polymers, including fibrin and collagen

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gels, while also eliciting enhanced axonal entanglement [85]. Silk hydrogels have now been manufactured as multifunctional scaffolds to enable neurogenesis for brain and nerve tissue repair [39]. Considering its complicated physiology and limited regenerating ability, one of the most challenging problems in neural construction is encouraging the restoration of damaged nerve tissue. A study by Boni et al. used electrospun 3D nano scaffolds (3DNSs) from a biomaterial combination of silk fibroin (SF), PEG, and PVA to restore injured nerve tissue. The 3DNSs were studied to see if they were suitable for direct implantation into the CNS. The bioactivity of 3DNSs was studied in vitro with PC12 cells and their impact on reacting astrogliosis in vivo with a photothrombotic rodent model of ischemia. The percentage of SF directly impacted the biomechanical properties and structural system of the 3DNSs, with compositions manifesting as either a gel-like framework (SF 50%) or a nano-electrospun fibrous architecture (SF 40%). In vitro testing demonstrated enhanced cellular survival in the existence of 3DNSs, while in vivo testing revealed a significant drop in GFAP production in the peri-infarct region ( p < 0.001 for F2 and p < 0.05 for F4) following stroke, indicating that 3DNSs may be inhibiting reactive astrogliosis. These results improved our knowledge of the physicochemical interplay between SF, PEG, and PVA and the possibility of 3DNSs as a bioactive strategy for stroke rehabilitation, mainly when employed in conjunction with medication or cell therapy [86]. Peripheral nerve damage is a typical complication of low-regenerative-potential injuries. For neural tissue regeneration, electroconductive materials could offer optimal cellular proliferation microclimates and complementary cell-guided signals. Electroconductive conduits were made by combining PEDOT–PSS to electrospun silk scaffolds or DMSO-incorporated PEDOT–PSS to electrospun silk conduits in a study by Magaz et al. (as displayed in Fig. 5 (i)). The coating quantity may be adjusted, and DMSO treatment may increase conductance even further. All scaffold types retained cellular viability while exhibiting significantly higher metabolism rate and multiplication than neat silk. PEDOT–PSS doped scaffolds treated with DMSO surpassed their PEDOT–PSS equivalents in part. Differentiation tests revealed that such PEDOT–PSS constructed silk scaffolds might enable neurite outgrowth, implying that they have the potential to be utilized as a potential substrate to repair the electrochemical connections in the injury region and maintain healthy organ function [87]. Due to its self-assembled sheet design, silk fibroin (SF), a native protein, exhibits outstanding biological characteristics and has high mechanical strength. SF scaffolds may be molded into suitable forms using various processing processes and show remarkable promise in peripheral nerve regeneration. As a result, it is thought that combining PPy and SF produces a useful water-soluble conducting polymer with excellent electroactivity [88]. In a recent study, Zheng et al. stipulated a general approach to combining photoacoustic (PA) stimulation of nerve into hydrogel construct employing a nanocomposite hydrogel method. Particularly, an extremely effective photoacoustic agent, polyethylene glycol (PEG)-functionalized carbon nanotubes (CNT), was included in silk fibroin to create a material possessing biocompatibility and soft

Fig. 5 (i) (a) Schematic representation of electrospun silk fibroin constructs pursued via PEDOT–PSS or DMSO-Treated PEDOT–PSS conjugation and (b) Schematic representation of the change in the conformation of the PEDOT–PSS following the DMSO treatment, from benzoid toward quinoid, reproduced with permission from [87], copyright 2020, American Chemical Society; (ii) Photoacoustic CNTs embedded silk constructs for stimulation and regeneration of neurons, reproduced with permission from [89], copyright 2022, American Chemical Society; (iii) Spidroin silk-associated electrospun mats modified with

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acoustic property (as depicted in Fig. 5 (ii)). The reports identified that the developed scaffolds permit neuron nongenetic activation, having a spatial accuracy determined via the area of light illumination, enhancing regeneration of neurons. The photoacoustic stimulation enhanced the growth of neurites by 1.74-fold compared to the unstimulated group in a rat model. Thus, the overall results demonstrated that the developed scaffolds could be a new approach to the non-pharmacological regeneration of neurons [89]. In another investigation, Revkova et al. endeavored to determine the impact of ECM’s peptide motifs (RGD, IKVAV, VAEIDGIEL) modified spidroin-associated electrospun material on neural precursor cell (drNPCs) adhesion, proliferation, and differentiation (as depicted in Fig. 5 (iii)). The scaffolds demonstrated orientation uniaxially and elastic modulus in the swollen phase, which is similar to those of the dura mater. For the first time, it was reported that drNPCs on the developed scaffold primarily maintain their stemness in the growth as well as differentiation media with brain-derived neurotrophic factor and glial cell-derived neurotrophic factor. In contrast, the inclusion of ECM-peptide motifs might shift the balance toward the differentiation of neuroglia. Overall results will be helpful for the researchers to develop an in vitro model for the neuroglial stem cell niche having the prospective regulation over their differentiation [90]. During regeneration, aligned electrospun nanofibers offer the proper assistance and topographical signals for directing axonal and neurite development. The transmission of neural impulses is a requirement of a typical nerve [88, 91]. Combining biodegradable, reactive polymers with conductive polymers can provide the neuroconductive feature. This should add new elements, such as electrical signals, to the otherwise established topographic and biological prompts, making it a more versatile neuro regenerative method [92]. Nune et al. created randomized and oriented electrospun nanofibrous composites substrates using a mixture of silk fibroin and melanin (as demonstrated in Fig. 5 iv)). They evaluated their antioxidant capabilities as well as their physicochemical characteristics. Human neuroblastoma cells (SH-SY5Y) were used to assess their neurogenic potentiality of cell viability, growth, adherence, and maturation. Nanofibrous scaffolds with appropriate physicochemical parameters were designed to facilitate neuronal regeneration as neural platforms. They increased the adhesion and vitality of neuroblastoma cells, demonstrating that they are biocompatible. The radical scavenging activity of silk/melanin composites substrates has shown a strong antioxidant nature. Furthermore, the cells’ development into neuronal cells and alignment through their axis aided the melaninloaded oriented silk fibroin electrospun scaffolds [93].

 ⁄ Fig. 5 (continued) ECM-peptide motifs for nerve tissue engineering, reproduced with permission from [90], copyright 2021, American Chemical Society; (iv) Melanin loaded electroactive and antioxidant silk fibroin constructs developed via electrospinning for neural tissue engineering, reproduced with permission from [93], copyright 2019, Elsevier

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Silk fibroin has also been used to prepare electrospun composites with PLA/PCL/ silk fibroin/PPy (PLCL/SF/PPy) for guided nerve tissue engineering. An in vivo study of repairing 10 mm sciatic nerve defect and these conduits showed nerve repair within 4–12 weeks of implantation [94]. Whereas silk fibroin-based scaffolds were also prepared by Xu et al. for post-surgery nerve regeneration. In a rat model, these scaffolds exhibited comparable results with the gold standard for nerve repair and achieved results close to those of autologous grafts [95]. Similarly, electrospinningbased electroconductive scaffolds were prepared by Zhao et al. using PPy as the conducting polymer. Results indicated that these scaffolds could improve myelination and axonal regeneration in vivo in the presence of electrical stimulation [88].

4.1.6

Miscellaneous

Keratin Hair, wool, feathers, and horns all contain keratins, which are fibrous structural protein components. Keratins have exceptional biocompatibility, biodegradability, and minimal hypersensitivity because their inherent ability is equivalent to biological molecules or matrix proteins in the human body. Keratin is already being manufactured into nanoparticles, hydrogels, and nanofibers for tissue engineering applications such as wound healing, inhibition of platelet aggregation, osteogenesis, nerve healing, and renewal. As a result, incorporating keratin into polymeric nanofibers could improve their cyto-biocompatibility [96–98]. Human hair keratin has good qualities for stimulating nerve tissue repair and functional improvement. Schwann cells could be activated to help in this healing. Human hair keratin has the power to stimulate Schwann cells via chemotaxis, promote cellular proliferation, and promote the production of some critical genes simultaneously time. In an animal study of nerve damage, human hair keratin induces nerve healing and has a therapeutic benefit similar to autologous nerve transplanting [99]. Keratin has the ability to enhance polymeric nanofibrous biocompatibility and bioactivity. Nevertheless, due to keratin’s weak spinnability, adding it to the composite nanofibers reduced the mechanical performance of the nanofibers and generated nonuniform keratin dispersion within the nanofibers [96, 100]. As a result, surface-functionalized polymeric nanofibers with keratin-based nanoparticles would enhance hydrophilicity and also their mechanical properties. Keratose (oxidative keratin, KOS) nanoparticles-encapsulating PVA-based nanofibers (KNPs/PVA) were synthesized using electrospraying layering following electrospinning and operated on brain cells. KNPs/PVA nanofibers were studied for biochemical orientation, mechanical characteristics, and hydrophilicity. At the similar mass ratios of KOS to PVA, the KNPs/PVA nanofibers had greater hydrophilicity and mechanical characteristics than the KOS/PVA mix nanofibers. In addition, when evaluated against PVA nanofibers and KOS/PVA mix nanofibers, KNPs/PVA nanofibers showed higher compatibility with biological tissues regarding cellular shape,

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adhesiveness, and multiplication. Compared to composite nanofibrous scaffolds, these findings revealed that the polymeric nanofibers that were surface-treated with KOS nanoparticles could improve hydrophilicity, physicomechanical characteristics, and cytocompatibility [101]. Nerve regenerative conduits were made from elastic, porous, non-toxic, and biodegradable cylindrical keratin nanofibers. Keratin was isolated from leftover chicken feathers and mixed with polyvinyl alcohol before being electrospun into nanofibrous conduits that had a mean diameter of 170–234 nm. When the keratin concentration was raised, the mean diameter of the nanofibers shrank. The range of nanofiber size distributions, on the other hand, narrowed, implying that as nanofibers thinned, their quantity grew, minimizing the interfacial gaps between them. The existence of keratin protein in nanostructures was established, ensuring cytocompatibility and biodegradation. TGA revealed that keratin increased the nanofibers’ heat stability and hydrophilicity [102].

Hyaluronic Acid Hyaluronic acid (HA), a glycosaminoglycan abundant in extracellular structures throughout the human system, plays an essential function as a lubricant. Because of its tunable features, including biodegradability, cytocompatibility, bioresorbability, and hydrogel formation capability, HA has been extensively studied for tissue regeneration. HA hydrogels improve neural progenitor survival and growth, showing significant potential for peripheral nerve regenerative treatments and CNS therapeutic methods. HA hydrogels, for instance, exhibit mechanical characteristics that alter brain progenitor development, paving the way for new neurodegenerative therapeutics [39]. In the nerve cell bioengineering therapies for neurological-based illnesses, pain control, and tissue regeneration, electroactive polymers allow the administration of an electrical-induced input to control neural function. Steel et al. created a nontoxic electroconductive material hybrid with an ultra-low carbon nanotube (CNT) content with the HA nanostructures. They expected that electrical impulses via composite nanofibrous scaffolds would promote neurite proliferation by creating a platform conducive to neuronal outgrowth and electronic conductivity, allowing voltagesensitive regeneration pathways to be activated. A customized stimulation apparatus was built, and the electrical stimulus values needed to elicit enhanced neurite proliferation were examined using the charge-balanced biphasic square wave function. It was observed that at 200 mV/mm electrical impulse for 30 min, the HA/CNT nanofibers could promote neurite growth. Neurons cultivated on stimulated HA-CNT nanofiber substrates grew much faster than those produced on unstimulated HA-CNT and HA control platforms [103]. Chitosan and hyaluronic acid are two naturally occurring, oppositely charged sugars commonly employed to make nanofibers for bioengineering. Electrospinning and subsequent treatment were used to create composite fibers made of CS and HA. In this case by Bazmandeh et al., HA was used to coat CS nanofibers in two

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different ways. In the first procedure (HA/CS1), the electrospun nanofiber was neutralized first and then coated, but in a second way (HA/CS2), neutralization and deposition were done at the same time. Both treatments preserved the mats’ fundamental fiber structure during coating, and there was little discernible morphological change between the HA/CS1 and HA/CS2 specimens. HA/CS2 nanofibers, on the other hand, revealed stronger interactions between HA and CS. The wettability of the resulting fibers was determined using contact angle measurements. Even though both scaffolds were highly wettable, the HA/CS2 exhibited a lesser wettability than the HA/CS1. More significantly, the scaffolds differed in their cytocompatibility. Cell growth was improved on both HA-coated scaffolds. However, when HA was applied in a direct synchronous technique, cell growth and adhesion were increased [104]. Table 1 depicts the latest studies using natural electrospun polymers for neural tissue engineering.

4.2

Synthetic Polymers

Currently, there is a lot of focus on using natural and artificial polymers to make tissue engineering substrates. Electrospun scaffolds made of synthetic polymers have been shown to be processable, but their limited cell attachment potential due to their hydrophobic nature and lack of surface cell identification regions limits their use in biomedical sectors. On the other hand, natural polymers have appropriate biocompatibility and efficient cell interfaces because of their strong structural homology to natural ECM and high hydrophilicity. However, when it comes to natural polymers, one should remember that their low tensile characteristics and rapid disintegration rate make them unsuitable for tissue engineering. As a result, combining synthetic and natural polymers to address these drawbacks results in a scaffold that takes advantage of both polymers’ strengths [47].

4.2.1

PVA

PVA is an aqueous-soluble synthetic polymer with a strong spinning ability and cytocompatibility that has been extensively used to produce fibrous structures for designing a variety of tissues, including the neuron, heart, and vasculature. For neural tissue regeneration, electrospinning could be used to create a PVA fibrous substrate with an average fiber size of 410–1,062 nm [115]. The studies showed that PVA fibers have a beneficial effect on PC12 cell attachment, multiplication, migration, and development [115]. The lack of immunogenicity and low cell affinity of synthetically-derived polymeric materials like PVA, on the other hand, limited their use in pure applications [116]. In an investigation by Sadeghi et al., the electrospinning process was used to assess new PVA/Gel/chondroitin sulfate (PVA/GE/Cs) nanocomposites conduits

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Table 1 Recent investigations on electrospun natural polymers utilized for neural tissue engineering In vitro cell line Mouse bone marrow mesenchymal stem cells

In vivo cell line –

Poly (hydroxybutyrate)/ chitosan

Rat neuronallike cells (B65 cell line)

–

Chitosan/collagen

RSC96 cells

Sprague Dawley rats

Device Scaffold

Polymer Gelatin nanofibers/ polyaniline/ graphene nanoparticles

Scaffold

Scaffold

Results The softwaredefined process variables (voltage 13 kV, flow rate 0.1 cm3/h, and PAG weight % of 1.3) were used to electrospun an optimal framework with the highest electroconductivity (0.031 0.0013 S/ cm), cell suitability, and an appropriate diameter Arranged PHB, PHB/CTS 80:20, and PHB/CTS 85: 15 fibers had mean diameters of 675 nm, 870.74 nm, and 740.3 nm, correspondingly, that were smaller than random fibers. The introduction of CTS reduced the water droplet’s contact angle in randomized and oriented PHB/CTS fibrous scaffolds from 124.79
to 43.14
. Furthermore, fiber alignment dramatically increases the nanofibrous PHB/CTS scaffold’s wettability. The elastic modulus of aligned PHB/CTS 85:15 improved from 6.41 MPa to 8.73 MPa The surface fiberlike architecture of chitosan/collagen

References [105]

[106]

[107]

(continued)

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Table 1 (continued) Device

Scaffold

Polymer

PVA/sodium alginate/graphene

In vitro cell line

–

In vivo cell line

–

Results hybrid scaffolds was accompanied by internal porosity. The introduction of chitosan to a purified collagen scaffold reduced mean pore size, liquid uptake, and degradation rate while increasing the mechanical properties of the composite substrates. The scaffolds were cytocompatible and non-toxic, might increase Schwann cell adhesion, migration, and growth, and showed clearly controlled degrading behavior without triggering any inflammatory process The porous-natured sharper nanofibers with a size of 141 31 nm in the dynamic linear electrospun Gr-AP membranous scaffolds possessed a greater yield of 1.25 g/h. The shape, uniformity, lipophilicity, and thermal properties of Gr-AP nanofibrous membranes were all improved by Gr. The critical conductive cutoff for Gr is 0.075 wt. %, which offers nanofiber membranes an equal

References

[108]

(continued)

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Table 1 (continued) Device

Polymer

In vitro cell line

In vivo cell line

Scaffold

Alginate-poly (vinyl alcohol) nanofibers

Pluripotent embryonic stem cells

–

Scaffold

Alginate/PVA/ graphene

PC12 cell lines

–

Results diameter variation, optimal conductance, sufficient hydrophilicity, adequate high surface area, and optimal thermostability After 24 h, the ALPA nanofibrous mat’s water solubility and swellability proportion reached 32.23  0.95 and 295  4.2, respectively. SEM analysis of the microstructure indicated the production of irregular and smooth threads. Furthermore, ALPA-nfs’ biocompatibility investigation stated that it has a cytotoxicity-free effect on pluripotent ESCs in vitro, allowing them to proliferate at a rate equivalent to those of the natural, indicating its usefulness as a biodegradable polymer scaffold The oriented fibrous Gr-AP composites closely resembled the anisotropic architecture of the original sciatic nerve. Compared to randomized fibrous composites, aligned fibrous Gr-AP scaffolds considerably improved mechanical characteristics and cell–scaffold interaction.

References

[109]

[110]

(continued)

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Table 1 (continued) Device

Polymer

In vitro cell line

In vivo cell line

Scaffold

Gold nanoparticledoped electrospun PCL/chitosan

–

Neonate Wistar rats

Scaffold

Chitosan/PCL

–

Neonate Wistar rats

Results Furthermore, electrical activity enhanced PC12 cell growth greatly. They also demonstrated superior structural and mechanical characteristics that improved neural cell–substrate interrelations The number of gold ions captured by electrospun nanofibers was directly proportional to their chitosan concentration, according to the UV–Vis spectra. The addition of AuNPs considerably improved the conductivity of the composites, according to electrical testing. Ultimately, the biological reaction of Schwann cells on AuNPs-doped biomaterials was greater than that on as-prepared substrates after 5 days of incubation in the form of enhanced cell adhesion and higher multiplication Utilizing the sacrificial fiber approach, the interconnecting porosity of the PCL/chitosan scaffolds was enhanced, leading to increased surface wettability.

References

[111]

[112]

(continued)

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Table 1 (continued) Device

Polymer

In vitro cell line

In vivo cell line

Scaffold

Genipin/chitosan

Culture dorsal root ganglia

–

Hydrogel

Aligned chitosan nanofiber

–

Sprague Dawley rats

Results The specimen with 40% PEO, with 75–80% porosity after PEO removal, had the maximum hydrophilicity amongst scaffolds with varying porosities. The reducing agents THPC and formaldehyde were also used to effectively boost the reduction speed of gold nanoparticles and scaffold electrical properties The physicomechanical characteristics of chitosan nanofibers were improved by genipin treatments. Schwann cell organization and development were aided by genipin therapy. When genipintreated fibers were being used, neurites grew rapidly ACG-RGI/KLT increased vascular invasion as well as nerve regrowth at an early stage of damage. ACG-RGI/ KLT aided nerve regrowth and improved function in rats after 12 weeks

References

[113]

[114]

utilizing carcinogen-free solvents. To generate a homogeneous combination with adequate solvent characteristics and ultimately non-toxic nanocomposite scaffolds, a solvent medium combining water and acetic acid was employed as a non-toxic solvent medium. The influence of the water-acetic acid ratio in the solvent medium,

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as well as polymer content (8–10 w/v %) on nanofiber structure, was studied. Cell toxicity tests revealed that the generated PVA/GE/Cs scaffolds are not hazardous to cells. After 24 and 48 h, SEM results showed that L929 mice fibroblast cells had a good interaction with the scaffold layer and adhered and propagated well on the produced scaffold, indicating that they could be used in tissue regeneration [66]. For neuronal tissue engineering, the construction and development of conducting scaffolds that generate appropriate intercellular connections by electrical impulses are crucial. Electrospun conducting PVA/PEDOT (poly(3,4-ethylene dioxythiophene) substrates in varying compositions were created in a study by Babaie et al.. These findings reveal that PEDOT-doped scaffolds outperformed pure PVA scaffolds with respect to physicochemical characteristics and cell survival. Real-time PCR analyses were conducted to examine the neural development of rat MSCs after scaffold tuning. When compared to TCP specimens, the platform samples, with or without initiation of electrical impulses, upregulated tubulin, nestin, and enolase. Furthermore, the nestin expression level was 1.5 times higher in scaffold materials with electrical activity than in scaffold materials. Ultimately, this research reveals that electrical activity combined with PVA/PEDOT conducting scaffolds can increase cellular responsiveness and neural development by imitating the features of natural neural tissue [117]. Because the regeneration of damaged nerve fibers is not always as predicted, the construction of nerve conduits has become fashionable in recent years. Conductivity and degradability are crucial parameters for perfect nerve conduits to allow proper delivery of brain messages and minimize secondary injury during nerve conduit elimination. In a study, the electrospinning process was applied to generate PVA/CNT electrospun fibrous films, after which their identification of the best nerve conduits using electrospinning, PVA, and CNT was conducted. The electrospun films with 0.25 wt% PVA had a lesser resistance up to 25.3 Ω, excellent fibers-like shape, and a diameter of 1 m, according to the findings. Furthermore, the electrospun films were non-cytotoxic and promoted cell proliferation. The MMT experiment showed that after 3 days of co-culture with cells, PVA/CNT electrospinning fibers had a cellular vitality that was 18.5-fold higher than the reference group on Day 1. PVA/CNT electrospun fibrous sheets are a suitable choice for the usage of nerve conduits based on all characteristic assessments [118]. Similarly, Bagheri et al. fabricated chitosan-aniline oligomer/polyvinyl alcoholbased scaffolds where 40% of the entrapped drug was released within 40 min. These conductive substrates had higher cellular activity than non-conducting ones [119].

4.2.2

PLGA

Polylactic-co-glycolic acid (PLGA) is an essential artificial FDA-approved biomaterial that can be infused explicitly into wounded locations. PLGA can fundamentally sustain stem cells following delivery due to particular and unique physicochemical properties. Regarding the possibility of PLGA nanoparticles as a therapy for brain illnesses, unaltered PLGA has several flaws, including a minus

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charged, a lipophilic backbone, and the presence of free glycolic monomers. As a result of its hydrophobicity, the PLGA nanoparticles struggle to engage effectively with various cells. Self-assembling micellar-based nanoparticles having hydrophilic shells and hydrophobic interiors can enhance the solubility of hydrophobic macromolecules. Furthermore, the presence of a nanoparticulate corona surrounded by hydrophilic units helps target molecules to maintain their shape in a variety of microhabitats [120]. Ischemic strokes are marked by severe neuron death, glial scar-like development, and neuronal tissue deterioration. These result in significant modifications in the extracellular environment, neuronal architecture, and long-term functional impairments. Although NSCs could help with some function-based deficits caused by strokes, recovery is never comprehensive, and tissue restoration is minimal. As a result, the goal is to utilize a mix of NSCs and biomaterials that are sufficiently enriched to keep these cells inside the infarct space and speed up the development of new tissue. To develop appropriate tissue engineering methodologies, researchers tested the regeneration ability of a PLGA-PEG micellar biopolymer loaded with Reelin and fetal NSCs on photothrombotic stroke mice models. In vitro studies revealed that PLGA-PEG, along with Reelin, increased the proliferation speed (Ki-67+ NSCs) and neurite propagation (axon and dendrite) of Ki-67+ NSCs ( p 0.05). Moreover, NSCs grown amid Reelin-incorporated PLGA-PEG micelles showed a high level of neural maturation (Map-2+ cells) ( p < 0.05). Following 1 month, double immunofluorescence labeling revealed that Reelin-loaded PLGAPEG micelles enhanced the number of migratory neural progenitor cells (DCX+ cells) and matured neurons (NeuN+ cells) near the lesion region relative to the control groups ( p < 0.05). In comparison to the other treatments, immunohistochemical results revealed that PLGA/PEG incorporated with Reelin dramatically reduced astrocytosis gliosis and boosted regional angiogenesis (vWF-loaded cells). The cavity width in the Reelin-based PLGA-PEG+NSCs sample was decreased due to these adjustments. Reelin-based PLGA-PEG+NSCs were found to improve neurological performance and motor improvement in neurobehavioral tests. These findings suggest that Reelin-loaded PLGA-PEG can promote NSC dynamic development, neuronal differentiation, and local revascularization after ischemia injury by creating a favorable milieu [120]. In another investigation, Pozzobon et al. developed a biodegradable neuronal conduit and tested its cytotoxic effects with stem cells as well as its regenerative healing capabilities in a rat model. The electrospinning process was used to create a PLGA conduit with aligned nanofibers, which was then modified with gelatin and implanted with either mouse embryonic stem cell lines (mESCs) or human mesenchymal stem cells (SHED). In vitro, cellular growth and survival were examined. In a rat model of sciatic neuron transection, the conduits were inserted. The Sciatic Functional Index (SFI) was utilized to track functional capacity after 8 weeks, and histology investigations were employed to determine tissue repair. The researchers created scaffolds of synchronized PLGA fibers with a mean diameter of 0.90  0.36 m and an orientation coefficient of 0.817  0.07. Gelatin injection raised fiber diameters to 1.05  0.32 m, decreased orientation coefficient to

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0.655  0.045, and made the substrate extremely hydrophilic. The stem cells continued robust and propagated after 7 days in incubation, according to the cell viability and live/dead experiment. Phalloidin/DAPI stain revealed that these cells attached and propagated broadly, indicating that they were thoroughly adapted to the biocompatible material. The SFI values of the conduit recipients were comparable to those of the control group. Eventually, conduits made of PLGA-gelatin nanofibers were created, which allowed the stem cells to communicate effectively. Even though in vitro research has demonstrated that this biopolymer is a potential material for nerve tissue repair, this graft’s in vivo trials have not shown substantial nerve regeneration advancements [121]. In neuronal bioengineering, the construction of scaffolds with the proper electrical characteristics is crucial. Since the alignment of fibers in scaffolds impacts cell growth and division, a study by Farkhondehnia et al. targeted to create oriented electrospun electroconductive nanofibers by electrospinning 1%, 10%, and 18% (w/v) doped polyaniline (PANI) with PCL/poly lactic-co-glycolic acid (PLGA) (25:s75) solutions. The width, wettability, and conductance of the fibers were all measured. The findings showed that the conducting nanofibrous substrates were suitable for nerve cell adhesion and growth. When relative to PLGA/PCL/PANI scaffolds that were not electrically stimulated, electrical activity increased neurite outgrowth. Electrical stimulus via nanofibrous PLGA/PCL/PANI scaffolds promotes cellular proliferation as the polyaniline ratio rises. On the other hand, polyaniline increases of more than 10% will cause cellular toxicity. Also, the conducting scaffolds with an acceptable PANI ratio, together with electrical activity, might be used to treat spinal cord injury [122]. In another study, Aval et al. prepared electrospun PLGA/Graphene microribbons, which greatly enhanced the tensile and elastic modulus compared to plain PLGA microribbons. Moreover, SH-SY5Y cells were shown to proliferate into mature neurons; hence they are promising candidates for neural tissue engineering [123].

4.2.3

PPy

Electrical stimulation (ES) with conducting polymers can boost neurite proliferation and accelerate brain regeneration significantly. However, in addition to ES, the functioning of nerve cells and their responsiveness to substrate conductivity are critical for the potential implementation of neural healing. As a result, the ES has been used in conjunction with appropriate materials, such as tissue scaffolds, to treat neurological injuries and has shown significant promise in peripheral nerve restoration. In this context, Zhao et al. utilized 3D bioprinting and electrospinning to create a conducting polypyrrole/silk fibroin (PPy/SF) hybrid nerve guidance conduit. ES of Schwann cells placed on such scaffolds resulted in increased viability, multiplication, and relocation, as well as elevated expression of neurotrophins. In addition, the PPy/SF conducting neural guiding scaffolds combined with ES were found to stimulate axonal growth and remyelination in vivo efficiently. Furthermore, they discovered that ES triggered the MAPK signal transduction pathway at the

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electroconductive conduits. The researchers also discussed the usage of PPy/SF NGCs with ES and the probable mechanisms of the fostered regeneration of nerves (as depicted in Fig. 6 (i)). These data showed the longitudinally guided PPy/SF conducting composite conduits possessed excellent features for clinical usage and stimulated nerve regrowth and functional status [124]. In another study, Shrestha et al. created a self-electrically driven double-layered NGC made of electrospun mats, with an ordered orientated interior core and a randomly oriented outside layer. After a homogeneous coating of PPy, bioinspired NGC was made from chitosan (CS)-based polyurethane (PU) nanofibers with welldeposited functionalized multiwall carbon nanotubes (fMWCNTs) (as demonstrated in Fig. 6 (ii)). The intertwined NGC internal structure displayed a cellular biomaterial functionality and improved physicochemical characteristics such as electrical properties, mechanical characteristics, and excellent biocompatibility, representing a natural hosting material for biological ECM for critical roles in peripheral nervous system tissue regeneration. Within in vitro cultured cells, the rebuilding, multiplication, and relocation of Schwann cells (S42), as well as the maturation of rat’s pheochromocytoma cells (PC12), were considerably boosted on the oriented mats relative to the randomly aligned mats. The phenotypic of nerve bundles and the shape of spontaneous outgrowth were both preferentially steered boosted along the axis of orientated nanofibers, indicating that axonal regrowth is very adaptable. CDNA gene expression was used to assess the proliferation of PC12 cells grown on as-fabricated NGCs. Thus, this will aid in the efficient deployment of developed NGCs, and they will be used in treatment interventions for treating wounded locations and stimulating nerve cell regeneration [125]. NGCs can create a favorable milieu for nerve restoration while encouraging Schwann cellular proliferation. In a study by Pan et al., researchers used polypyrrole-coated polycaprolactone nano yarns (PPy-PCL-NYs) as fillings in NGCs. PCL-NYs with an orientated architecture were generated using a doubleneedle electrospinning technique, and PPy was subsequently deposited on the PCL-NYs using an in situ chemical polymerization method. PCL nanofibers were then gathered as the outermost surface around nano yarns using a traditional electrospinning procedure to create PPy-PCL-NY nerve guiding conduits (PPy-PCLNY NGCs). The results revealed that PPy was coated on the exterior of PCL-NY in a homogenous and uniform manner. Young’s modulus and strain–stress curves of PPy-PCL-NYs were comparable to those of uncoated PCL-NYs. PPyPCL-NY NGCs were shown to be more favorable to SC development than PCL-NY NGCs in tests on biocompatibility with SCs. In conclusion, PPy-PCL-NY NGCs have a good prospect in neural growth and repair [126]. In another study, PPy/PLLA scaffolds were prepared by depositing PPy NPs over PLA electrospun fibers to prepare a conducting fiber film (conductivity of 10S/cm). The ECM components such as collagen, fibronectin, and laminin were then coated onto the scaffolds, which significantly improved the cell adhesion and proliferation rate of PC12 cells. The researchers also proposed a principal mechanism of the synergistic interactions among ECM-CFF, growth cone, and electrical stimulation (ES) (as displayed in Fig. 6 (iii)) [127]. Similarly, PPy/PCL electrospun fibers were

Fig. 6 (i) Schematic representation of the utilization of PPy/SF NGCs with ES and the possible mechanisms of the fostered regeneration of nerves, reproduced with permission from [124], copyright 2020, Elsevier; (ii) Electrodeless coating PPy on CS-grafted PU with functionalized multiwall CNTs electrospun construct for neural tissue engineering, reproduced with permission from [125], copyright 2018, Elsevier; (iii) Schematic representation of PC12 cells cultured on ECM-CFF (a) without ES or (d) with ES. (b, c) Arrangement of cytoskeleton in NGF/adhesion-guided neurites and extension of pseudopodia without ES. (c, f) Strong actin and tubulin’s dynamic rearrangement on the growth cone’s leading end following ES was employed through ECM-CFF onto the neurite, ensuing in prompt neurite and pseudopodia growth along the axis of the fiber reproduced with permission from [127], copyright 2017, Elsevier

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obtained in another investigation, followed by vapor phase polymerization. The results demonstrated neurite outgrowth on L929 cells after 4 days of culture [128]. In another study, electrospun cellulose/poly N-vinylpyrrole (PNVPY) /EC/ poly(3-hexylthiophene) (P3HT) nanofibrous scaffolds were obtained by in-situ polymerization. The prepared scaffolds displayed higher electrical conductivity than plain EC or EC/P3HT mats and could thus be employed for neural tissue engineering [129].

4.2.4

PCL

PCL, a semi-crystalline hydrophobic sequential polymer with recurring O-(CH2-)5CO- subunits, is popular for its ability to keep the scaffold structurally sound due to its high mechanical capacity and low long-term breakdown rate. Despite this, the lack of hydrophilicity limits cell adhesion, motility, proliferation, and development. In another investigation, Heidari et al. examined how graphene affected the biological characteristics of PCL/gelatin nanocomposites mats. Antibacterial features of electrospinning-based PCL/gelatin/graphene nanofibers-like mats were shown to be 99% effective toward both Gram-positive as well as Gram-negative microorganisms. Compared to PCL/gelatin nanofibrous scaffolds, drug release tests revealed that the π-π- stacking relationship between TCH and graphene resulted in considerably better regulated TCH release from electrospun PCL/gelatin/graphene. The nanomaterials are a good contender for use as conductive scaffolds in brain tissue regeneration and controlled drug administration due to these better qualities, as well as improvements in hydrophilicity and biodegradability [130]. Despite recent improvements, designing and fabricating transformative biomimetic neural conduits remains a challenging task. As a result, Liu et al. designed and built mechanically tunable neural conduits with bioinspired structural properties that can be used for nerve tissue regeneration. For constructing triple-layered conduits (as displayed in Fig. 7 (i)), researchers used a multimodal strategy that included electrohydrodynamic (EHD) jet printing, electrospinning, and dip-coating procedures. As the inner layer, high-resolved EHD jet printing-based PCL nanofibers with variable directional cues were used, accompanied by a dip-coating of the gelatin hydrogels as the intermediate layer, and finally, the channels were enveloped in electrospun PCL nanostructures as the outside layer. The findings of the investigation showed that the innovative technology has a better chance of fabricating mechanically adjustable triple-layered circuits that are compatible with neural precursors and vascular cells [131]. In another investigation, Entekhabi et al. created a tissue-constructed nerve graft consisting of an electrospun PCL conduit packed with collagen-hyaluronic acid (COL-HA) sponge in various COL-HA weight proportions as 100:0, 98:2, 95:5, and 90:10 (as shown in Fig. 7 (ii A, B). The impact of HA on sponges’ porosity, mechanical characteristics, water absorption, and degradation rate was investigated. There was strong cohesiveness between the electrospun PCL nanomaterials and the COL-HA sponges in all sponges with varying HA concentrations. The maximum

Fig. 7 (i) The preparation method of triple-layered NGC: (a) Diagrammatic representation for the EHD jet printing arrangement; (b) Fabrication of the microfiber 3D framework with porous structure; (c) The creation of the layers: inner and middle. (d) The diagrammatic representation of the outer layer’s fabrication, reproduced with permission from [131], copyright 2021, Elsevier; (ii) (A) Diagrammatic presentation of fabrication of 3D constructs, (B) (a) Optical microscopy of the NGC constituting inner COL-HA sponge and encompassing PCL layer prepared by electrospinning, (b, c) FE-SEM images of the NGC displaying the inner COL-HA sponge with microporous structure and outer electrospun PCL layer of nanofibers, (d) FE-SEM images of the lumen’s inner part depicting a good cohesion between the PCL electrospun nanofibers and sponges of COL-HA, reproduced with permission from [132], copyright 2021, Wiley

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tensile strength of the PCL nanofibrous layer was 2.23  0.35 MPa at a 35% elongation, which was comparable to the rat sciatic nerve. Compared to purely collagen sponge, increasing the HA concentration led to better water absorption, bigger pore sizes, porosity, and a reduction in Schwann cell growth; however, the drop in cell growth was not statistically significant. The higher disintegration rate and porosity of the COL-HA sponges were implicated in the reduced Schwann cell growth. In addition, the designed 3D construct dramatically boosts axon development in a dorsal root ganglia experiment. These findings imply that the manufactured 3D composite framework can promote axon formation by providing a conducive atmosphere for Schwann cell growth and differentiation [132]. In another study, Habibizadeh et al. create a unique three-dimensional hybrid scaffold by modifying the surfaces of PCL/CS nanofiber/net with an alginate-based hydrogel microlayer in order to obtain the benefits of both nanofibers and hydrogels at the same time. Electrospinning was used to create a bead-free irregularly arranged nanofiber/net (NFN) structure made of CS and PCL. The NFN arrangement provided reduced rough-like surface character, excellent hydrophilicity, and high porosity. After that, a microlayer of alginate expressing NT-3 and conjunctiva mesenchymal stem cells (CJMSCs) were deposited on the PCL/chitosan nanofiber/net as a new stem cell resource. The scaffold comprised a bi-layer architecture having interlinked pores in the range of 20 μm in diameter. Their findings demonstrated that alginate hydrogel microlayer surface functionalization of nanofiber/net generated reduced inflammatory reaction and more significant CJMSC development than the untreated scaffold. The initial rapid discharge of NT-3 was 69% in 3 days, accompanied by a 21-day continuous release. MAP-2, tubulin III, and Nestin gene function were raised 6, 5.4, and 8.8-times, respectively. The morphological alterations in SEM micrographs depicted that following the culturing on PCL/chitosan/Alg/Cell scaffold containing NT-3, CJMSCs differentiated to neuron-like cells (as shown in Fig. 7 (iii)). According to the findings, the surface-functionalized biomimetic scaffold improved biocompatibility and successfully developed CJMSCs into nerve cells [133]. In another study, Chen et al. fabricated polycaprolactone/gelatin electrospun fibers/melatonin using various ratios of melatonin. The fibrous scaffolds of 2% melatonin had a proliferative effect on PC12 cells [134]. An injectable anisotropic PCL nanofibers/alginate hydrogel scaffold was prepared by centrifugal electrospinning containing SPOINs and were then cut into micro-fibers and  ⁄ Fig. 7 (continued) Online Library; (iii) The SEM micrographs of the neuron-like cell, CJMSC, and PCL/chitosan/Alg constructs 7 days following incubation in DMEM at 37
C. Morphology of Neuron-like cell following induction with NT-3 (a-1), CJMSCs morphology (a-2), PCL/chitosan/ Alg construct with no seeding of cells (a-3), the red arrows and cycles were indicated to CJMSCs and neuron-like cells. Neuron-like cells having simple bipolar and several branched appearance (b), CJMSCs with the round impression (c), the blue arrows were indicated to neuron-like cell’s branched framework, reproduced with permission from [133], copyright 2021, Wiley Online Library

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incorporated into alginate hydrogels. An aligned orientation of olfactory ectomesenchymal stem cells was observed after 14 days of incubation [135]. Similarly, PCL/Gel electrospun fibers/HA/Cs nanoparticles (NPs) NP-coated fibers exhibited a higher proliferation duration (72 h) than NP-blended ones [136]. In another study, polycaprolactone nanofibers obtained by wet electrospinning showed the GAG profile of the proliferating REN-VM cells for the first time [137].

4.2.5

PGS

Polyglycerol sebacate (PGS) is a relatively new biopolymer that combines glycerol and sebacic acid in a simple polycondensation. Its outstanding biocompatibility and physical qualities (stiffness 0.056–1.5 MPa after curing) rendered it a popular biomaterial for biological applications such as the delivery of drugs, implanted devices, and TE substrates. For cardiac tissue engineering, PGS has been sculpted into microfabricated anisotropic accordion-like honeycomb microstructures, microfibers, and core-shell fibers. Electrospinning appears to be a viable method for processing PGS. Uncured PGS, on the other hand, is a viscous, sticky solution that cannot form a three-dimensional framework. It crosslinks into an elastomeric form after curing, making it difficult to dissolve in non-aqueous solvents [138]. In an investigation, Saravani et al. designed and fabricated nanocomposite substrates made of polyglycerol sebacate for nerve tissue regeneration. Under varied circumstances, the produced PGS was electrospun along with gelatin and chitosan in varying quantities. Electrospinning semi-crystalline PGS/CS/Gel yielded fibers with a mean diameter of 80 nm, according to the findings. MTT assays were carried out using the PC12 cell lines; after 3 days of cell culture, it appeared that the polyglycerol sebacate/Chitosan/Gelatin nano-composite shows promise for neural bioengineering [75]. PGS is a thermosetting, biologically degradable elastomeric polymer that has the potential to be used in nerve treatments. PGS production, on the other hand, is effort and resources intensive. The PGS pre-polymer (pPGS) was produced in a study by Saudi et al., employing three different synthesis duration of 3, 5, and 7 h at 170
C. Electrospinning was utilized to create oriented PVA-PGS fibers with varied ratios (60/40, 50/50, and 40/60) cross-linked using the thermal crosslinking technique. The findings showed that the pPGS synthesized in 3 h at 170
C is the best chemical process sample. All of the scaffolds were bead-less and had a consistent fiber diameter. The Elastic modulus of chemically bonded PVA-PGS fibers (50:50 and 40:60) was found to be within the predicted region for nerve treatments. PVA-PGS (50:50 and 40:60) fibers may improve cellular proliferation, according to cell culture experiments. The findings indicated that PVA-PGS (50:50 and 40:60) is a promising and prospective biopolymer for fabricating nerve regeneration platforms [139]. In another study by Saudi et al., for nerve tissue regeneration, aligned electrospun PVA-PGS fibers with varied concentrations of lignin (0–5 wt%) were created. The results demonstrated that the manufactured fibers were smooth with consistent diameter and that the increasing concentration of lignin lowered the fiber diameter

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from 530 to 370 nm. By elevating the lignin %, the elastic modulus rose from 0.1 to 0.4 MPa. The PC12 culturing revealed that lignin accelerated cell growth. Gfap, -Tub III, and Map2 mRNA expression levels, as well as immunohistochemistry (Map2), demonstrated that lignin had a beneficial effect on brain cell development. Finally, the findings point to PVA-PGS/5% lignin as a potential neural tissueengineered substrate [140]. Applying the electrospinning approach, oriented PCL/PGS fibers containing various concentrations of multi-walled carbon nanotubes (MWCNTs; 0–1.5 wt%) were generated by Saudi et al.. The SEM results showed that all fibers were oriented and homogeneous in mean diameter, with a drop in mean diameter from 833 to 476 nm as MWCNTs were increased. The incorporated MWCNTs particles in fibers were aligned along the electrospinning axis, according to TEM pictures. In contrast to the degradation rate, the scaffolds’ Young’s modulus, maximum tensile modulus, hydrophilicity, and uptake capability improved as the MWCNT content rose. The vitality and adherence of Rat pheochromocytoma cells (PC12) were examined, and MWCNTs were found to have a beneficial influence on cell-scaffold association. Ultimately, the findings revealed that PCL/PGS/MWCNTs scaffolds could be a viable biodegradable biopolymer for nerve tissue regeneration [141]. In another investigation, two elastomeric polymers, poly(ɛ-caprolactone)/poly (glycerol-sebacic acid) incorporating graphene oxide or nanoclay, were fabricated for soft tissue regeneration exhibiting proliferation and cellular adhesion similar to that of culture plates [142]. Similarly, Atari et al. prepared 2:1 blended poly(glycerol sebacate)/poly(ɛ-caprolactone) scaffolds in different solvent systems for neural tissue regeneration purposes. CF-AC improved the mechanical properties such as homogenous fiber diameters, viscosity, and spinnability [143]. PGS/polymethyl methacrylic acid-based electrospun nanofibers were blended with gelatin, which improved their biocompatibility and hydrophilic character. These scaffolds enhanced the proliferation of rat PC12 cells [138]. Similarly, electrospinning was used to develop PGS/PLA non-woven scaffolds and was pre-heated in a vacuum for the cross-linking of PGS and PLA to take place. The blending of PLA increased the spinnability of PGS, but the enhancement of elastomeric properties of nonwoven PGS still remains a challenge [144].

4.2.6

Miscellaneous

PLA Due to their exceptional physical and chemical capabilities, many artificial polymeric materials have been selectively used in neural scaffolds. Nanofibrous frameworks have been made from PVA, polyvinyl pyrrolidone (PVP), PL, PCL, and PVA, for instance, has already been widely used for brain tissue regeneration as an FDA-approved component. While the polymeric scaffolds were nontoxic and had excellent physicochemical qualities, they failed to adhere to the injured filopodium of the cells and decomposed into acid-based compounds, which were not beneficial

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to cellular survivability. Hence, certain attempts have been devoted to increasing the cytocompatibility of polymeric scaffolds, such as incorporating physiologically active compounds [101]. Surface treatment of electrospun nanofibers can be an excellent way to improve a scaffold’s electroconductivity and cell–cell adhesion properties. A neural tissueengineered scaffold incorporating PLA nanofibers, gelatin, and polypyrrole (PPy) was created in an investigation by Imani et al. to combine the topographical benefits of electrospun nanofibers with the versatility of gelatin and PPy. For this, a conducting copolymer was created by chemical means via anchoring different weight proportions of pyrrole on gelatin strands and afterward functionalized on the exterior of PLA nanofibrous scaffolds to promote conductance and bioactivities. The findings showed that scaffolds comprising 15 and 20% PPy might provide favorable circumstances for nerve cell attachment and development and could be used as a viable substrate in biomedical nerve uses [145]. In an investigation, Mahdi et al. created a biodegradable 3D drug-incorporated scaffold with core-shell organized fibers for brain tissue engineering applications utilizing coaxial wet-electrospinning. Wet-electrospun PLA served as the core, while cellulose acetate served as the fibril’s shell. The scaffold was subsequently covered with citalopram-incorporated gelatin nanocarriers (CGNs) with a mean size of 950 nm, which were made using the nanoprecipitation process. The conduit could not achieve the recognized ideal porous ratio of over 80%, and the observed porosity ratio was 60%. The CGN enveloping made the scaffold very hydrophilic with 0
contact angle. The weight of the untreated scaffold was maintained substantially consistent during in vitro breakdown in phosphate buffer. Following 40 days, the CGNs-loaded scaffold, on the other hand, had lost 45% of its weight. The CGNscoated scaffold had a greater cell survival than the bare scaffold using rat Schwann cells. Ultimately, the scaffold was transformed into a nerve guiding conduit and physically placed in a Wistar rat sciatic nerve deficit. The results of the functional sciatic indicator, hot plate latency, and weight-loss percent of the wet gastrocnemius muscle showed that the citalopram-loaded scaffolds might improve the functionality of sciatic nerve-affected models, indicating that they could be used in neural tissue regeneration [146]. In another study, Kang et al. prepared electrospun PLA/collagen membranes, which showed enhanced hydrophilicity and mechanical characteristics with the multilevel structure. Moreover, it showed cell attachment and guided proliferation in mouse fibroblast cells [147]. Gangolphe et al. fabricated micropatterned electrospun collectors composed of PLA-derived co-polymers (square- or honeycomb-like) to enhance their anisotropic properties during protraction. These electrospun scaffolds had better proliferation than the standard electrospinningbased mats [148]. Fang et al. prepared PLA/ γ-PGA nanofibers loaded with coumarin-6 and rhodamine for visualizing the structure of the core-shell nanofibers, where it demonstrated 90% reepithelialization in vitro [149]. Similarly, electrically polarized PLA was fabricated by Barroca et al. as smart platforms for nerve tissue engineering. The polarization on the scaffolds remained stable for up to 6 months [150].

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PGA Polyglycolic acid (PGA) is a biodegradable aliphatic polyester with good biocompatibility and high mechanical strength that has a lot of potential for tissue engineering applications. The effectiveness of PGA-based, easily absorbed sutures demonstrated that PGA-containing polymers might be used safely in soft tissue engineering [151]. PGA can even be used in bioengineering techniques like vascular tissue regeneration. The stiff architecture of the purest version of PGA has hampered the application of this substance in the implantation procedure [152]. Combining PGA polymer with extracellular and endogenous matrix-based polymers like collagen may improve this polymer’s characteristics, allowing it to be used in vivo. Extracellular matrix components such as elastin are commonly employed to enhance the rate of cell adherence to conduits in most investigations [153]. The integration of synthetic materials into biopolymers like collagen improved the chemical characteristics of the conduit’s surfaces and improved cellular responsiveness [154]. Because of their instability, PGA-based nanoconduits can only cross a modest nerve distance in most circumstances. According to such findings, researchers prefer PLA or PGA copolymers because they are more structurally stable [39]. An electrospinning process was used by Dehnavi et al. to make a unique conduit built on polymer blend scaffolds of PG, collagen, and nanobioglass (NBG), which was then contrasted to PGA/collage/PGA scaffolds developed in earlier investigations. Concerning other polymers, the tensile, chemical, cytocompatibility, and biodegradability features of PGA/collagen/NBG conduits were shown to be superior. Nanofibrous electrospun PGA/collagen/NBG conduits were highly suited to cellular attachment and multiplication than either PGA or PGA/collagen scaffolds, according to the MTT experiment and DAPI labeling technique, and may have promise for nerve regeneration [154]. Traditional microfibers have acceptable tensile qualities such as strength, elastic nature, and stiffness, whereas electrospun nanofibers offer a vast surface area, high pore volume, and controlled orientation. As a result, the mixing of nanofibers and microfibers can be used to create bioinspired frameworks for tissue regeneration. By electrospinning PCL nanofibers onto polyglycolic acid (PGA) microfibers, a coreshell organized fibrous architecture with controlled surface topography was generated in a study by Liu et al.. The resulting core-shell structure was described in terms of surface shape, surface hydrophilicity, and material performance. Based on FE-SEM images, a fiber collector and rotating discs can control the orientation of PCL nanofibers on the yarn interface. Furthermore, A-PCLs stimulate the adherence and multiplication of BALB/3T3 (mouse embryonic fibroblast cell line) and controllably align cellular proliferation following the biotopographic signals of PCL nanofibers. The created core-shell yarn showed considerable potential in the construction of diverse tissue scaffolds because it offered both the desirable surface characteristics of PCL nanofibers as well as the mechanical behavior of PGA microfibers [155].

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PEDOT For the ability of neural cells to connect with several other cell varieties, the nervous systems rely on a complex network of electronic transmission. Electrical features in neural scaffolds might be ideal for enhancing neuron proliferation and migration. As a result, highly conductive polymers that mimic brain tissue and allow for the restoration of neural synapses would be optimal [39]. Owing to its tunable qualities, such as high stability, high conductivity, and the capacity to encase and deliver substances, electroconductive polymeric materials have drawn interest in neural tissue regeneration. Their electronic and physicochemical characteristics can also be altered to suit a particular purpose. Nevertheless, inadequate bioactivity related to their incapability to a breakdown in vivo, which might stimulate the inflammatory process and immunostimulatory responses, necessitating increased treatments and medical interventions, is a critical problem in the use of conducting polymers as biopolymers for neural tissue regeneration [39, 156]. PEDOT is an intriguing, highly electronically conductive polymer with great stability, low redox potential, medium bandgap, and optical transmittance in its conductive form. PEDOT has a variety of uses in neural tissue regeneration, particularly as a material for microelectrodes for brain electrical impulses and monitoring [39]. Due to its efficiency and capacity to significantly increase scaleup into a continuous cycle, the electrospinning method has developed an effective and viable approach for fabricating fibers with dimensions in the submicrometer to micrometer region. It is feasible to make a randomly distributed nanofibrous mat with a highly porous surface area utilizing electrospinning [157]. In a study by Bhatnagar et al., PEDOT nanoparticles were electrospun with PLA from a solution of DMF and acetone. 10% and 15% PEDOT loading in PLA gave electrospun fibers having diameters up to 95.3 nm. PEDOT concentration in the resulting nanofibrous mats was less than the threshold for percolation. The highest conveying efficiency of PEDOT was 35.77% when measured by the matrix dilution technique. Thus, co-electrospinning seems to be a good strategy to obtain PEDOT nanofibers which could be further used for neural tissue engineering applications [157]. PEDOT: PSS-based poly[2,20 -m-(phenylene)-5,50 -bibenzimidazole] (PBI) electrospun scaffolds were prepared in a study conducted by Sordini et al.. PEDOT: PSS doping increased the electrical conductivity of PBI by 105 and 106 times and was non-cytotoxic to hBM-MSCs, thus could potentially be used for nerve regeneration [158].

4.2.7

Carbon-Based Polymers

Graphene Graphene is a carbon-based allotrope composed of a continuous sheet of carbon atoms organized in a hexagonal arrangement in two dimensions. It is virtually

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translucent, antibacterial, and antiviral, and it is exceptionally biodegradable with low cytotoxic activity. Nonetheless, matrix biological hazard has been documented dependent on whether graphene is employed in 2D or 3D cell cultures, with 3D systems being superior for brain cell development and multiplication [159, 160]. Furthermore, 3D graphene platforms can include a variety of nanostructures, including gold nanoparticles, enhancing neurogenesis, and directing axonal orientation [161]. Graphene has also been employed in various ways for neural tissue regeneration, including foams and graphene-based nanogrids [39]. To create electrospun fibers, Ginestra used varying proportions of graphene disseminated in a PCL-cyclopentanone mixture. The structure and geometry of the fibers and the physical behavior of the electrospinning structures were studied to see how graphene content affected the fibers’ performance. Due to the graphene %, a considerable dimensional variation between the fiber radii was achieved. The elastic modulus raised from 5.6  2 MPa in the absence of graphene to 22.5  5 MPa in the presence of 2% graphene. As a result, the existence of graphene was discovered to alter the mechanical characteristics of fibrous structures [162]. Considering the paucity of comparative investigations on the impact of graphenerelated substances on neuroprotection, a study by Magaz et al. comprehensively analyzes the function of graphene oxide (GO) and reduced GO (rGO)/silk-based composites micro/nanofibrous frameworks in controlling neuronal cell activity in vitro (as shown in Fig. 8 (i) a). Fibrous substrates are possible options for tissue regeneration in peripheral nerves because they can imitate the morphology of the original ECM. To investigate a series of highly conductive unwoven silk/rGO scaffolds, electrospinning-based silk/GO micro/nanofibrous conduits with GO loadings of 1 to 10% were electrospun, which were then subsequently post-reduced in situ (as shown in Fig. 8 (i) b). Within the dry state, the ionic conductivity of up to 4  105 S cm1 was measured, which increased to 3  104 S cm1 after water intake. On all surfaces, Neuronoma NG108–15 cells attached and remained functional. On the GO-based scaffolds, increased metabolic activity and multiplication were reported, and all these cell reactions were significantly boosted by electroactive silk/rGO. On several conducting substrates, neurite outgrowth of up to 100 m was attained by day 5, with a maximal expansion of up to 250 m. Such electrically active composite fibrous frameworks have the potential to improve neuronal cell responsiveness and could be useful support platforms in brain tissue regeneration [163]. The alignment and elongation of neurites perform particular roles in treating neuronal disease, designing tissue-engineered implants, and in applications of bioelectrodes. Numerous investigations have been inspected to produce 2D patterned substrates or 3D constructs having aligned topographical frameworks to govern the growth of axons. But, the majority of the techniques are either complex/difficult in method or time-/cost-forfeit. In this context, Wang et al. developed an efficient and facile one-step dimensionally confined hydrothermal (DCH) method to fabricate reduced GO fibers (rGOFs) (as demonstrated in Fig. 8 (ii)). The prepared scaffold demonstrated substantial capability in governing the differentiation of neural stem cells toward neurons. The developed constructs revealed excellent biocompatibility with trigeminal neurons and showed a high degree of topographical

Fig. 8 (i) (a) GO and electroactive reduced GO-based composite fibrous constructs for engineering excitable neural tissues, (b). (A) Diagrammatic representation of the silk fibroin’s extraction and synthesis, (B) The Manufacturing procedure of silk-based constructs: (B1) preparing silk-based dopes for electrospinning, and (B2) diagrammatic representation of the SF/GO solutions electrospinning and in situ post-reduction into SF/rGO, reproduced with permission from [163], copyright 2021, Elsevier; (ii) Schematic representation of rGOF preparation procedure via a one-step DCH technique, reproduced with

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cues alignment in neurite guidance over high electrical conductivity. Thus, the overall results showed that the scaffolds could be utilized in neural tissue engineering and other biomedical applications [164]. Usually, peripheral nerve injuries cripple significant functions of the nervous system. Thus, there is a pressing demand to prepare tissue-engineered products that will alleviate the oxidative insults and recover bioelectrical signals. In this regard, Jiang et al. developed multilayered melatonin (MLT) loaded-reduced graphene oxide (RGO)/PCL composite construct with beaded nano-frameworks to enhance the attachment and proliferation of cells (as shown in Fig. 8 (iii)). The developed scaffold enhanced the recovery of sensory and locomotor function via walking track analysis and evaluation of electrophysiology, improved the production of ATP for the supply of energy, and demonstrated high elastic moduli required for the structural integrity of the nerve. The overall results revealed recovery of functions and morphology by the developed scaffold and suggested its capability for translational utilization [165]. Graphene oxide has presently been employed in peripheral nerve construction; however, it has some drawbacks, including toxicity and a loss of electrical properties, which are both critical in controlling nerve-related cell activities. Fang et al. used electrospinning to create reduced GO–GelMA–PCL nanofiber neural guiding conduits. The electrical conductance and cytocompatibility of the hybrid materials were significantly improved when rGO was introduced into the GelMA/PCL matrix. Furthermore, hybrid nanomaterials with small doses of rGO (0.25 and 0.5 wt%) might dramatically boost Schwann cell (RSC96) growth. More crucially, rGO/GelMA/PCL hybrid nanomaterials could trigger Schwann cell gene expression associated with the epithelial-mesenchymal transition (EMT) and were found to enhance both motor and sensory neural regeneration and function restoration in rat models [166].

CNTs CNTs are carbon-based allotropes with a cylindrical structure that exhibit exceptional thermal conductivity as well as excellent mechanical and electrical capabilities. Due to their biocompatibility, conductivity, and non-biodegradability, carbon nanotubes are ideal candidates for neural tissue engineering [167]. CNTs are primarily used as implantation in situations when long-term stimuli for nerve expansion are required, such as rejuvenation following spinal cord or brain trauma. In brain tissue development, both SWCNTs and MWCNTs have also been employed.

 ⁄ Fig. 8 (continued) permission from [164], copyright 2021, American Chemical Society; (iii) A multi-layered ATP-generating reduced graphene oxide-based scaffold that rectifies neuronal damages by enhancing the functions of mitochondria and recovering bioelectricity conduction, reproduced with permission from [165], copyright 2022, Elsevier

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SWCNTs serve as platforms for modulating and stimulating neural cells via conductance variations, such as lateral currents, to treat neurologic and brain-related ailments [168]. In a recent study, Xia et al. recently reported using a multivalent polyanion-based CNT to fabricate an electrospun fibrous framework for neural tissue formation. CNTs are coated onto electrospun PCL nanofibers after being non-covalently modified by polymerized polyglycerol sulfate. The substrates not only provided a suitable environment for pluripotent stem cells (IPS) to attach and multiply, but they also induced enhanced IPS maturation. The average neurite length on the PCL/CNT/hPGS scaffold was 101.0  20.9 m after 7 days, whereas the average length was 72.3  5.9 m for tissue cultured polystyrene. Furthermore, the linked fibers can direct the alignment of the formed neurites [169]. In another investigation by Nazeri et al., laminin, a neurite-growth promotor protein, was used to alter electrospun PLGA/CNT nanofibrous composites using either a mussel-based poly(dopamine) (PD) covering or direct physisorption as a straightforward way for biomaterial fabrication. Following that, various scaffold parameters such as deterioration duration, connected laminin quantity, and CNT release profile was studied. SEM and confocal imaging were used to investigate the combinatorial influence of topographical and physiological cues on PC12 cell adhesion, growth, and division. The findings of the deterioration investigations revealed that the laminin-based scaffolds were biodegradable and had high structural stability for 4 weeks. The amounts of laminin linked to the PLGA/CNT and PLGA/ CNT-PD scaffolds were 3.12  0.6 g per mg and 3.04  071 g per mg, correspondingly. Neurite connections on the PLGA/CNT substrate amended via PD coating were considerably more prolonged than those on the PLGA/CNT scaffold amended via physisorption and unaltered scaffolds, even though laminin-modified scaffolds improved cell proliferation in the same way [170]. A study by Zadeh et al. looked into the impact of raw CNTs on the ultimate characteristics of polyurethane (PU)/CNT hybrids with biomedical uses in mind. Electrospinning was used to make clean PU and PU/CNT composites with varying quantities of CNTs (0.05–1%). The electrospinning conditions were tuned to produce a bead-free architecture with no substantial change in mean fiber diameter and porous fraction. The addition of CNTs increased crystallinity, water uptake ratio, Young’s modulus, hardness, conductance, deterioration duration in an accelerating media, clotting duration, and the adherence of human umbilical vein endothelial cell types, according to the findings. However, there was no evidence of a direct link between CNT proportion and calcium adsorption. However, the more the negative charge on the CNTs, the more Ca adsorption was on the samples. PU/1CNTs showed lower young’s modulus than PU/0.05CNTs and PU/0.1CNTs because of the formation of CNT agglomerates at high CNT levels. Furthermore, 7-day extracts of all samples showed no substantial damage to the cells; hence these biocompatible composites could be used for soft tissue engineering applications [171]. In the biomedical sector, electroconductive nanomaterials offer a variety of advantages. It is consequently critical to creating safe electro-conductive polymer composites. Blend electrospinning, concurrent PLGA electrospinning, CNT

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electrospraying, and ultrasound-based CNT binding on electrospun PLGA nanostructures were all investigated by Nazeri and coworkers to create electroconductive nanocomposites mats of PLGA/CNT. SEM and TEM were used to examine the shape and dimension of fibers, with the PLGA/MWCNT blend nanostructures having the smallest average widths of 477  136 nm. When contrasted to unprocessed PLGA scaffolds, MWCNT-sprayed PLGA samples had a substantially reduced water contact angle (83
), electrical resistance (3.0  104 Ω), and mechanical characteristics (UTS: 5.50  0.46 MPa) [172]. In a recent study, Zhang et al. fabricated electroconductive scaffolds by electrospinning PCL/CNT at different rotational speeds of 0, 500, 1,000, and 2,000 rpm. The scaffolds exhibited appropriate tensile strength at 1,000 rpm and were aligned across the parallel and perpendicular axes, which are beneficial to the differentiation of neural cells in vitro and restoration of injured cells in vivo [173]. Table 2 denotes the recent studies focusing on the utilization of synthetic electrospun polymers for neural tissue engineering.

5 Conclusion and Future Perspectives In summary, this book chapter emphasizes the most recent studies on the application of electrospun scaffolds in engineering neural tissues. Considering the devastating consequences of neural injuries on the quality of life, its functional rehabilitation is pivotal. Regrettably, existing treatment procedures are inadequate to recover neural tissues from the injury to CNS and PNS. Therefore, modern tissue engineering and drug delivery advances are imperative. In this context, the fiber-associated NGCs offer new aspirations in treating injured nerves, particularly over large defects, by furnishing nano-engineered conditions and presenting biomimetic frameworks equivalent to natural ECM. The required cell feedback could be achieved by altering the design factors of electrospun fibers like fiber diameter, fiber alignment, density, biocompatibility, surface morphology, and chemistry of the surface. The developed NGCs also act as a carrier for delivering cells and bio-factors to produce a nutritious environment leading to improved recovery of nerve functions. The materials, morphological framework, topographical characteristics, and an adhesive surface of NGC are vital factors for the designing and fabrication of NGC. Desired characteristics of NGC can also be accomplished by supplying them with suitable cells and neurotrophic factors combination, as well as biochemical and physical signals. NGCs have experienced enormous progress across the last two decades and are now furnishing the probability of offering a substitute to the present treatment choices. With our enhanced comprehension of the biology of neural damage and repair, novel designs meeting the pressing demands of molecular and cellular procedures are being developed. Nevertheless, there is a substantial amount of research required in the various features of cellular systems, engineering tissue, and surgical methods to determine the advantage of NGC over traditional autografts to treat largegap nerve injuries. Despite that, much doubt still prevails about whether NGCs can

Polymer Polypyrrole/chitosan/collagen

PCL/PANi/urethane

Polycaprolactone/poly(glycerol sebacate)/hydroxyapatite

Poly(ε-caprolactone)/type I collagen

Device Conduit

Conduit

Conduit

Conduit

–

Sprague Dawley rats

–

–

In vivo cell line –

PC12 cell line

Pheochromocytoma (PC12) cell lines

In vitro cell line –

Table 2 Recent investigations on synthetic polymers for neural tissue engineering Results The addition of polypyrrole in fibers increased their conductivity to 164.274  103 s/m, putting them in semiconducting and conducting polymeric ranges. Nanofibers containing 10% polypyrrole had higher cell attachment, growth, and proliferating capabilities than nanofibers containing other formulations, according to MTT and SEM tests The decomposition rate of nanocomposites grew dramatically during a 50-time period, and composite fibrous substrates with UPCL/PCL/PANI45:20:35 proportions provided the most balanced qualities, matching each needed standard for neuronal cells, and could be used in neural bioengineering The fiber diameter fell from 831 to 382 nm when the HAp particle loading was increased from 0 to 15%. The needle-shaped HAp particles were scattered along the fibers. All fibers had Young’s strength of 0.16–0.3 MPa, appropriate for nerve tissue regeneration. The vitality and adhesion of PC12 cells were improved by HAp particles No significant hypersensitive reactions were found in the hind limbs of rats treated with absorbing electrospun biopolymer nanofiber scaffolds. The topography of myelin sheaths in the damaged sciatic nerves was similar to normal. The CD4 hypersensitivity was significantly lower, and the sciatic nerve functioning restoration was higher than in rats repaired with either poly(-caprolactone) or silicone [177]

[176]

[175]

References [174]

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Poly (ε-caprolactone)/collagen/ NBG

Poly (ε-caprolactone/collagen/ NBG)

PCL/PGS/bioactive glass

PLA/PGA

PVA/gellan gum nanofibers

Conduit

Scaffold

Conduit

Conduit

Scaffold

Murine embryonic stem cells

RSC96 cells

–

–

–

Rat model

–

Bone marrow-derived stromal cells (SC-2 cells)

Rat model

– In the conduit cell group, H&E photographs of the extracted regeneration nerve and immunohistochemical data revealed that regenerating nerve fibers had grown and were supported by neovascularization. This electrospinning nerve conduit might find more implementation in cell therapy for tissue repair in the future due to the benefits of a greater surface area for cell adhesion which could improve longer nerve defect reconstruction In the conduit + NGF cohort, histology and immunohistochemical investigations demonstrated reduced fibrosis and higher levels of CD31 and NF-200 protein levels at the crush location. Also, they showed nanometer-scale characteristics, neurotrophic activity, good mechanical qualities, and biocompatibility which might help rats regenerate their sciatic nerve The electrospun fibers were hydrophilic, and the addition of BG did not affect wettability. However, the mechanical properties did not improve because of the poor interface between the polymeric fibers and the BG particles Compared to ordinary NGCs, the innovative GD-NGC deteriorated slowly and progressively from the interior to the outermost layer and had comparable or better properties, including porosity, hydrophilicity, mechanical behavior, and biocompatibility The degradation experiment indicated that the manufactured PG-NFs were robust in an aqueous environment and did not lose any weight. PG-NFs had a 40% apparent porosity and a conductivity of 126.93 pS. The PG-NFs were also non-cytotoxic and biocompatible, while cultured murine embryonic stem cells (ESCs) grew similarly to controls [182]

[181]

[180]

[179]

[178]

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supply superior results over conventional therapeutic procedures. While topographical characteristics from sub-micron range to nano range fibers, aligned or randomly directed, certainly influence the morphology and function of the neural cells, comprehensive investigations are still needed. Specifically, our present knowledge of the impact of the fibrous structure on the functions of the cells is still restricted to the assessment of alterations in cell morphology and viability. In-depth mechanistic assessment of functions of cells like myelination of Schwann cell, formation of the glial scar, formation of neuron synapse, signal transduction of neural cell, and differentiation of neural stem cell is vital to promote the accurate control of the functions of the cell via nanofibrous topographical cues. To overcome the cell’s poor endurance inside NGCs following transplantation, approaches are required to create a cell niche for promoting interaction, adhesion, and survival of cells. Also, most of the NGCs were demonstrated to be risk-free in vitro studies, but there is still a significant gap in the research on assessing the toxicity of NGCs in vivo, especially in the nervous system. Concerning in vivo studies, depreciation of the variability in test procedures employed by various groups will also permit more productive comparisons between test results. NGCs fabrication that can correspond to the patient-particular demands in damaged nerves, local vasculature, and structure of fasciculi could turn into a principal track in transplantation approaches in the future. Regeneration of nerves is a complicated procedure and a difficult area for researchers. With restricted knowledge of the nervous system to the accomplishment of recovery of partial neural functions, magnificent advancements have been achieved during the previous decades. Nevertheless, a lot remains undiscovered, and the expedition to consummating the ultimate target of completely regenerating functional nerves persists. Acknowledgments Author SP would like to thank the Indian Institute of Technology Madras for providing financial assistantship and resources. Conflict of Interest The authors declare no conflict of interest. Funding Not applicable.

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175. Nazarpak MH, Entekhabi E, Najafi F, Rahmani M, Hashjin MS (2019) Synthesis and characterization of conductive neural tissue engineering scaffolds based on urethanepolycaprolactone. Int J Polym Mater Polym Biomater 68:827–835 176. Saudi A, Zebarjad SM, Salehi H, Katoueizadeh E, Alizadeh A (2022) Assessing physicochemical, mechanical, and in vitro biological properties of polycaprolactone/poly(glycerol sebacate)/hydroxyapatite composite scaffold for nerve tissue engineering. Mater Chem Phys 275:125224 177. Yen CM, Shen CC, Yang YC, Liu BS, Lee HT, Sheu ML, Tsai MH, Cheng WY (2019) Novel electrospun poly(ε-caprolactone)/type I collagen nanofiber conduits for repair of peripheral nerve injury. Neural Regen Res 14:1617–1625 178. Mohamadi F, Ebrahimi-Barough S, Nourani MR, Mansoori K, Salehi M, Alizadeh AA, Tavangar SM, Sefat F, Sharifi S, Ai J (2018) Enhanced sciatic nerve regeneration by human endometrial stem cells in an electrospun poly (ε-caprolactone)/collagen/NBG nerve conduit in rat. Artif Cells Nanomed Biotechnol 46:1731–1743 179. Mohamadi F, Ebrahimi-Barough S, Nourani MR, Ahmadi A, Ai J (2018) Use new poly (ε-caprolactone/collagen/NBG) nerve conduits along with NGF for promoting peripheral (sciatic) nerve regeneration in a rat. Artif Cells Nanomed Biotechnol 46:34–45 180. Luginina M, Schuhladen K, Orrú R, Cao G, Boccaccini AR, Liverani L (2020) Electrospun PCL/PGS composite fibers incorporating bioactive glass particles for soft tissue engineering applications. Nanomaterials 10:978 181. Yu L, Zhang W, Jiang Y, Guo C (2020) Gradient degradable nerve guidance conduit with multilayer structure prepared by electrospinning. Mater Lett 276:128238 182. Aadil KR, Nathani A, Sharma CS, Lenka N, Gupta P (2019) Investigation of poly(vinyl) alcohol-gellan gum based nanofiber as scaffolds for tissue engineering applications. J Drug Deliv Sci Technol 54:101276

Adv Polym Sci (2023) 291: 287–312 https://doi.org/10.1007/12_2022_121 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 28 May 2022

External Stimuli Responsive Nanofibers in Biomedical Engineering Hamid Hamedi, Sara Moradi, and Alan E. Tonelli

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 External-Responsive Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Thermo-Responsive Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Magnetic-Responsive Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3 pH-Responsive Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4 Electrically-Responsive Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.5 Biomolecule-Responsive Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.6 Multi-responsive Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Biomedical Applications of External-Responsive Nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1 Wound Dressings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 Drug Delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3 Diagnosis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.4 Scaffolds for Cell Culture and Delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract Nanofiber technology has attracted great attention in many research and applications area because of its unique physicochemical properties and characteristics such as high surface area compared to bulk material. This property provides better cell adhesion and drug and protein loading. In addition, their fabrication from a wide range of polymers with different properties makes them excellent candidates for smart delivery systems. In this chapter, first different external stimuli-responsive

H. Hamedi and A. E. Tonelli (*) Textile Engineering Chemistry and Science, Fiber and Polymer Science Program, Wilson College of Textiles, North Carolina State University, Raleigh, NC, USA e-mail: [email protected] S. Moradi Department of Chemical Engineering, Faculty of Engineering, Arak University, Arak, Markazi, Iran

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nanofibers such as thermo-, magnet-, pH-, electrically-, biomolecule-, and multiresponsive are explained. Then their biomedical applications such as wound dressings, drug delivery systems, cell scaffolds, and diagnosis are discussed. Keywords Cell capture and release · Diagnosis · Drug delivery · Stimuli-responsive nanofibers · Wound dressings

1 Introduction Nanofibers have been widely studied and developed as they have advantages of nanomaterials for various applications. Nanofibers can be produced by different methods from biodegradable and biocompatible polymers with unique properties such as great porosity, mechanical properties, ability to load and deliver drugs [1] and imitating the extracellular matrix network of the skin [2] making them ideal for biomedical applications. Preparation methods of nanofibers include drawing, template, self-assembly, phase separation, and electrospinning. Template processes cannot produce continuous fibers and drawing is only applicable for viscoelastic materials, so phase separation, self-assembly, and electrospinning are the most important methods [3]. Among these three methods, phase separation has a long processing time, a complicated process, and self-assembly produces thinner fibers in comparison with electrospinning [3], so electrospinning is considered the most popular method to fabricate nanofibers because of its simplicity and costeffectiveness. Electrospinning can be regarded as a form of electrospray process. Small droplets are formed in electrospray processes due to low viscosity of the solution, while in electrospinning a firm fiber is fabricated after evaporation of the solvent [4]. Nanofibrous materials can be developed to own biological activity in their structure or have ability of loading and delivering various drugs in a controlled way. Recently, smart delivery systems which can change their behavior in response to environmental stimuli have been developed. The stimuli can be classified into internal such as pH, enzymes, and external such as temperature, ultrasound, and electromagnetic radiation. In this review chapter, external-responsive nanofibers and their applications in biomedical are discussed. First, different external-responsive nanofibers and their preparation are explained. In the second section, their application in various biomedical applications such as wound dressings and delivery systems are summarized.

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2 External-Responsive Nanofibers As mentioned above, external-responsive nanofibers change their behavior when triggered by temperature, magnetism, and light. Following, their preparation methods, characterization, and properties are discussed in detail.

2.1

Thermo-Responsive Nanofibers

Most thermo-responsive polymers have hydrophobic groups like ethyl, propyl, and methyl and their interactions with water molecules at different temperatures affect their thermo-responsive properties [5]. Hydrophilic and hydrophobic groups must be balanced in the structures of thermo-responsive polymers. Because of the balance between hydrophobic and hydrophilic groups, a small alteration of polymer solution’s temperature can lead to a new adjustment of the hydrophobic and hydrophilic interactions between the water molecules and polymer blocks. Mixing free energy, enthalpy or entropy of the system can affect this phenomenon as well. So, thermoresponsive nanofibers may be applied for formation of injectable or sustained drug release systems [6]. For instance, poly N-isopropylacrylamide (PNIPAAm) is a wellknown temperature-responsive polymer with a sharp phase transition at critical solution temperature (LCST) of 32
C between hydrophilicity and hydrophobicity. By increasing the temperature higher than the LCST, PNIPAAm quickly switches from its hydrophilic status to hydrophobic [7]. Different preparations of thermo-responsive poly N-isopropylacrylamide, poly ε-caprolactone, and egg albumen loaded with gatifloxacin hydrochloride, Gati as a drug model were fabricated for wound healing applications [8]. SEM photos showed that the polymer concentration, drug, and polymer mixture composition influenced the morphology of the non-woven fibers. Drug release showed primary fast release and then controlled and slow release afterward. The nanofibers had antibacterial activity against Staphylococcus aureus, good cell viability and rat wound healing. In another study by Young et al. [9], thermo-responsive and biocompatible PNIPAAm nanofibers were fabricated by electrospinning with the ability to imitate native extracellular matrix, providing chemical and physical signals to force cell phenotype and function. Release of L929 fibroblasts was calculated after 16 h of initial attachment to the PNIPAM nanofiber, tissue culture plastic (TCP, negative control), and UpCell surfaces. Prepared thermo-responsive nanofiber and UpCell surface cell release was in the same range and remarkably more than that of TCP. Core-shell nanofibers with stretchable polymeric cores and thermo-responsive polymeric shells for non-vascular nitinol stents were fabricated which generate heat in the stent for hyperthermia therapy by applying an alternating magnetic field (AMF) [10]. In this study, PNIPAAm was copolymerized with N-Hydroxymethyl acrylamide (HMAAm, a hydrophilic monomer) for increasing its LCST point above normal body temperature. Different ratios of PNIPAAm/HMAAm were prepared

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Fig. 1 Index of cell proliferation against ESO26 and OE21 cancer cell lines (Permission/license is granted for reusing the material from Elsevier, License No# 5218510922421) [10]

and nanofibers with different LCST were obtained which can be applied for specific drug release. Nanofibers were compatible with NIH-3T3 fibroblast cells and had cancer cytotoxic properties against ESO26 and OE21 cancer cells by combination of drug and hyperthermia therapy (Fig. 1). As it is clear in Fig. 1, the cell proliferation indexes for all treatment groups except for the control group decreased showing toxicity to cancer cell lines. Prepared drug delivery carriers released two various chemotherapeutic drugs by using specific temperature range, and generated heat from the stent was cytotoxic to cancer cells.

2.2

Magnetic-Responsive Nanofibers

Magnetic-responsive nanofibers respond to external magnetic fields and have broad applications in localized drug delivery. There are some limitations in applying free magnetic nanoparticles such as low tumor targetability, high irregularity in the administrated number of magnetic nanoparticles to the tumor, and the chance of the magnetic nanoparticles entering the normal tissues near the tumor. To avoid this, magnetic nanoparticles are applied in electrospun nanofibers for cancer treatment and drug delivery applications. On the other hand, magnetic nanoparticles are not toxic in the human body and can degrade in a reasonable period time [11]. Iron oxide was incorporated by two different methods into silk fibroin scaffold to prepare magnetic-responsive nanofibers for biomedical applications [12]. Silk fibroin extracted from Bombyx mori cocoon is a biodegradable, biocompatible biomaterial with good mechanical properties. In the first (encapsulation) method, iron oxide nanoparticles (IONs) were loaded into the electrospinning solution while in another method (dip-coating), nanofibers were soaked in the aqueous IONs solution. Magnetic-responsive nanofibers prepared with the encapsulation method had lower magnetization than those of the dip-coating method. The reason is because

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IONs in the encapsulation group are inside the nanofibers while they are on the surface of the nanofiber in the dip-coated method. The results showed that nanofibers prepared with the first method may be applied as a magnetic-responsive bio-interface for in vivo regenerative medicine applications and a tissue engineering scaffold. Nanofibers produced by the second method are more suitable as an interface for stimulating in vitro stem cell differentiation or biosensors usages. Poly D,L-lactide (PLA) and iron oxide nanoparticles (IONPs) were applied to fabricate magnetic-responsive nanofibers [13] to evaluate their influence on cell behavior of osteoblast and 3T3 fibroblast. The IONPs loaded nanofibers were highly magnetic-responsive and biodegradable. Cytotoxicity test showed that in all nanofibers, proliferation of osteoblast cells in the presence or absence of static magnetic field (SMF) was enhanced by increasing the culturing days and cell proliferation of nanofibers without SMF was higher than those triggered by SMF for the same cultured days. This may mean that moderate strength of SMF (10 mT) can have slight cell growth cytotoxicity. Increasing the amount of IONPs higher than 6% led to decrease of the cell proliferation showing signs of toxicity of the nanoparticles on the cultured cells. The results exhibited that nanofibers loaded with IONPs can be a candidate for magnetotherapy, however, more studies are needed to realize the interaction mechanism between the magnetic field and cells. Hou et al. [14] synthesized magnetic-responsive cellulose nanofibers functionalized with heparin by wet-wet electrospinning from 1-methyl-3methylimidazolium acetate (a room-temperature ionic liquid). Fe3O4 nanoparticles as superparamagnetic magnetite were loaded into three different types of nanofiber structures: cellulose/Fe3O4/heparin single filament fibers, cellulose/Fe3O4/covalently immobilized heparin core-shell fibers, and cellulose/Fe3O4/physically immobilized heparin core-shell fibers (Fig. 2). Although nontoxicity of the Fe3O4 has been proved, its injection into the bloodstream can cause cytotoxic effects. So, in this study, Fe3O4 nanoparticles were encapsulated inside a cellulose shell through coaxial electrospinning. All fibers showed great magnetic properties in both wet and dry states, but monofilament composites were attracted more to the magnet than core-shell nanofibers due to the presence of nanoparticles on the fiber surface. Magnetic nanofibers from self-assembling “Taylor cones” of polyvinylpyrrolidone/Fe3O4 ferrofluid (PFF) were prepared under coincident magnetic and electric fields through a needleless electrospinning method by Huang et al. [15]. The nanofibers fabricated with the needleless electrospinning had the same morphology compared to conventional electrospinning. The Fe3O4 particles accumulated and distributed on the surface and inside of the nanofibers and the nanofibers had good magnetic and ferromagnetic responsive behavior.

2.3

pH-Responsive Nanofibers

Demirci et al. [16] developed a pH-responsive nanofiber composed of poly 4-vinylbenzoic acid and ar-vinylbenzyl trimethyl ammonium chloride for controlled

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a

b

Cellulose-Fe3O4-heparin monofilament fibers

Cellulose-Fe3O4 core-shell fibers with heparin blended into shell

c

Cellulose nanofibers Magnetite core Magnetite nanoparticles Heparin

Cellulose-Fe3O4 core-shell fibers with covalently immobilized heparin

Fig. 2 Different structures of nanofiber (Permission/license is granted for reusing the material from ACS publications) [14]

release of ciprofloxacin. Results of in vitro release experiments suggested that the nanofibers could release ciprofloxacin in a controlled manner depending on the pH. By increasing the pH from 5.2 to 8.8, the total amount of released drug decreased due to the increase of electrostatic interactions. As shown in the SEM photos (Fig. 3), increasing the pH coagulated the structure of the nanofibers making it harder for the drug to release. Injectable nanofibers for anticancer drug delivery systems in different pH environments were designed by Wu et al. [17]. Thixotropic silk nanofiber hydrogels in an aqueous solution were prepared and used to locally deliver doxorubicin (DOX). Increase of pH decreased the amount of released drug (the most release drug was at pH 4.5 and the least was at pH 7). By adjusting the amount of silk in the hydrogel’s structure, DOX release can be tuned. Hydrogels with and without DOX had the same influence on stopping growth of tumor in the first 3 weeks, but after that remarkably better inhibition of tumor growth was observed in hydrogels loaded with DOX. The weight and volume of the tumor inside the control group mice (hydrogel without DOX) increased significantly after 3 weeks, but it stayed unchanged in hydrogel loaded with DOX showing that prepared nanofiber hydrogels had better therapeutic effect. Han and Steckl [18] developed Eudragit materials based core-sheath fibers. Soluble Eudragit polymers in various pH solutions ranges have been widely used

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Fig. 3 SEM images of nanofibers after release; (a) acetate buffer solution (pH ¼ 5.2), (b) phosphate buffered saline (pH ¼ 7.4), and (c) Tris-buffered saline (pH ¼ 8.8). (d) Release profiles of ciprofloxacin from nanofibers in different pH solutions (Permission/license is granted for reusing the material from ACS publications) [16]

in fabrication of oral drugs. No obvious release at pH 5 was observed while Eudragit nanofibers completely dissolved at pH 7. The coaxial electrospinning technique was used for preparing the nanofibers (Fig. 4c). Eudragit S100 (ES100) and Eudragit L100 (EL100) polymers (anionic copolymers prepared from methyl methacrylate and methacrylic acid) are dissolved at pH 6 or higher and pH 7 or higher, respectively. By combination of these polymers into both core-sheath layer fibers (either in core or sheath), various release kinetics and dissolution at different pH solutions can be gained. There was no release of Eudragit and incorporated material at pH 5 for core-sheath fibers composed of ES100 sheath and EL100 core (Fig. 4a), as both polymers were insoluble at pH 5. At pH 6, the EL100 core was solved, and core material had slow-release behavior due to the ES100 sheath layer protection. At pH 7, the remaining EL100 core and ES100 sheath with encapsulated molecules were fully dissolved and released. When core and sheath’s material was switched between core and sheath, the release pattern will also be switched as shown in Fig. 4b. These multi-pH response nanofibers can be used for applications ranging from biomedical to sensors.

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Fig. 4 pH response of nanofibers in different pH values (a ES100 sheath and EL100 core (b EL100 sheath and ES100 core, and (c) schematic of coaxial electrospinning instrument (Permission/license is granted for reusing the material from Royal Society of Chemistry, License ID# 11726581) [18]

2.4

Electrically-Responsive Nanofibers

Electrically-responsive nanofibers are also one of the external-responsive systems as the electric field is an efficient method to enhance the amount of release with precise control. Electro-responsiveness is gained by molecules turning their dipoles on their

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own with the applied electric field. However, this response most achieved using molecules undergoing an electrically induced redox reaction. Most of the prepared electro-responsive delivery systems since now are based on inherently conductive polymers like polypyrrole (PPy) [19]. Yun et al. [20] developed electro-responsive transdermal nanofibers based on poly vinyl alcohol/poly acrylic acid/multi-walled carbon nanotube. Dispersion of hydrophobic carbon nanotubes in hydrophilic hydrogel network was improved by surface treatment of nanotubes by oxyfluorination. The swelling and drug release behavior of nanofibers depended on the amount of carbon nanotubes, oxyfluorination situation, and electric voltage. Conductivity of the nanofibers enhanced by increase of carbon nanotube content and oxyfluorination with higher oxygen content. Jou et al. [21] prepared an electro-responsive microfiber for smart drug delivery based on poly ε-caprolactone incorporated with curcumin and poly 3,4-ethylenedioxythiophene nanoparticles. Prepared microfibers showed extracellular cell matrices behavior which eases cell spreading and increases cell proliferation. The routine release of curcumin from microfibers was very slow and applying external electric field could increase and tune its release. Release of curcumin led to the volume changes at the surface of the microfibers and inside of the encapsulated nanoparticles. Application of potential pulses made changes at the microfibers because of the migration of nanoparticles from inside to the surface. Their results showed that loading of isotropic actuators, such as nanoparticles, into biodegradable fibers can be a new approach for fabrication of programmable drug delivery systems.

2.5

Biomolecule-Responsive Nanofibers

This group of smart nanofibers could find several applications in diseases diagnosis and therapy. Heo et al. [22] prepared glucose-responsive hydrogel-fibers for investigating long-term glucose monitoring. Fluorescence intensity changes in presence of a glucose molecule. The results indicated that polyethylene glycol/polyacrylamide fibers decreased inflammation in comparison with polyacrylamide alone fibers and responded continuously to the concentration changes of blood glucose for up to 140 days proving their ability to be used in long-term in vivo glucose monitoring. They could track the glucose level by the fluorescence intensity of hydrogel fibers constantly in both high and low concentration of glucose. The content of glucose increased to 300 mg dL1 by injecting glucose and decreased to 140 mg dL1 by injecting insulin. Glucose-responsive polyvinyl alcohol/β-cyclodextrin/glucose oxidase nanofiber hydrogels were prepared as biosensor for continuous interstitial glucose level monitoring [23]. The biosensors showed great performance for the glucose concentration ranging from 0.1 to 0.5 mM with a sensitivity of 47.2 μA mM1, detection limit of 0.01 mM, and fast response time (lower than 15 s). Their results proved that the fabricated biosensor is able to measure the glucose concentration in human serum. Wang et al. [24] fabricated pH- and glucose-responsive boronic acid-based

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Fig. 5 Capture and release of lectins on nanofibers (Permission/license is granted for reusing the material from Royal Society of Chemistry, License ID# 1172662-1) [24]

nanofibers for reversible capture and release of lectins (FITC-Jacalin and FITCConA). Prepared nanofibers could capture ConA and Jacalin in a significantly higher amount under alkaline condition compared to the pristine nanofibers. Lectins could be captured on the nanofiber surfaces containing galactose and glucose while no lectin could be adsorbed on the pristine nanofibers. In acidic solution, nanofibers released both lectins and glycopolymers (Fig. 5). Wade et al. [25] designed protease-responsive electrospun hydrogels through chemical modification of hyaluronic acid (HA – a linear polysaccharide composed of alternating D-glucuronic acid and N-acetyl-D-glucosamine). They formulated HA-based macromers containing fluorescent peptides and protease-cleavable with the ability to be electrospun and then crosslinked through Methacrylated Peptide (MePs) which are sensitive to the incorporated amino acid sequence. Enzyme-linked immunosorbent assay (ELISA) is the most famous optical immune-assay (fluorescent or calorimetry) that is commonly used for detection of pathogens (bacteria, virus), protein biomarkers (such as antibodies), and smaller molecules existed in environmental samples. Mahmoudifard et.al [26] developed a new immunoassay platform based on protein molecules and polyethersulfone electrospun fibrous membrane. Their results indicated the immobilization of protein molecules on nanofibers successfully increased the ELISA signal and oxygen plasma treatment enhanced the amount of antibody immobilization. An intracellular restructured reduced glutathione-responsive peptide nanofiber loaded with doxorubicin was designed for tumor chemotherapy. Micelles endocytosed with tumor cells degraded after long blood circulation at high glutathione concentrations leading to more doxorubicin release and accumulation at the tumor site [27].

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297

Multi-responsive Nanofibers

Huang et al. [28] fabricated a magnetic-thermo-responsive fiber from PNIPAAm and Fe3O4 nanoparticles by an electrospinning process and UV-curing technique. Loading Fe3O4 nanoparticles in the structure had a slight impact on thermo-responsivity of the fibers. Tests of thermo-dependent magnetic behavior of the nanofiber were done under and above the LCST which in both cases fibers were absorbed to the left side of the vials, but the adsorbed amount decreased by increasing the magnetic distance. More fibers were attracted on the left end above LCST than those were under LCST, showing that thermo-responsive fibers have better magnetic attraction ability. Drug-release behavior of vitamin B12 as the model drug was faster (burstrelease) at 37
C in comparison with a gradual diffusion at 25
C and it was found that the prepared fibers can release more drug due to their thermo-responsive properties. Chitosan-poly(N-isopropylacrylamide copolymers with or without bovine serum albumin (BSA) were fabricated through electrospinning to obtain temperature- and pH-responsive nanofibers. Hydrogel nanofibers showed controlled protein release which could be regulated by changing the medium temperature and the pH. The amount and rate of released BSA were remarkably higher at 25
C than at 37
C and release of drug in acidic medium was much higher than in a basic medium. No obvious cytotoxicity of the nanofibers in the present of L929 cell proliferation was observed [29]. pH- and photothermal-responsive nanofibers based on porous carbon nanofibers loaded with doxorubicin (anti-tumor drug) were fabricated for synergistic chemophotothermal therapy and drug delivery applications. Prepared nanofibers had higher drug release at acidic pH (tumor tissue) than neutral medium (normal tissue) which is beneficial for cancer therapy efficiency. Also, a fast rate of drug release was observed under NIR irradiation which can be related to the accelerated rate of molecular motion made by the high temperature after irradiation (suspensions were irradiated under 808 nm laser for 10 min before the test for evaluating photothermalresponsivity of the nanofibers). Biocompatible carbon nanofibers were able to kill a great number of Mg-63 cancer cells in the exposure to 808 nm NIR irradiation for 10 min while few dead cells were observed in the control group (only NIR irradiation without treatment). It was also proved that NIR laser did not affect the normal cell proliferation [30]. Temperature- and pH-responsive polyethyleneimine-N-isopropylacrylamide (PEI-NIPAM) polymer was grafted on cellulose nanofibers for sustained drug release (Fig. 6). Doxorubicin release rate of the nanofibers reduced with increasing pH from 3 to 7.4 and increased with temperature from 25 to 37
C. In vitro and acute systemic toxicity, skin irritation, and sensitization tests were done to investigate the biocompatibility of the nanofibers. Results showed that nanofibers were not toxic and potentially cytotoxic. Extremely mild skin irritation and no obvious allergic symptoms were observed in New Zealand rabbits and albino guinea pigs, respectively [31].

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Fig. 6 Synthesis of cellulose nanofibers; CNF-COOH: Cellulose nanofiber, DOX: Doxorubicin (Permission/license is granted for reusing the material from Elsevier, License No#5218320283897) [31]

3 Biomedical Applications of External-Responsive Nanofibers 3.1

Wound Dressings

One of the potential applications of the stimuli-responsive nanofibers is wound dressings. Smart wound dressings such as stimuli-responsive with the ability to control drug release have been developed to avoid excessive usage of antibiotics and antimicrobial drugs and promote the wound healing process. A temperatureresponsive methacrylate gelatin nanofibrous hydrogel loaded with fatty acids/aspirin (ASP) encapsulated polydopamine (PDA) was developed by Zhang et al. [32] for accelerating wound healing. The ASP release at 40
C was remarkably faster than that of 25
C and 37
C. L929 and HaCaT cells were selected to evaluate biocompatibility of the prepared nanofibers and the results showed that the viability of the L929 and HaCaT cells were over 95% after 1 day. CCK-8 assay revealed that loading PDA led to the assisting cell proliferation and adhesion. Also, great in vivo wound healing effect, higher re-epithelialization in comparison with the control group and collagen fibers deposition were observed.

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Bioburden-responsive Fusidic acid loaded with Poly lactic-co-glycolic acid (PLGA) ultrafine fibers (UFs) was developed and characterized for wound healing applications [33]. Bioburden-responsiveness of UFs was defined as a polymer-based drug release system in response to naturally microbial stimuli in wounds which depends on enzyme-sensitive biodegradability of the polymer substrate and microbial enzymes with the ability of guiding the interactions leading to the macroscopic polymer transitions. Fusidic acid-loaded UFs with relevant antimicrobial release design were effective in healing of both lightly and heavily contaminated Staphylococcus aureus infected rat wounds. Nanofibers of chitosan/polyethylene oxide as shell and polycaprolactone as core were designed and loaded with lidocaine hydrochloride in the shell for pain relief and curcumin in the core as an anti-inflammatory agent. Sodium bicarbonate as the acid sensitive material was added to the core layer to provide wound microenvironment sensitivity to react with hydrogen ions and form CO2 under acidic pH (Fig. 7). Lidocaine hydrochloride released under acidic condition (because of protonation of chitosan and formation of –NH3+) and simultaneously reaction of sodium bicarbonate with hydrogen ions led to generation of CO2 and many holes in the fiber surface making it easier for the curcumin to release. The released rate of curcumin can be adjusted by changing the concentration of sodium bicarbonate. The prepared nanofibers had great cytocompatibility, hemocompatibility, and blood coagulation [34]. Chen et al. [35] synthesized and characterized reductant-responsive N-maleoyl functional chitosan/PVA nanofibers as wound dressings. Tetracycline hydrochloride was loaded into the nanofibrous structure which prevents different types of bacterial infections. In comparison with other stimuli, change of redox-potential by biocompatible reductants is more appropriate to be used as a triggered signal for wound dressings. In this regard, loading reductant-labile bonds like disulfide bonds into nanofibers has obtained many attractions. Nanofibers had excellent water stability and swelling degree, low toxicity, and reductant-responsive function. For proving reductant-responsive function of nanofibers, glutathione (GSH) as biocompatible reductant was added to the release medium. About 90% of the loaded tetracycline hydrochloride was released after 24 h and the cumulative amount of release was reduced rapidly after 24 h and reached 97% at 120 h while the amount of released drug without GSH was 53% at 24 h. pH-responsive chitosan/polyvinyl alcohol/graphene oxide electrospun nanofibers loaded with extract of garlic and allicin were developed for preparation of wound dressings with sustained-release and strong antibacterial activity properties [36]. The amount of drug release can be regulated with the amount of graphene oxide and nanofibers with graphene oxide had better antibacterial activity in comparison with those without graphene oxide. The amount of released drug decreased by increasing the pH of the release medium. Good hydrophilicity and moisture retention capacity of the nanofibers indicated that they are able to provide a suitable moisture level of the wounds. Temperature-responsive MXene-based nanobelt fibers loaded with vitamin E were prepared by Jin et al. [37] with the ability of controllable release for wound

Fig. 7 Wound healing schematic by microenvironment-responsive dual-drug-loaded wound dressing (Permission/license is granted for reusing the material from Royal Society of Chemistry, License ID#1172868-1) [34]

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Table 1 Nanofibers for wound dressings Ingredients PVA/PAA/ BTB a P(DEEA-coBA-coDMNOBAco-ABP) b PCL/PES d PU/Pro-Cip/ H-Cy e PDEGMA/P (LLA-CL)/ CIF f PNIPAAm/ chitosan/PPg PVA

Potential applications Delivery of ciprofloxacin and wound healing Potential wound dressing for controlled drug release Wound dressing Theranostic wound dressing Wound dressings

Wound dressings Wound dressings

Type of stimulus pHresponsive

Preparation method Heat physical crosslinking of electrospun

Reference [39]

UV-triggered CO2responsive Bacteriaresponsive Lipaseresponsive Thermoresponsive

Via “active ester-amine” chemistry reaction based on P (PFPA-co-ABP) c as a precursor Single and core-shell nanofibers Electrospinning

[40]

[42]

Electrospinning

[43]

Thermoresponsive NIRresponsive

Non-woven fabric

[44]

Electrospinning (nanofibrous hybrids)

[45]

[41]

a

PVA: Polyvinyl alcohol, PAA: Polyacrylic acid, BTB: Bromothymol blue DEEA: N,N-diethyl ethylenediamine, BA: benzylamine, DMNOBA: methanol. N,N-dimethyl-N(2-nitrobenzyl)-ethane-1,2-diamine, ABP: 4-acryloyloxy benzophenone c PFPA: Pentafluorophenyl acrylate d PCL: Polycaprolactone, PES: Poly(ethylene succinate) e PU: Polyurethane, Pro-Cip: Ciprofloxacin-based prodrug, H-Cy: Chromogenic probe (an ester derivative of a hemicyanine dye) f PDEGMA: Poly di(ethylene glycol) methyl ether methacrylate, P(LLA-CL): poly(l-lactic acidco-ε-caprolactone), CIF: ciprofloxacin g PNIPAAm: Poly(N-isopropylacrylamide), PP: polypropylene b

healing. The temperature could be controlled by exposure of near-infrared irradiation for promoting release of vitamin E. The nanofibers had great wound-healing ability and biocompatibility in both in vitro and in vivo tests. Wound dressings loaded with biocides can promote the wound healing process by stopping or healing infections. Core-shell polyhydroxy alkanoate (PHA)-based nanofibers were prepared by Li et al. [38] to investigate bacteria-triggered release of a biocide to the sites of bacterial infections. In the presence of pathogens (in this case Pseudomonas aeruginosa), the PHA-based shell was degraded and loaded biocide (in this case dodecyltrimethylammonium chloride) was released. Total amount of release from the core-shell nanofibers was bigger in supernatant than in PBS solution showing that nanofibers are bacteria-responsive. Some other studies are presented in Table 1.

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Drug Delivery

Excellent properties of nanofibers specially electrospun nanofibers such as high surface area, high loading capacity, great encapsulation efficiency, and simultaneous delivery of diverse therapies make them good candidates for drug delivery applications. Li et al. [46] took advantage of external stimuli responsive polymers to design a dual responsive electrospun nanofiber that is able to show a different release rate in response to the pH and temperature. In this design, they applied poly N-vinylcaprolactam and ethyl cellulose as a base polymer to produce a thermoresponsive fiber. Eudragit L100-based fibers were then generated to represent a pH-sensitive nanofiber. Dual electrospinning technique was used in this study. Thermo-responsive and pH-responsive nanofibers were collected on the same collector. These two fibers including KET as a drug were sprayed from two different syringe pumps. Their results showed that at the same temperature, increasing pH from 4.5 to 7.4 increased drug release 5 times (from 10% to 50%, respectively) after 60 h. In addition, it showed that at the same pH (7.4), increasing temperature from 25 to 37
C increased drug release from 50% to 80%. Moreover, temperature change was more significant at lower pH and at pH 4.5 increasing temperature from 25 to 37
C increased drug release from 10% to 60% which was much greater than drug release change at pH 7.4. Jiang et al. [47] designed pH-responsive electrospun nanofibers made of poly ε-caprolactone as the base polymer. Nanofiber mats were treated by air plasma and coated by polydopamine. R6G and DOX used as drug models and dissolved in aqueous solution with different pH. Then uptake and release capacities of treated nanofiber mats submerged in aqueous drug solutions were evaluated at different pH. As reported, at higher pH (9.0, 11.0) uptake capacity of nanofibers was better. While drug release showed opposite behavior and better release happened at lower pH (2.0, 5.0, and 7.0). Dai et al. [30] developed a pH/photothermal dual-responsive nanofiber for controlled drug release. In this design, they prepared porous carbon electrospun nanofibers based on polyacrylonitrile and polymethyl methacrylate and DOX was used as a drug model. The porous structure led to the higher drug absorption and lower rate of drug release which extended the delivery period. This electrospun nanofiber was able to convert photo-energy to heat in the near infrared region wavelength (700–1,100). In vitro and in vivo tests indicated that photothermaland chemo-therapy have higher tumor inhibition efficacy with no side effects on other organs. Near infrared-responsive nanofibers containing poly N-isopropylacrylamide and gold nanorods were designed for on-demand drug delivery systems. Both hydrophobic and hydrophilic drugs could be loaded into the designed platform. Heat generated from exposure of gold nanoparticles to near infrared light irradiation could control swelling and deswelling ratio of the nanofibers (as the base polymer of the nanofibers is thermal sensitive) resulting in controlling the drug release [48]. Table 2 presents other examples of applying stimuli-responsive nanofibers in drug delivery systems.

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Table 2 Nanofibers for drug delivery Ingredients MEL/Cypate/ HAa

Potential applications Delivery of anticancer drugs

Type of stimulus TME-NIRresponsive

Preparation method Physically assembled

Reference [49]

b

Controlled drug release Drug carrier in brain cancer treatment

pHresponsive pHresponsive

Polyelectrolyte complex of Ch/Alg on PLGA nanofibers Electrospinning & initiated chemical vapor deposition

P(AAm-coAAc)/PEG e

Drug delivery

pHresponsive

PAA/rGO f

On-demand antibiotics release On-demand antibiotics release On-demand release of ibuprofen On–off drug release devices

NIRresponsive

Electrospinning via heat treated to induce esterification crosslinking reaction Electrospinning

[53]

NIRresponsive

Electrospinning

[48]

Thermoresponsive

Electrospinning

[54]

Thermolightresponsive Magneticresponsive pHresponsive Infectionresponsive

Electrospinning

[55]

Co-axial electrospinning

[56]

Electrospinning

[57]

Electrospinning

[58]

Ch/Alg/ PLGAc PVA/P (4VP-coEGDMA)/RB

[50] [51]

d

PNIPAM/ GNRs g PNIPAM/ PCL h PHBV/CNCZnO i PCL/KCZ/ Fe3O4 j Chitosan/ pectin PCL/ polydopamine a

Antifungal drug release Drug delivery Drug delivery

[52]

MEL: Melittin, HA: Hyaluronic acid TME: Tumor microenvironment, NIR: Near-infrared laser irradiation c Ch: Chitosan, Alg: Alginate, PLGA: Poly(lactic-co-glycolic acid) d PVA: Poly vinyl alcohol, P(4VP-co-EGDMA): Poly(4-vinylpyridine-co-ethylene glycol di-methacrylate), RB: Rose Bengal (4,5,6,7-tetrachloro-20 , 40 ,50 ,70 -tetraiodofluoresceindisodium) e P(AAm-co-AAc): Poly(acrylamide-co-acrylic acid), PEG: Polyethylene glycol as cross-linker f PAA: poly(acrylic acid), rGO: reduced graphene oxide g PNIPAM: Poly (N-isopropylacrylamide), GNRs: Gold nanorods h PCL: Poly (ε-caprolactone) i PHBV: Poly (3-hydroxybutyrate-co-3-hydroxyvalerate), CNC-ZnO: Cellulose nanocrystal-zinc oxide j PCL: Polycaprolactone, KCZ: Ketoconazole, Dimethyl silicone oil as inner core of co-flowing solutions b

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Diagnosis

The size scale and surface area of nanofibers give us better biomaterial availability and consequently a more accurate diagnostic structure. Diagnosis applications of smart nanofibers is one of the most interesting aspects of these materials. Histamine in urine is a very important characteristic parameter for cancer detection. An electrospun nanofiber based on dendritic zinc porphyrin was designed by Seong et al. [59]. The fluorescence spectra change after reaction with various histamine solutions as in the lower histamine solution, the fluorescence intensity decreased with increase of histamine content. Bacteria-responsive and color-changing dressings have the ability to monitor wounds continuously for early detection of any bacterial infections. A polyurethane wound dressing loaded with hemicyanine-based chromogenic probe which can be split enzymatically by bacterial lipase released from methicillin-resistant Staphylococcus aureus (MRSA) and Pseudomonas aeruginosa was developed by Currie et al. [60] (Fig. 8). The hemicyanine-based chromogenic probe (HCy) was incorporated into the shell of coaxial polyurethane nanofibers to localize dye at the surface of the fibers and increase the rate of color changing response. The core-shell nanofibers loaded with HCy responded five times faster to 1010 CFU/cm2 P. aeruginosa than without HCy and had a uniform color change from yellow to green in less than 2 min. Threshold of detection can be tuned by adding polyvinylpyrrolidone to increase transfer of charge in HCy which enable bacteria detection after 2 h exposure at concentrations of 2.5  105 CFU/cm2 P. aeruginosa and 1.0  106 CFU/cm2 MRSA. Infection or inflammation of wounds can generate high level of hydrogen peroxide (H2O2) leading to prevention of the wound healing process. So, monitoring the amount of produced H2O2 can be a way for evaluating wound’s healing status. Wu et al. [61] fabricated a H2O2-responsive Europium(III) Coordination/polyacrylonitrile dressing for monitoring H2O2 level through color changes. Rat wound test showed that nanofibers could promote wound healing through stimulating neovascularization and are not toxic to human umbilical vein endothelial cells (HUVEC) and fibroblast cells (L929). Also, with increase of H2O2 concentration and quench time, the color of nanofibers changed from bright to weak (Fig. 9). Fabricating new platforms for continuous monitoring of small molecules and electrolytes like creatinine, glucose, and urea would be a great achievement in biomedicine. Balaconis et al. [62] fabricated plasticized polycaprolactone nanofibers loaded with boronic acid derivatives and alizarin to be used as glucose sensor with higher sensitivity to glucose and in vivo stability in comparison with their own previous study. According to the higher residency time of nanofibers and sensitivity of the new boronic acids, this sensor can be a potential for continuous monitoring of glucose and other analytes.

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Fig. 8 Schematic of performance of bacteria-responsive nanofibers (Permission/license is granted for reusing the material from ACS publications) [60]

3.4

Scaffolds for Cell Culture and Delivery

Kim et al. [63] developed electrospun thermo-responsive nanofibers based on copolymers of N-isopropylacrylamide with a UV-reactive benzophenone conjugated co-monomer to capture and release cells. The prepared nanofibers can capture, encapsulate, and release cells by dynamical transformation of their fibrous structure into structures like hydrogel by wrapping, swelling, and deswelling processes in response to external temperature change. The released cells had great proliferation and viability. Maeda et al. [64] designed temperature-responsive nanofibers using N-isopropylacrylamide for cell storage applications. Capturing and releasing ability of the nanofibers was evaluated by putting a glass bead on the fiber mesh and then

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Fig. 9 H2O2-responsive nanofibers performance (Permission/license is granted for reusing the material from Elsevier, License No#5218500402425) [61]

Fig. 10 Temperature-sensitivity behavior of nanofibers on a glass bead (Permission/license is granted for reusing the material from MDPI publications) [64]

adding hot water. The fiber mesh captured the glass bead and released it by decreasing the temperature below the LCST (Fig. 10). The authors also showed that cryopreservation of mammalian cells within the nanofibers was possible and cell viability of entrapped cells in the nanofiber mesh was higher than that of encapsulated ones.

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Depsipeptides were grafted onto a polypeptide backbone with pH-responsive nanofibers as supramolecular gelators. The mechanically-stable polypeptide nanofiber had great cytocompatibility, controlled enzymatic degradation and culture stability of primary human umbilical vein endothelial and neuronal cells which make them a potential for tissue engineering applications such as cell matrices [65]. Core-sheath nanofibers made of poly N-isopropylacrylamide (sheath), poly caprolactone, and nattokinase solution were developed by Shi et al. [66] for cell capture and release. Hydrophobicity and hydrophilicity of the core-sheath nanofibers can be switched by temperature change. Nattokinase was released from nanofibers to enhance platelet adhesion resistance on the nanofiber surface, making it easier to capture and isolate red blood cells directly from the blood in a temperature-sensitive manner. Release of efficiency up to 100% suggests nanofibers as effective scaffolds for capture and release of non-adherent cells. Temperature-responsive shape-memory nanofibers were designed for smooth muscle cells (SMCs) culture. The scaffolds were made of an outer layer of poly lactide glycolide trimethylene carbonate for programming the deformation from planar to small molecular tubule architecture and an inner layer of poly lactide glycolide/chitosan nanofibers to adjust cell morphology, adhesion, and proliferation. The prepared temperature-responsive platform can be deformed at 20
C for culture and release of SMCs and self-rolled immediately into 3-D tube at 37
C [67]. Niiayma et al. [68] developed temperature-responsive nanofibers made of poly ε-caprolactone-based polyurethane/hexamethylene diisocyanate/1,4-butanediol with shape-memory properties. Mechanical and shape-memory properties could be changed by altering molar ratios of the components. After 400% deformation, over 89% recovery was achieved. As the nanofibers were temperature-sensitive, biocompatible, and had the ability of controlling the cell alignment, they could be used in biomedical applications such as cell culture platforms. Light-responsive nanofibers composed of poly N-isopropylacrylamide/silicacoated gold nanorods/polyhedral oligomeric silsesquioxanes with the ability to deliver therapeutic drugs and cells were fabricated and characterized. Nanofibers could entrap, adhere, and proliferate the NIH3T3 fibroblast cells and also release cells with undisturbed cellular function upon NIR irradiation [69].

4 Conclusions In the past decade, many efforts have been made to preparate and characterize smart fibers and polymers for biomedical applications. In this chapter, some of these studies on external stimuli-responsive nanofibers for different biomedical applications such as wound dressing, drug delivery, diagnosis, and scaffolds for cell culture are discussed. Excellent properties of external stimuli responsive nanofibers make them great candidates for different biomedical applications.

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Adv Polym Sci (2023) 291: 313–334 https://doi.org/10.1007/12_2022_136 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 11 October 2022

Electrospun Antimicrobial Polymeric Nanofibers in Wound Dressings Yunfan Shi, Chenzi Zhang, Feng Jiang, Liuzhu Zhou, Ling Cai, Hongjie Ruan, and Jin Chen

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Antibacterial Constituent . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Antibiotics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Nanofibers Loaded with Metal Nanoparticles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Yunfan Shi, Chenzi Zhang, and Feng Jiang contributed equally to this work. Y. Shi and F. Jiang Center for Global Health, School of Public Health, Nanjing Medical University, Nanjing, China The First Clinical Medical College of Nanjing Medical University, Nanjing, China C. Zhang Center for Global Health, School of Public Health, Nanjing Medical University, Nanjing, China The Second Clinical Medical College of Nanjing Medical University, Nanjing, China L. Zhou and L. Cai Center for Global Health, School of Public Health, Nanjing Medical University, Nanjing, China H. Ruan (*) Women’s Hospital of Nanjing Medical University, Nanjing Maternity and Child Health Care Hospital, Nanjing, China J. Chen (*) Center for Global Health, School of Public Health, Nanjing Medical University, Nanjing, China The Key Laboratory of Modern Toxicology, Ministry of Education, School of Public Health, Nanjing Medical University, Nanjing, China Jiangsu Province Engineering Research Center of Antibody Drug, Key Laboratory of Antibody Technique of National Health Commission, Nanjing Medical University, Nanjing, China e-mail: [email protected]

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2.3 Plant Extracts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4 Biomacromolecules . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Summary and Prospective . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract The efficient wound management highly depends on the dressing materials, which involves the processess of surface protection, anti-inflammation, and antibacterial properties. Due to the inherent porous structure of electrospun nanofibers, various bactericidal agents can be blended to obtain the antimicrobial nanofibrous membranes as wound dressings. The formed polymeric dressings show desirable biocompatibility and vapor permeable property, low cytotoxicity, and enhanced drug delivery at the wound site, which is beneficial for the wound healing. This review mainly summarizes the recent progress of electrospun antimicrobial polymeric nanofibers as wound dressings with a special focus on the incorporation of bactericidal components including antibiotics, metal nanoparticles, plant extract, and biomacromolecule. Keywords Antibacterial activity · Drug release · Electrospun nanofibers · Wound dressing

1 Introduction Skin is the first line of defense system to protect the body against foreign invaders including microbial infection, wound, and harmful substances. However, when severe damages occur to the skin such as wound, the risk of infection by external pathogenic bacteria is significantly increased, which may delay the healing process and worsen the prognosis of wound treatment. Therefore, it is demanding to develop the effective tissue engineering materials to promote the healing while alleviating the bacterial infection to a great extent. Currently, the commercial wound dressings used for clinical practices include synthetic and natural materials [1]. For example, cotton gauze is commonly used for the wound healing, which acts as a physical barrier to achieve hemostatic effects. However, cotton gauze owns drawbacks such as the inherent non-antibacterial nature and the secondary injury as a result of dressing adherence to the wound, which make it not desirable for the prevention of wound infections [2, 3]. Recently, hydrogels show promising potentials as wound dressings due to their good biocompatibility, biodegradability, hemostatic performance, and antibacterial activity [2, 4, 5]. However, the poor mechanical performance of hydrogels may limit its clinical applications. In clinical practice, antibiotics with broad-spectrum antibacterial efficacy are often administered to prevent bacterial infection. However, the misuse of antibiotics may lead to considerable side effects and drug resistance [6]. It was reported that the antibiotic-resistant bacteria infections have caused ~700,000 deaths each year

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around the world, which is estimated to be 10 million by 2050 [7]. Therefore, the emergence of drug-resistant bacteria has posed the increasing threat to the public health. By efficient incorporation of the antibacterial agents into the biodegradable polymer material on the basis of minimized cytotoxicity, we can not only reduce the risk of bacterial infection, but also improve the antibacterial efficacy of active components. Recently, great attentions have been made in the preparation of various nanofibrous membranes by the electrospinning methods. The fabricated electrospun nanofibers show high porosity, large surface area, excellent vapor permeable property, and structural adaptability, which are suitable for medical uses such as tissue engineering scaffold, drug delivery matrix, and wound dressing [8]. As summarized in Table 1, we have listed the antibacterial electrospun polymeric nanofibers loaded with antibiotics, metal nanoparticles, plant extract, and biomacromolecule, which hold therapeutic potential to combat bacterial infection. The preparation of antibacterial agent loaded nanofiber dressings is shown in Fig. 1. Owing to the built three-dimensional (3D) drug delivery system of electrospun nanofibers, antibiotics can be imbedded efficiently into the biodegradable polymers so that the formulated dressings can achieve increased drug concentration yet minimized cytotoxicity at the wound site. In particular, after metal nanoparticles of broad-spectrum antibacterial activity or plant extracts were loaded, the constructed electrospun nanofibers show promising potential to deal with emerging antibiotic resistance during the long-term therapy.

2 Antibacterial Constituent 2.1

Antibiotics

Infection caused by pathogens such as E. coli and P. aeruginosa is one of the main factors to affect the wound healing [10]. Broad-spectrum antibiotics have been frequently used in the clinical treatment of wound infection [11]. Therefore, the antibiotics-blended wound dressing fabricated by electrospinning was applied to the topical treatment of infected wounds. In addition to the reduced usage of administered antibiotics by this way, the constructed wound dressing that enables the sustained drug release is beneficial for the practical uses to avoid frequent changing of dressings. DCH, a kind of broad-spectrum antibiotics, shows efficient inhibition of activity of matrix metalloproteinase inhibitor even at the level of sub-antimicrobial dose, which has been approved by Food and Drug Administration (FDA) for the clinical treatment of periodontitis. It was found that topical use of DCH can promote the healing of chronic wound without any pronounced side effects but the local drug concentration and sustained drug release on the wound site remain difficult to control. In addition, the photosensitive nature of DCH may limit its long-term use. Therefore, DCH-encapsulated PLA electrospun nanofibers were constructed for the treatment of chronic wounds based on the type I diabetic rat models [15]. A high

Polygalacturonic/hyaluronic acid nanofiber loaded with AgNPs Electrospun CA/ZnO/AgNPs composite nanofibers

DCH/PLA nanofiber mats Electrospun PCL/mupirocin and CS/lidocaine hydrochloride multifunctional double layer nanofibrous scaffolds Vancomycin-loaded electrospun CS/PEO nanofibers

Electrospun CIP-loaded EC nanofibers Electrospun CIP-loaded PVP nanofibers Ciprofloxacin HCl and quercetin functionalized electrospun nanofiber membrane Core-shell structured antimicrobial nanofiber dressings

Name CIP-loaded CS/polyethylene oxide/ silica nanofibers CS/PVA/GO nanofibrous membrane with CIP

Quercetin

PCL/gelatin laden/Mino/ gelatin infused with G. sylvestre extracts PLA PCL and CS

Ciprofloxacin HCl

Mino

AgNPs

AgNPs

Vancomycin

Polygalacturonic/ hyaluronic acid ZnO nanoparticles, CA

CS/PEO

PVP

CIP

DCH Mupirocin and lidocaine hydrochloride

EC

Other components/carriers CS/polyethylene oxide/ silica CS, PVA, and GO composites

CIP

CIP

Active constituents CIP

Table 1 Antibacterial agents-loaded electrospun polymeric nanofibers

+ +

+

G+, GG+, GG+, G-

+

-

+

+ E. coli and S. aureus

+

+

Methicillinresistant S. aureus G+, G-

+ -

+ +

E. coli S. aureus, E. coli, and P. aeruginosa

+

+

+

+

+

+ +

+

+

+

-

-

+ -

+

Controlledrelease +

-

Accelerate healing +

+

Antibacteria +

E. coli, S. aureus, and B. subtilis G+, G-

Bacteria G+, G-

[18]

[17]

[6]

[15] [16]

[11]

[14]

[13]

[13]

[12]

Reference [10]

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+ + + +

+

E. coli S. aureus and E. coli G+ G+, GG+, G-

Pullulan, CS, and citric acid PCL/CS

Calendula officinalis extract Melilotus officinalis extract PCL/CS

PCL/chitosan oligosaccharides CS/PEO

Aloe vera extract Aloe vera extract Quercetin

+

G+, G-

Alginate/PVA

Honey

+

G+

PCL/gelatin

Au_MBA NPs

+

E. coli

PCL and gelatin

AgNPs

+

G+, G-

Cellulose

AgNPs

+

+

+

+

+

-

+

+

-

-

-

+

-

-

[27]

[26]

[25]

[24]

[23]

[22]

[21]

-

[20]

[19]

+

-

CA Cellulose acetate, CIP ciprofloxacin, CS chitosan, DCH doxycycline, EC ethyl cellulose, GO graphene oxide, MBA mercaptophenylboronic acid, Mino minocycline hydrochloride, PCL Polycaprolactone, PEO polyethylene oxide, Nps nanoparticles, PLA polylactide, PVA poly(vinyl alcohol), PVP Poly (vinylpyrrolidone), P. aeruginosa Pseudomonas aeruginosa

PCL/CS/Melilotus officinalis extract electrospun nanofibers

Silver nanoparticles-covered 3D cellulose Harnessing biocompatible nanofibers and silver nanoparticles MBA-activated gold nanoparticles as nanoantibiotics Honey-loaded alginate/PVA nanofibrous membrane Aloe Vera extract-based composite nanofibers Electrospun PCL/CS/Aloe vera blended nanofiber membranes Electrospun CS oligosaccharide/ polycaprolactone nanofibers CS/PEO nanofibers containing Calendula officinalis extract

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Fig. 1 Various antibacterial agents including antibiotics, metallic particles, plant extracts, and biological macromolecules blended in the electrospun nanofibers as the wound dressings. The nanofibers shown in the figure were adapted from [9] with permission

loading content of DCH up to 20% was achieved in the formed nanofibers, which is more advantageous than topical coating of DCH solution. Li et al. have incorporated ciprofloxacin (CIP), a highly used fluoroquinolone, into the electrospun fibers composed of ethyl cellulose (EC) or poly (vinylpyrrolidone) (PVP). The release of CIP from EC fibers was found much slower than that of PVP as a result of combination of polymer swelling and drug diffusion. The CIP-loaded nanofibers collected on foil and on gauze were effective against the growth of G+ and G- bacteria, suggesting their promising potential as wound dressing materials [13]. CIP was co-electrospun with chitosan (CS) and graphene oxide (GO) using poly(vinyl alcohol) (PVA) to obtain drug-loaded CS/PVA/GO nanofibrous membrane of excellent cytocompatibility. The presence of GO can increase the distance between nanofibers (Fig. 2), which contributes to the increased drug release ratio to achieve desirable bactericidal effect toward E. coli, S. aureus, and B. subtilis (Fig. 3) [12]. Ajmal et al. mixed CIP with quercetin (Que) and subsequently co-electrospun under appropriate parameters to produce nanofiber membrane [14]. Note that Que mainly acts as a scavenger of reactive oxygen species to curb inflammatory responses. The high entrapment efficiency of CIP (~92%) and Que (~94%) and prolonged in vitro drug release (1 week) of obtained nanofibrous membrane demonstrated its therapeutic efficacy to suppress probable infection and oxidative damage. The built dressing showed accelerated wound healing (Fig. 4).

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Fig. 2 Scanning electron microscopy (SEM) images of (a) CS/PVA nanofibrous membrane, (b) CS/PVA/GO nanofibrous membrane, (c) CS/PVA/CIP hydrochloride (CIPHCl) nanofibrous membrane, (d) CS/PVA/GO/CIPHCl nanofibrous membrane, (e) CS/PVA/CIP nanofibrous membrane, (f) CS/PVA/GO/CIP nanofibrous membrane. Reproduced from [12] with permission

Hashemikia et al. have introduced a hybrid organic/inorganic material for wound dressing by incorporation of CIP into the chitosan/PEO/silica nanofibers. To improve the spinnability of the chitosan/silica mixture, PEO as co-electrospun homopolymers was used. The obtained nanofiber demonstrated high stability, biocompatibility, and flexibility, which accelerated and enhanced tissue regeneration according to in vitro and in vivo preclinical tests (Fig. 5) [10].

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Fig. 3 Radius statistical diagram of inhibition zones of CS/PVA nanofibrous membrane, CS/PVA/ GO nanofibrous membrane, CS/PVA/CIP nanofibrous membrane, CS/PVA/GO/CIP nanofibrous membrane, CS/PVA/CIP nanofibrous membrane, and CS/PVA/GO/CIP nanofibrous membrane. Reproduced from [12] with permission

Fig. 4 (i) Images and (ii) wound area closure rate of full-thickness wounds after treatment with gauze, PCL, PCL/CHL, PCL/CHL/Que-nanofiber at different time intervals. Reproduced from [14] with permission

For the clinical treatment, it is ideal that wound dressing may possess multifunctional properties including good biocompatibility and comfortability, avoiding infection, high patient compliance, and high absorptivity to exudates. The double layer electrospun nanofibrous dressing shows improved therapeutic efficacy due to its layered microstructure. Li et al. have reported a CS/PCL double

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Fig. 5 (a) Images and (b) closure rate of the wounds after the treatment with CS/PEO, CS/PEO/ SiO2, and CS/PEO/SiO2/CIP. Reproduced from [10] with permission

layer nanofibrous scaffold for multifunctional wound dressing practices [16]. The outer layer was PCL mixed with mupirocin, a kind of potent and broad-spectrum antibiotics, which exerts antimicrobial activity to the developed wound dressing. The inner layer without direct contact with wound site is composed of electrospun CS nanofibers loaded with lidocaine (LID) as pain-relieving components. The overall dressing exhibited improved hydrophilicity, cytocompatibility, and antibacterial performance, which serves as an ideal multifunctional wound dressing material. In particular, by employing coaxial electrospinning technique, the coresheath nanofiber dressing can be easily fabricated, which provides a promising solution for the controllable drug release for the wound management [28]. Moreover, Li et al. have compared the biocompatibility and antibacterial performances between the antibiotics-loaded nanofibers using hydrophobic and hydrophilic matrix [13]. It was shown that CIP-loaded PVP electrospun nanofibers using hydrophilic polymers as matrix showed a relatively fast drug release, while

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Fig. 6 In vitro release profiles of CIF from (a) PVP fibers and (b) the EC fibers as compared with that of the gauze. Reproduced from [13] with permission

CIP-loaded EC electrospun nanofibers, using hydrophobic polymers as matrix, exhibited an improved biocompatibility, which promoted the growth and proliferation of fibroblasts in the dressing environment (Fig. 6). Unfortunately, with the extensive use of antibiotics, invalidation of antibiotic therapy has become a serious issue for the public health. The emergence and progress of a series of drug-resistant strains such as P. aeruginosa have posed an increasing difficulty for the treatment of bacterial infections [29]. In the development of antibiotics resistance, the formation of the biofilm often plays an important role. Through deliberate structural design, electrospun nanofiber as wound dressing could exert remarkable inhibitory effect against the bacterial biofilm formation. Ramalingam et al. prepared the core-shell structured antimicrobial nanofiber dressings containing antibiotics and herb extract combination with Mino loaded PCL/gelatin as the shell and gelatin infused with herbal extract as the core. The core-shell structured nanofiber dressing displayed sustained release of the bioactive constituent, which was effective to prevent the colonization of bacteria and stimulate the skin cell migration [11].

2.2

Nanofibers Loaded with Metal Nanoparticles

With the advancement of nanotechnologies, metallic nanoparticles as typified by silver, titanium oxide, and zinc oxide nanoparticles hold promising potentials as compared with conventional antibiotics in dealing with emerging drug resistance in the environment as well as healthy facilities. The major advantages of metallic particles with nanoscaled size are that they displayed prominent and broad-spectrum antibacterial effect. As metal nanoparticles can be synthesized in a controllable and cost-effective manner, it has attracted many attentions to develop biosafe and effective antibacterial agents based on the metallic nanoparticles during the past decades [17, 18, 30, 31].

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The antibacterial mechanism of nanoparticles may involve multiple pathways as a result of the surface properties, released metallic ions, membrane disruption, and reactive oxygen species (ROS) upon contact with the bacteria [30, 32]. For example, silver nanoparticles have the ability to bind strongly to sulfur-containing proteins, leading to the structural disruption of bacterial cell membranes. Consequently, silver nanoparticles may enter the cell to perturb the cellular function and eventually cause cell death. Notably, increasing silver ions released from nanoparticles may induce considerable ROS generation, which is indicative of the cellular toxicity of silver nanoparticles [30, 33]. Through optimized microstructure, electrospun nanofibers loaded with metallic nanoparticles can reduce their inherent cytotoxicity while the antibacterial properties are preserved to a great extent. Roughly, the metallic nanoparticles can be loaded into the nanofibers in two ways: embedded or coated on the surface of nanofibers. The mixture containing the metallic nanoparticles can be electrospun to produce metallic nanoparticles embedded in nanofibers. As the outer layer of polymers start to degrade, the nanoparticles are released gradually, which result in effectively attenuated metal-mediated cytotoxicity. El-Aassar et al. produced polygalacturonic/hyaluronic acid nanofiber loaded with AgNPs, in which a sustained release of highly stabilized silver nanoparticles was achieved. The obtained dressing also showed desirable cytocompatibility and superior antibacterial performance. Moreover, the presence of hyaluronic acid in the dressing can contract the wound, which accelerates the wound healing [17]. Jatoi et al. blended ZnO/AgNPs in CA to produce CA/ZnO/AgNPs composite nanofibers. With AgNPs attached onto the surface of ZnO that embedded inside the nanofibers, the fabricated electrospun dressing exerted stable and long-term antibacterial effect [18]. The electrospun nanofibers loaded with metal ions can be achieved via immersion method, in which the metal nanoparticles are simultaneously formed on the nanofibrous scaffold under mild reaction conditions. For example, Moon et al. have reported a one-pot synthesis of silver nanoparticles-decorated fibers by simultaneous reduction of silver ions using sodium borohydride during the 3D foaming of cellulose matrix [19]. The presence of silver nanoparticles was found to remarkably improve the thermal stability of the produced fibers, which showed good antimicrobial activity with low cytotoxicity. The simple fabrication process of silver nanoparticles-decorated nanofibrous scaffold may allow for scale-up manufacture of antibacterial wound dressing. Besides silver nanoparticles, many metals including gold, zinc, and copper have been studied due to their prominent antibacterial activity. Compared to that of silver, gold nanoparticles can kill the bacteria without the generation of ROS and thus exhibit relatively low cytotoxicity. To deal with multidrug-resistant (MDR) bacterial infection, Wang et al. prepared PCL/gelatin/Au MBA NPs fibrous membrane as wound dressing. Both Au NPs and MBA cannot act as antibiotics when stand alone. However, the MBA attached Au NPs exhibit potent antibacterial effect against G+ MDR clinical isolates as compared with many clinically used antibiotics. Therefore, the prepared dressing shows great antibacterial property with low cytotoxicity

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[21]. Zhang et al. also confirmed that zinc, as an antimicrobial material, can promote the wound healing [34]. Recently, owing to their manipulable hollow structure and high specific surface area, metal-organic frameworks (MOFs) that made of metal ions in coordination with organic ligands have attracted great attention in the field of drug delivery. Consequently, some studies have focused on immobilizing MOFs on fibers for treating wound infection. Basically, there are mainly two methods to immobilize MOFs on fibers: Doping MOFs into fibers to produce the doped composite fiber or growing MOFs on the fiber surface through surface chemistry. Yang et al. revealed the differences of properties between nanofibers prepared by these two methods, using ZIF-8@gentamicin nanoparticles and polyacrylonitrile/gelatin (PG) [9]. Their study showed that ZIF-8@gentamicin coated on PG can increase the drug loading (Fig. 7), which resulted in superior antimicrobial performances and accelerating the wound healing. While gentamicin was efficiently encapsulated into ZIF-8, a synergistic antibacterial effect of blended nanocomposite was identified (Fig. 8). As ZIF-8@gentamicin embedded in nanofibers ensures the stable and long-term drug release, the produced dressing without the need of frequent change is suitable for the treatment of chronic wounds. Consistent with in vitro antibacterial results, the introduction of gentamicin sulfate in the dressing exhibited satisfying therapeutic effect to promote the wound healing [9].

2.3

Plant Extracts

Plant extracts also show excellent antibacterial property. Compared to metal particles, plant extracts have low cytotoxicity and high capability to promote wound healing, which is desirable for the formulation development of wound dressing. The antimicrobial activities of plant extracts originated from their active ingredient including flavonoid, phenolic acid, and anthraquinone. Que is a bioactive flavonoid found in many fruits and vegetables, which functions not only capable of blocking the synthesis of nucleic acids, but also inhibiting the cell wall synthesis. Thus, Que possesses good antimicrobial activity against G+ bacteria [25]. Kharat et al. reported Calendula officinalis extract (CO) that rich of quercetin in chitosan/ PEO nanofibrous scaffolds by electrospinning [26]. The incorporation of CO extract improved mechanical properties of nanofibrous membrane, which showed excellent antibacterial properties and wound healing ability via improving collagen synthesis, re-epithelialization, and remodeling of the tissue (Fig. 9). As a type of succulent plant, Aloe vera (A. vera) has been used in clinical practices for centuries due to its healing activities. In addition, the antibacterial property of A. vera is believed to be ascribed to anthraquinone group, which act as tetracycline that may inhibit the bacterial growth by blocking the ribosomal A site to interrupt cellular protein synthesis [23]. Moreover, the active ingredients of A. vera, including ascorbic acid and acetyl mannans, also exert strong antimicrobial effect [23, 24]. Yin et al. have used a sloping free surface electrospinning device to prepare

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Fig. 7 The morphological features of (a) ZIF-8 nanoparticles and different nanofibers observed by SEM: (b) PG + ZIF-8@gentamicin, (c) PG, (d) PG/ZIF-8@gentamicin; (e–g) EDS mapping of PG/ZIF-8 nanofibers. Reproduced from [9] with permission

large quantities of PCL/CS/A. vera nanofiber membranes for short-term dressing or acute wounds. The addition of A. vera effectively improves the hydrophilic and antibacterial properties of the constructed nanofibrous membranes [24]. Nowadays, an increasing number of plant extracts have been found to possess antibacterial activity. Shahrousvand et al. revealed that Melilotus officinalis extracts can inhibit the growth of broad-spectrum bacteria, with a stronger antimicrobial effect against Bacillus than Shigella [27]. Chelidonium majus L. has been proved to inhibit the growth of S. aureus and P. aeruginosa [35]. Centella exhibits a strong inhibition against S. aureus, E. coli, and P. aeruginosa, which are commonly found

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Fig. 8 (a) Images of healing efficacy of dressing-treated wounds based on the rat model; (b) Inhibition zones of different fibers against E. coli and S. aureus. Reproduced from [9] with permission

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Fig. 9 The antibacterial performance of the CS/PEO and CS/PEO/CO membranes against (a) S. aureus and (b) E. coli on 7 × 108 CFU/ml bacteria. Reproduced from [26] with permission

in wound dressing. And Hinokitiol has become well-known for its inhibitory effect on the growth of fungi, bacteria, and insect [36]. Some essential oils also have good antibacterial activity. When preparing nano dressings containing essential oils, a surfactant is needed to increase the hydrophilicity. Both hydrodistillate of lavender oil and clove essential oil have the ability to fight against S. aureus and E. coli infection, which have been proven experimentally [37, 38]. Moreover, elicriso oil is effective against C. albicans by Aderibigbe et al. [39]. Apart from antibacterial property, other properties of plant extracts were summarized in Table 2, which are suitable for the wound healing.

2.4

Biomacromolecules

For the effective wound management, functional dressings based on protein, polysaccharide, and biopolymers are promising options due to their excellent biocompatibility and biodegradability. The beneficial topographical features of these biomacromolecules resembling extracellular matrix (ECM) can provide an optimized physiological microenvironment for wound healing. Proteins like collagen, gelatin, and silk protein, either animal or vegetal derived, have been successfully electrospun into nanofiber membranes. Organic solvents are necessary for electrospinning process [44]. For example, collagen must be dissolved in hexafluoroisopropyl alcohol (HFIP), and gelatin can be dissolved in trifluoroethanol or formic acid.

High biocompatibility, ECM-like Low cytotoxicity Anti-inflammation; antioxidant

Quercetin [14, 25, 40, 41] + – +

Chelidonium majus L. [35] +

– –

Table 2 Beneficial properties of plant extracts for the wound healing

– –

Centella [36] + + +

Calendula [26] + + –

Papaya extract [42] –

– +

Melilotus officinalis [27] –

– +

Artemisinin [43] –

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In the body, collagen is the most abundant protein involved in many cellular functions including shape maintenance, differentiation, migration, and protein synthesis. It also plays an important role in all stages of the wound healing cascade (hemostasis, inflammation, proliferation, and reconstruction) [45]. Collagen can preserve white blood cells, macrophages, fibroblasts, and epithelial cells. Simultaneously, it stimulates cell activity, which in turn attracts cells to migrate to the wound site and promotes the deposition of new collagen matrix. In the stage of chronic wound reconstruction, the overexpression of matrix metalloproteinase (MMP) activity, the imbalance of transforming growth factor, and other factors can lead to matrix protein overgeneration and fibrosis [45]. Collagen can effectively absorb exudate from the wound, bind and protect the growth factors, and inactivate the overactivated MMP in the wound exudate, so as to promote wound healing. Previous studies showed that the combination of epidermal growth factor and collagen could effectively reduce wound pain and improve wound healing rate [46]. Gelatin is a kind of denatured protein obtained by hydrolysis of collagen. Compared with collagen, gelatin has weaker antigenicity and can activate macrophages to achieve hemostasis. Due to its RGD (Arg-Gly-Asp) motif, the gelatin is able to bind integrin and effectively promote cell adhesion to biomaterials [47]. Ajmal et al. developed PCL-gelatin based nanofiber membrane, in which the addition of gelatin enhanced the hydrophilicity and biodegradation rate of the nanofibers. The results confirmed the use of the dressing contributed to the accelerated closure of full-thickness wound [40]. CS is a partial deacetylated product of chitin, which can be extracted from the shell of shrimp and crab in large quantities. Besides its antibacterial, adhesive, and hemostatic properties, CS can open the tight junctions of epithelial cells and enable other biological macromolecules with antibacterial and hemostatic functions to reach the lesion through epithelial tissues [48]. The electrospinnability of CS can be increased by adding poly(ethylene glycol) (PEG) or PVA into the acetic acid/ water mixture. Nanofiber membranes can also be prepared by dissolving CS in aqueous solution of trifluoroacetic acid or concentrated acetic acid [49]. Bacterial cellulose, as a new natural biological material characteristic of microfiber, possesses good mechanical, emulsification, and gelation properties. In addition, it has simple production process, low production cost, and the activity to promote tissue regeneration and wound healing, which has gained increasing attention in the field of tissue regeneration. Solway et al. compared the healing rate of diabetic foot ulcer (DFU) treated by bacterial cellulose wound dressings and Xeroform vaseline gauze and pointed out that microfibrin could shorten the wound healing time and reduce the contamination of external bacteria [50]. Alginate fiber dressings have been used for wound care since the 1980s. Nowadays, alginate fiber dressings with various functional properties can be prepared by adding some special metal ions (Ag+, Cu2+, Zn2+, etc.) or other functional substances in the spinning process. Compared with traditional gauze dressings, alginate dressings have many advantages [51]. When used on the wound surface, alginate dressings can quickly absorb water and maintain a wet and sealed environment, promote the growth of new granulation tissue, avoid direct exposure of nerve endings to the

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air, so as to achieve the analgesic effect. Moreover, it plays the role of autolysis debridement, avoiding wound infection and inflammation, accelerating wound healing, and achieves better clinical effect [52]. Many biocompatible polymers have been applied for wound management typified as PVA. PVA is a kind of polymer hydrolyzed from polyvinyl acetate with plenty of hydroxyl groups on the molecular chain, which has good film forming and bonding properties [53]. Moreover, PVA shows good biocompatibility, biodegradation, and hydrophilicity, high mechanical strength, and low toxicity. Therefore, when blended with other polymers like CS, the presence of PVA in the wound dressing contributes to its improved solubility, viscosity, thermal and mechanical properties [54]. Compared with other biomacromolecules, silk protein is cheap, widely available, and easy to obtain. Silk protein can promote the differentiation and reproduction of human epidermal cells and fibroblasts, so as to promote wound healing. Moreover, it is easy to process and compound with other materials, so it has been extensively used in the electrospun fiber dressing for wound healing [55].

3 Summary and Prospective Skin wound healing involves complex processes in human body, which propels the development of dressing materials with multifunctional properties. The ideal wound dressings should have excellent antibacterial property, biocompatibility, and the activity to promote tissue regeneration. Among various tissue regeneration scaffolds, electrospun nanofibers show unique advantages. Nanofibers possess relatively small pore size and high surface area to volume ratios with ECM-mimicking architecture beneficial for cell adhesion and proliferation and tissue regeneration, which can effectively isolate pathogenic microorganisms in the external environment. Owing to these structural topographical properties, the produced wound dressings loaded with antibacterial active material including antibiotics, metal nanoparticles, and plant extracts can rebuild the skin barrier, promote antibacterial effect, and accelerate wound healing. Moreover, the systemic diffusion of antibacterial compounds such as antibiotics is reduced, so as to improve the bactericidal efficacy and reduce the cytotoxicity. Therefore, antibacterial nanofiber dressings have been widely used in local anti-infective treatment of trauma patients. In addition, the smart design of drug delivery may endow the antibacterial constituent blended nanofiber stimuliresponsive attributes [56], which make it possible for the targeted release of antibiotics at the infection site and thereby reduce its usage. The intelligent molecular design of nanofibrous dressing will provide promising solution to deal with antibiotic resistance. In recent years, more and more attention has been paid to functional dressings based on biological macromolecules, in which proteoglycan and polymer are the main components. These biological macromolecules have excellent biocompatibility and biodegradability, which are essential for wound healing. Such dressings can

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establish an optimized physiological microenvironment for wound healing, so electrospinning fiber scaffolds based on biological macromolecules have been widely used in the biomedical field. Among them, biopolymers are relatively simple to prepare and can be modified with a variety of substances [57]. In comparison, protein-based materials have better biocompatibility and biodegradation safety. Carbohydrate-rich materials have better water absorption and can absorb wound exudate. In a long run, the further molecular understanding of these therapeutic proteins associated healing process will help to develop potent and diverse application specific wound dressings. Acknowledgments We thank the financial support of National Natural Science Foundation of China (U1703118), Natural Science Foundation of Jiangsu Higher Education Institutions of China (No. 19KJA310003), Natural Science Foundation of Jiangsu Province (No. BK20181364) and a project funded by the Priority Academic Program Development of Jiangsu Higher Education Institutions (PAPD) and the Open Project of Jiangsu Biobank of Clinical Resources, No. SBK202005001.

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Adv Polym Sci (2023) 291: 335–360 https://doi.org/10.1007/12_2022_139 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 11 October 2022

Application of Electrospun Polymeric Fibrous Membranes as Patches for Atopic Skin Treatments Urszula Stachewicz

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1.1 Challenges in Skin Treatment . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1.2 Electrospinning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Electrospun Patches . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Patches from Biodegradable Polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Patches from Non-biodegradable Polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract All over the world people suffer from many skin problems including eczema, atopic dermatitis (AD), and various allergy-related issues, which generally cause skin dryness, redness, and itching. To protect the skin barrier, increasing its moisture and hydration is highly recommended, which helps to prevent further bacterial infection and external irritation. Electrospun membranes are ideal material for skin patches allowing topical drug delivery and reducing transepidermal water loss on the skin due to their high surface area to volume ratio. Therefore, in our research we focus on electrospun polymer membranes with a porosity above 90% to deliver gamma-linoleic acid (GLA) from natural oil to increase skin hydration. Various geometries of membranes (aligned, random, nano- and microfibers) and polymers (PVB, PCL, PHBV, PI, PA6, PS) are used to design patches with controlled oil release over a few hours. The experimental studies in vitro and in vivo have been confirmed with the theoretical modeling of oil transport through the porous patches. The results indicate the beneficial effects of the electrospun patches

U. Stachewicz (*) Faculty of Metals Engineering and Industrial Computer Science, AGH University of Science and Technology, Kraków, Poland e-mail: [email protected]

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and their high biocompatibility. We show that oil spreading and transport in patches depend on the sizes of the fibers and pores, their surface properties, and porosity. The controlled transport over 6 h of oil containing GLA is directly correlated with the increased skin hydration level. Additionally, electrospun fibers and membranes can be easily modified by blend electrospinning of polymers with oils, or coating fibers with urea or chlorine used in treating AD. The electrospun patches can be simply applied overnight for protecting the skin and improving the comfort of people suffering from various acne and eczema problems. Keywords Essential oils · GLA · Membranes · Patches · Polymer fibers · Porosity

1 Introduction 1.1

Challenges in Skin Treatment

Environmental changes have a huge effect on us especially from healthcare. The increasing and uncontrolled pollutions and lifestyle result in increasing allergies especially in young generation and kids. Skin is our largest organ with the protecting functions. Always, the skin has the first contact with the environment as it is the most exposed. It causes various skin problems, which are expanding continuously. Atopic dermatitis (AD) is the most common form of eczema and causes a red rash and is common among people with a personal or family history of hay fever, food allergies, or asthma, as these conditions commonly occur together. Skin redness, dryness, and itchiness are the most frequent symptoms of dermatitis, but scaling, flaking, and blisters can also occur, see Fig. 1, which continuously reduces the quality of life [1]. Acute AD produces weeping, oozing plaques of very itchy skin. Itching is a characteristic symptom of eczema, a part of transepidermal water loss, swelling,

Fig. 1 (a) Example of atopic skin with the characteristic redness symptoms [5], (b) Pictures of AD patients with wet wrap therapy [6]

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redness, and dry skin [2]. Currently used skin treatments have difficulties to provide a long-term cure in eczema, therefore new ways for skin regeneration and keeping it hydrated are needed. Currently, all the AD treatment is focused around increasing skin moisture with standard emollients, often topical steroids are used, which helps in reducing inflammation and itching [3, 4]. Skin is basically built of three main layers: epidermis, dermis, and subcutaneous layer. The epidermis mainly consists of keratinocytes, which are playing an important role in keeping the skin moisture and heat to protect us, and in our immunology during the wound healing processes [7]. In epidermis, keratinocytes proliferate and start gradual differentiation, when they significantly change their morphology and start to produce keratin, interleukins, cytokines, and growth factors. Keratin is a filament-based protein, which has a profound function in cell signaling and intracellular vesicle transport [8]. Keratin acts as an essential part in constructing natural protection for many animals in the form of hooves [9, 10], wool [11], horn [12], hair and feather [13]. In dermis, another important protein is present called collagen, which provides strength to the skin [14] not only to humans but to other species too [15]. Other parts of skin are obviously cells, elastin, and extrafibrillar gel-like substances. However, the most exposed epidermis part is the sub-layer of Stratum Corneum (SC) built of lipid matrix containing ceramides, free fatty acids, and cholesterol, cells such as corneocytes. The SC lipids are often disorganized which causes reduction of the skin barrier against pathogens, chemical and environmental conditions, resulting in several skin inflammatory diseases such as AD. The structure of SC is disturbed and requires bringing back the skin protection and moisture. Here a long-term and stable treatment is very important, such as wet wraps that can be used to prolongate the treatment and drug activity [4, 16]. However, the drug release from standard bandages is challenging to control, as the geometrical design and properties of materials have a significant effect on drug delivery systems [17]. When we look for the best patches for AD treatment, we think about something similar to bandages and having properties of wound dressings which is suitable for both air and water vapor exchange. Importantly, the patches should be easy to handle and made of flexible and elastic, durable, comfortable and soft for the skin material [18]. Sometimes the night dressing in the form of wet wraps is applied, which is made from textiles soaked in water with medications such as topical steroid creams, and applied to the affected skin on the body, see example in Fig. 1b. Wet wraps are applied in the evening after bathing with pajamas dressing on top [19]. Additionally, the water vapor transmission rate is an important factor and has a great impact on skin dressings [20]. Skin lipids and lipids metabolism are important in transepidermal water loss and improving skin barrier [21]. Cholesterol (CHOL), ceramides (CER), and free fatty acids (FFA) maintain skin hydration. The FFA absorbed from the food are linoleic acid (LA) and α-linoleic acid (ALA) as they cannot be synthesized by the human body [22]. However, metabolite of LA, ƴ-linoleic acid (GLA) can be found in natural oils such as evening primrose [23] and borage [24] and importantly was proved to be effective in AD treatment. When the catalyst of GLA synthesis from LA so-called δ-6-desaturase is decreased, its deficiency causes dry skin. Both LA and

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GLA are present in evening primrose oil. A part of evening primrose oil, oils from blackcurrant seed (Ribes nigrum) [25], borage [26], hempseed [27], and black cumin (Nigella sativa) [28, 29] contain GLA. Various pharmacological and non-pharmacological therapies using antibiotics and topical calcineurin inhibitors, emollients, and topical corticosteroids are used but they show adverse side effects [30, 31]. These types of treatment can be supported by functionalized textiles [6] or nanobandages [32]. One of the candidates for the patches are electrospun fibers in the form of membranes that are reported to be a favorable material for wound healing and tissue engineering [33, 34]. Numerous research studies of electrospun fibers are devoted to skin substitute and wound healing processes [35]. Here from materials perspective one can look for the biodegradable materials that can be replaced by the native tissue [36, 37]. Many of these aspects can be neglected when we try to design the bandages or patches used on the skin. The non-degradable and synthetic polymers can be easily applied including other regenerative medicine areas too [38, 39]. The advantage of porous membranes produced via electrospinning is the high surface area and porosity [40, 41] giving not only the breathability and desired wetting properties [42, 43], but also the possibility to incorporate drugs and oils treating many skin problems [44].

1.2

Electrospinning

In general, electrospinning is a simple and cost-effective method to produce polymer fibers and membranes. It is electrohydrodynamic process [45] where we apply a high voltage typically to the nozzle with passing polymer solution. Polymer solution is delivered through pumps or pressure-driven systems. When the voltage is applied with positive or negative polarities [46] to the nozzle, the potential difference is created between it and the grounded collector. The charges accumulate on the pendant droplet of polymer solution which is hanging the tip of the nozzle. The electrostatic forces overcome surface tension of polymer solutions and from the formed cone, the jet ejection starts toward the collector. Depending on the environmental conditions during electrospinning solvents present in the polymer solution evaporate rapidly and cause instabilities called whipping motion [47]. The conditions such as humidity allow controlling fibers morphology and porosity [48] as well as solvents and non-solvents used for polymer solution preparation [49]. The electrospun fibers can be deposited on various collectors [50]. The stationary collectors allow obtaining random orientation of fibers and the rotating drums cause the alignment of fibers [51, 52]. A part of that various patterns of electrospun fibers was obtained [53, 54], often to enhance the 3D structure of scaffolds used for tissue engineering [55, 56]. In electrospinning various types of nozzles can be used, for example co-axial [57] or side-by-side [58], see example in Fig. 2. The system with the multiplied nozzles [59] or free surface electrospinning without any nozzle [60] can be applied to

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Fig. 2 Schematic of electrospinning setups with (a) standard nozzle, (b) side-by-side nozzle, (c) co-axial nozzle, from [71], (d) with two nozzles with the possibility to electrospun two polymer solutions at the same time to produce composite meshes, from [39]

increase the efficiency of fiber production. The electrospinning with two nozzles is able to create the composite meshes with at least two different polymers [61] or have the electrospraying combined with electrospinning, at the same time, or in the layer by layer system [62]. The electrospun membranes are often used in smart textiles applications [63] for energy [64–66] and water harvesting [67, 68]. Electrospun fibers are also used in many commercially available products because electrospinning is a cost-effective technology [69], that is promising and useful in constructing multifunctional medical patches [70].

2 Electrospun Patches 2.1

Patches from Biodegradable Polymers

Generally, the electrospun scaffolds are produced from biodegradable polymers. Mostly they are based on poly(vinyl alcohol) (PVA), poly(ethylene oxide) (PEO), polycaprolactone (PCL), poly(lactic acid) (PLA), poly(glycolic acid) (PGA), and poly(lactic acid-co-glycolic acid) (PLGA), there is also a large group of polyhydroxyalkanoates (PHA) polymers used in biomedical applications [33, 72]. For electrospun patches we will focus on just two biodegradable polymers, namely, poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV) and PCL.

2.1.1

Poly(3-Hydroxybutyrate-Co-3-Hydroxyvalerate)

PHBV is a thermoplastic aliphatic polyester which is produced by bacteria and gaining a lot of interest as biomaterial in many medical applications. Especially

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electrospun fibrous scaffolds showed great biocompatibility and applicability for culturing osteoblasts and fibroblasts, indicating universal application for different types of tissue regeneration [73]. PHBV is often combined with the ceramic particles such as hydroxyapatite or titanium dioxide [37, 74], zinc oxide [75], strontium carbonate [76], or pearl powder [77]. Here, surface properties play important role in all cell material interaction [78], that may be affected by material degradation and stability related to aging of biodegradable materials. Therefore, the physicochemical properties of PHBV scaffolds were verified over time by zeta potential tests. Cell adhesion, proliferation, and differentiation can be controlled via surface charges of the biomaterial. The important parameter referring to contact with an aqueous electrolyte solution such as zeta potential is here very useful to measure to be able to understand many surface activities [79]. Moreover, surface charges control the adsorption of proteins affecting this way of cell attachment and filopodia formation by cells [78]. The zeta potential of electrospun PHBV fibers and films presented in Fig. 4 was studied over 8 weeks of aging to evaluate their stability. The PHBV fibers were produced with positive and negative voltage polarity as it has been shown before the voltage polarity is able to control surface properties of electrospun fibers [38, 66, 80]. The zeta potential of PHBV fibers and films was tested in aqueous KCl solution and the isoelectric point (IEP) was measured. In the range of pH 3–6, a linear profile of the zeta potential on pH is observed. The difference in zeta potential between electrospun PHBV+ and PHBV- membranes and films was noticed; however, after 4 weeks of storage the differences start to disappear and PHBV+ fibers and films seem to undergo some changes. After 8 weeks the zeta potential results for all samples become almost undistinguishable, while the IEPs of PHBV+ and PHBV- settled at pH 3.5 ± 0.1, the IEPs of the PHBV film also achieved this pH after 8 weeks of storage [81]. To assess the biocompatibility of PHBV samples confocal laser scanning microscopy (CLSM) was used together with the standard proliferation tests. Fibroblasts attached to the surrounding PHBV fibers in all directions. Importnatly, cells stretch, elongate and adjust to the PHBV fibrous scaffolds. Regardless of the surface chemistry and zeta potential values, or material topography, all PHBV samples indicate excellent cellular response suggesting the perfect biocompatibility.

2.1.2

PHBV Fibers Blend with Evening Primrose Oil

The electrospinning of PHBV polymer solution blended with 5% and 10% of evening primrose oil (EPO) was performed to enhance the mechanical properties and wettability of electrospun membranes. The additions of EPO cause the changes in the morphology of PHBV fibers that were observed using SEM, see Fig. 3. The fiber diameter increases with added oil in the PHBV polymer solution, it was 2.46 ± 0.31 μm, for PHBV with 5% of EPO was 3.80 ± 0.42 μm, whereas for the fibers with 10% of EPO was 5.92 ± 0.76 μm. The addition of EPO to electrospinning of PHBV fibers increases also the maximum strength of PHBV meshes in the tensile tests, Fig. 3a. Additionally, the water contact angle was measured on all the PHBV

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Fig. 3 SEM micrograph of electrospun (a) PHBV fibers and (b) blend PHBV fibers with 5% of EPO and (c) 10% of EPO. (d) The mean values of fibers’ diameter (Df) of PHBV fibers presented in the box chart. (e) The representative stress–strain curves from tensile tests. (f) The water contact angles values within 15 s, on PHBV fibers and blends of PHBV fibers with 5% and 10% of EPO, from [82]

membranes, see Fig. 3b. No changes in the contact angle were noticed within 15 s of the measurement on PHBV fibers, however on PHBV with EPO the contact angle started to decrease. For PHBV +5% EPO from 124.3 ± 1.2° to 117.7 ± 1.1°, then to 40.5 ± 1.3° in 12 s. For PHBV fibers with 10% EPO the contact angle decreases from 114.4 ± 0.9° to 63.4 ± 1.1°. The electrospinning PHBV solution with the addition of EPO changed the hydrophobic properties of PHBV membranes to hydrophilic [82].

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The dressing or patch application requires appropriate stretchability and wettability of the electrospun material, especially if they should be applied on the skin. Additionally, the oil spreading in PHBV samples was investigated on top and bottom, see Fig. 10a. The differences between the top and the bottom of oil spreading are much smaller for PHBV+EPO, which indicates more oil delivery through the membrane. Moreover, PHBV blends with EPO showed also a decrease in water vapor transition rate (WVTR) of about 30% and 40%. In Fig. 4, the decrease of 50% was noticed for pure PHBV fibers. The thermal imaging points out that the fibers are an excellent heat transfer barrier, see Fig. 4c, d. PHBV + EPO loaded with additional oil is effective as patches for increasing hydration by 15–20% during skin moisture tests on volunteers [82].

Fig. 4 (a) The graph comparing the oil spreading area on PHBV electrospun fibers and blends with EPO over 3 h. (b) Water vapor transmission rate in g m-2 day-1 and (c, d) thermal camera images showing the heat transmission rates through PHBV, PHBV +5% EPO, PHBV + 10% EPO electrospun fibers and samples with additional 30 μl of EPO, from [82]

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343

Polycaprolactone

PCL is a biodegradable polymer belonging to the group of saturated aliphatic polyesters that are extensively explored for medical applications [83]. The electrospun PCL scaffolds were studied in terms of their biocompatibility for various tissues [49, 84, 85]. Via electrospinning PCL fibers morphology can be easily controlled [86, 87] and PCL fibers were applied as patches, containing Spirulina and alginate used in many cosmetic products [88]. In particular, PCL fiber membranes were also combined with hemp oil to construct skin patches for AD treatment. Four types of electrospun PCL fibers smooth, porous, random, and aligned (Fig. 11) were investigated to design various patches according to their wettability and mechanical performance [89]. The hemp oil spreading and transport through the electrospun PCL membrane were investigated on the skin model based on the gelatin, see Fig. 5. Within 6 h the patches from porous PCL fibers (pPCL) showed similar spreading area both for random and aligned fibers but larger than for non-porous, so called smooth, sPCL fibers. The porosity of the PCL fibers significantly increased the hemp oil spreading area [89]. We observed the largest spreading area for sPCL aligned fibers due to its short distance between fibers and elongated pore shape, but eventually it inhibited oil penetration deep through the patch to the skin. Therefore, the patches with the aligned fibers had the lowest oil release. The volunteers’ skin hydration level was measured after 6 h of patch application. Here we observed the beneficial role of hemp oil, as 20% increase in skin moisture was recorded. The regeneration of the damaged skin barrier is much faster when the transepidermal water loss is controlled to maintaining sufficient hydration. Further the tested PCL patches consisted of aligned porous PCL and random porous PCL fibers showed up to 55% oil release during the 6 h tests. The electrospun PCL membranes are able to provide long-term, controlled oil release, which suggests their suitability for delivering treatment for AD patients.

2.2

Patches from Non-biodegradable Polymers

Poly (vinyl butyral-co-vinyl alcohol-co-vinyl acetate) (PVB) is a hydrophobic polymer used as a component in laminated safety glass in the automotive industry [90], coating [91], and electrodes [92]. PVB was electrospun using a portable device for wound healing treatment [93] and using more standard setups [94]. Using low (LMw) and high molecular weight (HMw) PVB nano- and microfibers can be produced via electrospinning [95], as presented in Fig. 6. Both types of PVB membrane showed good mechanical properties with the higher extension for microfibers during tensile testing. The biocompatibility of both types of electrospun PVB membranes was verified in cell culture study, together with the cytotoxicity tests. The following oils: evening

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Fig. 5 The images showing the hemp oil spreading in 6 h tests for (a–c) smooth sPCL random, (d–f) smooth sPCL aligned, (g–i) porous pPCL random, (j–l) porous pPCL aligned, and (m) graphical representation of the spreading area of hemp oil measured every 30 min during 6 h tests. (n) Schematic of the layered sPCL/pPCL patches with hemp oil, from [89]

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Fig. 6 Macroscopic picture of PVB (a) – nanofibers mat and (b) – microfibers. (c, d) – SEM micrographs and images of water droplets on electrospun PVB membranes, respectively, for nanoand micro-fibers. (f) Exemplary stress–strain characteristics from tensile tests of PVB membranes based on nano- and micro-fibers. The 3D reconstruction from FIB-SEM of PVB membranes for: (e) – nanofibers and (g) – microfibers, from [95]

primrose, black cumin seed, and borage were incorporated between PVB fibers to test patches as natural oil carriers for atopic skin treatment. In Fig. 7, the wetting and spreading of oils in patches were performed to utilize them later in vivo test on skin. Nanofibers are better carriers for low viscosity oils, as they penetrate the 3D network of fibers easier than microfibers, as indicated by the oil spreading tests. Additionally, we performed the numerical flow simulations of the EPO through the PVB nano- and microfiber-based patches, Fig. 8. The velocity profile of oil in nanofibers was greater than for microfibers, as the capillary pressure is higher in smaller pores. The numerical modeling results were confirmed with the simple experiments where the time needed for oil to travel through the membranes with the known thickness was measured. The setup with the high-speed camera was placed at the bottom of the patches and allowed to observe the differences in time after the oil deposition on the nano- and microfibers networks. For PVB nanofibers, evening primrose oil reaches the bottom of the samples twice faster than for microfibers. The results in Fig. 7 were the first step to show that PVB patches able to deliver natural oils to keep the skin moisturized. The biocompatibility studies were extended to keratinocytes. Additionally, more studies related to the fibers’ geometrical arrangements and their change in oil transport and spreading were performed [96]. The application of the patches around the body was considered taking into

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Fig. 7 Images of oil spreading on PVB nano- and microfibers. (a–f) Borage oil on nano- and microfibers (g–l), black cumin seed on nano- and microfibers (m–s) EPO on nano- and microfibers (j–l), The surface area of oils spreading on PVB nano- and microfiber for all tested oil (t, u) the comparison of maximum spread of oil in PVB patches during 60 min, from [95]

account the stretchability of patches, when their pore size and shape are changing and affecting the oil spreading [96]. In the elongated patches with the aligned fibers, the shape of spreading area of oil is changing to ellipsoidal as oil flows along fibers in the elongated PVB samples. In elongated electrospun samples also the distance between the fibers is reduced, which limits pore sizes [97]. The pore diameter is usually larger when the membrane is constructed from microfibers [98]. In Fig. 9, the results from the skin hydration tests are presented from 6 volunteers with random nano- and micro-fiber-based PVB patches with EPO and stretched, so elongated patches. The skin hydration level was measured before and after placing the patch on the skin, and as a control sample just PVB fibers were used. Both types, as-spun and elongated patches show the significant increase in the skin hydration of the 6 h tests. Over 100% increase in moisture of skin was reported for volunteers with the initial dry skin conditions, below 20% hydration. Such a low skin hydration represents the dry skin condition typical for atopic skin. Notably, the skin test with EPO indicated the increase in skin hydration when patches with PVB nanofibers were used.

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Fig. 8 The results from numerical simulation of EPO flow through the PVB (a) nano- and (b) microfibers; (c, d) cross-sectional view (ZX plane) of the velocity profile for nano- and microfibers; (e) velocity profiles for PVB patches from the cut lines driven across the presented ZX slices in (c, d), from [96]

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Fig. 9 Skin tests of PVB patches with EPO on six volunteers (a–f) as-spun random nano- and micro-PVB (i–n) stretched up to 140% and elongated nano- and micro fibers. (g, o) The examples of patches on the skin just after oil deposition and (h, p) after 6 h tests, from [96]

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349

PVB Membranes and Urea

The electrospun PVB can be utilized as multifunctional dressing by itself or with the additional modification using other skin curing substances. Urea is one of them, which is broadly used in dermatology [99]. The formulation containing 4–10% showed improved skin moisture in AD treatment [100, 101]. Therefore, various methods of adding urea to the PVB fibers by blend electrospinning and electrospraying, containing 5% and 10% of urea, were tested. The morphology of produced PVB patches was characterized, see Fig. 10, including their chemical composition and wetting properties. The PVB patches with urea were biocompatible with keratinocytes. Moreover, release tests of urea from the patches were verified. Next, the cytotoxic was tested of PVB patches with urea via induced by cells’ direct contact with urea, confirming the suitability of the urea-based patches for topical application. In Fig. 10g, the results from the cumulative release of urea are presented showing the highest release for the blend electrospinning of PVB with 5% of urea. The low release level from the samples produced via electrospinning and electrospraying at the same time was limited due to the small amount of urea deposition. The study showed that the immediate effect (in 30 min) of urea release can be achieved with PVB fibers blended as indicated in Fig. 10g. During 6 h treatment, which is the typical overnight time, the release of urea from PVB patches is constant proving and easy and comfortable way of using them as night dressing.

2.2.2

PS and PA6 Composite Membranes

To improve the skin barrier in eczema treatment the patches based on electrospun hydrophobic polystyrene (PS) and hydrophilic nylon 6 (PA6) with oils were created to improve skin moisture. The membranes were produced by electrospinning: PS, PA6, composite PS–PA6, and sandwich system combining PS with PA6 layers [103]. Hydrophilic polymeric carriers used in the drug delivery systems showed a faster drug release than hydrophobic, thus often the combination of both is applied [104]. Here three types of oils were investigated for patches: evening primrose oil, borage, and black cumin seed. Oil spreading in the membranes, fluid uptake ability (FUA), and WVTR were studied in great detail to verify the effect of fibers diameter and their wettability on designing patches. The WVTRs were measured after 24 h for PS, PA6, composites PS–PA6, and gauze as a reference material. Adding oils to electrospun membranes during WVTR tests decreased it significantly for PA6 and composite PS–PA6 samples, but for PS and gauze it remained the same. WVTR in electrospun membranes depends on their porosity and pore size [105], the larger pores the gas or moisture is easier transferred through the membrane [106]. The hydrophilic PA6 patches were characterized with the largest oil spreading than PS and composites based on PS-PA6 fibers. Regarding the in vivo tests, for volunteers with the dry skin conditions, repressing patients with AD, the greatest increase in skin moisture was for patches based on PA6 fibers, composite PS–PA6 and

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Fig. 10 SEM micrographs of electrospun, (a) PVB, (b) blend PVB with 2% of urea, (c) blend PVB with 5% of urea, (d) electrosprayed urea on PVB fibers. (e) Histogram of fiber diameter distribution (f) Example of PVB patches with 5% of urea on the skin in a dry form and sprayed with water. (g) The release curve of urea from blend PVB fibers and electrosprayed urea on fibers, from [102]

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sandwich layers system based on PS + PA6. In case of hydrophobic PS patches the results are strongly dependent on volunteers’ skin type. Additionally, the sandwich system showed the importance of including hydrophilic and hydrophobic membranes in designing patches for well-controlled oil release over a few hours.

2.2.3

PI Membranes

The electrospun polyimide membranes were previously investigated in terms of biocompatibility [107], however we further investigated it with fibroblasts and additional in vivo skin test for skin patches and dressing. PI is resistant to many chemical reagents and PI membranes, which are highly porous (95.6%) and have great stretchability, see Fig. 11. Here the PI patches with blackcurrant seed oil rich in GLA were investigated [108]. The addition of blackcurrant seed oil in PI patches reduces the permeability thus WVTR values, see Fig. 11a. The electrospun membranes used as patches act as excellent oil reservoirs, according to the applied pressure oil is released, which enhances the GLA delivery to the skin apart from the standard oil diffusion process through the porous membrane. The GLA mass transport through the PI membrane by using with blackcurrant seed oil was modeled. We used the 3D reconstruction of the electrospun PI membrane, showed in Fig. 11d to model the patches. The skin model consisted of the top two layers: SC and epidermis. In the numerical simulations the mass transport by diffusion was used, however, the concentration gradient between the blackcurrant seed oil and the skin was the main driving force for GLA transport. GLA moved across the membrane by decreasing the concentration gradient and it started to be distributed along the skin layers. The modeling proves that PI patches are suitable for long-term oil release and are able to deliver GLA to the skin. Simply, the longer the patch is applied, the more GLA can be delivered to the atopic skin. In vivo experiments showing the direct increase in the skin hydration are correlated with the delivering GLA in numerical simulation results.

2.2.4

PI Membranes with Chlorine

Often patients with AD have an increased amount of Staphylococcus aureus on the skin surface. To reduce inflammation on skin the PI nanofibers modified with sodium hypochlorite (NaOCl) were designed in the form of patches [109]. Taking a bath in diluted bleach is commonly applied in clinical practice as they are able to reduce AD severity [110, 111]. The electrospun PI membranes were immersed in 3.7–4.5% NaOCl (Avantor, Poland) solution for 30 min, according to the previous protocols [112]. The active chlorine concentration was determined just before the preparation for the PI membranes bathing with a standard iodometric/thiosulfate titration method. The presence of NaOCl in PI patches was confirmed with FTIR and XPS. Additionally, the zeta potential was measured and its reduction was observed in the range from -38.7 to -30.7 mV for the whole tested pH range. The low value

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Fig. 11 (a, b) SEM micrographs of electrospun PI membrane (c) histogram of PI fiber diameter; (d) 3D reconstruction based on FIB-SEM tomography of electrospun PI membrane; (e) the stress– strain curves of PI fibers

of zeta potential is able to enhance the antimicrobial activities of surfaces [113, 114]. The antimicrobial efficacy assay against the Gram-positive model organism S. aureus and the Gram-negative model, Escherichia coli was performed. The PI patches with active chlorine demonstrated effective bacterial killing, showing their excellent antimicrobial properties [109].

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3 Summary Within this chapter the examples of designing skin patches based on the electrospun polymer membranes were presented. Our approach is to produce wet dressing for irritated skin based on oil droplets captured between fibers in electrospun highly porous membranes. The nanofiber patches are used as reservoirs for oil droplets to serve as a wet dressing to help hydrate skin, reduce itching and redness. The patches can be placed directly over the affected skin area and covered with a dry garment. These electrospun patches can be worn overnight or during the day for longer periods of time comparing to traditional wet dressing. Long and slow release of essential oils through proposed here patches reduces the need of very frequent application of body lotion and has remarkable effect on the skin condition. The biodegradables such as PCL and PHBV and non-biodegradables PVB, PI, PA6, and PS polymers were used to create skin patches. Various geometries of membranes with aligned and random fibers were used and also with the stretched random fibers. The PI patches with blackcurrant seed oil were tested up to 6 h with the skin hydration tests every 1 h. After 3 h, or usually from the measurements taken between 4 and 5 h the increase is reaching the maximum, which was at least 20% increase from the initial tests. After 6 h it was slightly reduced in comparison with the fifth h test. In nanofiber-based PVB patches with evening primrose oil after 6 h we could expect the increase event up to 30% for the dry skin. The studies indicated that oil spreading and transport in the PVB patches depend on the fiber diameter. The blend electrospun PHBV fibers with 5% of EPO with deposited EPO in the membranes we obtained the maximum 20% increase. Summarizing all the in vivo studies we showed 20% of skin hydration increase for the following patches after 6 h tests for the patient with the low initial skin hydration level, so-called dry skin conditions typical for patients with AD: • • • • •

Random PI nanofiber-based patch with blackcurrant seed oil Random PVB nanofiber-based patch with evening primrose oil Random PHBV + 5% EPO microfiber patches with evening primrose oil Random porous PCL microfiber patch with hemp oil Random composite membranes with PA6 nanofibers and PS microfibers with borage oil

The dry dressings used as a control sample did not affect skin hydration level. Importantly the WVTR is typically between 1,000 and 1,500 g m-2 days-1 for electrospun membranes soaked with oil. The WVTR for skin is 204 g m-2 days-1 [115] and for the wound dressing 2,000–2,500 g m-2 days-1 [116]. The electrospun patches are able to fulfill the gap in the current development of skin patches based on the presented research here. Electrospun fibers can be easily adapted and modified according to the patient needs and conditions and provide the long-term release of active molecules and drugs.

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Acknowledgments This research was part of the “Nanofiber-based sponges for atopic skin treatment” project carried out within the First TEAM program of the Foundation for Polish Science co-financed by the European Union under the European Regional Development Fund, project No POIR.04.04.00-00-4571/17-00.

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Adv Polym Sci (2023) 291: 361–408 https://doi.org/10.1007/12_2022_126 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 5 July 2022

Nanofibrous Scaffolds for the Management of Periodontal Diseases Alaa M. Mansour and Ibrahim M. El-Sherbiny

Contents 1 2 3 4

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Brief Overview of the Periodontium . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Periodontal Diseases . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Treatment Strategies . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1 Traditional Treatment Strategies . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.2 Regenerative Strategies . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 Periodontal Regeneration and Nanofibrous Biomaterials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.1 Nanofibrous-Occlusive Membranes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.2 Nanofibrous Scaffolds (Grafting Nanofibrous Biomaterials) . . . . . . . . . . . . . . . . . . . . . . . . 6 Implant-Related Nanofibrous Biomaterials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 Conclusions and Future Perspectives . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

362 362 363 364 364 365 366 366 380 395 396 397

Abstract Periodontium is an intricate complex system that consists of different types of tissues. There are several diseases that affect periodontium causing destruction and loss of its tissues. The goal for periodontal treatment is the reconstruction of the lost periodontal tissues. Periodontal regeneration is considered one of oral health care challenges, and it can depend on using nanostructured biomaterials. These nanostructured biomaterials simulate the microenvironment of the extracellular matrix (ECM) and act as a biomimetic platform to attract stem cells and stimulate their differentiation to specific lineages. There are different forms for nanostructured biomaterials such as nanofibers and nanoparticles. Nanofibers have a similar

A. M. Mansour Department of Oral Biology, Faculty of Dentistry, Mansoura University, Mansoura, Egypt e-mail: [email protected] I. M. El-Sherbiny (*) Nanomedicine Research Labs, Center for Materials Sciences (CMS), Zewail City of Science and Technology, Giza, Egypt e-mail: [email protected]

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structure and size to those of the natural collagen seen in the ECM of periodontal tissues. This chapter gives a brief overview of periodontium and periodontal diseases. Moreover, it discusses the different strategies for periodontal therapy including periodontal tissue regeneration and the recent nanofibrous biomaterials that can be used for periodontal regeneration. Keywords Guided tissue regeneration · Membrane · Nanofibrous · Periodontal and peri-implant diseases · Periodontium · Scaffold

1 Introduction Regenerating lost tissues is the main therapeutic goal in all medical fields. Tissue engineering is an interdisciplinary field that involves a biodegradable supporting matrix and bioactive component with cells to make a new tissue that replaces damaged tissue. In tissue engineering, biomaterials mimic the ECM and provide 3-D space for cells to attach, proliferate, and differentiate forming new tissues with adequate function and structure. Additionally, they can act as a carrier for cells or bioactive molecules [1–3]. Orofacial tissues are unique, especially in their development. For example, bones in orofacial region are derived from both paraxial mesoderm and neural crest cells, while skeletal bones are mesodermal in origin [4]. Moreover, orofacial tissues have limited regeneration capacity [5]. Periodontium (tooth-supporting tissues) is a complex structure and its regeneration requires the synergy of both cellular and molecular events [6]. The adequate understanding of periodontal biology combined with new advances in biomaterials will allow appropriate periodontal tissue reconstruction [7]. In this chapter, the periodontium, and its related disease as well as the treatment strategies of periodontal diseases, especially those depending on nanofibrous biomaterials, will be discussed. As the implant is an artificial tooth substitute and has related diseases similar to periodontal diseases, we will briefly discuss its related problems and how nanofibrous biomaterials are used to overcome these problems.

2 Brief Overview of the Periodontium Periodontium is an intricate system of different tissues encircling the tooth [8]. It consists of two hard tissues and two soft tissues. Hard tissues are alveolar bone (AB) and cementum (CM). The soft tissues are periodontal ligament (PDL) and gingiva [9] (Fig. 1). Each part of the periodontium has great importance in tooth stability and its anchorage to the jaw. AB is a dynamic mineralized connective tissue (CT) forming the socket which surrounds the root part of the tooth [10]. CM is a

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Fig. 1 Illustrations of healthy periodontium, periodontitis, and peri-implantitis. The healthy periodontium consisted of AB, PDL, CM, and gingiva (G). The gingiva is attached to the tooth through junctional epithelium (JE) as a part of the dentogingival junction. The periodontitis and peri-implantitis have accumulated plaque and calculus on the surface of the tooth or the artificial crown, osteoclastic activity leading to AB resorption that may extend to the furcation area (furcation involvement), and presence of large number of inflammatory cells

mineralized CT resembling the bone, except CM is not vascular and exhibits little turnover. It covers the roots of teeth and serves to anchor gingival and periodontal fibers. PDL is a fibrous CT connecting CM to the AB [11–15]. Gingiva is the soft tissue of epithelium and CT covering and protecting the underlying previous periodontal tissues. It has a unique structure called dentogingival junction [16]. The dentogingival junction exists at the interface between the gingiva and the tooth, [17] holds the gingiva to the teeth and protects the subjacent AB from microbes resident in the dental plaque [18, 19].

3 Periodontal Diseases The term “periodontal diseases” includes different forms of inflammatory conditions. Periodontal disease begins with gingivitis which is a reversible inflammatory condition restricted to the gingiva and initiated by bacteria from the dental plaque. When the gingivitis is untreated, the inflammation is progressed to other periodontal tissues causing irreversible tissue loss and called chronic periodontitis. Chronic periodontitis is one of the most widespread chronic inflammatory non-communicable diseases [20, 21]. Normally, bone undergoes a continuous remodeling based on the coordination between bone resorption and deposition that is performed by two important cell types: cells of the osteoblastic lineage and bone-resorbing cells [10]. However, in periodontitis, this balanced coordination is disrupted, leading to increase in the number and activity of osteoclasts [22, 23]. Loss of AB leads to periodontal pocket

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formation that acts as a reservoir of anaerobic bacteria such as Aggregatibacter actinomycetemcomitans, Porphyromonas gingivalis, and fusobacterium nucleatum. In addition, AB resorption affects gingival level causing G. recession. In multirooted teeth, this resorption may be extended to the furcation area, causing one of the challengeable periodontal conditions, which are furcation involvement (bone loss at the base of two or more roots meet) [24] (Fig. 1). There are acute forms of periodontal diseases called aggressive periodontitis and necrotizing ulcerative gingivitis and periodontitis. Aggressive periodontitis is mainly heritable condition and shows imbalance between anti-and pro-inflammatory cytokines [25]. The necrotizing ulcerative form is a rare and painful condition and typically presents in debilitated hosts. It is characterized by a rapid course. Its lesions consist of ulceration and necrosis of the interdental papilla with a whitish-yellow pseudomembrane surrounded by erythematous halo [26]. Periodontal diseases can negatively affect mastication, speaking, and aesthetics [27]. Periodontal diseases may also be related to systematic disorders [28–31]. They share many risk factors with other chronic inflammatory noncommunicable diseases, such as unhealthy diet, smoking, stress, genetic determinants or glycemic control [20, 32]. Under the COVID-19 pandemic, oral mucosa may have a role in the disease transmission and pathogenicity as oral mucosa and periodontal pocket cells have the virus receptors. This provides a way for the virus to spread inside the body [33, 34], as the periodontal pocket is considered a bidirectional linkage between the oral environment through crevicular fluid and blood circulation through gingival blood capillaries [35]. Recent research concluded that severe complications from COVID19 have been associated with patients with periodontitis [36]. This may be related to the increase in the inflammatory response with periodontitis, which might worsen clinical course of COVID-19 [37].

4 Treatment Strategies There are different strategies for periodontal treatment. All of them aim to (1) eliminate or even reduce the inflammation, (2) arrest disease progression, and (3) regenerate new periodontal tissues to maintain long-term healthy and functional dentition [9, 20]. There are several periodontal treatment modalities such as scaling and root planning, open flap debridement, guided regeneration, bone grafts, and tissue engineering approach [38].

4.1

Traditional Treatment Strategies

This strategy focuses on the removal of the source of inflammation that are plaque and calculus via scaling, root planning and surgical treatments. Gingivitis is a reversible condition, so mechanical removal of plaque and calculus with or without

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adjunctive systemic antibiotics can be effective. If the condition is extended to have AB loss, so it is an irreversible condition and mechanical debridement cannot remove all subgingival microorganisms. Additionally, this strategy cannot restore the lost tissues. Therefore, it is crucial to use alternative regenerative approaches [20, 39].

4.2

Regenerative Strategies

Regeneration of tooth-supporting structures is a challenging process that requires synergy of several cellular and molecular events. This may be related to the specific anatomical, morphological, and compositional characteristics of these complex tissues [40]. Besides, persistent periodontitis significantly decreases the regeneration ability of CM and PDL [41]. As a result of the complicated composition of periodontium that consisted of hard and soft tissues, scaffold-based regenerative approach is the main approach for PDL-CM-AB complex regeneration [39]. The ideal scaffold should have crucial requirements. It must be biocompatible without any mutagenic or cytotoxic effects [42] and be biodegradable and its degradation rate has to be coordinated with the rate of the new tissue formation [43]. Biodegradable periodontal biomaterials are very attractive choice clinically as they do not need another surgery for removing the biomaterial [44]. It also must possess adequate mechanical properties and appropriate porosity to enable cell attachment and proliferation and to facilitate nutrient exchange [45–47]. Other requirements should be taken under consideration during selecting the scaffold for periodontal regeneration and its method of fabrication. The synthesis of personalized biomaterials is to accommodate the periodontal pockets and defects that can vary in shape and size [48]. In addition, the scaffold needs to provide a structural guidance for proper regeneration [39] such as scaffold allows the formation of properly oriented PDL fibers extend from the newly formed AB to CM [40, 49, 50]. Moreover, the fibrous nature of the PDL with the nanoscale proteins suggests the selection of nanofibrous biomaterial. Moreover, nanofibrous scaffolds are better than solid scaffolds in promoting osteoblast differentiation [51]. For that reason, we suggest that nanofibrous biomaterials could be the proper scaffolds for periodontal regeneration [52]. Nanofibrous scaffolds consist of uninterrupted nanofibers with a short diffusional path and high porosity [53]. They can be fabricated by three fabrication techniques: phase separation, self-assembly, and electrospinning [54, 55]. Phase separation is a thermodynamic technique based on polymer separation into two phases: polymerpoor and polymer-rich phases. According to the method of separation, the phase separation can be divided into several types such as nonsolvent-induced phase separation and thermally induced phase separation (TIPS). The polymer-poor phase is then removed, leaving voids. Finally, the polymer-rich phases is lyophilized to form a porous nanofibrous scaffold [56]. The advantage of this technique is the ability to use molds with defined dimensions to fabricate custom 3-D scaffolds

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[57]. Self-assembly is a technique allowing the formation of 3-D structures from 1-D molecules. It organizes the components into well-defined aggregates. One of the main disadvantages of this technique is that the self-assembled nanofibrous scaffold usually has poor mechanical properties [58]. Electrospinning is a very desirable technique for fabricating nonwoven nanofibrous biomaterials that is nearly similar to ECM [59]. In this technique, the solution is forced through syringe’s nozzle producing a nonwoven mat of nanofibers under the influence of an electric field [55]. Although these electrospun nanofibers can be fabricated from different material types, polymers are the most commonly used type owing to their adequate physical properties [52]. The properties for electrospun materials can be controlled by adjusting the electrospinning processing parameters [60]. In support of the periodontal treatment, various nanofibrous biomaterials have been developed from a physical barrier to support periodontal healing to scaffold system to allow integrated periodontal regeneration [8]. According to that, there are two approaches for periodontal regenerative strategy: occlusive membrane and tissue engineering approaches. Therefore, we will cover the most recent nanofibrous biomaterials used for guided regeneration and tissue engineering employed for the last 6 years and their efficacy assessed in vitro and/or in vivo.

5 Periodontal Regeneration and Nanofibrous Biomaterials 5.1

Nanofibrous-Occlusive Membranes

The use of contact inhibition membrane at the interface between gingiva and PDL/AB tissues to prevent cells migration inside the periodontal defect and maintain the space for progenitor cells to recolonize and regenerate the periodontal tissue is called guided tissue regeneration (GTR) [9, 61]. GTR has been commonly used for periodontal regeneration in dental clinics. Its procedure consists of raising mucogingival flap, scaling and root planning, positioning the membrane under gingiva, then suturing the flap [39]. Many studies have been performed recently to develop a novel and more effective occlusive membrane, which have been summarized in Table 1. We generally divided those membranes into drug carrier, bioactive, and bilayered nanofibrous-occlusive membranes. It is worth mentioning that there is another application for guided regeneration over the deficient alveolar ridges or bone defects and so named guided bone regeneration (GBR) [6, 9].

5.1.1

Drug Carrier Occlusive Membranes

As chronic periodontitis is a persistent inflammatory condition, occlusive membranes with anti-inflammatory activity are preferred to alleviate the inflammation and stop disease progress and allow healing. This anti-inflammatory activity can be

Method of nanofibrous fabrication/parameters Electrospinning/voltage ¼25 kV, feed rate ¼ 2 ml/h, and syringe-collector distance ¼ 15 cm

Coaxial electrospinning/voltage ¼13.26 kV, syringe-collector distance ¼ 19 cm, and feed rate ¼ 0.5 ml/h

Occlusive membrane 1. PLA/CA with AgNPs and nHAP. Nanofibrous membrane 2. PCL with AgNPs and nHAPnanofibrous membrane

Nanofibrous PCL membrane with ibuprofen (IBU-PCL)

In vitro & in vivo

Type of study In vitro Morphological characteristics Nanofibers were interconnected randomly organized with smooth surface, most of the AgNPs and HANPs were embedded in the fibers with some of the nanoparticles distributed on the fiber surface The nanofibers were uniform in size and interconnected

Biodegradability/ hydrophilicity Biodegradable, but the addition of 10% HANPs resulted in a significant decrease in the degradation rate

–

Mechanical properties The addition of 10% nHAP enhanced the tensile strength while the addition of 20% nHAP decreased the tensile strength

–

Biocompatible/ the membrane reduced proliferation of Pg-LPSstimulated cells and decreased epithelial Cs migration

Cytotoxicity/cell proliferation and differentiation Biocompatible/ addition of nHAP and AgNPs enhanced the cell viability

In mouse periodontitis model, the membranes significantly enhanced the clinical attachment and decreased bone resorption

Other Sustained release of the loaded AgNPs till 35 days allowed antibacterial effect against E. faecalis and E. coli

Table 1 The occlusive membranes (GBR/GTR), methods of nanofibrous fabrication /parameters and type of the study and their characterization

(continued)

[63]

Reference [62]

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Electrospinning/voltage ¼15 kV, feed rate ¼ 2.5 ml/h, and syringe-collector distance ¼ 8–10 cm

Two-step electrospinning method • LL is formed by conjugated

Bilayered membrane of PLGA/ GEL/DEXloaded MSNs nanofibers

Method of nanofibrous fabrication/parameters Coaxial electrospinning/voltage ¼20–23 kV, feed rate ¼ 3 ml/h (shell) & 1.5 ml/h (core), & syringe-collector distance ¼ 18 cm Electrospinning/voltage ¼10.5 kV, feed rate ¼ 1 ml/h, & syringe-collector distance ¼ 14.5 cm

PCL nanofibers with PDA surface coating

(PCL-GELNBG) and their biomimetic transformation into HAs

Occlusive membrane PCL /GEL crosslinked with metronidazole (core/ sheath) nanofiber membranes

Table 1 (continued)

In vitro

In vitro & in vivo

In vitro

Type of study In vitro

Nanofibers with around 96% porosity and contained NBG agglomerates within their internal structure The membrane consisted of uniform and bead-free nanofibers with interconnected porous structures Both layers had electrospun nanofibers with interconnected pores

Morphological characteristics The core/sheath structure was observed using transmission electron microscope

Adequate mechanical properties. DCH increased the

The membrane had adequate mechanical properties that enhanced by the presence of the HA spherules PDA coating enhanced the mechanical properties

Mechanical properties The membrane had adequate tensile strength

Biodegradable/ LL is more hydrophilic than DL

Faster biodegradability than uncoated

Biodegradable/ the addition of GEL and NBG enhanced the hydrophilicity

Biodegradability/ hydrophilicity Biodegradable/ GEL coating improved hydrophilicity

Biocompatible/ the membrane allowed the attachment and proliferation of

Biocompatible/ the presence of PDA layer promoted osteogenic differentiation of PDLSCs

Biocompatible

Cytotoxicity/cell proliferation and differentiation Biocompatible/ GEL coating on the PCL fibers enhanced L929 cells adhesion and proliferation

Effective antibacterial potency against both

[67]

[66]

[65]

–

The membrane enhanced periodontal tissues regeneration

Reference [64]

Other Antibacterial effect

368 A. M. Mansour and I. M. El-Sherbiny

Bilayered nanofibrous membrane of PCL/PGS/CS (GTR layer) & PGS/PCL/β-TCP (GBR layer)

PLA/GEL nanofibrous with nMgO

(LL) & PLGA/ DCH (DL)

Two-step electrospinning method/ • GBR layer (voltage ¼18 kV, flow rate ¼ 1 ml/h, distance of 16 cm) collected on aluminum foil • GTR layer (flow rate ¼ 0.3 ml/h,

electrospinning method (voltages ¼ 10 kV, rate ¼ 1.0 ml/h) • DL by traditional electrospinning technique over prefabricated LL (voltage of 10 kV, rate ¼ 2 ml/h) Electrospinning/voltage ¼12 kV, feed rate ¼ 3 ml/h, syringecollector distance ¼ 8–10 cm The membrane composed of randomly oriented and densely packed nanofibers with inter-fiber space less than 10 μm

Nanofibers were smooth and homogeneous in both layers with β-TCP clusters appeared inside the nanofibers of GBR layer • Porosity

In vitro and in vivo

In vitro

LL had higher number of pores than DL

Young’s modulus and tensile strength increased with the increase in β-TCP amount till 10% wt but decreased when β-TCP amount was

The membrane had adequate tensile strength

brittleness of DL, while DEX@MSNs decreased the modulus of elasticity of LL

Biodegradable with degradation rate match the rate of periodontal regeneration/ addition of 2% nMgO showed less water contact angle than the membranes with different amount of nMgO Biodegradable/ GTR layer is more hydrophilic than GBR layer, which could be attributed to the presence of amine groups in the CS structure Biocompatible/ excess amount of β-TCP (15%) was not suitable for cell proliferation due to the agglomerations of β-TCP within the structure of nanofibers

Biocompatible/ nMgO promotes the adhesion, proliferation, and osteogenesis differentiation of BMSCs in a dosedependent manner but not for the fibroblasts

L929 cells and BMSCs. MSNs enhanced the osteoinductive capacity for BMSCs

[69]

–

(continued)

[68]

The membrane had antibacterial activity

S. aureus and E. coli

Nanofibrous Scaffolds for the Management of Periodontal Diseases 369

In vitro

In vitro

Electrospinning/voltage ¼ 16 kV, syringecollector

PCL/GEL nanofibrous loaded with CeO2 NPs

Type of study

Electrospinning/voltage ¼ 20 kV, and syringe-collector distance ¼12 cm

voltage ¼ 12 kV, distance ¼ 12 cm) collected on the GBR layer

Method of nanofibrous fabrication/parameters

Bilayered composite membrane of PA6CS nanofibers reinforced with nHA/PA6

Occlusive membrane

Table 1 (continued)

• PA6/CS layer has relatively uniform and smooth nanofibers • n-HA/ PA6 layer has a microporous and mesoporous structure with uniform pore distribution Smooth fibrous structure with black particles inside

of both layers was found in the range of 60–80%

Morphological characteristics

CeO2 NPs enhanced the mechanical properties

15% due to β-TCP nanoparticles agglomeration within the nanofibers Adequate mechanical properties

Mechanical properties

Biodegradable/ GEL enhanced the hydrophilicity

Biodegradable

Biodegradability/ hydrophilicity

Biocompatible/ the membrane promoted CCK-8 proliferation &

Biocompatible/ the membrane supported cells attachment and proliferation on both layers with osteoconductive effect

Cytotoxicity/cell proliferation and differentiation

The membrane had antioxidant capacity

–

Other

[71]

[70]

Reference

370 A. M. Mansour and I. M. El-Sherbiny

In vitro

Electrospinning/voltage ¼15 kV, syringe-

Nanofibers were smooth,

The membrane consisted of smooth and uniform nanofibers

In vitro

Electrospinning/ voltage ¼ 20–29 kV, cm & flow rate ¼ 0.1–0.8 ml/h

Hyaluronic acid derivatives nanofibrous membrane loaded with DEX

Both layers were integrated, and a nanofiber layer act as a barrier over the porous sublayer

Nanofibrous layer was fabricated by electrospinning/voltage ¼ 20 kV, syringecollector distance ¼ 10 cm & flow rate ¼ 3 ml/h

Bilayered membrane of CS/PEO (nanofibrous layer) & CS and Si-doped nHA (sublayer)

In vitro & in vivo

indicating the successful encapsulation of CeO2 NPs Both membranes showed nanofibrous structure after the postelectrospinning treatments

In vitro

Electrospinning/voltage ¼ 26 kV, flow rate ¼ 1 ml/h, and syringe-collector distance ¼ 15 cm

1. Chitosan membranes treated with TEA/tBOC 2. Chitosan membranes treated with BA

distance ¼ 15 cm & flow rate ¼ 1 ml/h

Biodegradable/ the incorporation of Dex, and the cross-linking reduced the membrane hydrophobicity Biodegradable with appropriate

–

Adequate mechanical

Biodegradable/ Si-nHAP increased the hydrophilicity

–

Si-nHap incorporation enhanced the physical and mechanical properties

–

Biocompatible/ Si-nHAP enhanced Saos-2 cells attachment, spreading and biomineralization on the membrane surface Biocompatible/ the controlled release of Dex stimulated osteogenesis of MC3T3 cells

Biocompatible

osteogenesis differentiation of PDLSCs

The membrane has

–

Both membranes enhanced bone regeneration in calvarial defects more than collagen membrane Increasing Si concentration enhanced the antimicrobial activity

(continued)

[75]

[74]

[73]

[72]

Nanofibrous Scaffolds for the Management of Periodontal Diseases 371

collector distance ¼ 15 cm and feed rate ¼ 2 ml/h

Electrospinning/voltage ¼16 kV, feed rate ¼ 0.6 ml/h, and syringe-collector distance ¼ 13–18 cm

Electrospinning/voltage ¼15 kV, feed rate ¼ 0.5 ml/h, and syringe-collector distance ¼ 15 cm

Electrospinning/voltage ¼15 kV, syringe-

Silver-modified/ col-coated PLGA/PCL electrospun membrane

Poly lactic acid (PLA)/calcium alginate nanofibrous membrane

Gelatin nanofibrous

Method of nanofibrous fabrication/parameters

Occlusive membrane CS/PCL/GEL nanofibrous membrane

Table 1 (continued)

In vitro

In vitro

In vitro

& in vivo

Type of study

Nanofibers were continuous with beadfree and uniform diameters that remain intact after soaking in water Nanofibers have rough

Nanofibers were randomly oriented with clearly distinguished interfiber spaces

dense with random orientation

Morphological characteristics

Biodegradable/ addition of calcium alginate reduced contact angles

Biodegradable

–

Biodegradable/ Ag impregnation and collagencoating enhanced hydrophilic properties

degradation rates which more closely resembled tissue regeneration process/GEL enhanced hydrophilicity

Biodegradability/ hydrophilicity

Addition of calcium alginate to PLA enhanced the mechanical properties

Silver impregnation and col coating did not cause affect the modulus of elasticity and the tensile strength

properties (increasing with more PCL)

Mechanical properties

Biocompatible

Biocompatible/ Ag impregnation and collagencoating enhanced MC3T3 cells attachment, proliferation, and osteogenic differentiation Biocompatible/ addition of calcium alginate enhanced BMSCs attachment, proliferation, and osteogenic differentiation

Cytotoxicity/cell proliferation and differentiation Biocompatible/ cells viability was enhanced

The scaffold suppressed

–

hemostasis effect After subcutaneous implantation, the scaffolds showed adequate cell barrier effects Silver ions exerted antibacterial activity against S. aureus and S. mutans

Other

[78]

[77]

[76]

Reference

372 A. M. Mansour and I. M. El-Sherbiny

collector distance ¼ 14 cm and feed rate ¼ 0.5 ml/h

Coaxial electrospinning/voltage ¼18 kV, syringecollector distance ¼ 20 cm, and feed rate ¼ 0.1 ml/h (core) and 0.8–1.6 ml/ h (shell)

Electrospinning/voltage ¼25–30 kV,

membrane loaded with SP600125 and SB203580 polymeric micelles

Nanofibrous membrane with dual drug-loaded core-shell structure (incorporating BMP-2 in the core and polymeric micelles containing SP600125 in the shell)

Aloe Vera containing PCL

In vitro

In vitro and in vivo

and in vivo

The membrane consisted of

Nanofibers were uniform with smooth surface

surfaces with obvious spindle-knots with uniformly dispersed micelles

The tensile properties are

–

Biodegradable/ the

Biodegradable

Biocompatible/ when LPS-induced PDLSCs co-cultured with the membrane, they expressed less inflammatory cytokines. The membrane had osteoinductive capacity on LPS-induced BMSCs Biocompatible/ the addition of

–

both MMP-2 and MMP-13 expression, controlled the inflammation of bone and promoted the periodontal regeneration in furcation defects In vivo, the membrane allowed more AB formation in acute alveolar bone defect

(continued)

[80]

[79]

Nanofibrous Scaffolds for the Management of Periodontal Diseases 373

Method of nanofibrous fabrication/parameters

syringe-collector distance ¼ 16–23 cm and feed rate ¼ 0.5 ml/ h

Electrospinning/ voltage ¼ 16–18 kV, syringe-collector distance ¼ 10–15 cm, and feed rate ¼ 1 ml/h

Occlusive membrane

nanofibrous membrane

Fish Col/bioactive glass (BG) /CS composite nanofiber membrane

Table 1 (continued)

In vitro and in vivo

Type of study

Nanofibers were smooth and bead-free

bead-free nanofibers with 85 to 74% porosity

Morphological characteristics

The membrane had certain mechanical strength

increased at first and then decreased

Mechanical properties hydrophilicity of nanofiber membranes increased with the incorporation of aloe vera Biodegradable/ the membrane is highly hydrophilic

Biodegradability/ hydrophilicity aloe vera at 70/30 ratio provided adequate support for 3T3 cell growth and proliferation Biocompatible/ the membrane enhanced the PDLSCs proliferation and osteogenic differentiation

Cytotoxicity/cell proliferation and differentiation

The membrane had antibacterial activity and promoted AB regeneration in dogs’ furcation defect with higher expression of osteocalcin

Other

[81]

Reference

374 A. M. Mansour and I. M. El-Sherbiny

Nanofibrous Scaffolds for the Management of Periodontal Diseases

375

gained from the polymers used for membrane fabrication such as using hyaluronic acid derivatives (HA) electrospun membranes. HA is a natural polymer with the bacteriostatic, anti-inflammatory, and osteoinductive activity [74, 82]. Other studies developed anti-inflammatory nanofibrous scaffolds by the addition of antiinflammatory drugs. The most used anti-inflammatory drugs were glucocorticoids [67, 74, 83] and non-steroidal anti-inflammatory drugs, e.g., meloxicam [84], aspirin [85], and ibuprofen [63]. When Ibuprofen was encapsulated within the polycaprolactone (PCL), it alleviated P.gingivalis-induced inflammation. In addition, the membrane showed a burst release effect for ibuprofen that decreased epithelial and CT cells proliferation and epithelial cells migration which is one of the main problems in periodontal healing, but it has no improvement on in vivo bone formation [63]. On the contrary, the addition of suitable amounts of dexamethasone (DEX) in nanofibrous scaffolds has an osteoinductive effect enhancing the healing capacity of periodontal tissue, as DEX is anti-inflammatory and induces osteogenesis-related genes expression [74]. The main cause of periodontal diseases is bacteria. For that, there are many studies concerned with developing multifunctional occlusive membranes with antibacterial activity either by the addition of antibacterial drugs or by the addition of nanoparticles with antibacterial effect. The addition of the drugs to electrospun nanofibrous scaffolds can be done by blending or coaxial electrospinning. The coaxial electrospinning leads to the formation of core/shell structure. The nanofibrous membranes with core/shell structure have a sustained release of the drugs, which lead to enhanced antibacterial and anti-inflammatory effects [86] (Fig. 2). He et al. [64] fabricated metronidazole-loaded PCL/gelatin (GEL) core/ sheath nanofiber. They used GEL coating to enhance the hydrophilicity of PCL and in turn improved the cell attachment and proliferation. As the GEL shell can be easily dissolved in water, cross-linker genipin can be used to crosslink the gelatin shell. This cross-linking did not cause a significant change in the nanofiber’s morphology nor the pore size. Drug release profiles of metronidazole content were characterized by an initial burst release that followed by sustained release. This allowed effective antibacterial activity against fusobacterium nucleatum [64]. Matrix metalloproteinases (MMPs) are proteolytic enzymes which considered the major cause for human periodontal tissues destruction especially MMP-13. They can cleave gelatin, elastin, fibronectin, collagen I, III, and IV [87–89]. The p38 MAP kinase and c-Jun N-terminal kinase inhibitors, such as SB203580and SP600125, have inhibition effects on specific types of MMPs. However, these drugs are insoluble in water, so they have impaired activity in moisture environments such as oral cavity. Micelles with hydrophobic core can improve the bioavailability of SP600125 and SB203580 by increasing their solubility [90]. Additionally, micelles are small enough to penetrate AB trabeculae, CT, and even the periodontal pocket. Micelles incorporation into GEL nanofibers permitted the sustained release of the drugs. Its application in class II furcation defects enhanced the periodontal tissue regeneration, including CM, AB, and well-oriented PDL fibers [78]. The combination of both coaxial electrospinning and polymeric micelles is a novel design made by Liu et al. [79]. They fabricated core/shell nanofiber membrane

376

A. M. Mansour and I. M. El-Sherbiny

Fig. 2 Schematic illustration of different strategies in electrospinning: (a) Post-electrospinning adsorption, (b) Blending before electrospinning, and (c) Coaxial electrospinning

with polymeric micelles containing SP600125 incorporated in the shell and BMP-2 inserted in the core. SP600125 release was detected firstly, whereas BMP-2 was released after 12 days. This time-programmed release allowed the suppression of pro-inflammatory factors, then enhancing osteogenesis. Using this nanofiber membrane with periodontitis inhibited the AB destruction, and bone defects healed within 2 months [79].

5.1.2

Bioactive Nanoparticles-Containing Occlusive Membranes

Bioactive nanoparticles are considered as attractive biomaterials in several medical applications. Nanoparticles can provide enhanced osteoinduction and osteoconduction [53, 91]. There are several bioactive nanoparticles that have been used in periodontal and bone defects healing such as gold nanoparticle (Au NP) [92], silver nanoparticles (AgNPs) [93], cerium oxide nanoparticles (CeO2 NPs) [94],

Nanofibrous Scaffolds for the Management of Periodontal Diseases

377

bioactive glass nanoparticles [95], magnesium oxide nanoparticle (nMgO) [68], graphene oxide (GO) [96] and nanohydroxyapatite (nHAP) [73, 97]. Several studies added these nanoparticles to nanofibrous scaffolds to have the advantages of both [65, 71]. Silver nanoparticles (AgNPs) have antifungal, antiviral, and antibacterial activities with a low potential to generate antibacterial resistance [93]. Several studies have been used AgNPs with nanofibrous scaffold either mixed with the polymer before electrospinning [62] or introduced into the nanofibrous scaffold after electrospinning [76]. Both methods confirmed the antibacterial action of these nanofibrous scaffolds against periodontal pathogens. Abdelaziz et al. [62] have incorporated AgNPs and nHAP in polylactic acid/cellulose acetate (PLA/CA) and in polycaprolactone (PCL) to add antibacterial activity and to enhance mechanical properties and bone regeneration activity for these polymers. Concern to antibacterial action, they concluded that the formula of 10% nHAP and 2% AgNPs has the highest release rate of silver ions [62]. Qian et al. [76] impregnated AgNPs on poly-lactic-co-glycolic acid (PLGA)/PCL electrospun membrane via in situ reduction after coating the fibers with polydopamine (PDA), and finally, the membrane coated by collagen I (Col-I). The Col coating caused sustained release of silver ions and in turn prolonged antibacterial properties. In periodontitis model, the membrane was effective and improved AB regeneration [76]. Magnesium oxide nanoparticle is a bioabsorbable light metal-based nanoparticle. It has also antibacterial effect via destroying the bacteria’s structural integrity. When nMgO-incorporated into PLA/GEL nanofibrous membranes, it induced bone marrow mesenchymal stem cells (BMSCs) differentiation to osteoblasts that confirmed by increasing alkaline phosphatase (ALP), Runx 2, Col-I, and osteopontin (OPN) expression. This effect was confirmed also by in vivo application of the membrane in surgical periodontal defects. This membrane promoted osteogenesis, and the regenerated AB showed higher densities of capillaries and osteoblasts [68]. El-Fiqi et al. [65] incoroporated nanobioglass (NBG) agglomerates in PCL/GEL (10 wt% PCL-10 wt% GEL/25 wt% of NBG) and then used electrospinning technique to form nanofibrous membrane (Fig. 3). After membrane fabrication, they made a transformation of PCL/GEL-NBG to PCL/GEL-hydroxyapatites (HAs) through biomimetic mineralization process using Kokubo’s SBF (Fig. 3). This led to growth of HAs and production of nano/micro-structured fibrous membrane. This membrane has several advantages, including nano/micro-topography, large specific surface area, calcium/phosphate ratio similar to bone, high protein adsorption capacity, and sustained release of beneficial ions (Ca2+, PO43, and SiO44) [65]. These ions are so important for bone regeneration as Ca2+ promotes osteoinduction and regulates osteoblastic functions [98, 99]. Additionally, PO43 regulates the differentiation of osteoblasts, while SiO44 stimulates both osteogenesis and angiogenesis [65]. Likely, the addition of CeO2 NPs to electrospun PCL/GEL nanofibrous membrane allowed periodontal ligament stem cells (PDLSCs) attachment and osteogenic differentiation. This effect was confirmed by in vivo implantation of the membrane in rat cranial defect as the defects were almost reconstructed with mature bone after 8 weeks [71].

378

A. M. Mansour and I. M. El-Sherbiny

Fig. 3 Electron microscope images: SEM for PCL/GEL-NBG membrane with NBG agglomerates along the nanofibers (a), transmission electron microscope image for a single nanofiber showing NBG agglomerates within the nanofiber (b), and SEM images with low (c) and high magnification (d) for PCL-GEL-HAs membrane with hydroxyapatite spherules [65]. Figures are reprinted with permission from Elsevier

5.1.3

Bilayered/Multilayered Occlusive Membranes

The bilayered membrane was designed basically to combine the advantages of different biomaterials or different fabrication techniques. This combination results in formation of multifunctional membranes [70, 75]. Niu et al. [70] combined electrospinning and solvent casting and evaporation to form polyamide-6(PA6)/ chitosan(CS) @ nHA/PA6 bilayered membrane. Firstly, the PA6/CS nanofibrous layer was fabricated by electrospinning, which had adequate mechanical properties. Then, porous nHA/PA6 layer was fabricated over the prepared PA6/CS layer by a combined method of solvent casting and evaporation. The nHA/PA6 layer was biocompatible and had osteoconductive effect [70]. Zhang et al. prepared a multilayered membrane of PCL/GEL electrospun layers glued together by CS. This formed a strong cell barrier membrane [75].

Nanofibrous Scaffolds for the Management of Periodontal Diseases

379

Another interesting bilayered membrane has been developed by Masoudi Rad et al. [69], Lian et al. [67] and Tamburaci et al. [73]. They designed bilayered membrane for periodontal application, whereas one side of membrane was AB-related side acts as osteoconductive layer, and the other side of membrane related to the gingiva acts as barrier layer. They fabricated the bilayered membrane through the two-step electrospinning method. Masoudi Rad et al. added β tri-calcium phosphate (β-TCP) to osteoconductive layer, while they added CS to the barrier layer to decrease fiber diameter, porosity, and contact angle. The porosity is crucial in the occlusive membrane to prevent fibroblasts migration into the periodontal defect as fibroblasts migration leads to excessive soft tissue formation that subsequently hinders hard tissue regeneration [69]. While Lian et al. [67] designed the bilayered membrane to be consisted of loose layer (LL) composed of electrospun PLGA/GEL nanofibers with DEX-loaded mesoporous silica nanoparticles (MSNs), and dense layer (DL) composed of PLGA electrospun nanofibers with the doxycycline hyclate (DCH). DCH is a broad-spectrum antibiotic. The LL had a loosely packed and porous structure to allow cells attachment and infiltration, while the DL has a denser structure to act as a barrier. Both DCH and DEX showed sustained release from the bilayered membrane, which enhanced the antimicrobial activity for the membrane. Additionally, The osteoinductive activity of DEX had been verified by increased ALP and osteocalcin (OCN) expression [67]. Depending on this concept, Tamburaci et al. [73] fabricated bilayered nanocomposite membrane. This membrane composed of CS/PEO (90:10) nanofiber upper layer with microporous sublayer of CS and Si-doped nanohydroxyapatite particles (Si-nHAP). Nanofibrous layers were fabricated by electrospinning technique, while microporous sublayer was fabricated by freeze-drying technique. The microporous layer was considered the osteoconductive layer, and the nanofibrous layer was considered the barrier layer. They found that Si-nHAP enhanced osteoblast attachment due to their hydrophilic property and surface roughness formed by spherical nanopatterns. Besides, they found that increasing Si concentration improved the osteogenic activity and antimicrobial activity against both E. coli and S. aureus [73]. The main problem of using CS is the TFA salts formed during the electrospinning process. The TFA salt is a toxic hydrophilic salt that leads to swelling and loss of nanofibrous structure in moisture conditions. Eliminating the TFA salts is needed to maintain the nanofibrous architecture in CS membrane. For that, electrospun CS membranes can be exposed to post-spinning processes using either butyrylanhydride (BA) or triethylamine/tert-butyloxycarbonyl (TEA/tBOC) modifications to remove the TFA salts. In calvarial defects, both BA treated and TEA/tBOC treated CS membranes group showed greater density for the newly formed bone without any inflammatory responses compared to the collagen membrane group [72]. Despite the GTR positive outcome, its regenerative effect is affected by many factors, such as smoking, diabetes, and tooth morphology [100, 101]. To overcome that, tissue engineering approach has been applied using biomaterial scaffolds with cells and/or bioactive factors in the periodontal defect either alone or with GTR/GBR [8].

380

5.2

A. M. Mansour and I. M. El-Sherbiny

Nanofibrous Scaffolds (Grafting Nanofibrous Biomaterials)

Periodontal tissue engineering strategy uses stem cells, scaffold, and bioactive molecules to create biomimetic structures for periodontal tissues formation. There are two main approaches for periodontal tissue engineering. The first one is by using acellular scaffolds with or without growth factors (GFs), so the cells within the body ingress into the implanted scaffold and form the new tissue throughout the scaffold matrices. The second approach depends on using scaffolds seeded with cells in vitro. These two approaches are not mutually exclusive and can be easily combined [2, 102]. The focus in most of studies is to regenerate AB, but soft tissue regeneration is also needed. Periodontal regeneration is considered successful, when there are observable cementogenesis, inserted PDL fibers, and vital newly formed AB [20]. There are several important aspects to design a scaffold for periodontal regeneration, such as architecture, composition, structure, and ease of use (injectability) [39]. The nanofibrous scaffolds developed in the last 6 years have been tabulated in Table 2.

5.2.1

Composite Nanofibrous Scaffold

The traditional biomaterials can be combined together to form composite biomaterials [39]. As periodontium is composed of both hard and soft tissues, scaffold for each tissue type needs specific composition and design. The scaffold for AB regeneration should stimulate osteogenesis and mineralized tissue formation, whereas the scaffold for PDL regeneration should stimulate soft tissue formation and inhibit mineralization. Polymeric biomaterials are suitable for PDL regeneration, while biomaterials containing inorganic components, e.g., calcium phosphate (CaP) and hydroxyapatite are suitable for hard tissue regeneration. For regenerating CMPDL-AB complex, composite biomaterials of both polymers and inorganic components can be appropriate choice [39, 50, 126]. Zein and GEL are biocompatible and biodegradable natural polymers [127– 129]. Electrospun zein/GEL membrane enhanced PDLSCs attachment and proliferation with weak ALP activity, so it is necessary to add osteoinductive component into the zein/GEL nanofibers [130]. The addition of nHAP to form zein/GEL/nHAp nanofibrous composite scaffold allowed osteogenic differentiation of PDLSCs and improved new bone formation when implanted in cranial defects [116]. PLA also is a biodegradable and biocompatible synthetic polymer that is commonly used in tissue engineering [131]. To enhance its mechanical and biological properties, several studies made a modification to be more appropriate for periodontal regeneration. Shen et al. [119] added an appropriate amount of CS nanoparticles that promote BMSCs attachment, proliferation, and osteogenic differentiation [119]. Hwang et al. [107] presented a new strategy to fabricate highly porous and fluffy PLA nanofibrous scaffold. This scaffold was formed of PLA and lactic acid (LA) and was fabricated

In vitro

In vitro

Method of nanofibrous fabrication/parameters

Electrospinning/the flow rate ¼ 1–1.2 ml/h, voltage ¼ 15–17 kV and syringe-collector distance ¼10–12 cm

Electrospinning/the flow rate ¼ 3 cm3/h, voltage ¼ 15 kV and syringecollector distance ¼10 cm

Nanofibrous scaffolds

PCL-PEG-PCL/ zeolite nanofibrous scaffolds

Metronidazole containing PLA-based, nanofibrous scaffolds (mats and disks)

Type of study Nanofibers were smooth and non-beaded. The incorporation of zeolite nanoparticles increased the average diameter of fibers • The mat consisted of loose nanofibers with considerable interconnected pores • The disk had more compact structure with smaller pores • Metronidazole crystals existed among the fibers and were more obvious in the disk

Morphological properties –

Biodegradable/ the disk was more hydrophilic than the mat

–

Biodegradation/ hydrophilicity

–

Mechanical properties

–

Biocompatible/ DPSCs attached and uniformly dispersed on the scaffold surface with the production of long cytoplasmic processes

Cytotoxicity

• Disk released the metronidazole much faster than the mat • Both of disk and mat had efficient antibacterial effect against F. nucleatum & P. intermedia

Zeolite enhanced osteogenic differentiation of DPSCs

Other

Table 2 The tissue engineering biomaterials, method of nanofibrous fabrication /parameters and the type of the study and their characterization

Reference

(continued)

[104]

[103].

Nanofibrous Scaffolds for the Management of Periodontal Diseases 381

Electrospinning then leaching of LA

In vitro

In vitro & in vivo

–

Fluffy-type low-density PLA

In vitro

Electrospinning/the flow rate ¼ 0.794 ml/h, voltage ¼ 28 kV and syringecollector distance ¼15 cm

1. GCL (GEL-low molecular weight CS) 2. GCH (GEL-high molecular weight CS) 3. GA (GEL-alginic acid sodium salt) PLLA spongy nanofibrous microscaffold with immobilized PLGA-MSNspeptide as a platform for BMP-2 and celecoxib

Type of study

Method of nanofibrous fabrication/parameters

Nanofibrous scaffolds

Table 2 (continued)

• The PLGA shell is porous microspheres and the PLGA microsphere contained MSN clusters • PLLA scaffold was spongy and had nanofibrous morphology The nanofibers had hollow and spongy

Nanofibers were uniform and smooth. The nanofibrous structure was preserved even after hydration

Morphological properties

Biodegradable

Biodegradable

–

Biodegradable/ all the scaffolds had low contact angles, indicating adequate hydrophilicity

Biodegradation/ hydrophilicity

–

Adequate mechanical properties

Mechanical properties Cytotoxicity

Biocompatible/ spongy architecture allowed the ingress of the cells within

Biocompatible/the scaffold enhanced the gene expression of ALP, BSP and RUNX2 and promoted the mineralization of BMSCs even under inflammatory condition

Biocompatible/all the scaffolds supported the growth of several cell types

Other

–

The scaffold enhanced bone regeneration in periodontitis model, that was similar to normal AB

All of the scaffolds had antimicrobial properties against Aggregatibacter actinomycetemcomitans and Streptococcus mutans, especially CGL

Reference

[107]

[106]

[105]

382 A. M. Mansour and I. M. El-Sherbiny

In vitro and in vivo

Electrospinning/the flow rate ¼ 0.05 ml/h, voltage ¼ 20 kV, and syringecollector distance ¼10 cm

Self-assembly

PVA/HA/CNT nanofibrous scaffold

PLLA NF-SMS loaded with PLGA MS and MSN

In vitro and in vivo

In vitro and in vivo

Electrospinning/the flow rate ¼ 1 ml/h, voltage ¼ 21  0.5 kV, and syringe-collector distance ¼17 cm (3-D scaffold was formed of 30 layers of nanofibers inserted in CS)

1. PCL–PEG nanofibers (aligned or random) embedded within CS-based scaffold 2. PCL–PEG nanofibrous scaffold (2D)

nanofibrous scaffold

Nanofibers were about 160 nm

Most fibers in 2D scaffold were aligned nearly in a single direction, while 3D scaffold showed concentrated distribution of nanofiber and some nanofibers appeared curved Homogenous cross-linking structure with some beads projecting from the surface

structure with nano pore

Mechanical strength of HA/CNT was improved compared to HA, which attributed to the nucleation of HA around the tubes –

–

Biocompatible

Biocompatible/the scaffold suppressed osteoclastogenesis and enhanced osteoblalastogenesis

Biodegradable

–

Biodegradable

the scaffold and accelerated cell proliferation Biocompatible/the viability of BMSCs on 3D scaffolds was significantly higher

In periodontitis model, the scaffold enhanced Treg-mediated immunity against the AB loss

It induced angiogenesis in a chick chorioallantoic membrane assay

The 3-D aligned scaffold showed more aligned and mature PDL fibers than 3-D random scaffold in periodontal defect

(continued)

[110]

[109]

[108]

Nanofibrous Scaffolds for the Management of Periodontal Diseases 383

In vitro and in vivo

Electrospinning/the flow rate ¼ 1 ml/h, voltage ¼ 12–15 kV, and syringe-collector distance ¼ 13–15 cm. For aligned nanofibers, the collector rotated at 800 rpm

Electrospinning for PCL/voltage ¼ 15–20 kV, and flow rate ¼ 2.5 ml/h

HMGB1 immobilized on aligned PLLA/ PCL nanofibers

CS/pDNA nanocomplexes immobilized to PCL nanofibrous scaffold

In vitro

In vitro and in vivo

Electrospinning/the flow rate ¼ 1 ml/h, voltage ¼ 11.5 Kv, and syringe-collector distance ¼15 cm. For oriented nanofibers, the speed of the rotating collector was 900 rpm

Composite PCL/GEL oriented nanofibrous scaffold

Type of study

Method of nanofibrous fabrication/parameters

Nanofibrous scaffolds

Table 2 (continued)

The scaffold had nanofibrous structure with rough surface due to

The nanofibers were basically aligned and had interconnected pores

Most of the fibers had the same direction, but some of them were randomly oriented due to processing limitation

Morphological properties The modulus of elasticity was higher in the oriented nanofibrous scaffold, while the maximum stress was lower than the random nanofibrous scaffold Combination between PCL and PLLA increased the modulus of elasticity and the ultimate tensile strength Adequate mechanical properties

Mechanical properties

Biodegradable

Biodegradable

Biodegradable

Biodegradation/ hydrophilicity Cytotoxicity

Biocompatible/ incorporation of MMP-sensitive peptide increased NCs release, which in turn directed

Biocompatible/ MSCs distributed uniformly throughout the aligned and the cells aligned like osteoblasts in native bone tissue

Biocompatible/ PDLCs aligned parallel to the nanofibers’ direction with increased PDL markers

Other

Reference

[112]

[113]

–

[111]

The scaffold enhanced vascularization, induced osteogenesis in a critical calvarial defect

With concomitant cyclic loading, PDLCs arranged parallel to the nanofibers with increased ALP activity

384 A. M. Mansour and I. M. El-Sherbiny

In vitro & in vivo

Nanofibrous PLLA was fabricated by TIPS and porogen leaching techniques

Electrospinning/voltage ¼ 20 kV and flow rate ¼ 0.6 ml/h

Electrospinning/for the blended scaffold: voltage ¼ 15 kV, flow rate ¼ 1.0 ml/h For core/shell scaffold: the syringes attached to a coaxial nuzzle, voltage ¼15 kV, & flow rate ¼ 1.0 ml/h and 0.2 ml/

Zein/GEL/nHA nanofibrous scaffolds

PLGA/gum tragacanth /tetracycline hydrochloride nanofibrous scaffolds (blended & core-shell structure)

In vitro

In vitro and in vivo

In vitro & in vivo

Electrospinning/the flow rate ¼ 3.5 m L/h, voltage ¼ 12–26 kV and syringe-collector distance ¼ 10 cm

PCL nanofibrous scaffold coated with crosslinked alginate and incorporating nHAP PLLA nanofibers /CaP composite scaffolds

Highly porous nanofibrous networks with relatively rough surfaces related to nHAP Smooth and non-beaded nanofibers with the pore size of the mat was decreased by addition of gum tragacanth and

Nanofibrous structure with flower-like structure

immobilized NCs The scaffold depicted a fine nanofibrous structure with interconnected pores

Incorporation of tetracycline decreased the tensile strength and elongation, while the elastic

nHAP enhanced mechanical properties

Coating PCL with alginate and nHAP cause enhanced mechanical properties –

Biodegradable t/nHAP enhanced hydrophilicity and slowed down the degradation rate Biodegradable/ tetracycline enhanced the hydrophilicity of the scaffold

Biodegradable

Biodegradable/ the coating increased the overall hydrophilicity

Biocompatible

Biocompatible/calcium release from the CaP enhanced osteogenic cells proliferation with optimal calcium concentration was about 70 μg/ml Biocompatible/ nHAp enhanced the attachment, proliferation, and osteogenic differentiation of PDLSCs

MSCs toward osteogenesis Biocompatible/the coating enhanced ADSCs, attachment and proliferation

The scaffold had antibacterial activity against S. aureus and Pseudomonas aeruginosa, but Pseudomonas aeruginosa has smaller inhibition ring

[115]

• There was sustained release of calcium over 3 weeks • Scaffolds seeded with BMSCs enhanced bone regeneration after subcutaneous implantation In rat cranial defect model, the scaffold enhanced osteogenesis, especially when seeded with PDLSCs

(continued)

[86]

[116]

[114]

The scaffold enhanced PDL regeneration in class II furcation defects

Nanofibrous Scaffolds for the Management of Periodontal Diseases 385

Emulsion electrospinning/ syringe-collector distance ¼ 15 cm, voltage ¼ 17 kV and the feed rate ¼ 0.012 ml/min.

CS nanoparticle/PLA nanofibrous scaffold

PCL scaffold discs (13%, 16%, and 20% W/V) (flat or cylindrical)

Electrospinning/feed rate ¼ 4 ml/h, and the voltage ¼ 15–20 kV

h, for the shell and core, respectively Electrospinning/voltage ¼ 20 kV, and flow rate ¼ 1 ml/h through a small metallic needle, and the collector was flat collector for flat disc and rod for cylindrical scaffold

PCL/PLLA/HA nanofibrous composite scaffold with incorporated simvastatin (SIM)

Method of nanofibrous fabrication/parameters

Nanofibrous scaffolds

Table 2 (continued)

In vitro

In vitro

In vitro

Type of study

The nanofibers were rough due to CS nanoparticles

tetracycline hydrochloride • 13% W/V PCL showed microparticles • 16% W/V PCL showed spindle-shaped beads • 20% W/V PCL showed smooth fibers The nanofibers were smooth and bead free

Morphological properties

Biodegradable/ the addition of CS nanoparticles increased hydrophilicity

Biodegradable

–

The combination of CS nanoparticles enhanced the mechanical properties of pure PLA

–

Biodegradation/ hydrophilicity

modulus was increased –

Mechanical properties

Biocompatible/ SIM/HA boost cell viability and enhanced the osteoblasts differentiation and mineralization Biocompatible/by adding a certain amount of CS nanoparticles, the BMSCs adhesion, osteogenic differentiation, and ECM

Biocompatible/20% W/V PCL scaffold was considered the suitable for the attachment and stability of cells

Cytotoxicity

[118]

[119]

–

[117]

Reference

–

–

Other

386 A. M. Mansour and I. M. El-Sherbiny

Thermal phase separation

PLGA/BG composite scaffolds

In vitro & in vivo

In vitro & in vivo

In vitro & in vivo

Self-assembling

Electrospinning/for PCL layer: voltage ¼ 25–27 kV, syringe-collector distance ¼15 cm and feed rate ¼ 3.5–4 ml/h. For CS/PEO layer: voltage ¼ 14–16 kV, flow rate ¼ 0.5–0.8 ml/h and syringe-collector distance ¼ 15 cm

In vitro & in vivo

Electrospinning/feed rate ¼ 0.5 ml/h, and voltage ¼ 12 kV

CS/PEO nanofibrous coating layer for PCL nanofibrous layer in sandwich-like composite scaffold with/without sildenafil

Bilayered nanofibrous scaffold of PVA/bromelain conjugated Mg-doped nHA coated with Col/sericin Self-assembling peptide RADDA-16 nanofibrous hydrogel

PCL nanofibers were bead free and randomly oriented with smooth surface, while CS/PEO nanofibers layer is smaller and denser than PCL layer Nanofibrous 3-D network structure with uniform pore distribution

The scaffold showed a nanofibrous structure with interconnected nanopores

The nanofibers were smooth, bead free and interconnected randomly with the presence of nanorods –

Biodegradable/ CS/PEO coating layer increased the scaffold hydrophilicity

Biodegradable

Presence of the CS/PEO layers enhanced mechanical properties

–

Biodegradable/ the coating increase the scaffold hydrophilicity

–

nanofibrous scaffold Adequate mechanical properties

• Biocompatible/the scaffold enhanced MG-63 cells adhesion, proliferation, and

Biocompatible/the hydrogel enhanced the PDL cells proliferation and increased VEGF and OPN expressions Biocompatible/the scaffolds supported cell attachment and proliferation

mineralization were promoted Biocompatible/the scaffold enhanced cell attachment and proliferation

It had capacity for vascularization, and tissue maturation in the transplanted area

The scaffold enhanced alveolar bone defect healing especially those with sildenafil

The scaffold enhanced the healing of criticalsized socket defect

Chicken chorioallantoic membrane and aortic ring assaies showed that the scaffold was hemocompatible and provoked angiogenesis

(continued)

[123]

[122]

[121]

[120]

Nanofibrous Scaffolds for the Management of Periodontal Diseases 387

Method of nanofibrous fabrication/parameters

TIPS with a porogen leaching technique and cross-linked with BBP

Coaxial electrospinning/ voltage ¼ 25 kV, flow rate ¼ 2 ml/h for the sheath & 0.6 ml/h for the core, and syringe-collector distance ¼ 15 cm

Nanofibrous scaffolds

BBPfunctionalized GEL nanofibrous scaffolds

PEI/pBMP-2 with PLGA (core/shell) nanofibrous scaffold

Table 2 (continued)

In vitro

In vitro

Type of study

The scaffold was highly porous with smooth nanofibers

Nanofibrous structure with interconnected pores

Morphological properties

Adequate mechanical properties

BBP conjugation to the scaffold did not affect its mechanical properties

Mechanical properties

Biodegradable

–

Biodegradation/ hydrophilicity Cytotoxicity

Biocompatible/the scaffold allowed PDLSCs attachment, proliferation, and osteogenic differentiation

osteogenic differentiation. The ALP activity of MG-63 cells increased over time with calcified nodules formation Biocompatible/the scaffold improved osteogenic differentiation and osteogenic gene expression, including BMP-2 and OCN

The large surface area of the nanofibrous structure together with the high affinity between BBP and BMP-2 improved capacity of the scaffold for BMP2 binding and protection from degradation –.

Other

[125]

[124]

Reference

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by electrospinning followed by selective leaching of LA. The selective leaching out of LA content was performed by agitating with distilled water. This leads to hollowing out of the fiber and the formation of sparsely distributed nanofibers and fluffy 3-D fibrous mesh due to the repulsion force between the functional groups of LA in the nearby nanofibers [107, 132]. Polycaprolactone is another commonly used biocompatible and biodegradable synthetic polymer with adequate mechanical properties. However, the surface of PCL can be considered hydrophobic, PDLSCs cultured on PCL nanofibrous scaffolds showed adequate proliferation and detected PDL-specific marker (periostin) [117]. Several studies modified PCL to enhance its surface hydrophilicity. Mansour et al. [114] developed a composite nanofibrous scaffold composed of PCL electrospun nanofibers coated by cross-linked alginate (Alg) with nHAP. Alg was used in the outer layers to decrease the water contact angle for the nanofibers, enhancing cells attachment and proliferation, while nHAP added to improve the PCL nanofibers’ mechanical strength and osteogenic property that enhanced the cells osteogenic differentiation. This scaffold had been applied in dogs’ class II furcation defects either alone or with adipose-derived stem cells (ADSCs). This composite scaffold enhanced the periodontal regeneration in the furcation defect that confirmed by higher amount of bone regeneration, aligned mature PDL fibers with inserted Sharpey’s fibers in both newly formed CM and AB [114]. Noteworthy, the combination between stem cells and nanofibrous scaffolds for treating defects showed better bone and periodontal regeneration than treating the defects with the composite scaffold alone [114, 116]. Alipour et al. [103] used PCL- polyethylene glycol (PEG)-PCL electrospun nanofibrous scaffolds incorporated with zeolite. When dental pulp stem cells (DPSCs) seeded into this scaffold, they infiltrated the scaffold, secreted a large amount of ECM, and produced multilayer cell sheets. Additionally, zeolite enhanced the osteogenic differentiation of DPSCs [103]. There are other polymers, such as polyvinyl alcohol, which is hydrophilic, but it has low mechanical and physical properties. For that, it is recommended to introduce other components such as carbon nanotubes (CNTs), CaP, and HA in its structure [133–136]. CNTs are unique tubular structures with a nanoscale dimension. They have as well excellent mechanical properties. Development of PVA/nHA/CNT nanofibrous composite scaffold showed adequate mechanical and physical properties, in addition, its biocompatibility and angiogenesis effects were confirmed in vivo [109].

5.2.2

Delivery Vehicle Nanofibrous Scaffolds

There are several biophysical and biochemical cues that is necessary to restore the hierarchical architecture of periodontal tissues. These cues are controlled by several biological factors. Nano scaffolds can stabilize the bioactive agents and control their release at the intended target [39, 137]. Based on that, nanofibrous scaffolds can be used to deliver drugs, biological factors, or genetic material [138]. They can also

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make sustained and controlled delivery for these factors due to its high surface area [53, 139].

Drugs Nanofibrous scaffolds have substantial drug loading capacity and effective in vivo drug reactivity [53, 139]. Mostly, drugs are mixed with the polymers to produce drug incorporated nanofibers, which lead to burst release effect [140]. The drugs can also be inserted into the core of the nanofibers fabricated by coaxial electrospinning [141, 142]. This core/shell structure allowed sustained drug release, which subsequently enhanced the drug activity [86]. Several studies have designed scaffold loaded with antibiotic to avoid systemic antibiotic and its unneeded side effects. Budai-Szűcs et al. [104] designed a PLA electrospun nanofibrous scaffolds in the form of mats and disks. The mat was obtained directly from the collector, while the disk was obtained by compressing about 15 mg of the fiber in a pellet die. Metronidazole was the antibiotic used and its concentration was either 12.2 or 25.7 wt%. They found that disks were much faster than the mats in releasing the drug. Moreover, the antibacterial activity increased with increasing metronidazole concentration [104]. Other drugs related to enhanced bone formation have been added to nanofibrous scaffolds, such as sildenafil. It is a phosphodiesterase-5 inhibitor that has a confirmed angiogenic effect and bone healing acceleration effect. It may enhance bone healing by increasing the micro-circulation and the mediatory concentration at the fracture site or by the cysteine-rich protein pathway [143–145]. Yahia et al. [122] developed a sandwich-like structure of electrospun sildenafil-loaded PCL (middle layer) and CS/polyethylene oxide (CS/PEO) nanofibers (coating layers). In vivo, the scaffolds were applied into rabbit’s mandibular defects. In both histological and radiographically findings, the addition of sildenafil improved the bone formation within the defect [122].

Genetic Materials Although gene delivery is one of the promising approaches to tissue regeneration, it needs an efficient carrier to deliver the genetic materials safely and locally. There are some considerations during gene transfer in periodontal applications, such as the target cells, way of gene delivery and the anatomy of periodontal defect as one- or two-walled will need supportive carrier (scaffold) [146]. Gene delivery has a wide variety of nonviral and viral carriers. Here, We are concerned with nonviral approach because of cost-effectiveness, ease of fabrication, immuno-safety, and lack of DNA insert size limitation compared to viral approach [147, 148]. The combination of nanofibrous scaffold with nonviral gene delivery approach represents a reasonable and efficient process for gene therapy-based tissue engineering [113].

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The target genes for periodontal regeneration are genes related to GFs such as BMP2, wingless and Platelet-derived growth factor or transcription factors such as Ostrix and Runx2 [149]. Plasmid-BMP-2 (pBMP-2) can be carried by nanofibrous scaffolds. The electrospun nanofibrous scaffolds can be fabricated by single axial electrospun and the pBMP-2 mixed with the polymer before electrospinning or by coaxial electrospinning forming shell/core scaffolds. The shell/core scaffolds showed adequate gene release behavior and prolonged gene expression with high transfection efficiency. The pBMP-2 was inserted in the core, this allowed release control by the shell to provide sustained release effect [125]. With nonviral gene delivery platforms, the most commonly used carriers are cationic polymers that interact electrostatically with negatively charged circular plasmid DNA (pDNA), forming nanocomplexes (NCs) [150, 151]. Physical properties of these NCs, e.g., size, compactness, and surface charge, significantly affect the gene transfection efficacy. Microfluidic devices are used to form NCs. However microfluidic techniques have many advantages, it has some limitaions such as the existence of laminar flow and slow diffusion kinetic of pDNA [152]. Therefore, Tesla micromixer (TMM) was developed as a microreactor to allow efficient mixing. The TMM has three inlets, the pDNA encoding BMP-2 is injected via central channels, while the polymer is injected through the lateral inlets. This allowed reasonable conditions for polymer and pDNA to form more compact, smaller NCs with enhanced stability [113].

Bioactive Factors Most synthetic polymers lack the ability to guide and promote cell differentiation, so these polymers can be used as the carrier for exogenous GFs or bioactive materials that promote GFs production, such as BG [53, 103]. Several experiments had shown that the combination between PLGA and BG accelerated BG degradation and promoted bone defect regeneration [123, 153, 154]. The BG degradation products can promote GFs production, cell proliferation, osteoblasts differentiation, and bone deposition [155]. One of the most used GF for mineralized tissue formation is BMPs due to their superior osteoinductivity. Despite their promising clinical potential, there are concerns on their safety, cost-effectiveness, and high-dosage-related side effects. These side effects may result from the burst release [156]. Hence, alternative sustained delivery systems are needed. One method for sustained release can be achieved by its coupling with electrospun nanofibrous composite scaffolds. This combination allows sustained release of BMP over 4 weeks and enhanced cell proliferation. In vivo, PLGA/Col/GEL electrospun nanofibrous scaffold coupled with BMP-2 created an osseous-like structure for osteogenic cells infiltration and enhanced the healing process of critical-sized socket defects [157]. BMP-binding peptide (BBP)-functionalized scaffolds can be considered another feasible strategy to avoid the administration of high doses of exogenous BMPs. Cross-linking the

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nanofibrous scaffold with BBP can stimulate endogenous BMP2 expression without affecting the macro/microstructure and mechanical properties for the scaffold [124]. One of the potential bioactive agents is called bromelain, which is plant proteolytic enzyme (proteinase inhibitors). It has an anti-inflammatory, antimicrobial agent, antiplaque agent, anticancer activities [158, 159]. It elicits the antiinflammatory effect through reducing prostaglandin E2 and cyclooxygenase-2 synthesis [160], so it is used to treat osteoarthritis and gingivitis [161]. One of the important properties for bromelain is the ability to modify the surface area of nanoparticles and promote their binding affinity [158]. The functionalization of polymeric nanofibrous electrospun scaffold with Mg-doped nHA and bromelain enhanced the scaffold’s mechanical, physical and biological properties. This combination allowed sustained release of both bromelain and the nanoparticles that lead to enhanced antibacterial potential, cell proliferation, angiogenesis and confirmed by enhanced oral wound healing [120]. After injury, high mobility group box 1 (HMGB1) is one of the first signals secreted by necrotic cells. HMGB1 chemoattracts for inflammatory cells [162] and mesenchymal stem cells (MSCs) [163]. HMGB1 also stimulate the osteogenic differentiation of MSCs and angiogenesis. Lv et al. [112] immobilized HMGB1 on the surface of PLLA/PCL nanofibrous scaffold via heparin. This scaffold induced osteogenic differentiation of MSCs. Additionally, the scaffold enhanced bone regeneration process after its implantation in the rat calvarial defect [112, 163].

5.2.3

Injectable Nanofibrous Scaffolds

Presently, scaffolds for periodontal regeneration present in three forms: particulate, solid and injectable. Since it is necessary for the scaffold to be properly adapted in the defects, the injectable scaffold would be perfect for periodontal regeneration, as periodontal defects have irregular shape and size [164, 165]. Self-assembling peptide nanofiber hydrogel increased PDLSCs proliferation and enhanced periodontal defects regeneration. Histologically, it enhanced new AB formation with regeneration of oriented PDL-like collagen bundles and angiogenesis [121]. Conventional injectable hydrogel usually loses its mechanical strength after injection, which may lead to delivery failure [166], so several studies prepared a self-healing hydrogel. Self-healing hydrogel can stay in a stable state after injection, which is perfect for AB regeneration. Additionally, it allows adequate support for gingival tissue repair [167–169]. Solid microspheres (SM) used within injectable scaffolds and act as a drug carrier on their spherical polymer matrix [170]. As the porosity is one of the main requirements, nanofibrous-microsphere (NF-MS) and nanofibrous-spongy microsphere (NF-SMS) were developed [171]. Comparing these three types of microspheres, all of them allowed DPSCs attachment, but NF-SMS allowed the cells to attach to both the surface and the interior areas. NF-SMS also enhanced the cells proliferation and odontogenic differentiation [172]. Additionally, NF-SMS has a structure of both a hollow microstructure and a fibrous nanostructure. In micrometer scale, it

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improves nutrition transportation, decreases the degradation products, and facilitates new tissue formation. In nanometer scale, it mimics the structure of ECM and enhances the cell interactions [171]. Nanofibrous-spongy microsphere can be functionalized with MSNs to develop advanced multibiologic injectable scaffold. MSNs have high surface area, welldefined pore structures, and adjustable pore size and can be loaded with GFs such as IL-2 and TGF-β to locally recruit T cells and induce their differentiation into regulatory T cells (Tregs). Whereas NF-SMS act as an injectable scaffold for Tregs attachment. Tregs play important roles in microenvironment modulation for periodontal tissue regeneration. This injectable scaffold suppressed the exacerbated immune response during periodontal diseases [110, 173]. Recently, Hao et al. [106] developed PLGA/MSNs core-shell porous microsphere immobilized PLLA spongy nanofibrous microscaffold. The PLGA/MSNs-PMS was designed as a carrier for multiple biomolecules such as BMP-2 and Celecoxib. BMP-2 was used to stimulate osteogenesis, while Celecoxib was used to inhibit the inflammation during AB regeneration [174, 175]. The PLGA/MSNs-PMS was attached to the PLLA scaffold to control the release of the loaded biomolecules. Additionally, the design of PLGA microspheres’ pores prevented the MSNs clusters to migrate outside the microspheres, while the release of BMP-2 and drug-loaded single MSNs had no hindrance. When this device was injected into in periodontitis model, the MMP cleaved the MSN clusters, and the BMP-2 and celecoxib were released. The regeneration of AB was significant and similar to normal AB [106].

5.2.4

Multiphasic Nanofibrous Scaffolds

The multilayered/multiphasic scaffold allowed simultaneous regeneration of two or more tissues which is the main goal in periodontal regeneration, as ideal periodontal regeneration requires AB and CM regeneration with inserted PDL fibers. Therefore, PDL, CM and AB regeneration should be considered during designing the multiphasic scaffolds for periodontal regeneration [126, 176]. Sowmya et al. [126] developed a trilayered scaffold, and each layer enhanced specific periodontal tissue regeneration. The first layer (CM layer) consisted of PLGA/NBG/cementum protein 1. The second layer (PDL layer) composed of chitin–PLGA/fibroblast GF-2. The third layer (AB layer) composed of chitin–PLGA/NBG/platelet-rich plasma derived GFs. This structure allowed complete periodontal regeneration of periodontal defects in vivo [126]. Additionally, the interface between hard and soft tissues with oriented PDL fibers is crucial in the regeneration of functional periodontal tissues [108]. There are several studies recently designed nanofibrous scaffolds with aligned and oriented fibers to originate properly aligned PDL fibers and to mimic the bone ECM [112]. Jiang et al. [108] fabricated 3-D multilayered scaffold by embedding aligned PCL/PEG electrospun nanofibrous mats into CS-geniping mixture, then freezedrying was performed and finally, washing was done to remove genipin residual. With rat’s periodontal fenestration defect, the scaffold was placed in cross-sectioned

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Fig. 4 Images for rat’s periodontal defects after 2 months: micro-CT images (a, b, c, & d) showing more bone regeneration around the tooth in the scaffold groups than untreated group, and H&E photomicrographs for PDL region (e, f, g, & h) with inserted angulation of the PDL fibers showing the fibers in aligned scaffold group had angles value closest to the natural PDL and had more parallel fibers than random scaffold group [108]. Figures are reprinted with permission from Elsevier

slices against the root and then packed with Bio-OssÒ membrane to stabilize the scaffolds. The aligned scaffold guided PDL orientation that is most like normal PDL with higher PDL marker periostin and higher Col I/III ratio, suggesting formation of more matured PDL fibers. The amount of regenerated AB was significantly regenerated with higher volume and density than randomly oriented scaffold (Fig. 4) [108].

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Fig. 5 SEM images and fluorescence microscopic images for cells on aligned and random nanofibers showing the cell adherence and cell morphology with arrowed dash lines indicate orientations of cell elongation on the aligned scaffold [108]. Figures are reprinted with permission from Elsevier

The aligned nanofibrous structure controls the aligned morphologies for most cell types via contact guidance (Fig. 5) [108, 177]. It also showed superior outcome in periodontal regeneration as it provides nanotopological cues [108, 111, 112]. As PDL has mechanical stressed conditions of both compression and tension, nanotopological and mechanical cues can be combined to create a native tissue environment for studying cell behavior properly. When cells were cultured into PCL/GEL nanofibrous scaffold with applied mechanical stress, the PDL markers such as periostin and tenascin were significantly expressed with reduction in osteogenesis, demonstrating the roles of nanotopological and mechanical cues in control the phenotype of PDL cells. Furthermore, the nanofiber/cell constructs engineered under mechanical stress showed enhanced bone regeneration with complete integration of the constructs within the defects [111].

6 Implant-Related Nanofibrous Biomaterials Dental implant is the most conservative artificial tooth replacement. The main difference between dental implant and natural teeth is the absence of PDL fibers around the implants. Lack of adequate AB support and presence of adjacent anatomic structures such as inferior alveolar canal and maxillary sinus are the most frequently encountered problems facing dentists during implant placement. In these situations, bone regeneration strategies like bone graft or osteogenic scaffolds with or without GBR are the best choice [178–180].

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The stability for the implant depends mainly on osteointegration. Osteointegration (Osseointegration) is tight connection between AB and the surface of the dental implant [181]. As the dental implants are mainly formed of titanium which has poor osteoinductivity and osteoconductivity, several studies were aimed to modify the implant surface by nanofibrous coating to enhance the osteointegration and prolong the dental implant’s life and functionality [182–188]. The coating can be simply performed by using the implant as the collector during electrospinning process [189]. The electrospun coating can be PLA/PCL nanofibrous coatings [190], Equisetum arvense containing polyvinyl alcohol-sericin electrospun nanofibrous coating [185], or combination of minerals substituted hydroxyapatite/PEG/Cissus quadrangularis extract [191]. There is another simple method of coating, which is electrospray deposition. When Col I nanofibers had been electrosprayed over dental implant, it allowed better gingival/implant contact than non-coated implants [187]. Other researchers added drugs to the nanofibrous coating to prepare multifunctional coating that enhance the implant osteointegration and eliminate the implant-related inflammation and infection. The drugs commonly used were antibiotics such as vancomycin and rifampicin [192, 193] and anti-inflammatory drugs such as aspirin and dexamethasone [183, 194, 195]. Other studies aimed to manage peri-implant diseases successfully and rescue the implant from failure [196–199]. About the peri-implant diseases, there are two periodontal diseases related to dental implant: peri-implant mucositis, which is an inflammation in the gingiva around dental implants, and peri-implantitis, which is a chronic inflammation around dental implants leading to periodontal tissues destruction (Fig. 1) [188]. The management of peri-implant diseases is similar to the treatment of periodontal diseases, except that the debridement instruments have to be coated with softer materials to avoid scratching the implant surface with metallic conventional instruments [20]. In vitro, less abrasive debridement instruments, such as airflow devices, have shown effective biofilm removal without scratching the implant surface [200]. The regenerative approaches for peri-implantitis also are similar to that for periodontitis which depend on using GTR with or without scaffold.

7 Conclusions and Future Perspectives Although periodontal regeneration is a huge field that cannot be simply covered in one article, this book chapter briefly discussed the recent types of nanofibrous biomaterials developed experimentally for periodontal regeneration and the concept for their fabrication. The chapter also showed how the scaffold architecture and composition can affect stem cells behavior and subsequently, affect the in vivo outcomes. This was obvious with aligned nanofibers that promoted alignment of the cells and the regenerated PDL fibers. For long time ago, it was noticeable that AB regeneration is the most concerned part in periodontal regeneration. Henceforth, researchers have a direction toward 3D multilayered and multifunctional scaffolds that concerned with the regeneration of the main three parts of periodontal tissues,

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which are AB, PDL and CM. There is also another technology called 3D bioprinting that is used for several medical studies. We expected that with more advances in 3D bioprinting, it will be widely used for optimal fabrication of nanofibrous-based 3D scaffolds for periodontal regeneration.

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Adv Polym Sci (2023) 291: 409–424 https://doi.org/10.1007/12_2022_141 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 3 January 2023

Recent Advances in Brain Tumour Therapy Using Electrospun Nanofibres Arathyram Ramachandra Kurup Sasikala

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Application of Electrospun Nanofibres in the Brain Tumour Research . . . . . . . . . . . . . . . . . . . . 3 Nanofibre as a Chemotherapeutic Delivery Platform . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1 Carmustine or BCNU . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 Temozolomide (TMZ) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3 Paclitaxel (PTX) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.4 Delivery of Other Therapeutics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Conclusion and Future Perspective . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

410 411 412 412 416 417 419 421 421

Abstract The chances of surviving a brain tumour have remained dismally low for decades. Several global research efforts are being carried out to accelerate the understanding and different ways to treat brain tumours but presently, most malignant brain tumours remain incurable. This is due to the high rate of tumour recurrence even after tumour resection as well as the presence of blood-brain barrier (BBB) which limits the efficacy of the treatment. Therefore, local intracranial delivery systems have emerged as viable alternatives for the possible delivery of localised drug doses immediately post-neurosurgical resection, to potentially kill the remaining infiltrative cells with minimum systemic exposure and adverse side effects. In this aspect nanofibre-based local delivery system has shown greater efficacy for the effective collocation to the irregularly shaped post-surgical resection margin for the successful delivery of anticancer drugs by reducing the successful drug diffusion distance from the resection site and the brain parenchyma beyond.

A. R. K. Sasikala (✉) Centre for Pharmaceutical Engineering Science, School of Pharmacy and Medical Sciences, University of Bradford, Bradford, UK e-mail: [email protected]

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This chapter will focus on the application of nanofibres as a localised therapeutic delivery platform for the brain tumour treatment. Keywords Brain tumour · Electrospun nanofibres · Localised drug delivery system · Tumour recurrence

1 Introduction The chances of surviving a brain tumour have remained dismally low for decades. Several research are being carried out to accelerate the understanding and different ways to treat brain tumours [1] but even at the moment brain tumour remains incurable [2, 3]. Amongst the brain tumours, Glioblastoma (GBM) is considered as one of the most aggressive and highly heterogeneous primary brain tumours in adults [4, 5]. The median survival rate of patients affected with GBM is below 15 months from the date of diagnosis showing the severity and unmet therapeutic need of this highly aggressive disease [4–7]. The conventional chemotherapeutic strategies such as systemic chemotherapy fails to show clinical relevance due to the presence of BBB that prevents therapeutics to enter the brain parenchyma and chemoresistance of tumour cells [8]. Several strategies [9, 10] have been devised to improve BBB crossing, but none are currently used in clinical settings due to their practical invasiveness [11], severe side effects, and procedural complications coupled with repeated drug administration [12]. To date, “the gold standard therapy for GBM” involves the surgical resection of the accessible tumours followed by standard radiotherapy with concomitant and adjuvant chemotherapy. However, most brain tumours, especially GBM, infiltrate the surrounding normal brain parenchyma to form satellite tumours and almost impossible for complete surgical ablation [13]. Therefore, local delivery system has emerged as a good alternative [14, 15] to prevent GBM tumour recurrence that often exist 2 cm in the interior of the post-surgical tumour resection lining [16]. The localised delivery (LD) of potential drug doses into the resection margin will help to kill the residual infiltrative cells with minimum systemic exposure [17]. To date, numerous localised drug delivery systems have been developed in the laboratory for the treatment of GBM, but only one of the localised delivery systems called Gliadel™ (chemotherapeutic impregnated polymeric wafers) has been approved by the Food and Drug Administration (FDA) and the National Institute for Health and Clinical Excellence (NICE). Gliadel™ wafers are implanted in the tumour resection cavity at the time of surgery to release carmustine, a chemotherapeutic agent to kill the residual cancer cells in the tumour resection cavity. Gliadel™ has shown a mild but significant survival benefit of 2.3–1.8 months median survival for newly diagnosed and recurrent high-grade gliomas, respectively [18, 19]. Some of the main drawbacks associated with Gliadel™ include the (1) rigid structure of the Gliadel™ wafers that restrict the effective apposition to the unevenly shaped

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neurosurgical residual cavity and therefore the drug cannot reach the entire cavity area, (2) due to the disc like shape, the wafers displace from the tumour resection cavity and reduce the effective drug diffusion to the brain parenchyma, (3) Moreover, the inability of the drug, carmustine to penetrate deep inside the infiltrative margins to eradicate residual tumour cells to prevent the tumour recurrence, and finally (4) the mono-therapeutic approach using carmustine alone is ineffective to treat highly heterogenous tumour like GBM [20]. This confirms the rationale of localised drug delivery system (LDDS) and the importance of developing novel LDDS to overcome these limitations and maximise the chance of clinical translation for the GBM treatment. OncoGel™ is a new experimental drug delivery system that allows the controlled release of an approved intravenous anticancer drug paclitaxel from a gel (ReGel™) [21]. Preclinical and early clinical investigations of OncoGel™ in brain tumour treatment demonstrated that the OncoGel™ enables the physical targeting of paclitaxel to brain tumour vial intralesional injection in the surgical tumour resection cavity with an acceptable safety profile and moderate increase in rat gliosarcoma model [22]. Thus, compared to systemic therapies, these LDDS-based approaches are expected to improve the efficacy of drugs by increasing the exposure time of tumour cells to drug from degradation and clearance by the immune system until its release from the drug delivery system and allow the oncological treatment maintained in the interval between surgery and radiotherapy the effective treatment of brain tumours. Therefore, numerous research is happening to develop the best possible localised drug delivery system for the treatment of brain tumours. In this chapter, we focus on the development of electrospun nanofibres-based localised drug delivery systems for the brain tumour related research.

2 Application of Electrospun Nanofibres in the Brain Tumour Research Electrospinning the most widely used and versatile technique to produce fibres from micro-to nanometres of diameter with precise surface morphology [23]. These fibres are fabricated by applying strong electric field on required polymer solution and melt spinning is done in the case of a polymer lacks a good solvent [23, 24]. By carefully tuning the parameters of electrospinning, the properties of nanofibres including the diameters can be modified. The unique properties of nanofibres such as enhanced surface area to volume ratio, modified surface functionalities including superior mechanical performances make them ideal candidates for many bio applications [23, 24]. Among them, the application of electrospun nanofibres in the drug delivery field is one of the most studied areas for the controlled delivery of various therapeutics such as anticancer drugs, antibiotics, macromolecules like protein and DNA [25].

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Correspondingly, in brain tumour research, electrospun nanofibres are gaining ample attention as a post-surgical drug delivery device. Since the therapeutic molecules are encapsulated in the nanofibres, the localised delivery of chemotherapeutics at the tumour resection margins is possible once it is implanted and thereby preventing systemic toxicity related to chemotherapy. Due to the versatility in designing various nanofibres, the nanofibres can deliver multiple therapeutics. One example is core–sheath electrospinning in which the core of the electrospun fibre might store one drug while the outer sheath stores another, offering a successive or overlapping release depending on the pharmacokinetics of the polymers [26]. Another research scenario that uses electrospun nanofibres in brain tumour research is to study the cell migration and apply potential anti-invasion therapies to halt the ability of malignant brain tumours to disperse through neural tissues. The unique topography of the nanofibres mimics the structure of nerves and blood vessels that brain cells normally use to invade other parts of the brain [27]. Researchers are developing various nanofibre morphologies such as aligned [28, 29], fibres with different densities [30] to discriminate the migration potential of different brain tumour stem cells as well as drug screening. In this chapter, I am focusing on the former application of nanofibres as a postsurgical therapeutic delivery platform for the brain tumour treatment.

3 Nanofibre as a Chemotherapeutic Delivery Platform Chemotherapy is a traditional and less invasive compared to surgery but in the case of aggressive cancers like brain tumour they are only used as adjuvant therapy. So, the invasiveness of placing a nanofibre-based drug delivery system can be overcome by placing it during the time of surgery in the post-surgical tumour resection cavity. Over the past few years, different chemotherapy drugs such as carmustine (BCNU), paclitaxel (PTX), and temozolomide (TMZ) have been explored to prepare nanofibres-based localised therapeutic system for brain tumour treatment. Table 1 gives the list of natural/synthetic polymer nanofibres-based therapeutic systems for controlled and localised delivery of anticancer therapeutics for brain tumour treatment both in vitro and in vivo. Among them, some of the uniquely interesting studies are discussed in detail in the following sections.

3.1

Carmustine or BCNU

BCNU is one of the most common drugs utilised for GBM treatment [31, 32]. BCNU can penetrate the BBB, but its intravenous administration is associated with a list of side effects including bone marrow suppression, hepatic dysfunction, and pulmonary fibrosis [31–34]. One of the FDA-approved localised delivery systems Gliadel™ showed some improvements of 2 months in the patient

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Table 1 Electrospun nanofibres-based therapeutic systems for localised brain tumour therapy Polymer PLGA

Anticancer agent BCNU

PLGA-PLLA

BCNU

PCPP-SA and PCL

BCNU

PLGA-PLA-PCL

TMZ

PLGA

O6 BG, TMZ, BCNU

PCL

TMZ, NGF

PCL

TMZ

Highlights BCNU was released from the NFs for more than 6 weeks after placing in the surgical cavity of mouse Sustained release of CBNU in its active form for 72 h from the NFs for enhanced anticancer effects compared to the pristine BCNU which failed to show the anticancer effect after 48 h due to its instability and short half-life Consistent and extended release of BCNU from the NanoMesh compared to the consistent thick solid disc with the enhanced in vivo anticancer efficacy and long-term biocompatibility Controlled TMZ release from flexible polymeric nanofibres for a longer period when implanted in orthotopic rat GBM for controlling glioma recurrence Sequential and sustainable delivery of TMZ, BCNU, and O6BG for more than 14 weeks when surgically implanted on the brain surface of the rat bearing F98 tumour and 9L tumour for better therapeutic advantages for GBM treatment Nanofibre-based device integrating TMZ doped PCLS NFs NGF-loaded PCL membranes using sodium alginate hydrogel as a post-surgical implant for glioma treatment by inhibiting glioma cell growth inhibition and in the meantime supporting neural cell differentiation TMZ-loaded PCL nanofibres showed enhanced anticancer effect against U87 glioma cells due to the upregulation of p53, Bax, and Bcl2 genes expression

Reference [15]

[36]

[26]

[37]

[38]

[39]

[40]

(continued)

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Table 1 (continued) Polymer Poly (L-lactide-co-ε caprolactone) (PLACap), poly (L-lactide-co-trimethylene carbonate) (PLAGATMC)

Anticancer agent TMZ, nimorazole

Poly (-caprolactone diol) based polyurethane (PCL-diol-b-PU)

TMZ

PLGA

PTX

PLGA

PTX

Ca-alginate microparticles and polypropylene carbonate (PPC)

PTX, TMZ

PLGA

BCNU, irinotecan, cisplatin.

Highlights A three phased release profile of TMZ from PLACap and PLAGATMC core shell nanofibres produced by a combination of coaxial electrospinning and electrospray techniques TMZ-loaded chitosan nanoparticles incorporated PCL-diol-b-PU nanofibres further coated with 18 nm gold nanoparticles. The NF composite showed anticancer effects by inhibiting the growth of U-87 GBM cells in vitro The PTX-loaded PLGA NFs showed a sustained release of drug even after 42 days with a drug penetration up to 5 mm from the implant site and resulted in a substantial tumour reduction in an intracranial U87 MG-luc2 glioblastoma model PTX-loaded PLGA submicron fibres showed sustained PTX release over 80 days in vitro with small burst release compared to the PLGA sheets and PLGA microfibres for the enhanced glioma inhibition in vivo Biodegradable implant was developed using PTX-TMZ nanofibres containing PTX-loaded Ca-alginate microparticles for the prolonged release of PTX and TMZ with a reduced initial burst release. The weight ratio of PTX and TMZ was 1:1, optimal synergistic effect was obtained for the inhibition of glioma C6 cells All the three therapeutics were released in high concentration from the PLGA NFs for more than 8 weeks in the cerebral cavity of rats without antiinflammatory reactions and with enhanced GBM treatment efficacy

Reference [41]

[42]

[43]

[44]

[45]

[46]

(continued)

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Table 1 (continued) Polymer Poly (lactic acid) (PLA), polyethylene oxide (PEO)

Anticancer agent Rapamycin

PCL-poly (ethylene glycol)PCL (PCEC)

Curcumin

Starch

Carvacrol

PVP, PCL

Mycophenolic acid (MPA)

PLGA

Salinomycin

PEI, PLGA

MMP-2 mRNA, PTX

PLA

TRAIL

Highlights PEO-PLA NFs showed sustained release of rapamycin and induced cytotoxicity in both U251 and U87 GBM cells Curcumin-loaded PCEC nanofibres showed excellent antitumour activity against rat glioma 9L cells by inducing apoptosis The carvacrol-loaded starch nanofibres demonstrated up to 50% reduction in C6 rat glioma cells MPA, an FDA-approved immuno-suppressant was loaded in the coaxial nanofibres formed by PVP core and PCL sheath exhibited controlled release of MPA for in vitro GBM cell growth inhibition The Sali-PLGA NFs showed 30 days stability and a prolonged release of Sali for 2 weeks. The Sali-PLGA NFs exhibited 50% cytotoxicity against human glioblastoma (U-251) cells The sustained release of MMP-2 suppressing gene and the anticancer drug paclitaxel (PTX) from PLGA nanofibres exhibited significant antitumour efficacy against intracranial xenograft tumour models in comparison with pristine PTX and PTX NFs Human mesenchymal stem cells (hMSCs) releasing antitumour protein TRAIL-loaded NFs exhibited threefold reduction in volume of GBM xenografts

Reference [47]

[48]

[49]

[50]

[51]

[52]

[53]

survival rate by utilising BCNU as an anticancer agent which released over a period of 2–3 weeks after placing in the tumour post-surgical cavity [35]. To enhance the efficacy and long-term release of BCNU, Tseng et al. [15] fabricated biodegradable PLGA (poly[(D,L)-lactide-co-glycolide]) nanofibres. They demonstrated that the high concentration of BCNU was released from the PLGA NF for greater than 6 weeks when placed in the cerebral cavity of the mouse. Compared

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c

fiber membrane

b release

solid flim

d

time

release

short term drug release long term drug release

a

time

Fig. 1 (a) Drug release comparison from the solid film, (b) drug release kinetics over time from the solid film, (c) Drug release comparison from the multilayered fibre membrane wafer, (d) the drug release kinetics over time from the multilayered fibre membrane wafer. Reproduced with Creative Commons CC BY licence [26]

to the disc shaped Gliadel wafers, the flexible PLGA NF could completely cover the tumour resection cavity and better conform to the geometry of the tumour resection margins for better drug transport. Moreover, there were no signs of inflammation reaction of the brain tissue after placing the PLGA NFs further evidencing its therapeutic efficacy for effective GBM treatment. Inspired from the design of Gliadel, Han et al. [26]. developed a 3-dimensional disc (NanoMesh) composed of core–sheath nanofibres. The NanoMesh was fabricated by encapsulating BCNU in a polyanhydride poly-[bis( p-carboxyphenoxy) propane-sebacic acid] (PCPP-SA) nanofibre core with a polycaprolactone (PCL) sheath initially, and then they were folded multiple times to form the discs. They demonstrated the consistent long-term delivery of BCNU from the NanoMesh compared to the consistent thick solid. This is due to the uniform diffusion length of drug from multilayered porous membranes of the NanoMesh as shown in Fig. 1. When implanted in the F344 rats with intracranial 9L gliosarcoma, the Nanomesh showed in vivo anticancer efficacy by exhibiting a median and overall survival compared to untreated controls. To improve the GBM treatment efficacy, Tseng et al. [46] further developed a PLGA-based biodegradable nanofibrous membrane containing three different chemotherapeutic agents such as BCNU, irinotecan, and cisplatin. High concentration of all the three therapeutics was released for more than 8 weeks from PLGA NFs in the cerebral cavity of rats without any inflammatory reactions and with enhanced GBM treatment efficacy.

3.2

Temozolomide (TMZ)

TMZ is one of the most common oral anticancer drugs used in the treatment of GBM, but the treatment efficacy of TMZ is very limited due to the presence of BBB.

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Poor stability and short plasma half-life (1.8 h) of TMZ are also major hurdles to overcome. Thus, several studies have explored the localised formulation of TMZ into nanofibres to achieve high drug concentrations for the treatment of GBM in comparison to the oral administration. Ramachandran et al. [37] developed composite nanofibres of PLGA, PLA, and PCL to encapsulate TMZ in which TMZ release was controlled from days to months by switching the fibre-by-fibre composition as shown in Fig. 2a. They showed highly controlled drug release for a longer period from these nanofibres when implanted in orthotopic rat GBM for controlling glioma recurrence. In an interesting study, Liu et al. [38] developed PLGA-based hybrid structured nanofibrous membranes (HSNMs) through coaxial electrospinning for the delivery of multiple therapeutics such as TMZ, BCNU, and an antineoplastic agent O6 Benzylguanine (O6 BG) which promotes the cytotoxicity of the alkylating chemotherapeutics. They demonstrated sequential and sustainable delivery of O6BG TMZ, BCNU for greater than 14 weeks from HSNMs when they were surgically implanted on the brain surface of the rat bearing F98 (Rat malignant glioma) tumour and 9L (rat gliosarcoma). Compared to several experimental groups and control groups, HSNM treated groups showed retarded and restricted tumour growth, prolonged survival time, and attenuated malignancy as shown in Fig. 3. Huang et al. [39] explored the dual role of nanofibre-based device as a postsurgical implant for glioma treatment by stopping the glioma cell growth and in the meantime neural cell differentiation for tissue reconstruction. For this, they integrated TMZ doped PCL NFs and neuron growth factor (NGF) coated PCL membranes using sodium alginate hydrogel. The NGF device showed enhanced efficacy in inhibiting the growth of C6 glioma cells by releasing TMZ. Moreover, the sufficient release of NGF induced differentiation of neuronal cells over 4 weeks. Thus, the study put forward a new concept in the post-surgical glioma treatment.

3.3

Paclitaxel (PTX)

PTX is an anticancer drug widely used for the treatment of many types of cancers, such as breast cancer [54], ovarian cancer [55], and lung cancer [56]. Apart from the presence of BBB, the systemic administration of PTX for GBM treatment is also hindered by other reasons such as its hydrophobic nature and fast clearance from plasma. To circumvent these limitations, several studies have reported for the local delivery of PTX alone or in combination with other chemotherapeutic agents for the treatment of GBM. Therefore, several nanofibre formulations have been tested for the localised delivery of PTX. In one study, Ranganath et al. [43] developed a PLGA nanofibre loaded with PTX and tested its anticancer effects in an intracranial U87 MG-luc2 glioblastoma mice model. The NF-PTX composite showed a controlled release of PTX for 42 days post implantation with a drug penetration capability of up to 5 mm from the implant site. The results were compared with the PLGA-loaded microfibres and they found that

Fig. 2 (a) Schematic showing the tuning of formulation parameters for developing the TMZ nanofibre implant, (b) Schematic showing the overall idea of the work, (c) Photograph of the developed nanofibre implant showing the flexibility, (d) SEM image showing the smooth fibre morphology of 20 wt% TMZ-loaded implant, (e) EDS mapping showing the uniform distribution of drug across the nanofibres. Reproduced with Creative Commons CC BY licence [37]

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Fig. 3 (a) MRI of brain taken 4 weeks after the treatment in each subgroup listed in the table in b, (b) Table showing the list of subgroups where F98 and 9L represent the rat malignant glioma and rat gliosarcoma, respectively. In the MRI image, central necrosis and tumour are presented with thick and thin arrows, respectively, (c) Photographs showing the surgical procedure (1) craniectomy (~10 mm × 10 mm), (2) one-twentieth Gliadel wafer, and (3) HSNAM (10 mm × 10 mm) surgically implanted onto the brain surface of the rats. Reproduced with Creative Commons CC BY licence [38]

due to the low PTX release rate from the microfibres, there was not much penetration of PTX towards the resection cavity. Moreover, there was a substantial tumour inhibition (~30-fold) occurred in the case of PTX-PLGA NFs treated groups compared to the sham and placebo controls after 41 days of treatment. In a different study, a biodegradable implant was developed for the local delivery of both PTX and TMZ. For that, the PTX-loaded Ca-alginate microparticles (MPs) were developed first and then added to the polypropylene carbonate (PPC) emulsion containing different weight of TMZ to carry out the electrospinning to produce fibres consisting of beads-in-string structure [45]. The obtained PTX-TMZ fibres showed sustained release of these chemotherapy agents with a reduced initial burst release. The cytotoxicity assays of the nanofibre with different PTX:TMZ ratio demonstrated that the optimal synergistic effect for the inhibition of glioma C6 cells was obtained when the weight ratio of PTX and TMZ was 1:1.

3.4

Delivery of Other Therapeutics

In case of GBM, conventional chemotherapeutics inhibit the tumour growth by halting the cell proliferation, but its malignancy is mostly caused by local recurrence greater than 90%. So extensive efforts have been devoted to developing new therapeutics for the treatment of GBM. Some of the recent works using rapamycin

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[47], curcumin [48], carvacrol [49], mycophenolic acid [50], etc., are listed in Table 1. In an interesting study, Norouzi et al. [51] developed salinomycin (Sali)-loaded PLGA NFs for the treatment of GBM. Salinomycin is an antibiotic which is introduced recently as an anticancer drug. The Sali-PLGA NFs showed 30 days stability with a prolonged release of Sali for no less than a 2-week period. The SaliPLGA NFs exhibited 50% cytotoxicity against human glioblastoma (U-251) cells by inducing the intracellular reactive oxygen species (ROS) lead apoptosis as shown in Fig. 4.

Fig. 4 Flow cytometric analysis of cell apoptosis/necrosis of U251 after 48 h of the treatment, stained with Annexin V-FITC and PI (a) control, (b) Sali, (c) NFs + Sali treated cells, (d) Fluorescence microscopy images of U251 showing the antitumour effect of NFs + Sali compared to Sali and controls after 48 h of the treatment. Reproduced with Creative Commons CC BY licence [51]

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Lei et al. [52] developed a gene/drug dual delivery system for the synergistic therapy of glioma. Here they utilised an essential proteinase regulating brain tumour invasion and angiogenesis called Matrix metalloproteinase-2 (MMP-2) as one of the therapeutic targets. So they first developed an RNA plasmid that can suppress MMP-2 expression in tumour cells and then it is complexed with polyethyleneimine (PEI). The multiple therapeutic cargoes such as the gene and an anticancer drug paclitaxel (PTX) were further loaded into PLGA nanofibres for achieving a sustained release of the anticancer drug and the RNA simultaneously. As a result of this they exhibited significant tumour inhibition on intracranial xenograft tumour models due to the synergistic therapeutic effects of the gene and PTX. Recently, Engineered Stem Cells (ESC)-based therapy is emerged as an effective approach for GBM therapy [53, 57–59]. Stem Cells’ unique tumour-homing capacity enables them to migrate to both local and invasive GBM foci to be used as an ideal drug carrier for various antitumour agents [60]. Retaining the cytotoxic SCs in the tumour resection cavity is one of the main challenges in the GBM treatment. For addressing this challenge, Bago et al. developed drug releasing hMSCs-loaded PLA nanofibres that offered fivefold increase in hMSCs retention in the post-surgical tumour cavity to result in an increased release of antitumour protein TRAIL. Thus, the hMSCs-loaded nanofibres exhibited a threefold decrease in tumour volume and prevented the regrowth of residual GBM foci for an increased survival time from 13.5 to 31 days in mice.

4 Conclusion and Future Perspective Electrospun nanofibres have been shown greater efficacy in treating brain tumours by offering the potential delivery of high local drug doses directly to the neurosurgical resection to kill the remaining infiltrative cells with minimum systemic exposure. Nanofibres can carry multiple therapeutics including a combination of chemotherapeutics, stem cells, and genes and the sustained delivery of these therapeutic cargoes is tuned according to the fibre fabrication techniques and polymer properties. Thus, the nanofibre-based platform enables the simultaneous and sustained delivery of therapeutics to act synergistically towards tumour cell eradication at the edge of tumour and beyond. More clinical and preclinical studies are required to translate the nanofibre-based combined therapeutic strategies in the next decades. One major disadvantage is the invasiveness of its applications as this system can be used only as an adjunct to the modern surgical strategies rather than a substitute for surgical therapy.

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Adv Polym Sci (2023) 291: 425–468 https://doi.org/10.1007/12_2022_124 © The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 Published online: 16 June 2022

Layered Fibrous Scaffolds/Membranes in Wound Healing Ayşe Günyaktı, Tuğrul Tolga Demirtaş, and Ayşe Karakeçili

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Skin Structure, Function and Wound Healing Process . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Acute Wound Healing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Chronic Wounds (Longer Than 12 Weeks) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Polymers as Wound-Healing Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1 Natural Polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 Synthetic Polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Cutaneous Scaffolds with Added Therapeutic Agents for Wound Treatment . . . . . . . . . . . . . 4.1 Growth Factors (GFs) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.2 Antibiotics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.3 Natural Substances . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.4 Antimicrobial Peptides (AMP) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.5 Metal Nanoparticles (MNPs) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.6 Metal-Organic Frameworks (MOFs) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 The Conceptual of Design of Layered Scaffolds/Membranes for Wound Healing . . . . . . . . 6 Processing Techniques for Fibrous/Layered Wound Dressings . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.1 Multi-layered Electrospun Membranes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.2 Combination of Three-Dimensional (3D) Porous Scaffolds with Nanofibers for Preparation of Layered/Fibrous Wound Dressings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.3 Combination of Hydrogel Scaffolds with Nanofibers for Preparation of Layered/Fibrous Wound Dressings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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A. Günyaktı Biotechnology Institute, Ankara University, Ankara, Turkey T. T. Demirtaş Department of Basic Pharmaceutical Sciences, Faculty of Pharmacy, Erciyes University, Kayseri, Turkey A. Karakeçili (*) Chemical Engineering Department, Faculty of Engineering, Ankara University, Ankara, Turkey e-mail: [email protected]

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6.4

Combination of Three-Dimensional (3D) Printed Scaffolds with Nanofibers for Preparation of Layered/Fibrous Wound Dressings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 455 7 Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 456 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 457

Abstract Skin is the largest organ of the human body acting as a barrier to protect the body from external effects and trauma. As a result of external physical damages or physiological disorders such as diabetes, skin tissue is disrupted, and cellular integrity is lost in the wounded site. Design and production of functional bioactive wound dressing matrices to protect the injured area, assist the wound healing process and guide the regeneration of healthy tissue are of utmost importance. Considering the complexity of the wound healing process and challenging requirements to fulfill the clinical need in terms of both healing and regeneration, multi-layered fibrous membranes/scaffolds offer an effective strategy for the design and development of multi-functional wound-healing matrices. Such matrices act to stimulate the woundhealing cascade by combining different materials with different physicochemical and structural properties in each layer and integration of various bioactive molecules and therapeutic agents. Keywords Antibacterial activity · Bioactive agents · Electrospinning · Electrospun nanofibers · Multi-layered structures · Wound-healing

1 Introduction Skin is the main barrier separating the body from its surroundings, consisting of three different layers of varying thickness, namely epidermis, dermis, and hypodermis, which have different cellular and biochemical compositions [1]. Due to this multi-layered complex architecture, it exhibits anisotropic, heterogeneous, and multi-dimensional properties. Cutaneous wounds are generally defined as the deformations that occur in the biochemical, cellular, and anatomical integrity of the skin due to the damage caused by the physicochemical, pathophysiological, or thermal factors. These deformations can be superficial to affect the epithelial cells of the skin, or they can reach the muscle fascia and even to bone tissue. Cutaneous wounds are classified as acute or chronic according to the severity of deformation and duration of the healing time. While the healing process in acute wounds occurs with regular progression and minimal complications, deviations in the natural physiological course of the regeneration process are experienced in chronic wounds and the damaged tissue cannot accomplish anatomical and functional integrity by itself. According to data published by the World Health Organization (WHO), infected or chronic wounds affect approximately 14 million people worldwide each year [2]. Although there is a wide range of treatment options (surgical procedures,

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systemic or topical agent, hyperbaric oxygen, graft or flap, etc.) targeting these patient groups, wound management is unfortunately one of the few areas where a limited decrease in incidence has been observed [3]. This situation demonstrates that the development of new approaches for the treatment of scar tissue is of critical importance. Tissue engineering allows the production of ECM analog scaffolds that shape the physical form of dam aged tissue and guide new tissue formation. Up to date, cellular or non-cellular skin substitutes of different thicknesses (epidermal, dermal, dermoepidermal), produced with various engineering approaches (Integra®, Biobrane®, Alloderm™, Matriderm®, GammaGraft™, OrCel®, Transcyte™, Dermagraft®, Apligraf®, Hyalomatrix®, DenovoDerm™), have reached the market. However, these commercially available scaffolds have some specific disadvantages such as high cost and limited exudate absorption, microbial contamination, and permanent scar formation [4]. Hence, there is an urgent need for new scaffolds that can overcome these disadvantages and support regeneration. In this essence, fibrous structures produced by electrospinning have remarkable advantages such as ECM-like architecture, high surface area-to-volume area ratio and flexibility to provide a suitable platform for skin regeneration. Besides, owing to their nanoporous structures (0.1–10 μm), electrospun fibers can act as a barrier against the penetration of exogenous microorganisms without interfering with cellular respiration [5]. Nonetheless, single layer fibrous scaffolds/membranes have limited thickness for applications to deep and/or exuding wounds and cannot fully host the complex cell population of the skin in their structure. Hence, a uniform morphological structure cannot effectively provide the bio-physico-chemical ques required for damaged tissue reconstruction. Consequently, layer-by-layer approach provides the utmost advantages for the preparation of layered fibrous scaffolds/ membranes for wound healing applications. Hybrid structures produced by using different techniques in combination, offer the superior properties of each layer to be combined in a single scaffold [6]. In general, multi-layered scaffolds/membranes consist of a compact outer layer that acts as a barrier against a variety of exogenous factors and a looser, large-porous inner layer that promotes exudate absorption, infiltration of dermal cell groups into the structure and accumulation of granulation tissue. In addition, various therapeutic agents can be added to different layers of the scaffold to enable modulation and manipulation of the microenvironment in chronic wounds. This chapter begins with an overview of structure and function of skin followed by the basic biology of wound healing cascade. Types of polymeric materials used in the production of fibrous scaffolds for skin tissue regeneration and some therapeutic agents that will add bioactivity to the structure are summarized. Finally, layered hybrid scaffolds/membranes developed by electrospinning alone and/or in combination with different processing techniques are discussed.

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2 Skin Structure, Function and Wound Healing Process Skin, the largest organ of the body with a surface area of 1.6–2 m2, provides both physical and immunological (skin-associated lymphoid tissue (SALT)) defense against a wide variety of stress, including physical, chemical, and biological by forming a protective barrier between the living organism and external environment [7]. In addition to its barrier function, skin also has homeostatic and sensory roles in many different processes such as maintaining fluid and electrolyte balance, thermal regulation, regulation of blood pressure, energy storage, immune control, and vitamin D synthesis [8–11]. Besides its critical role in maintaining homeostasis, skin has a multi-layered (2–3 mm) anatomical structure consisting of epidermis, dermis, and hypodermis/subcutaneous tissue [11]. The epidermis is an impermeable outer layer that separates the internal environment of the body from the external environment. Beneath this layer is the dermis, which consists of a dense fibrous connective tissue system that allows the penetration of vascular tissues, lymphatic vessels, and nerve fibers, also providing flexibility and tensile strength to skin (4). Unlike the compact epidermis layer, the dermis exhibits a porous architecture of approximately 60–90% [12]. The hypodermis is the lowest layer of skin tissue and contains partially dermis-like cell types but contains larger blood vessels and nerve fibers compared to the dermis (8). Despite this unique multi-layered anatomical structure, skin is very sensitive to environmental damage due to its location on the outermost surface of the human body. Therefore, skin injury is one of the most common forms of wounds. Lesions that may occur in the skin tissue due to thermal, irradiation, physical, chemical factors, metabolic diseases (immune dysfunction, diabetes, obesity), genetic disorders (sickle cell disease) or pathophysiological conditions associated with aging is generally referred as “cutaneous wound” [13]. Such lesions may damage the barrier property of the skin and may cause dehydration, infection, pain, subcutaneous tissue damage (vessels, nerves, tendons, muscles, and bones), sepsis, gangrene, malignant transformation, amputation and even mortality depending on the severity/depth of the wound [14]. Cutaneous wounds are commonly divided into two categories as acute and chronic. Mechanical traumas, surgical wounds, or mild burns are “acute wounds” in which the skin regains its normal anatomical structure and function in a short period of time (1–12 weeks) without leaving any scars [15, 16]. In contrast, chronic wounds show slow (more than 12 weeks) or no healing due to changes in the biochemical components of the wound bed [17, 18]. Therefore, the healing time of a cutaneous wound and the biochemical state of the wound bed is closely related to acute or chronic state.

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Acute Wound Healing

When any defect occurs in the skin, the repair response kicks in, and four distinct stages, hemostasis, inflammation, proliferation, and remodeling, occur in a specific temporal order (partially overlapping each other). An acute wound in a healthy host heals quite rapidly with the continuous operation of these phases (Fig. 1). Hemostasis (immediate): It is the first response to dermal vascular damage and consists of vasoconstriction (vascular spasm), primary and secondary hemostasis stages [19]. Vasoconstriction, which is reflexive contraction, occurs in the first

Fig. 1 Schematic of wound healing cascade after an injury. Hemostasis is the first step in which platelet plug formation occurs followed by secretion of important cytokines, chemokines, growth, and coagulation factors. Formation of hemostatic clot protects the wound bed against microorganism invasion, functions as a cytokine and growth factor reservoir and provides a temporary scaffold for the migration of cells to the area. Hemostasis is followed by Inflammation in which the immune cells infiltrate into the injured area. The wound bed is cleared from necrotic cell debris, pathogenic microorganisms and foreign matter to provide a sterile microenvironment. After the resolution of this phase, Proliferation phase takes place in which the new granulation tissue formation and regeneration of damaged vascular network take place. Subsequently, Maturation and Remodeling of granulation tissue occur characterized by collagen reorganization and wound contraction

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minutes of hemostasis. During vasoconstriction, the extravascular blood loss in the region is limited for a while by providing reflex contracture of the damaged vascular smooth muscle [19, 20]. The primary hemostasis stage is characterized by the formation of the platelet plug. Platelets interacting with von Willebrand factor (VWF) become activated, undergo morphological changes and form a temporary platelet plug on the damaged endothelium [21]. Some important cytokines, chemokines, growth and coagulation factors that induce vasoconstriction and inflammatory response are secreted from the platelets in the aggregate structure ultimately forming the plug. In the secondary hemostasis stage, also called the coagulation cascade, insoluble fibrin fibers polymerize together with the platelet plug and form a mechanically stronger and more stable hemostatic clot (eschar formation) [22, 23]. When a sufficient clot is formed, excessive thrombosis is prevented by stopping the clotting process. Hemostatic clot also protects the wound bed against microorganism invasion, functions as a cytokine and growth factor reservoir for attracting skin and immune system cells to the area and provides a temporary scaffold for the migration of these cell groups [24]. Inflammation (days 4–6): It is the infiltration of immune cells into the damaged area to clear the wound bed of necrotic cell debris, pathogenic microorganisms, and foreign matter to provide a sterile microenvironment [25]. Various proinflammatory cytokines, chemokines, and growth factors are produced by the thrombocyte in the hemostatic clot or by the damage-related molecular motifs (DAMP), pathogenassociated molecular models (PAMP) and, chemotaxis of neutrophils is provided to the wound bed [26]. Neutrophils, as the first group of cells to reach the wound, release signal molecules that promote the increase of immune response, such as necrosis factor (TNF-α) and interleukin-1, -6, and -8 (IL-1, IL-6, and IL-8) [27, 28]. Additionally, these cells can destroy foreign materials by direct phagocytosis, production of antimicrobial agents (such as free radicals, antimicrobial proteins, caustic proteolytic enzymes) or chromatin and protease traps in wound bed (NETs) [29]. After neutralization of foreign materials, there is a decrease in neutrophil count and activity (apoptosis or necrosis) due to the decrease in the signal coming from the pathways activated by microbial contamination. Depending on the release of phosphatidylserine (DAMP) on the surfaces of neutrophil residues in the wound bed, adaptive immune system cells produce signal molecules such as interferon-gamma (IFNγ) and TNF-α. This allows monocytes to be drawn from the peripheral blood to the wound area [30]. Monocytes differentiate into M1 macrophages (pro-inflammatory) in response to the local environment and clean the wound area by efferocytosis of neutrophil residues. With wound bed cleansing, there is a decrease in the levels of some proinflammatory cytokines such as IL-23 and IL-7A. Accordingly, the release of more neutrophils from the bone marrow to the wound area is prevented (negative feedback) [30, 31]. After the elimination of cellular and tissue fragments, the simulators are extinguished. Subsequently, polarization is provided from the M1 phenotype to heterogeneous M2-macrophage phenotype (anti-inflammatory) [32, 33]. Proliferation (day 4-week 2): Proliferation is a constructive repair phase characterized by the “simultaneous” formation of new granulation tissue and regeneration

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of the damaged vascular network in the debrided wound area. After the resolution of inflammation, fibroblasts that are responsible for the synthesis of new connective tissue proteins begin to migrate and proliferate from the incision edges of the wound via signal molecules (platelet-derived growth factor (PDGF), insulin-like growth factor (IGF-1), transforming growth factor-β (TGF-β), fibroblast growth factor (FGF)) released by platelets and macrophages [34]. With the progression of proliferation and accumulation of hyaluronan, proteoglycans, fibronectins, and collagen (predominantly type III but less type I collagen), the temporary hemostatic clot/ matrix is replaced by newly synthesized granulation tissue ( fibroplasia) [29]. New blood vessels are formed from pre-existing vessels (angiogenesis) to ensure proper blood flow in the granulation tissue, which is relatively avascular and is fed only by the diffusion from undamaged capillaries around the wound bed. Toward the last stages of the proliferative phase, keratinocytes migrate along the surface of the granulation tissue with the effect of keratinocyte growth factor (KGF) expressed by fibroblasts and initiate regeneration of the basement membrane between the epidermis and dermis [35–37]. The differentiation of wound fibroblasts into contractile myofibroblasts is induced by TGF-β1 released from M2 macrophages. With the expression of α-smooth muscle actin (α-SMA), myofibroblasts reduce the wound size (contraction) by pulling the borders of the lesion toward the center. Thus, the distance to be covered by the migrating keratinocytes in the granular matrix is reduced and re-epithelialization is facilitated [38]. Maturation and Remodeling (week 3-year 2): Maturation of granulation tissue is the process of increasing its mechanical strength and is characterized by collagen re-organization and wound contraction [39]. This phase begins as soon as the granulation tissue is present and therefore overlaps with the proliferation phase. The granulation tissue accumulating in the wound area exhibits higher content of collagen type III, hyaluronan, and fibronectin. Therefore, it is weaker than the intact/ undamaged tissue in terms of matrix composition and organization [40]. Type III collagen, which has a soft and gelatinous structure during the restructuring process, is degraded and removed by matrix metalloproteinases (MMPs) [1]. Synthesis of stronger and firmer type I collagen and active remodeling of connective tissue is accomplished by resident cells of the wound bed [41]. Type I collagen fibers expressed and deposited by these cell groups are cross-linked via lysyl oxidase for increased strength and organization [40]. The maturation of connective tissue takes place under the control of a wide variety of anti-inflammatory growth factors such as TGF-β1 and FGF [42]. The second characteristic feature of this stage is that, after the connective tissue is sufficiently perfused and restored there is a decrease in the number of cells through apoptosis and this increases the tensile strength of the maturing connective tissue [43]. It should be noted that hair follicles, sweat glands and nerves do not completely heal and at the end of the maturation and remodeling phase. The skin tissue can retain only 80% of its original strength [44, 45].

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Chronic Wounds (Longer Than 12 Weeks)

Generally, chronic wounds develop as a symptom of an underlying condition or comorbidity such as malnutrition, pressure, vascular insufficiency, infection, edema, sickle cell disease, and diabetes [46]. Such factors trigger the development of chronic wounds by causing deviations in the natural physiological progress of the acute wound healing process. Chronic wounds can be commonly categorized as diabetic, venous, arterial (ischemia) and pressure ulcers [47]. Although they show differences in etiology, all chronic wounds show plateau or stall in an extremely inflammatory state depending on the disruption of the delicate balance between pro- and anti-inflammatory signals [48] (Fig. 2). This uniformity is due to the constant components of the multifactorial pathogenesis of chronic wound such as local tissue hypoxia, high microbial colonization, necrotic burden, recurrent ischemia and cellular senescence [49]. Low oxygen levels (about 5 mmHg), moisture from wound exudate and necrotic cell debris facilitate microorganism colonization in the form of planktonic or biofilm [46]. Microorganisms cause a destructive microenvironment by producing proteases and exotoxins that degrade proteins important for wound repair [44]. Depending on the increased MMP expression in chronic wounds, growth factors and their receptors that are important in the transition to the proliferation phase are degraded and this causes low mitogenic activity and senescence of epithelial tissue cells [49–51]. The impaired functions of immune system cells also prevent the resolution of inflammation, despite an increased infiltration into the wound bed. All these events lead to chronic degradation of ECM components and connective tissue, resulting in

Fig. 2 Schematic depiction of the timeline for (a) acute wound and (b) chronic wound site during the healing process

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complete loss of epidermis, dermis, subcutaneous fat, and even damage of lower extremities [52]. Because of the severity of chronic wounds, it takes years for them to regain their anatomical and functional integrity, or they never heal. Therefore, therapeutic strategies that will resolve the inflammation stage and progress the chronic wound to the proliferation-remodeling stage are of great importance.

3 Polymers as Wound-Healing Materials For the effective and rapid treatment of acute or chronic wounds, an ideal dressing is essential to create a favorable environment to provide protection and moisture for wound bed, remove excess exudate and prevent infection. Several polymeric materials have attracted considerable attention to be used as artificial dressings. In comparison to traditional ones, polymeric wound dressings have remarkable advantages like moist retention, mechanical compliance, carrier for growth factors/antimicrobial agents, ease of application, inherent antimicrobial properties and efficient contact with wound-site for proper healing. They can be prepared in various forms such as films, foams, hydrogels, and scaffolds. Recently, micro-nano fiber forms prepared by electrospinning and/or a combination of various forms in layered structure stand-out to mimic the wound site. Polymeric materials used in the preparation of wound dressings can be classified as; natural and synthetic polymers [47, 53] (Fig. 3).

3.1

Natural Polymers

Natural polymers are attractive in wound dressings due to their non-toxic structure, biocompatibility and biodegradability features, as well as their unique advantages such as inherent antimicrobial capabilities.

3.1.1

Alginate

Alginate is a heteropolysaccharide made from two uronic acid linear co-polymer of -D-Mannuronic acid and -L-Glucuronic acid that is derived from a variety of brown seaweed species. To formulate alginate fibrous membranes, the ability of uronic acids to bind metal ions such as sodium and calcium is critical. Calcium alginate form is insoluble in water and can be woven into fibers and formulated as a patch, or it can be used as a loose fiber ribbon or robe that can be used to fill wound cavities due to their highly absorbent character and structural integrity. The polysaccharide basis promotes wound-healing by reducing the inflammatory phase while combination with antibacterial and enzymatic components can help to remove necrotic tissue and microbial bodies.

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Fig. 3 Natural or synthetic polymers can be used alone or in combination for the preparation of wound dressings. They possess remarkable advantages like moist retention, mechanical compliance, ease of application, inherent antimicrobial properties, carrier for bioactive agents, and efficient contact with the wound site for proper healing

3.1.2

Chitosan

Chitosan is a deacetylated form of chitin that is found in crustacean exoskeletons and some fungi. It is a linear polysaccharide made up of the irregular distribution of D-Glucosamine and N-Acetyl-D-Glucosamine units. Most of the physical properties of chitosan depend on the degree of acetylation that is inversely related to solubility, viscosity, and biodegradability. Being one of the most abundant biopolymer in nature after cellulose, chitosan possesses many properties that are advantageous for wound-healing. Besides its enhanced microbial efficacy against bacterial and fungal pathogens, good water-absorbent properties make it an ideal candidate for wound dressings.

3.1.3

Collagen

Collagen is a glycoprotein and it is the main constituent of connective tissue. Even though there are 29 different forms of collagen, only a few of them are used in the development of collagen-based biomaterials. For bioengineered skin substitutes, collagen can be employed in numerous ways. When it comes to treating ulcers, Apligraf® (Organogenesis, Canton, MA, USA) is one of the most utilized products. Collagen is a useful natural polymer to heal and replace the defeated extracellular matrix structure in chronic wounds.

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Gelatin

Gelatin is a collagen-derived product with interesting medical applications. It is possible to make gelatin by partially denaturalizing collagen from connective tissues, skin, or bone. Type A or Type B gelatins can be formed by acid-base or alkalibased hydrolysis, respectively by treating these raw materials with either acid or alkali. A variety of dressings, including microspheres, sponge-like microspheres, and electrospun fibrous membranes, have been used for cutaneous tissue applications and burn wound treatment. DuoDerm®/Granuflex® (ConvaTec) is a hydrocolloid dressing that contains pectin and carboxymethylcellulose in addition to gelatin that is also used as a gelatin-based hemostatic sponge. The main advantage of gelatin in wound dressings is the beneficial effect in re-epithelization.

3.1.5

Hyaluronic Acid

Hyaluronic acid is composed of glucuronic acid and acetyl-D-glucosamine units. It is found in the vitreous humor, synovial fluid, articular cartilage, and the skin’s dermis and epidermis. As one of the main elements in ECM, hyaluronic acid displays several advantages in tissue engineering applications. In the healing process of acute and chronic wounds, hyaluronic acid plays multiple roles as promoting early inflammation, increasing cell infiltration, enhancing granulation tissue formation and cell migration. It does not cause an immune response and it is angiogenic. The good absorbent properties allow the formation of a hydrated gel-like structure.

3.1.6

Silk Fibroin

Fabric and stitching have long relied on silk fibroin, a protein produced by bees, spiders, lacewings, and silkworms with a complicated structure. It is particularly useful in biomedicine because of its remarkable mechanical qualities, biocompatibility, biodegradability, flexibility, water vapor permeability, and minor antibacterial capabilities. Silk is often incorporated with other polymers like alginate and prepared in the form of nanofibers by using electrospinning. Keratinocyte proliferation and migration are promoted in silk-based wound dressings, improving wound closure.

3.2

Synthetic Polymers

Synthetic polymers have the advantage of homogeneous physicochemical properties when compared to natural polymers. They are usually mechanically stable, degrade in a controlled manner and biologically inert. However, they do not possess inherent therapeutic properties and may present toxicity risk.

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Poly(Lactide-Co-Glycolide) (PLGA)

It is the copolymer of polylactic acid (PLA) and polyglycolic acid (PGA) approved by FDA to be used in biomedical field. The degradation rate and mechanical properties can be controlled by adjusting the monomer ratios. It is a biocompatible structure and usually combined with other synthetic or natural polymers to prepare fibrous membranes by electrospinning.

3.2.2

Polyethylene Glycol (PEG)

Polyethylene glycol is a stretchable ether-based polymer that is hydrophilic, biocompatible, and nonimmunogenic. Because of these characteristics, it is a popular synthetic material for wound dressings. PEG’s crystallinity, thermal and mechanical properties are improved by combining the polymer with other synthetic and natural polymers. These PEG-based membranes/scaffolds have been used to treat ulcerative diabetic scars and known to promote healing by stimulating collagen production as well as ECM deposition.

3.2.3

Polyurethane (PU)

The organic units of polyurethane polymer are linked together by carbamate bonds. Semi-permeable structures prepared from PU help to maintain a moist environment [54]. The good barrier properties and oxygen permeability make PU a favorable choice in wound dressing applications [55]. However, it is a soft and hydrophobic polymer that obstructs the cell interaction, removal of the exudate and ease of application. To overcome these problems, PU fibrous membranes are usually prepared in combination with other synthetic or natural polymers for improved cell interaction and hydrophilicity [56].

3.2.4

Polyvinylpyrrolidone (PVP)

Using both bulk and solution polymerization, polyvinylpyrrolidone (PVP) is synthesized from the hydrophilic monomer vinylpyrrolidone (NVP). PVP is soluble in water and most organic solvents and has a low toxicity. It has been widely used as a wound dressing biomaterial due to its outstanding biocompatibility. In vitro studies have shown that the PVP electrospun fibrous membranes incorporating antibacterial agents like curcumin or silver nanoparticles have remarkable antimicrobial activity against S. aureus and Candida albicans.

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Polycaprolactone (PCL)

Polycaprolactone is a linear aliphatic polyester comprised of repeated units of hexanoate. PCL has been widely investigated in preparation of membranes/scaffolds in wound healing due to its remarkable mechanical properties and biodegradable structure. It has a hydrophobic structure with excellent blend compatibility with other polymers and can be easily processed. PCL has been used in wound dressings in fibrous structure usually prepared by electrospinning for the effective delivery of growth factors, essential oils, antimicrobial agents alone or in combination with other synthetic or natural polymers. Numerous patents have been reported for fibrous PCL-based wound dressings emphasizing the importance of PCL in the avenue of wound-healing [57].

4 Cutaneous Scaffolds with Added Therapeutic Agents for Wound Treatment During delayed healing process, the cellular organization and the endogenous molecules produced by these cells become insufficient in terms of type, function or quantity for a successful regeneration process. Localized applications of exogenous bioactive/therapeutic agents such as growth factors, cytokines, antibiotics, antimicrobial peptides, various phytotherapeutics, metal ions or nanoparticles are frequently used methods to assist the modulation and regeneration process of the damaged microenvironment [58]. Considering the morbidity status and different needs of the chronic wound, the best strategy in the treatment process is multifunctional approaches that require a combination of bioactive agents that can simultaneously target the various pathways involved in the healing response. However, most of these therapeutic agents have disadvantages such as short pharmacokinetic effect, poor solubility and cytotoxicity that leads to further complications in wound care [59]. Various methods such as encapsulation and chemical conjugation are used to increase the bioavailability of therapeutic agents. In this way, they can be included in wound dressing materials prepared using tissue engineering approach [17]. In fact, it is possible to develop double barrier-controlled release systems that require the integration of nanotechnology. All these approaches provide biophysical and biochemical cues for regenerative cells that provide a reconstruction of damaged tissue, as well as obtaining biomimetic scaffolds possessing the characteristics of ECM (Fig. 4). In this section, various bioactive and therapeutic agents with pharmaceutical, medical, biotechnological potential for applications in wound-healing/skin tissue regeneration, are summarized.

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Fig. 4 Bioactive and therapeutic agents involved in wound/healing process for enhanced healing and tissue regeneration

4.1

Growth Factors (GFs)

GFs are endogenous signaling molecules that are expressed in various skin and inflammatory cell groups in response to tissue damage. They trigger and coordinate numerous cellular/molecular events in each phase of the wound-healing process [60]. The roles and functions of some important endogenous growth factors in wound healing and remodeling have been discussed in Table 1. Owing to these important functions, changes in the expression or activity of GFs may adversely affect the normal wound-healing process [61]. Application of externally localized GFs is suggested as a promising approach to re-organize the levels of these signaling molecules, modulate impaired cell functions, and regenerate damaged skin tissue [62]. Although many of these growth factors have been clinically tested to treat chronic or degenerative wounds (topical or intralesional injection), the application of externally localized GFs has many limitations including, lack of penetration due to exudate in the wound bed, degradation due to intense protease activity, loss of activation due to short-term half-life, limited diffusion, low absorption into the deeper layers of the skin, and supra-physiological dose-related cytotoxicity [63]. These emerging side effects clearly demonstrate the importance of spatio-temporal control in exogenous GF treatments. Therefore, delivery systems that control the stability and release of GFs in wound bed are key to safe, effective and successful treatment. In this sense, the most common polymeric scaffolds used

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Table 1 Important GFs involved in the wound-healing cascade Growth factor PDGF

Cell source Platelets Macrophages Endothelial cells Fibroblasts

Target and effect Neutrophil and macrophages: chemotaxis Fibroblasts: proliferation and differentiation Other growth factors: up-regulation such as TGF-β Fibroblasts: proliferation and migration Endothelial cells: proliferation Keratinocytes: proliferation, migration, differentiation, and inhibition of apoptosis that induced by ROS

FGF-2 FGF-7 FGF-10

Fibroblasts Keratinocytes Endothelial cells

EGF

Platelets Macrophages Fibroblasts

Keratinocytes: proliferation

HBEGF

Keratinocytes Macrophages

Keratinocytes: proliferation and migration

TGF-α

Platelets Lymphocytes Macrophages Keratinocytes Fibroblasts Macrophages, epidermal cells

Keratinocytes: pro-motility

TGF β1 TGF - β2 TGF β3

Platelets Macrophages Fibroblasts Keratinocytes

Macrophages: chemotaxis Fibroblasts: migration, proliferation and differentiation Keratinocytes: migration

IGF-I IGF-II

Liver’s cell Neutrophils Macrophages Fibroblasts

Fibroblast, keratinocyte, and endothelial cells: activation and proliferation

VEGF

Endothelial cells: chemotaxis and proliferation Other growth factors: up-regulation

Function in wound healing Initiation of the inflammatory response [69] Synthesis of granulation tissue and contraction of wound [70–72] Re-epithelialization [73] ECM synthesis and remodeling Wound vascularization [74] Re-epithelialization [74– 76] ROS detoxification [77] Synthesis of granulation tissue and re-epithelialization [78– 80] Collagen deposition wound contraction re-epithelialization [81– 83] Re-epithelialization (early role) [84, 85]

Angiogenesis Lymphangiogenesis [86, 87] Enhancing of granulation tissue [87, 88] Inflammation (early stage) [89] Regulation of M1/M2 polarization [90] ECM synthesis and remodeling [90–93] angiogenesis [90, 94] Fibrosis, tensile strength, re-epithelialization (latestage) [91, 95–98] Anti-scarring [99] ECM synthesis and collagen remodeling, angiogenesis [70] (continued)

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Table 1 (continued) Growth factor CTGF

Cell source Fibroblasts

Target and effect Fibroblast: proliferation, differentiation Endothelial cells: proliferation, migration, and differentiation

Function in wound healing ECM synthesis, re-epithelialization, angiogenesis [100–104]

PDGF platelet-derived growth factor, FGF fibroblast growth factor, HB-EGF heparin-binding EGF-like growth factor, TGF-α transforming growth factor-α, VEGF vascular endothelial growth factor, TGF-β transforming growth factor-β, IGF insulin-like growth factor, CTGF connective tissue growth factor, ROS reactive oxygen species, ECM extracellular matrix

in the literature are electrospun nanofibers, hydrogels, sponges, and cryogel-based structures [64–68]. If the manufacturing techniques used to obtain these structures are chosen in a way that preserves the activity of GFs during or after the process, it is possible to both limit the disadvantages arising from the direct application of exogenous GFs and produce biomimetic scaffolds similar to the chemical composition of the skin ECM. In general, GFs can be incorporated into the structure by blending directly into polymeric solutions during the fabrication of tissue scaffolds [103, 105–109], by post-production surface modification with various functional groups [110–116], or by loading/conjugating to a secondary carrier system such as micro/nanoparticles (block or hybrid systems) [117–120]. Each method has its own advantages and disadvantages. While the direct blending of GFs with the polymeric solution used in the production of scaffolds can cause structural modifications or burst releases [68, 111, 121], the techniques based on surface modification have limitations for a controlled release [110, 122, 123]. Block systems produced by integrating secondary carrier and release systems into the scaffold structure can be a useful approach in maintaining the bioactivity of GFs and providing a long/effective release profile. Therefore, in the production of GFs-containing biomimetic scaffolds, care should be taken to select a method, (1) that does not cause structural changes (denaturing, misfolding, aggregation) that may compromise the tertiary structure of these molecules, (2) that provides protection against intense protease activity in the wound bed, (3) that offers a prolonged/effective release profile.

4.2

Antibiotics

Antibiotics are antimicrobial agents that show their inhibitory activities via four different mechanisms: inhibition of (1) protein, (2) cell wall, (3) nucleic acid synthesis, and (4) suppression of metabolic pathways of bacteria [124]. However, systemic or topical long-term and high-dose antibiotic use may cause adverse effects such as cytotoxic effect, development of antibiotic resistance, and loss of beneficial microbial flora. Antibiotics of different spectra can be incorporated into electrospun fibrous wound dressings (mostly by direct blending) to limit their side effects and

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increase bioavailability [125–131]. It has been reported that these antimicrobial dressings show a lower cytotoxic effect on skin cells, higher antibacterial activity, and skin regeneration compared to free antibiotics, due to partially controlled and prolonged drug release. The burst release that may occur at the beginning can be controlled by modulation of the hydrophilic or hydrophobic polymer ratios in the fibrous scaffold composition [132], as well as by integrating a secondary release system into the structure.

4.3

Natural Substances

Natural substances are attractive alternatives for the treatment of skin lesions to overcome both the possible side effects of antibiotics and antibiotic resistance [133]. In addition, natural ingredients such as essential oils, phenol, polyphenol, or phytochemicals offer a versatile therapeutic strategy that combines antimicrobial, antioxidant, and anti-inflammatory properties. However, these structures have some disadvantages that limit their bioavailability, such as high volatility, low solubility due to hydrophobic nature, and sensitivity to environmental stimuli (light, heat, and humidity). Therefore, current trends point out the combination of natural therapeutic agents with engineered scaffolds [134, 135]. Although the hydrophobic nature and high volatility of these phytochemicals make polymeric scaffold fabrication somewhat difficult, they have been successfully integrated into electrospun [136–138], wet spun [139] and hydrogel-based wound dressing structures [140, 141]. Incorporating these bioactive agents into polymers with both hydrophilic and hydrophobic properties or modifying it with structures such as cyclodextrin (CD) are wellestablished methods used to overcome the dispersion problems [142–144]. These emerging hybrid scaffolds not only increase the stability and therapeutic efficacy of natural ingredients but also provide an ideal bioactive wound dressing for tissue regeneration.

4.4

Antimicrobial Peptides (AMP)

AMPs (LL-37, lactoferrin, papiliocin, dermaseptin, sapecin B, bufforin II, etc.) are molecules that are important part of innate immunity and generally have a cationic nature [145, 146]. AMPs show very potent inhibitory activity against a wide variety of pathogens, including viruses, parasites, fungi and antibiotic-resistant bacteria. The basic mechanism of antibacterial activity is the electrostatic interaction between AMPs and the negatively charged phospholipid layer of the bacteria that causes infection [147]. This interaction disrupts the permeability barrier of the bacterial membrane that leads to subsequent cell lysis. In addition to their antibacterial effects, AMPs support wound repair by different mechanisms such as immunomodulation (inhibits inflammatory cytokine production), induction of epithelial cells

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proliferation, and differentiation [148]. However, the stability of these peptides in vivo is very low and requires very frequent dosing at high concentrations for desired therapeutic efficacy, which can cause toxic and hemolytic effects [149, 150]. To overcome this drawback and increase their clinical utility for the treatment of infected wounds, AMPs have been integrated into various wound dressing materials such as pressurized gyration-produced fibers [151], electrospun fibers [152–158], and hydrogels [149, 159, 160]. Nevertheless, it should be noted that these peptide structures are non-stable in organic solvents and irregular in aqueous solutions, which makes them difficult to use in the production of antimicrobial scaffolds [161]. Surface modifications of scaffolds and the use of secondary carrier/release systems stand out as useful pharmaceutical formulations to overcome problems such as loss of activity, aggregation, and cytotoxic effect.

4.5

Metal Nanoparticles (MNPs)

Metal nanoparticles (MNPs) are attractive alternatives that are often used to reduce surveillance of unnecessary antibiotic consumption, prevent infection, and improve sanitation. Metal NPs target multiple microbial mechanisms thanks to their properties such as low particle sizes, high stability, surface energy, and photocatalytic activity [162, 163]. These properties cause microbial death by increasing oxidative stress or cell permeability and by changing various cellular structures (DNA, protein, enzyme) and metabolic pathways (respiratory chain, protein synthesis), accordingly [131, 164]. However, metal ions released from metal nanoparticles can cause adverse effects such as allergic reactions, irritation, argyria, cytotoxic activity on mammalian cells. For this reason, literature reports focus on MNPs that can release metal ions in a long period of time and at lower doses. Most of these reports have focused on selenium (Se), silver (Ag), zinc (Zn), and copper (Cu) metal NPs to impart antibacterial function to cutaneous wound dressing materials [165]. In addition to the antimicrobial effect, some MNPs, also promote skin tissue regeneration by showing an anti-inflammatory effect (Zn), promoting cell proliferation (Se and Cu), and acting as a co-factor (Zn) for lysyl oxidase, an important enzyme of the remodeling phase. Despite the broad-spectrum bactericidal activity and various other therapeutic effects offered by metal ions or metallic nanoparticles, they can easily cross various biological barriers in the human body and damage the normal biochemical processes of cells [166]. In addition, the aggregated nanoparticles formed in biological systems can cause cellular toxicity. Integration of MNPs into scaffolds in various forms designed for the treatment of infected and/or chronic wounds can allow controlled release of these structures, thereby limiting the side effects [167– 171]. These integrated systems also allow the design of high-performance hemostatic dressing materials targeting traumatic, surgical, or deep pathological wounds.

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Metal-Organic Frameworks (MOFs)

MOFs are a class of nanocrystalline hybrid materials formed by the bonding of metal ions and multitopic organic ligands to form a two or three-dimensional structure. Compared to metal nanoparticles, MOFs offer unique structural diversity, a controlled degradation in response to endogenous and exogenous stimuli, and on-demand bioactivity. MOFs also stand out as carriers and controlled release systems for a wide variety of bioactive agents, due to their high surface area, adjustable pore diameters, and modifiable frame structures. MOFs not only function as drug release systems but also create an ionic microenvironment in the wound bed for many therapeutic effects providing antimicrobial, antioxidant, and antiinflammatory activity through endogenous metals such as copper [172, 173], zinc [174], and cobalt [175]. Similarly, the use of natural bioactive molecules as structural blocks during synthesis may confer additional biological properties to MOFs. Although direct use of MOF constructs with or without bioactive agents for the treatment of infected chronic wounds is a convenience, MOF constructs are not able to absorb the dense exudate of the wound bed and provide a temporary extracellular matrix. In addition, these nanocrystal structures in powder or bulk form, which come into direct contact with the wound surface, can further trigger inflammation and thus inhibit the regeneration process [174, 176]. Therefore, integrating MOFs into scaffolds designed using natural or synthetic polymers provides useful properties that are difficult to achieve with a single component. In this context, improved MOF-polymer hybrid wound dressing systems, by incorporating various MOFs such as ZIF-8, ZIF-67, Cu-BTC, HKUST-1, UiO-66, Ag2[HBTC][im] into electrospun [174–179], film [180], hydrogel [181, 182], or hydrogel fiber [183] structures have been reported. These hybrid scaffolds provide a dual cooperative controllable release system that provides sustained release of metal ions and other bioactive agents through the barrier effect of both the MOF structures and the molecular chain in the polymeric structures.

5 The Conceptual of Design of Layered Scaffolds/Membranes for Wound Healing Despite many clinical approaches (Integra®, Biobrane®, Aalloderm™, Matriderm®, GammaGraft™, OrCel®, Transcyte™, Dermagraft®, Apligraf®, Hyalomatrix®, DenovoDerm™) in practice and technological advancements in bioengineered acellular or cellular skin constructs [53, 184, 185] the limited success rate still encourages researchers to strive against the impediments and challenges. Although growth factor-based therapies, gene therapies, antibiotic therapies for infection management, and application of natural substances with healing capacity have been widely explored, the degradation, toxicity, and instability of these bioactive substances due to improper delivery to the wound site remains to be a critical problem. To aid the

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wound-healing process and repair of damaged tissue, fabrication of scaffolds/membranes from various materials as wound dressings has been a promising strategy. These scaffolds/membranes would not only act as a substrate mimicking the structural functionality of the natural ECM but also serve as a carrier for bioactive agents that play a major role in the wound-healing process. There are a number of key factors directing the design and fabrication of scaffolds/membranes in wound healing strategy. Besides supporting the delivery of bioactive agents like antibiotics and growth factors, these fabricated structures are expected to be biocompatible and biodegradable with a proper degradation profile concomitant to the wound-healing period. An ideal wound dressing is expected to act as a shield against external contamination to prevent infection, provide the moist environment, be flexible and permeable to air, hold an appropriate drainage capacity and absorb wound exudate, permit the flow of cell nutrients and products, guide cellular functions by exerting mechanical and biological stimuli and mimic the native ECM to stimulate cell attachment and migration. Hence, an effective wound dressing or skin graft should act as a functional bioactive matrix. Various types of platforms have been designed and fabricated to assist woundhealing process in forms of hydrocolloids, hydrogels, films, foams, membranes using natural, synthetic and composite polymers. Different fabrication techniques can be applied-either conventional or advanced-for producing wound dressing matrices. Considering the complexity of the wound healing process and challenging requirements to fulfill the clinical need in terms of both healing and regeneration, single-layer scaffolds/membranes would be a rather poor approach due to their individual characteristics. Saliently, multi-layered scaffolds/membranes would provide a more effective strategy for the design and development of functional woundhealing matrices. This strategy would enable the use of different materials with different physicochemical properties in each layer. Besides, each layer can be prepared in a different structure that can provide several advantages. As mentioned earlier, the epidermis layer of the skin tissue has a lower cell content than the dermis, and barrier lipids prevent dehydration while acting as a shield against exogenous pathogen penetration through antimicrobial peptides and lipids [186, 187]. In this more compact superficial layer (75–600 μm thick), there is no ECM accumulation and no blood vessels [188]. The dermis is thicker (1.5–2.0 mm) compared to the epidermal layer and is surrounded by the ECM, the complex three-dimensional structure that supports the neural network, blood and lymphatic system, immune cells, and fibroblasts [189]. Due to this special structure, the regeneration of the dermis is more complex and less efficient than the surface layer [104]. Although monolayer dermal scaffolds offer improved clinical outcomes, the lack of the epidermis layer makes the wound bed susceptible to microbial contamination. These scaffolds may require secondary dressing or surgical intervention as they do not support simultaneous regeneration of different skin layers. In this sense, multilayered structures developed with different production techniques can successfully simulate the natural structural architecture of the skin to accelerate the regeneration process of a full-thickness wound [190]. Typically, layered skin scaffolds consist of a thin upper layer similar to the compact structure of the epidermis and a porous and

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less dense lower layer similar to the structure of the dermal ECM. The top layer can prevent the risk of wound infection by preventing foreign matter penetration and dehydration with its low porosity and dense structure that does not compromise gas exchange [191, 192]. This fibrous layer is generally prepared by electrospinning of hydrophobic polymers. The bottom layer with its high/large porosity and loose structure can promote the transport of various nutrients and metabolites, the absorption of excess exudate, cell adhesion, infiltration, proliferation and migration, vascularization, and accumulation of granulation tissue [193, 194]. In addition, three-layered artificial skin tissue samples in which the basement membrane or hypodermis between the epidermis and dermis are mimicked have also been reported in the literature [195, 196]. Although these multi-layer scaffolds are attractive structures reflecting the natural structural architecture of skin tissue, the cellular organization of non-healing wounds and the endogenous molecules produced by these cells are not sufficient in terms of type, function, or quantity for a healthy regeneration process. Integration of various bioactive molecules into tissue scaffolds makes it possible to obtain biomimetic substrates with some advantageous properties of natural ECM and to provide mechanical support as well as biochemical support to the damaged tissue. In this sense, antimicrobial properties can be gained by loading various biocidal agents to reflect the characteristics of the natural epidermis on the upper layer, while the regeneration process can be supported by integrating GFs into the dermal layer [197, 198].

6 Processing Techniques for Fibrous/Layered Wound Dressings Fabrication and processing techniques have crucial impact on scaffold/membrane properties. Conventionally, solvent casting/particulate leaching, freeze-drying, molding, gas foaming, thermogelling, and supercritical processing techniques have been applied to produce scaffolds in a porous structure including hydrogels, hydrocolloids, films, foams, and sponges to support the physiological environment of the wound tissue. Among the different structures prepared by various techniques, an increased interest has been directed toward scaffolds/membranes possessing fibrous structures prepared by electrospinning. Micro/nanofibers have been formulated and fibrous matrices have been used either alone or in combination with other morphologies to form layered fibrous scaffold/membrane architecture. In preparation of symmetric or asymmetric layered fibrous membranes/scaffolds, one of the most challenging problem is the delamination of the layers in in vitro and/or in vivo conditions. The overcome this problem, co-solvent systems shall be used while preparing the electrospinning solution for different types of polymers [199]. Moreover, the layered structures can be inter-crosslinked following the electrospinning process or self-adhesive polymers like gelatin can be applied between the layers during the construction of the layered structures.

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Multi-layered Electrospun Membranes

Electrospinning has become one of the most cost-effective, simple, and flexible fabrication techniques for the production of ultrathin polymer fibers with diversified properties. The versatility of the technique enables the control over fiber production and fiber shape and structure can be tuned (random, aligned, or core-shell) by manipulating the process parameters. Electrospinning set-up requires relatively less capital, and the operation is relatively simple. A wide variety of polymers synthetic, natural or a combination of both can be used which makes the technique flexible. Eminently, the nano-sized fibers generated by electrospinning provide essential advantages for the production of wound dressing membranes/mats. As the technique is compatible with both natural and synthetic polymers, multifunctional wound dressings with different physicochemical and biological properties can be prepared. Additionally, bioactive and/or therapeutic agents can be incorporated in electrospun nanofibers in several ways such as blend electrospinning, co-axial electrospinning, emulsion electrospinning or grafted on nanofibers post-electrospinning by non-covalent or covalent immobilization techniques (Fig. 5). Electrospinning technique has a prominent advantage for the development of multi-layered fibrous membranes, composed of different layers. Each layer can be prepared in a specific character to reproduce skin anatomy for enhanced wound healing. Multi-layered membranes can be bilayer or more than two layers. Asymmetric membranes can be accomplished by sequential electrospinning of different polymeric solutions in a layer-by-layer assembly (Fig. 6a). The layered fibrous structures are usually designed in a way that mimic the epidermis and dermis layers of the skin. In this approach, dense and porous layers can be acquired by utilizing different solutions and process parameters. The dense layer with smaller-sized hydrophobic nanofibers represents the epidermis which would avoid fast dehydration and protect the wound site against bacterial invasion. The porous interconnected hydrophilic nanofibers mimic the dermis with high absorption capacity for fluid exchange and promote cell adhesion and proliferation for regeneration of healthy tissue. The sequential electrospinning would allow the use of different types of polymers to accomplish different morphology and character on each layer. Alves et al. [191], used this approach to assemble a bilayer asymmetric membrane. In their design, the protective layer was composed of a blend of synthetic biodegradable polymers–polycaprolactone and polylactic acid-to act as a cover to prevent microbial contamination and dehydration. The hydrophilic bottom layer was designed to promote cell adhesion and proliferation and high porosity. It was prepared by electrospinning a bioactive blend composed of methacrylated gelatin/methacrylated chitosan. In vitro studies revealed that the bilayer membranes promoted fibroblast adhesion and proliferation and showed hemocompatibility that is essential for controlling the hemostatic phase of healing process [191]. The use of natural polymers such as chitosan runs several advantages such as, biocompatibility, antibacterial effects, non-toxicity, and bioactivity. However, the disadvantages in

Fig. 5 A variety of approaches for incorporating bioactive and/or therapeutic agents into electrospun nanofibers. (a) Blending the bioactive agents with polymer solution during electrospinning, (b) Core-Shell electrospinning, (c) Electrospinning with an emulsion-based polymer solution with added bioactive agents, (d) Covalent or non-covalent surface immobilization after electrospinning

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Fig. 6 Different approaches for the production of fibrous/layered wound dressings. (a) Multi-layered Electrospun Membranes. Layered fibrous membranes are prepared by sequential electrospinning of each layer to construct asymmetric membranes. (b) Combining porous 3-D scaffolds with nanofibrous membranes to produce layered/fibrous structures. Solvent casting-particulate leaching, wet spinning, or freeze-drying processes can be combined with electrospinning for production of multi-layered scaffolds with fibrous layer. (c) Combining hydrogels with nanofiber mats. Fiber-hydrogel composites can be used by combining a variety of methods, including electrospinning and electrospraying in tandem, as well as direct polymerization of hydrogels on nanofiber matrices produced by electrospinning. (d) Combining 3D-printed structures with electrospun fibrous membranes. Nanofibers can be incorporated in 3D-printing solution for layer-bylayer production of scaffolds incorporating fibrous structures or nanofibers can be electrospun on 3D-printed scaffolds for preparation of layered fibrous scaffolds

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spinnability and low mechanical strength impair the use of chitosan alone to produce fibrous structures. To overcome this drawback chitosan is usually blended with biodegradable synthetic polymers to achieve fibrous structures by electrospinning. A bilayer chitosan-polycaprolactone-hyaluronic acid (CS-PCL-HA) was prepared by Chanda et al. [200], to obtain a mechanically stable and cytocompatible structure. Adequate fiber diameters of CS-PCL-HA scaffold were acquired to mimic the natural ECM. For the preparation of bilayer structure, CS/PCL membranes were fabricated first and then HA was electrospun on the pre-formed CS/PCL fibrous matrix (in the presence of polyethylene oxide) to form the second layer. Two layers were then exposed to glutaraldehyde vapor for interlayer crosslinking. The increased swelling, suitable degradation profile, and decreased antibacterial activity pointed out improved physicochemical and biological characteristics over PCL or CS/PCL fibrous membranes. Prevention from bacterial colonization and dehydration of the wound site are important requirements in the design of wound dressings. Obtaining increased hydrophobicity with high biocompatibility and permeability are reported as the major challenges in the preparation of layered fibrous membranes. Yu et al. [201], have combined the electrospinning with micropatterning to attain highly hydrophobic nanofibrous PCL membranes as a protective outer layer. In their study, they prepared an asymmetric bilayer membrane by electrospinning PCL on a micropatterned nylon mesh to obtain higher hydrophobicity on the outer layer. The hydrophilic inner layer was constructed by electrospinning gelatin on PCL layer to promote cell proliferation and angiogenesis. Genipin was applied to crosslink gelatin. They evaluated the healing capacity of the bilayer membrane in vivo on a full-thickness skin wound model on db/db mice (type 2 diabetes) and STZ rats (type 1 diabetes). The results showed that the developed dressing promotes wound healing by stimulating cell proliferation, angiogenesis, collagen deposition, and re-epithelialization. Hence, the wound dressing can be used as a promising candidate for diabetic wound healing. Miguel et al. [202], have used silk fibroin (SF) as a fibrous protein to increase the biocompatibility of their asymmetric membrane. They have developed a bilayer structure of polycaprolactone/silk fibroin (PCL/SF) as the outer layer to represent the epidermis with hydrophobic, waterproof, and mechanical resistance character. The inner layer consisted of a blend of SF and hyaluronic acid (HA) to represent the dermis and was designed to promote cell adhesion and proliferation with high hydration capacity. Human dermal fibroblasts showed enhanced cellular adhesion and proliferation on the bilayer membrane. The porosity, wettability, and mechanical properties showed suitable characteristics to support wound healing process. In a study by Qi et al. [203], a uni-directional multi-layered nanofibrous membrane with water-transport property was designed to diminish the excessive wound exudate that causes infection and hinders wound repair and healing. The tri-layered membrane was featured with hydrophobic to hydrophilic gradient that enabled a self-pumping of the wound exudate for a spontaneous flow from inside to outside of the wound site in a uni-directional manner. The multilayered composite was prepared by electrospinning of a hydrophobic PU as inner layer, hydrophilic PU/PAN-SPA as mid-layer and hydrophilic PAN-SPA as the outer layer in a sequential electrospinning. Polyhexamethylene guanidine

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hydrochloride was incorporated in the inner and outer layers to serve as an antibacterial agent. The resulting tri-layered fibrous dressing showed excellent antibacterial activity, suitable biocompatibility, moist permeability and exhibited superior water absorption with successful uni-directional water-transport function. Ma et al. [204], have reported a tri-layered nanofiber scaffold consisting of radiallyaligned PCL nanofibers at the bottom, PCL nanofiber membranes with square arrayed microwells and nanostructured cues at the top, and micro-skin tissues in between. Their design offered the combination of tissue engineering strategy and autologous skin micrografts as a viable clinical approach. Besides the structural advantages of layered electrospun membranes, various bioactive agents and/or nanoparticles can be introduced during electrospinning process for enhanced wound healing and antibacterial efficacy. The high surface area to volume ratio, high drug loading capacity and tunable release of the active agent highlight the electrospun membranes as excellent drug carriers. Layer-by-layer assembly via electrospinning enables the use both hydrophobic and hydrophilic polymers and different types of bioactive/therapeutic agents with adjacent characters can easily be incorporated into different layers of nanofibrous membrane during electrospinning. Hence, layered fibrous membranes can carry both antimicrobial agents and biomolecules at the same time. The release profile of the bioactive agents is mainly affected by hydrophobicity/hydrophilicity, fiber structure, porosity, biodegradability, polymer/bioactive agent ratio and the interactions between polymer/ bioactive agent/solvent. Meanwhile, multi-layered structures also enable sustained release of the bioactive agents. The bioactive agents can be blended in the polymer solution before electrospinning where the biomolecules/drugs can be uniformly distributed within the fiber structure (Fig. 5a). Burst release can be avoided by encapsulating the drug/bioactive agent in a core-shell structure by co-axial spinning where the drug in the core is enclosed in a polymeric shell layer (Fig. 5b). Loading of bioactive agents can also be performed post-electrospinning via physical adsorption or covalent bonding (Fig. 5d). The non-covalent interactions (electrostatic and/or hydrophobic) in physical adsorption lead to a faster release of the biomolecule, whereas covalent binding offers a sustained release. The selection of these approaches depends on the role of the bioactive agent-whether it should be released in a faster manner or requires a controlled long-time release profile. Moreover, in multi-layered fibrous structures, physical adsorption and covalent bonding can be used simultaneously to achieve sequential delivery of bioactive agents. Jafari et al. [205], have proposed a PCL/gelatin bilayer membrane to overcome the challenges in the treatment of diabetic, chronic, and full-thickness wounds. They combined the mechanical properties of PCL, biological cues of gelatin with antimicrobial effect of amoxicillin and bioactivity of zinc oxide (ZnO) nanoparticles. The top layer was prepared from PCL/gelatin with the addition of Amoxicillin. For the bottom layer, ZnO nanoparticles were included in the PCL/gelatin solution prior to electrospinning of the bottom layer onto the top layer. ZnO nanoparticles acted as a stimulator for angiogenesis. They confirmed the steady release of amoxicillin in vitro after 1 day and the bilayer nanofibrous scaffolds improved wound contraction rate with enhanced collagen deposition, neovascularization, and reduced scar formation

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in an in vivo full-thickness wound model. In another study by Chen et al. [206], curcumin-loaded sandwich-like nanofibrous membranes were prepared using sequential electrospinning. The hierarchical tri-layered membrane was designed to coordinate the healing stage and accelerate the healing process. The sub-layer consisted of gelatin, chitosan and PCL acted as a hemostatic nanofibrous structure to stop bleeding, absorb the exudate and keep the moisture in the wound site. The mid-layer was designed to release Curcumin to reduce wound oxidative stress and inflammation. PCL was used as the carrier polymer and PCL/curcumin mixture was electrospun to compose the sub-layer. The top layer for antibacterial effect was prepared by incorporating Ag nanoparticles in a mixture of silk-fibroin/PCL polymer solution prior to electrospinning. This tri-layered structure showed excellent antibacterial, hemostatic and antioxidant properties in vitro. Curcumin showed a burst release during 24 h related to the surface drug molecules and then released in a controlled manner up to 120 h. In vivo studies on a rat model revealed decreased inflammatory response, enhanced collagen deposition and epidermal regeneration. CD31 and TGF-β expressions were facilitated in the early stage of wound pointing out an accelerated wound healing. Chogan et al. [207], have also used a similar approach and investigated the effect of local administration of metformin-HCl on fibrosis and wound healing via a three-layer nano fibrous scaffold. In their design, polyvinyl alcohol (PVA)-metformin hydrochloride (metformin-HCl) fibrous layer was in the middle and supported by polycaprolactone (PCL)-chitosan layers on the sides. The results obtained from their in vivo studies on Wistar rats demonstrated that the slow-releasing anti-fibrogenic scaffold had a positive impact on alleviating scar formation and accelerating wound healing. In another study by Eskandarinia et al. [54], polycaprolactone/gelatin (PCL/Gel) scaffold was electrospun on a dense membrane composed of polyurethane incorporating extract of propolis as an antimicrobial agent. The bilayer wound dressing exhibited significant antibacterial activity against Staphylococcal aureus, Escherichia coli and Staphylococcus epidermidis. Moreover, the bilayer membrane significantly accelerated the wound closure and collagen deposition in the Wistar rats’ skin wound model. Layer-by-layer assembly using electrospinning also enables the covalent modifications of separate layers. In a study by Yüksel et al. [208], the covalent immobilization of antimicrobial peptide Magainin II was realized on poly(lactic-co-glycolic acid (PLGA) fibrous membrane prepared by electrospinning. A bioactive layer composed of PLGA/gelatin incorporating epidermal growth factor (EGF) was then electrospun on the PLGA/Magainin II fibrous membrane to obtain a bilayer antimicrobial-bioactive fibrous membrane. Asiri et al. [68], have loaded EGF or FGF in polyvinyl alcohol (PVA) electrospun membranes and prepared single layer and bilayer nanofibrous structure containing either EGF or FGF in each layer. They have pointed out the superiority of bilayer scaffolds and release of multiple growth factors over single layer nanofibrous membranes and concluded that bilayer structures provided better wound repair and enhanced tissue regeneration in in vivo studies on a mammalian wound rat model. Golchin and Nourani [116], have proposed a bilayer electrospun fibrillar scaffold immobilized with EGF. Polycaprolactone (PCL) was used as the upper layer and chitosan (CS)/polyvinyl

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alcohol (PVA) as the lower layer of wound dressing. They found that EGF-immobilized wound dressing efficiently accelerated wound closure and improved histological healing in an in vivo full-thickness wound healing mouse model. Overall, layered fibrous structures prepared by sequential electrospinning having asymmetric/heterogeneous characters have revealed the beneficial effects on wound healing process.

6.2

Combination of Three-Dimensional (3D) Porous Scaffolds with Nanofibers for Preparation of Layered/Fibrous Wound Dressings

Layered/hybrid 3D porous/nanofibrous tissue scaffolds produced by a combination of different techniques are increasingly getting attention since a uniform morphological structure cannot effectively provide the spatial properties required for skin reconstruction [209]. In general, such an architecture can be achieved through the combination of electrospinning with different techniques such as solvent casting/ particulate leaching, wet spinning, or freeze-drying (lyophilization). These gradient structures offer several advantages that facilitate the biomimicry with specific features of different layers of the skin tissue (Fig. 6b). For example, the dense epidermal (top) layer produced by using methods such as electrospinning provides a relatively low-porous fibrous structure. Such a low-permeable structure can cover the wound bed and protect the traumatic skin tissue from secondary damage caused by exogenous microorganisms or other external agents [194]. Also, the use of hydrophobic polymers (PCL, PLA, PLGA, polyurethane (PU) in the production of this layer provides epidermis-like properties such as mechanical strength and waterproofing [202]. In addition, the barrier functions of these layers can be further strengthened by various bioactive agents such as Ag nanoparticles, essential oils, and propolis integration [54, 209–211]. However, structures with small pores allow only superficial migration (epidermal) of cells such as fibroblasts and keratinocytes and inhibit cell proliferation in the deep wound area (dermal). Therefore, the inner (lower) layers of the 3D scaffolds should consist of a looser and wider/higher porous structure than the outer layer, which is mostly produced by using hydrophilic polymers [104]. Ramanathan et al. [209] developed a bilayer scaffold that carries the advantages of electrospun nanofibrous cellulose acetate membrane layered on the porous 3D collagen matrix produced by freeze-drying. The bioactive latexcontaining nanofibrous outer layer served to protect the wound area from infection, while the increased porosity and 3D structure in the inner layer served for proper cell adhesion, proliferation, and exudate absorption. While small pores (dense layer) in such composite scaffolds are particularly useful for cell attachment and intracellular signaling, medium-large pore sizes (porous layer) provide special micro-niches for cell proliferation that is important in the ECM accumulation, transport of nutrients,

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and metabolic wastes [212]. In this sense, the temperature of pre-freezing used in the lyophilization technique provided convenience in controlling the pore sizes of this layer [213]. In another study by Li et al. [215], electrospun core-shell nanofibers were freeze-dried to obtain 3D multi-layer patterned structures (3D-PT-P/GM) for promoting diabetic wound healing with improved angiogenesis. The nanofibrous scaffolds prepared by electrospinning had unique core–shell architectures with the shell being Gelatin methacryloyl (GelMA) hydrogel and the core being Poly (D, L-lactic acid) (PDLLA). GelMA hydrogel coated on the surfaces of nanofibers would allow the absorption of the exudate and provide a moist environment for wound repair. Multi-layered structure was achieved after freeze-drying process for high porosity to promote cell infiltration, migration, and 3D vascular network formation. The in vivo studies demonstrated that the scaffolds significantly promoted the formation of a 3D network of capillaries and healing of diabetic wounds was accelerated with enhanced angiogenesis, granulation tissue formation, and collagen deposition. Apart from freeze-drying technique, a three-dimensional porous dermis-like layer can also be produced using different techniques such as wet spinning or solvent casting/particulate leaching [214, 215]. Dalgic et al. [215], have reported a bilayer scaffold consisting of a protective electrospun Pullulan (PUL) membrane that limits bacterial penetration and water loss, and a wet spun Poly(hydroxybutyrate-co-hydroxyvalerate) (PHBV) porous structure that promotes cell proliferation and migration. In this scaffold, mouse fibroblast cells (L929) proliferated superficially on the nano-sized electrospun PUL fibrous layer, while in the 3D large-pore wet-spun PHBV layer, proliferation was reported up to a depth of approximately 120 μm throughout the matrix. Meanwhile, the dense PUL layer inhibited bacterial contamination during the 14-day test period and exhibited epidermal layer-like properties. Such scaffolds are asymmetrical structures that support the regeneration process of different skin layers by simulating the structural organization of the cutaneous tissue, to ensure the optimum biomimetic performance of the ECM. Although the fabrication of three-dimensional cutaneous tissue scaffolds is mainly focused on the epidermal and dermal layer, there are also biomimetic threelayered scaffolds targeting the hypodermis, as well. In a study by Haldar et al. [196], a gradient scaffold was designed with increased hydrophobicity and porosity toward the inner layers. Each layer exhibited a different microarchitecture, thus allowing simultaneous regeneration of three layers of skin tissue. Polycaprolactone (PCL) was used as a common polymer in the production of the outer and middle layers of this multi-layered scaffold and the difference in the fabrication techniques provided specific structural properties to the epidermis and dermis. The first layer produced by solvent casting offered a non-porous and highly hydrophobic structure, while the electrospun-middle layer similar to collagen fibers provided larger pores and higher hydrophilicity. In the third layer, a highly porous and hydrophilic gelatin-based lyophilized structure was reported, which prevented maceration by allowing the adsorption of wound exudate. This multi-layer scaffold exhibited a high wound closure rate (90%) compared to groups of animals treated with either no treatment or a standard topical ointment, i.e., Neosporin. Considering the anatomical, mechanical, physicochemical and biochemical properties or cellular composition of the

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damaged tissue, each designed layer can provide an effective regeneration as long as the scaffold stays as a whole single structure. Haldar et al. [196], have electrospun the middle layer on the semi-dry outer layer and positioned the obtained double layer on the gelatin polymer solution with self-adhesive properties, to keep the different phases of the 3D composite structure together and to preserve the integrity and consolidation. This pre-lyophilization action triggered the formation of an absorbent pressure toward the upper layers, resulting in physical adhesion between all layers of the scaffold. Consequently, combination of electrospinning with 3D porous scaffold fabrication techniques and integration of various molecules/agents provides a useful approach to bio mimic the structural, cellular and biochemical properties of skin tissue for near-ideal scaffold development.

6.3

Combination of Hydrogel Scaffolds with Nanofibers for Preparation of Layered/Fibrous Wound Dressings

Hydrogels demonstrate a class of materials widely used in wound-healing applications. They possess a highly hydrated 3D polymeric network and can absorb severalfold more water of their dry weight maintaining a high level of moisture at the wound site and can be molded in different dimensions and shapes. It is possible to include a wide range of bioactive agents into hydrogels, which facilitates their use as wound dressing materials. On the other hand, because of their limited mechanical qualities and susceptibility to moisture loss and bacterial penetration, the use of hydrogel structures alone is not preferable. Chemical or physical alterations have been tried in an effort to solve these difficulties. As an alternative approach, hydrogel scaffolds can be combined with fibrous structures to obtain layered matrices (Fig. 6c). This approach not only provides advantages to overcome the inherent drawbacks of hydrogels but also enables to mimic the heterogeneous structure of skin. Combination of nanofibers with hydrogels offer several advantages including improved mechanical properties, biocompatibility, bioavailability, and wound-healing capability. The manufacture of fiber-hydrogel combination has been accomplished using a variety of techniques, including sequential electrospinning/electrospraying and direct polymerization of hydrogels on nanofibrous mats. Kim et al. [216] prepared a bilayer scaffold composed of human hair keratin/ chitosan nanofiber mat and gelatin methacrylate (GelMA) hydrogel by using electrospinning and photopolymerization techniques. The nanofiber layer mimicking the epidermis layer was prepared by electrospinning the human hair keratin and chitosan, followed by cross-linking with glutaraldehyde. The bilayer scaffold was prepared by photopolymerization of GelMA to mimic the dermis, under the crosslinked nanofiber mat. Fibroblasts encapsulated in the hydrogel matrix and HaCaT cells cultured on the nanofibrous layer continued to proliferate in a co-culture for 10 days. Asadi et al. [217], have used a similar approach to design a multifunctional

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bilayer scaffold combining electrospun chitosan-polycaprolactone nanofibrous mat with tannic acid reinforced methacrylate gelatin-alginate hydrogel. Their results obtained from in vivo experiments on a full-thickness wound model demonstrated the outstanding performance of the layered fibrous scaffolds with an enhanced wound closure rate, effective collagen deposition and quick re-epithelialization. In another study, Zandi et al. [218], designed a niche for endogenous tissue regeneration. They presented a biomimetic bilayer scaffold composed of gelatin nanofibers operating as dermis layer and photo-cross-linkable composite gelatin metacryloyl hydrogels modified with silicate nanoplatelets and loaded with epidermal growth factor to serve as the epidermis layer for the full-thickness wound healing application. They targeted to overcome the challenges of transdermal delivery of EGF, including short half-life and lack of efficient formulation. In an established excisional full-thickness wound model, an enhanced wound closure (up to 93.1  1.5%) after 14 days was demonstrated. They have concluded that adhesive and hemostatic bilayer scaffolds with sustained release of the growth factors have the potential to stimulate complete skin regeneration for full-thickness wound-healing.

6.4

Combination of Three-Dimensional (3D) Printed Scaffolds with Nanofibers for Preparation of Layered/Fibrous Wound Dressings

3D printing is a highly individualized, adaptable, and precise technology and reported to be particularly suitable for the preparation of wound dressings. Many types of biodegradable, multi-material wound dressings were prepared using 3D printing including cell-laden and drug-eluting constructs, simulating dermis and epidermis. However, in the absence of a vascular network, diffusion restrictions adversely affect the viability of the 3D printed scaffolds and tissue construction by direct deposition or aggregation of living cells is not suitable for in vivo applications. Meanwhile, most 3D printed polymer constructs have limited mechanical strength, which restricts their applications in tissue engineering. To overcome these challenges, 3D printed constructs are combined with fibrous layers-usually prepared by electrospinning-either to reinforce the constructs mechanically or prepare heterogeneous layered structures (Fig. 6d). In wound dressing applications using this approach, the 3D printed layer is usually designed to represent the dermis layer and electrospun nanofibers represent the epidermis to prevent bacterial invasion and maintain the moisture content of the underlying 3D printed hydrogel. Wang et al. [219], have used this approach to design a bilayer scaffold for the treatment of fullthickness skin wounds. In their design, the outer barrier was composed of electrospun poly (lactic-co-glycolic acid) fibrous layer and 3D-printed alginate hydrogel represented the dermis. He and Molnár [220], proposed the combination of electrospinning and 3D printing in such a way that poly(lactide acid) electrospun layers were located between the 3D printed poly(lactide acid) filament layers as

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interleaves. They pointed out the positive effect of increasing nanofiber content on the mechanical strength of the structures. In another study by Miguel et al. [198], an asymmetric bilayer skin construct was developed by combining 3D printing and electrospinning. Chitosan/sodium alginate hydrogel was 3D-printed to generate a dermis-like layer confirming a moist environment and polycaprolactone/silk blend sericin was electrospun as the top layer to mimic the epidermis and act as a protective barrier against dehydration. Apart from the use of electrospinning and 3D printing in combination, nanofiber structures can also be incorporated into the ink to generate scaffolds containing fibrous structures with improved mechanical qualities and precise control over the shape, size, and microstructure of the constructs. In a study reported by Chu et al. [221], 3D printed gelatin methacrylol (GelMa) scaffolds were generated incorporating a proangiogenic self-assembling peptide nanofiber (SLg). The interwoven gelatin and peptide were used to create interpenetrating polymer networks (IPNs), which improved absorbency and elasticity. In their in vivo investigation, 3D printed GelMA/SLg scaffolds showed best collagenous fibrous structure, stimulated rapid revascularization, and skin regeneration repair cycles. In another study, Clohessy et al. [222] used 3D printed hydrogels as a sacrificial layer to connect electrospun fibrous layers. In their study, a blend of poly(glycolic acid) and poly(ethylene glycol) was electrospun as part of a custom fabrication method that incorporated 3D printed poly(vinyl alcohol) sacrificial elements. After removing the sacrificial templates from the layer-by-layer structure, an interconnected void within the electrospun fibers was produced. They have concluded that, when the construct was tested in vivo a full-thickness excisional skin wound, quality of healing was improved, and neovascularization was increased.

7 Summary Multi-layered fibrous scaffolds/membranes provide an effective strategy for the design and development of functional and bioactive wound-healing matrices. These structures offer remarkable advantages to mimic both the structural entities in the native ECM of skin tissue and biochemical cues during wound-healing process. Bilayer/tri-layer nanofibrous membranes prepared by sequential electrospinning of different types of polymers provide asymmetric structures to mimic the dermis and epidermis layers with dense/hydrophobic and porous/hydrophilic nanofibrous structures, respectively. Moreover, the opportunity to combine electrospinning with different techniques provides spatial morphology required for skin reconstruction. Hydrogels, 3D porous scaffolds and 3D-printed scaffolds can be combined with electrospinning to achieve multi-layered structures. This strategy would enable the use of different materials with different physicochemical properties to be constructed in different morphology to best mimic the hierarchical skin tissue. In addition to structural advantages, each layer may carry bioactive/therapeutic agents with different functions and activities that make these structures multifunctional. Multi-layered fibrous structures would be an intervening innovative

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approach to overcome the complex and challenging wound-healing process requirements.

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