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2021 37 Javier Aragón Fernández Nanofibrous membranes obtained by electrospinning for bone tissue engineering and wound dressing applications EXTRACTO Departamento Director/es Ingeniería Química y Tecnologías del Medio Ambiente Irusta Alderete, Silvia Ricardo, Ana Aguiar De Bartolo, Loredana © Universidad de Zaragoza Servicio de Publicaciones ISSN 2254-7606 En cumplimiento del artículo 14.6 del Real Decreto 99/2011, de 28 de enero, por el que se regulan las enseñanzas oficiales de doctorado, los autores que puedan verse afectados por alguna de las excepciones contempladas en la normativa citada deberán solicitar explícitamente la no publicación del contenido íntegro de su tesis doctoral en el repositorio de la Universidad de Zaragoza. Las situaciones excepcionales contempladas son: •Que la tesis se haya desarrollado en los términos de un convenio de confidencialidad con una o más empresas o instituciones. • Que la tesis recoja resultados susceptibles de ser patentados. •Alguna otra circunstancia legal que impida su difusión completa en abierto. El presente documento es un extracto de la tesis original depositada en el Archivo Universitario. Javier Aragón Fernández NANOFIBROUS MEMBRANES OBTAINED BY ELECTROSPINNING FOR BONE TISSUE ENGINEERING AND WOUND DRESSING APPLICATIONS Director/es Ingeniería Química y Tecnologías del Medio Ambiente Irusta Alderete, Silvia Ricardo, Ana Aguiar De Bartolo, Loredana Tesis Doctoral Autor 2019 Repositorio de la Universidad de Zaragoza – Zaguan http://zaguan.unizar.es UNIVERSIDAD DE ZARAGOZA [Extracto] Nanofibrous membranes obtained by electrospinning for bone tissue engineering and wound dressing applications. A thesis submitted to obtain the degree of doctor, presented by Javier Aragón Fernández Zaragoza, 2018 Nanofibrous membranes obtained by electrospinning for bone tissue engineering and wound dressing applications. A thesis Prepared in the framework of Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) to obtain multiple Doctoral degrees issue by Universidad de Zaragoza, Departamento de Ingeniería Química y Tecnología del Medio Ambiente Università della Calabria, Istituto per la Tecnologia delle Membrane (ITM) Universidade Nova de Lisboa, Facultade de Ciência e Tecnologia Supervisors: Dr. Silvia Irusta Alderete, Associate Research Professor, Departamento de Ingeniería Química y Tecnología del Medio Ambiente, Universidad de Zaragoza, Spain Dr. Loredana De Bartolo, Senior Researcher, Istituto per la Tecnologia delle Membrane, Italy Dr. Ana Isabel Aguiar-Ricardo, Full Professor, Faculdade de Ciências e Tecnologia of Universidade Nova de Lisboa, Portugal INDEX SUMMARY AND OBJECTIVES 1 CHAPTER I 21 Introduction 21 I.1 Biomaterials 23 I.1.1 Biomaterial classification 25 I.1.1.1 Metallic biomaterials 25 I.1.1.2 Ceramic biomaterials 27 I1.1.3 Polymeric biomaterials 29 I.1.1.4 Composite biomaterials 31 I.2 Drug delivery system 33 I.3 Wound dressing materials 37 I.3.1 Wound dressing classification 40 I.4 Tissue engineering 47 I.4.1 Bone tissue engineering 50 I.4.1.1 Classifications of scaffolds in bone tissue engineering 54 I.4.1.1.1 Metallic scaffolds in bone tissue engineering. 55 I.4.1.1.2 Ceramic scaffolds in bone tissue engineering. 55 I.4.1.1.3 Polymeric scaffolds in bone tissue engineering. 56 I.5 Techniques to produce materials for biomedical applications 58 References 63 CHAPTER II 79 Composite scaffold obtained by electro-hydrodynamic technique for infection prevention and treatment in bone repair 79 II.1 Introduction 82 Objective 84 II.2 Preparation of polymeric particles and membrane 84 II.2.1 PLGA particles production 84 II.2.2 Electrospun scaffolds production 85 II.3 Results and discussion 85 II.3.1 Characterization of composite scaffolds 85 II.3.2 MIC and MBC determination 94 II.3.3 In vitro cell studies 95 II.4 Conclusions 98 References 99 CHAPTER III 103 Laser-treated electrospun fibers loaded with nano-hydroxyapatite for bone tissue engineering. 103 III.1 Introduction 106 Objective 108 III.2 Synthesis of inorganic nanoparticles and electrospun scaffolds 108 III.2.1 Synthesis of hydroxyapatite nanoparticles 108 III.2.2 Preparation of electrospun scaffolds 108 III.3 Results and discussion 109 III.3.1 Hydroxyapatite characterization 109 III.3.2 Scaffolds characterization 111 III.3.2.1 As spun scaffolds 111 III.3.2.1 Laser treated scaffolds 114 III.3.3 In vitro bioactivity 116 III.3.4 In vitro cell morphology and viability 117 III.4 Conclusions 125 References 126 CHAPTER IV 130 Polymeric electrospun scaffolds for bone morphogenetic protein 2 delivery in bone tissue engineering. 130 IV.1 Introduction 133 Objective 134 IV.2 Scaffold fabrication 135 IV.3 Results and discussion 137 IV.3.1 Scaffold characterization 137 IV.3.2 In vitro protein release 141 IV.3.3 In vitro enzymatic degradation 144 IV.3.4 Cell viability and morphology 148 IV.3.5 Osteogenic, osteoinductive, and osteoconductive activities of scaffolds 150 IV.4 Conclusions 155 References 156 CHAPTER V 161 Electrospun asymmetric membranes for wound dressing applications 161 V.1 Introduction 164 Objective 166 V.2 Membrane preparation 167 V.3 Results and discussion 167 V.3.1 Membranes characterization by SEM and FTIR 167 V.3.2 Mechanical properties 171 V.3.3 Fluids handling properties 172 V.3.4 Carvacrol release 175 V.3.5 Antimicrobial properties 177 V.3.6 Cytocompatibility 178 V.3.7 Cell scratch assay 181 V.4 Conclusions 183 References 185 CHAPTER VI 190 GENERAL CONCLUSIONS 190 APPENDIX 1 202 Materials and methods 202 A.1.1 Materials 205 A.1.2 Physical-chemical characterization 206 A.1.3 In vitro studies in simulated body fluid (SBF) on “Laser-treated Electrospun fibers loaded with nano-hydroxyapatite” (Chapter III) 207 A.1.4 Mechanical properties207 A.1.4.1 Mechanical properties of “Composite scaffold loaded with RFP” (Chapter II) 207 A.1.4.2 Mechanical properties on “Electrospun asymmetric membranes” (Chapter V) 207 A.1.5 Membranes permeability in “Polymeric electrospun scaffolds for bone morphogenetic protein” (Chapter IV) 208 A.1.6 Encapsulation efficiency 208 A.1.6.1 Encapsulation efficiency in “Composite scaffold loaded with RFP” (Chapter II) 208 A.1.6.2 Encapsulation efficiency in “Polymeric electrospun scaffolds for bone morphogenetic protein” (Chapter IV) 209 A.1.6.3 Encapsulation efficiency in “Electrospun asymmetric membranes” (Chapter V) 209 A.1.7 In vitro release study and kinetic modeling 209 A.1.7.1 In vitro release study and kinetic modeling of “Composite scaffold loaded with RFP” (Chapter II) 209 A.1.7.2 In vitro release study and kinetic modeling of “Polymeric electrospun scaffolds for bone morphogenetic protein” (Chapter IV) 210 A.1.7.3 In vitro release study and kinetic modeling of “Electrospun asymmetric membranes” (Chapter V) 210 A.1.8 Drug release kinetics 210 A.1.9 Swelling studies of “Electrospun asymmetric membranes” (Chapter V) 212 A.1.10 Water vapor transmission of “Electrospun asymmetric membranes” (Chapter V) 212 A.1.11 Biodegradation studies 213 A.1.11.1 Enzymatic degradation in “Polymeric electrospun scaffolds for bone morphogenetic protein” (Chapter IV) 213 A.1.11.2 Biodegradability in a mimic real wound environment of “Electrospun asymmetric membranes” (Chapter V) 213 6 RESUMEN Y OBJETIVOS | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) RESUMEN Y OBJETIVOS Esta tesis doctoral se ha realizado dentro del marco de un acuerdo de cotutela entre la Universidad de Zaragoza (Universidad de origen), la Universidad de Calabria (Universidad anfitriona) y la Facultad de Ciencias y Tecnología de la Universidad NOVA de Lisboa (FCT NOVA) (Universidad anfitriona). El trabajo de investigación se ha llevado a cabo dentro del programa de Doctorado en Ingeniería de Membranas Erasmus Mundus (EUDIME), (FPA 2011-0014), financiado por la Unión Europea. La tesis se centró principalmente en el uso de la técnica de electrohilado para producir diferentes tipos de membranas que puedan ser utilizadas en distintas aplicaciones biomédicas. Se sintetizaron y produjeron nanopartículas orgánicas e inorgánicas para ser utilizadas como rellenos o como portadores (sistema de administración de fármacos), así como membranas nanofibrosas electrohiladas. Este trabajo se llevó a cabo en el Instituto de Nanociencia de Aragón (INA), específicamente en el grupo de Nanostructured Films and Particles (NFP) bajo la supervisión de la profesora Silvia Irusta y la Dra. Gracia Mendoza. Una parte importante de la caracterización físico-química se realizó en el INA. En la Universidad de Calabria se trabajó bajo la supervisión de la Dra. Loredana de Bartolo en el Instituto de Tecnología de Membranas (ITM). Allí se utilizaron técnicas específicas tanto para la caracterización como para estudiar diferentes señales biológicas producidas por las membranas sintetizadas, bajo la supervisión. Por otro lado, la movilidad llevada a cabo en la Facultad de Ciencias y Tecnología (FCT NOVA) de la Universidade NOVA (FCT NOVA) bajo la supervisión de la profesora Ana Isabel Aguiar-Ricardo, permitió realizar una caracterización completa de dos membranas asimétricas siguiendo diferentes Normas Internacionales que establecen diferentes ensayos a realizar en apósitos primarios utilizados en heridas. El desarrollo de nuevos scaffolds cargados con proteínas morfogenéticas o antibióticos es de gran interés en el campo de la ingeniería de tejidos óseos. Scaffolds electrohilados con una microporosidad mejorada puede ser beneficioso Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | RESUMEN Y OBJETIVOS 7 para mejorar la viabilidad celular debido a que una alta porosidad junto a la presencia de microporos puede proporcionar un entorno tridimensional (3D) que no solamente facilita la siembra y difusión celular sino también proporciona una mejor difusión de los nutrientes y residuos a través del scaffolds. La adición de cerámica de fosfato de calcio ha sido ampliamente investigada para fabricar scaffolds altamente porosos para la ingeniería de tejidos óseos debido a que presentan una composición muy similar al hueso, incluyendo excelentes propiedades de biocompatibilidad, osteoinductivas y osteoconductoras. Partículas cargadas con proteínas morfogenéticas de hueso distribuidas homogéneamente en el scaffolds podrían asegurar una liberación continua del factor de crecimiento proporcionando de esta forma las señales bioquímicas necesarias para la reparación y regeneración ósea. Los scaffolds cargados con antibióticos pueden proporcionar una liberación sostenida del fármaco en el sitio de interés, así como el mantenimiento de propiedades osteogénicas mejoradas para la regeneración exitosa del hueso. Evitando de esta forma que se alcancen niveles de toxicidad o niveles ineficaces en la zona de interés, así como la aparición de efectos secundarios indeseados en los pacientes que provocan un rechazo a los tratamientos prolongados de fármacos por vía sistemática (vía oral e intravenosa). Otra aplicación biomédica interesante de las membranas electrohiladas es la fabricación de apósitos inteligentes eficientes para el tratamiento de heridas. Para lograr una curación rápida de la herida es necesario desarrollar membranas apropiadas con poros interconectados capaces de prevenir la deshidratación rápida y la penetración de bacterias. Para mantener un ambiente húmedo en el lecho de la herida se necesita una alta capacidad de absorción y una adecuada transmisión de vapor de agua. Además, si la membrana electrohilada presenta propiedades bactericidas facilitará el proceso de curación. El objetivo principal de esta tesis fue el desarrollo mediante electrohilado de membranas fibrosas con las características apropiadas para ser utilizadas en la ingeniería de tejidos óseos o como apósito para heridas. En los Capítulos II al V se plantean una serie de objetivos específicos con el fin de cumplir el objetivo principal. 8 RESUMEN Y OBJETIVOS | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Este documento de tesis se dividió en las siguientes secciones:  CAPÍTULO I, corresponde a la introducción general donde se describen los conceptos de biomateriales, scaffolds, ingeniería de tejidos y el objetivo principal de los sistemas de liberación de fármacos. Así como, la clasificación de los biomateriales y la ingeniería de tejidos según el origen de los materiales. Además se ponen de manifiesto todos los factores que deben tenerse en cuenta para desarrollar y aplicar adecuadamente los apósitos para heridas. Se mencionaron las diferentes técnicas utilizadas en la literatura haciendo énfasis en el uso de electrohilado y electropulverización para producir scaffolds o membranas para su uso en la ingeniería del tejido óseo y como apósitos para heridas.  CAPÍTULO II, se enfoca en el desarrollo y mejora de andamios 3D capaces de promover una eficiente regeneración ósea junto con la liberación de antibióticos dirigidos para prevenir la colonización de bacterias. El objetivo de este trabajo fue sintetizar y caracterizar un sistema de liberación de fármacos que consiste en nanofibras electrohiladas de policaprolactona (PCL) decoradas con partículas de poli (ácido láctico-coglicólico) (PLGA) cargadas con rifampicina (RFP). Este material debe promover la reparación ósea evitando el deterioro del scaffolds provocado por una infección. Se realizó la evaluación in vitro de la capacidad bactericida del material electrohilado sintetizado contra bacterias Gram positivas (Staphylococcus aureus) y Gram negativas (Escherichia coli), así como su citocompatibilidad en cultivos 3D con osteoblastos humanos. Estos resultados se enviaron a la Revista de farmacia “International Journal of Pharmaceuitics” para su publicación en formato de artículo y está bajo revisión.  CAPÍTULO III, se describe la síntesis y caracterización de membranas con estructura de núcleo-envoltura de PCL y acetato de polivinilo (PVAc) obtenidas por electrohilado. Las fibras se cargaron con nanopartículas de hidroxiapatita sintética (HAn) para aumentar la bioactividad de los materiales. Los scaffolds desarrollados se trataron con ablación láser para crear características topográficas deseadas a nivel micrométrico con el objetivo de favorecer la adhesión y crecimiento celular. Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | RESUMEN Y OBJETIVOS 9 Todas las membranas obtenidas presentaron una estructura de poros tridimensionalmente interconectados y el tratamiento con láser provocó un aumento en la viabilidad y densidad celular. Además, el aumento en la biocompatibilidad de los scaffolds sugiere que los microporos pequeños favorecen la adhesión y proliferación celular. Estos resultados fueron publicados en el artículo titulado “Laser-treated electrospun fibers loaded with nanohydroxyapatite for bone tissue engineering”. Javier Aragon, Nuria Navascues, Gracia Mendoza, Silvia Irusta. International Journal of Pharmaceutics 525,112–122, 2017. DOI:10.1016/j.ijpharm.2017.04.022.  CAPÍTULO IV, se refiere al desarrollo de un scaffold electrohilado compuesto por fibras con estructura de núcleo-cubierta de PCL o PCL/PVAc cargado con HAn sintética. Estas fibras se decoraron con partículas de PLGA cargadas con proteína morfogenética ósea 2 (BMP2) mediante el uso simultaneo de electrohilado coaxial y electropulverización. El objetivo de este trabajo fue evaluar las propiedades estructurales y físico-químicas así como el proceso de biodegradación de los nuevos scaffolds desarrollados y su capacidad para abordar las características arquitectónicas, bioquímicas y funcionales del tejido óseo. Para esto, se probó la bioactividad del scaffold mediante el cultivo de osteoblastos humanos sobre ellos y se monitoreo de la viabilidad celular durante 4 semanas. Se evaluó la actividad osteogénica in vitro de las células sembradas sobre los scaffolds determinando la actividad de la fosfatasa alcalina (ALP) y la expresión de osteocalcina (OCN) y osteopontina (OPN) como proteínas osteogénicas. Estos resultados fueron publicados en el artículo titulado “Polymeric electrospun scaffolds for bone morphogenetic protein 2 delivery in bone tissue engineering”. Javier Aragón, Simona Salerno, Loredana De Bartolo, Silvia Irusta and Gracia Mendoza. Journal of Colloid and Interface Science, 531 (2018) 126–137. DOI:10.1016/j.jcis.2018.07.029.  El CAPÍTULO V, describe la síntesis de un apósito antimicrobiano para heridas, con una resistencia mecánica adecuada que es capaz de absorber exudados y evitar la deshidratación rápida de una herida. Se prepararon membranas asimétricas de PCL/PVAc cargadas con carvacrol (CRV) mediante el 10 RESUMEN Y OBJETIVOS | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) uso simultáneo de electrohilado y electropulverización. Las membranas constan de dos capas; la primera es una capa de PCL electrohilado; la segunda, una lámina de PVAc que estaría en contacto con la piel liberando a su vez el compuesto antimicrobiano. Se demostró que el uso de diferentes disolventes pueden dar lugar a la obtención de diferentes morfologías de la capa PVAc-CRV. Los valores obtenidos de elongación máxima de las membranas antes de romperse son adecuados para ser utilizados como apósitos para heridas ya que están en el mismo rango reportado de elongaciones en la piel humana. Las membranas presentan una tasa óptima de Transmisión de vapor de agua (WVTR) con valores que se encuentran en el rango requerido para mantener un buen balance entre humedad y pérdida de agua en la herida. En la primera semana, se liberó más del 60 % del CRV cargado, mientras que después de tres semanas, las membranas liberaron entre el 85 y el 100 % del CRV cargado mediante la contribución de un proceso de difusión de tipo Fickiano y la relajación delas cadenas poliméricas. Las membranas sintetizadas son candidatas potenciales para ser utilizadas como apósitos para heridas. El manuscrito que resume estos resultados se envió a la revista “Materials Science and Engineering C” y está bajo revisión (MSEC_2018_3013).  CAPÍTULO VI, resume las conclusiones generales del trabajo de tesis.  APÉNDICE 1, describe las principales técnicas de caracterización y los métodos para evaluar diferentes propiedades en función de las posibles aplicaciones.  APÉNDICE 2, resume los artículos publicados y la participación en foros científicos durante el período de tesis. Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | SOMMARIO E OBIETTIVI 11 SOMMARIO E OBIETTIVI La presente tesi di dottorato è stata svolta in base ad un accordo di cosupervisione tra l'Università di Saragozza (Università di appartenenza), l'Università della Calabria (Università ospitante) e la Facoltà di Scienze e Tecnologia dell'Università NOVA di Lisbona (FCT NOVA) (Università ospitante) . Questa ricerca è stata condotta all'interno del programma Erasmus Mundus Doctorate in Membrane Engineering (EUDIME), (FPA 2011-0014), finanziato dall'Unione Europea. Questa tesi si è concentrata principalmente sull'uso della tecnica di elettrofilatura per produrre diversi tipi di membrane per applicazioni biomediche. Sintesi e produzione di nanoparticelle inorganiche e organiche da utilizzare come fillers o vettori (sistema di somministrazione di farmaci) e produzione di membrane nanofibrose elettrofilate. Questo lavoro è stato condotto all'interno dell'Institute of Nanoscience of Aragon (INA), in particolare nel gruppo Nanostructured Films and Particles (NFP) sotto la supervisione della Professoressa Silvia Irusta e della dott.ssa Gracia Mendoza. All’interno dello stesso istituto INA si è svolta una parte importante della caratterizzazione fisico-chimica. Lo studio di diversi segnali biologici e l'uso di tecniche specifiche per la caratterizzazione delle membrane è stato effettuato presso l'Università della Calabria sotto la supervisione della dott.ssa Loredana de Bartolo presso l’Istituto per la Tecnologia delle Membrane (ITM). Infine, la mobilità svolta presso la Facoltà di Scienze e Tecnologia (FCT NOVA) dell'Universidade NOVA (FCT NOVA) sotto la supervisione della prof.ssa Ana Isabel Aguiar-Ricardo, ha permesso una caratterizzazione completa di due membrane asimmetriche in accordo a diversi standard internazionali sull’esecuzione dei test per il trattamento primario di una ferita. Lo sviluppo di nuovi scaffold caricati con proteine morfogenetiche o antibiotici sono di grande interesse nel campo dell'ingegneria del tessuto osseo. Gli 12 SOMMARIO E OBIETTIVI | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) scaffold elettrofilati con una maggiore porosità o pori microscopici potrebbero essere utili per promuovere la vitalità cellulare e la produzione di matrice extracellulare. Infatti un incremento della porosità e della dimensione dei pori potrebbe fornire un ambiente tridimensionale agevolando la semina/diffusione cellulare e migliorando la diffusione di nutrienti e rifiuti attraverso gli scaffold. L’addizione di ceramiche di fosfato di calcio è stata ampiamente studiata per la fabbricazione di scaffolds altamente porosi utili per l’ingegneria del tessuto osseo alla composizione chimica simile al tessuto osseo, l’eccellente biocompatibilità, le proprietà osteoinduttive e osteoconduttive. Una distribuzione omogenea delle particelle caricate con la proteina morfogenetica ossea lungo l'intero scaffold potrebbe garantire un rilascio continuo del fattore di crescita, stimolo biochimico necessario per la riparazione e la rigenerazione ossea. Gli scaffold caricati con antibiotici possono fornire un rilascio mirato e prolungato di farmaci, evitando la somministrazione multifarmaco orale e endovenosa a lungo termine, che implica effetti collaterali tossici, bassa erogazione al sito target e bassa aderenza al trattamento da parte del paziente, oltre a sostenere un incremento dell’osteogenicità utile per la rigenerazione dell'osso. Un'altra interessante applicazione biomedica delle membrane elettrofilate è la realizzazione di medicazioni intelligenti per il trattamento delle ferite. Una guarigione rapida delle ferite richiede lo sviluppo di membrane appropriate con pori interconnessi per prevenire la rapida disidratazione e la penetrazione dei batteri. Un'elevata capacità di assorbimento e una trasmissione adeguata del vapore acqueo saranno necessari per mantenere un ambiente umido nel letto della ferita. Inoltre, proprietà battericide miglioreranno il processo di guarigione. L'obiettivo principale di questa tesi è lo sviluppo di membrane fibrose mediante elettrofilatura per applicazioni nell'ingegneria del tessuto osseo o per la medicazione delle ferite. Per raggiungere questi obiettivi, sono stati definiti diversi obiettivi specifici, descritti nei Capitoli da II a V. Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | SOMMARIO E OBIETTIVI 13 Questo documento è stato diviso nelle seguenti sezioni:  CAPITOLO I, è un'introduzione in cui vengono descritti i concetti di biomateriali, scaffold e ingegneria tissutale e l'obiettivo principale dei sistemi di somministrazione di farmaci. Inoltre, viene descritta anche la classificazione dei biomateriali e dell'ingegneria tissutale in base all'origine dei materiali. Vengono inoltre elencati tutti i fattori per lo sviluppo e l’applicazione di una medicazione. Sono stati menzionati diversi tipi di tecniche utilizzate in letteratura per produrre scaffold o membrane per l'ingegneria del tessuto osseo e le medicazioni delle ferite, concentrandosi sull'uso di electrospinning ed electrospray.  CAPITOLO II, si concentra sullo sviluppo di impalcature 3D avanzate in grado di promuovere una rigenerazione ossea efficiente insieme a rilascio di antibiotici mirati per prevenire la colonizzazione dei batteri. Lo scopo di questo lavoro è sintetizzare e caratterizzare un sistema di somministrazione di farmaci costituito da nanofibre elettrofilate di policaprolattone (PCL) decorate con rifampicina (RFP) caricate in particelle di acido poli(lattico-co-glicolico) (PLGA). Questo materiale promuoverebbe la riparazione ossea evitando il danneggiamento dello scaffold a causa dell'infezione. La capacità battericida del materiale elettrofilato sintetizzato è stata valutata in vitro contro i batteri Gram positivi (Staphylococcus aureus) e Gram negativi (Escherichia coli), nonché la citocompatibilità nelle colture umane di osteoblasti 3D. Un manoscritto con questi risultati è stato presentato all'International Journal of Pharmaceutics ed è in fase di revisione (Ms. Ref. No.: IJP-D-18-01794).  CAPITOLO III, descrive la sintesi e la caratterizzazione delle membrane core-shell di PCL e polivinilacetato (PVAc) ottenute per elettrospinning. Le fibre sono state caricate con nanoparticelle di idrossiapatite sintetica (HAn) per aumentare la bioattività dei materiali. Gli scaffold preparati sono stati quindi trattati mediante ablazione laser per creare le caratteristiche topografiche su microscala desiderata al fine di favorire l'adesione e la crescita delle cellule. Tutte le membrane ottenute hanno mostrato una struttura reticolare dei pori interconnessa tridimensionalmente e il trattamento laser ha causato un aumento della vitalità e della densità cellulare. Inoltre, un aumento della biocompatibilità 14 SOMMARIO E OBIETTIVI | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) degli scaffold suggerisce che dimensioni di micropori più piccole favoriscono l'adesione e la proliferazione cellulare. Questi risultati sono pubblicati nell'articolo "Laser-treated electrospun fibers loaded with nano-hydroxyapatite for bone tissue engineering". Javier Aragon, Nuria Navascues, Gracia Mendoza, Silvia Irusta. International Journal of Pharmaceutics 525, 12-122, 2017. DOI: 10.1016 / j.ijpharm.2017.04.022.  CAPITOLO IV, si riferisce allo sviluppo di un'impalcatura elettrofilata composita di fibre core-shell PCL o PCL / PVAc caricate con HAn sintetico. Queste fibre sono state decorate con la proteina morfogenetica ossea 2 (BMP2) caricata in particelle di PLGA mediante elettrofiltrazione simultanea ed elettrofilatura coassiale. Lo scopo di questo lavoro è di valutare le proprietà strutturali e fisico-chimiche e i processi di biodegradazione degli scaffold recentemente sviluppati e la loro capacità di affrontare le caratteristiche architettoniche, biochimiche e funzionali del tessuto osseo. A tale scopo, la bioattività dello scaffold è stata testata coltivando gli osteoblasti umani sugli scaffold e monitorando la vitalità cellulare fino a 4 settimane. L'attività osteogenica in vitro delle cellule seminate sugli scaffold è stata esaminata valutando l'attività della fosfatasi alcalina (ALP) e l'espressione delle proteine osteogeniche osteocalcina (OCN) e osteopontina (OPN). Questi risultati sono pubblicati nell'articolo "Polymeric electrospun scaffolds for bone morphogenetic protein 2 delivery in bone tissue engineering". Javier Aragón, Simona Salerno, Loredana De Bartolo, Silvia Irusta e Gracia Mendoza. Journal of Colloid and Interface Science, 531 (2018) 126-137. DOI: 10.1016 / j.jcis.2018.07.029.  CAPITOLO V, descrive la sintesi di un materiale per la medicazione delle ferite antimicrobiche, con un'adeguata resistenza meccanica che evita la rapida disidratazione e l'assorbimento degli essudati. Le membrane asimmetriche PCL / PVAc caricate con carvacrolo (CRV) sono state preparate mediante elettrospinning e elettrospray simultaneamente. Le membrane sono costituite da due strati; il primo è un foglio elettrofilato di PCL; il secondo, un foglio di PVAc che durante il contatto con il tessuto epidermico dovrebbe rilasciare il composto antimicrobico. L'uso di diversi solventi determina morfologie differenti per lo strato PVAc-CRV. I valori di deformazione fino a rottura sono nella gamma della Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | SOMMARIO E OBIETTIVI 15 pelle umana, essendo adeguati per essere depositati su una superficie della ferita. I campioni presentano valori di trasmissione del vapore acqueo (WVTR) nell'intervallo richiesto per mantenere un buon equilibrio idrico con la perdita di acqua dalla ferita alla frequenza ottimale. Nella prima settimana, oltre il 60 % del CRV caricato è stato rilasciato mentre dopo tre settimane le membrane sono state rilasciate tra l'85 e il 100 % del CRV caricato attraverso una diffusione di Fick e una diffusione dovuta al rilassamento del polimero. Le membrane sintetizzate sono potenzialmente candidabili per le applicazioni di medicazione. Il manoscritto che riassume questi risultati è stato presentato alla Materials Science and Engineering C ed è in fase di revisione (MSEC_2018_3013).  CAPITOLO VI, riassume le conclusioni generali del lavoro di tesi.  APPENDICE 1, descrive le principali tecniche di caratterizzazione e i metodi per valutare le diverse proprietà in base alle possibili applicazioni.  APPENDICE 2, riassume gli articoli pubblicati e la partecipazione a forum scientifici durante il periodo della tesi. CHAPTER I .................................................................................................................................................................... 21 Introduction ............................................................................................................................................................ 21 I.1 Biomaterials ...................................................................................................................................................... 23 I.1.1 Biomaterial classification .................................................................................................................... 25 I.1.1.1 Metallic biomaterials .................................................................................................................... 25 I.1.1.2 Ceramic biomaterials ................................................................................................................... 27 I1.1.3 Polymeric biomaterials ................................................................................................................ 29 I.1.1.4 Composite biomaterials .............................................................................................................. 31 I.2 Drug delivery system .................................................................................................................................... 33 I.3 Wound dressing materials .......................................................................................................................... 37 I.3.1 Wound dressing classification .......................................................................................................... 40 I.4 Tissue engineering ......................................................................................................................................... 47 I.4.1 Bone tissue engineering ...................................................................................................................... 50 I.4.1.1 Classifications of scaffolds in bone tissue engineering .................................................. 54 I.4.1.1.1 Metallic scaffolds in bone tissue engineering. ........................................................... 55 I.4.1.1.2 Ceramic scaffolds in bone tissue engineering. .......................................................... 55 I.4.1.1.3 Polymeric scaffolds in bone tissue engineering. ...................................................... 56 I.5 Techniques to produce materials for biomedical applications ................................................... 58 References ................................................................................................................................................................ 63 Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 23 I.1 Biomaterials The loss of a body part or an organ generates, besides the loss of function, social and psychological disorders. This has determined not only the interest of finding new materials, but constantly new technologies are developed to improve them and provide materials that meet the most modern requirements in this field. For more than a century, efforts have been devoted to finding materials with the appropriate characteristics to solve the problems related to the reconstruction of tissues and organs injured due to aging diseases or due to trauma from accidents or falls. The widespread definition of biomaterials was suggested by the American National Institute of Health, "biomaterial is any substance or combination of substances, other than drugs, synthetic or natural in origin, which can be used for any period of time, which augments or replaces partially or totally any tissue, organ or function, in order to maintain or improve the quality of life of the individual" (1). Gold and ivory were the first materials used by Egyptians and Romans as biomaterials for replacements of cranial defects (1). Amniotic membrane (innermost layer of the placenta) was the first biological material documented for its use as a surgical material in skin transplantation with better results when compared to xenograft or cadaveric coverings (2). Polymethyl methacrylate (PMMA) is a nondegradable polyacrylate and was one of the first polymers accepted to be used in orthopedics applications in the mid-1950s (3). The tissue biocompatibility of this polymer became further apparent when Plexiglas fragments were accidentally implanted in eyes and other body tissues of World War II fighter pilots during aircraft crashes (3). The advance in modern medicine has implied the increase in human life and, thus, a greater use of biomaterials. Nowadays, biomaterials play an essential role in medicine facilitating healing and reestablishing the main functions of tissues after injury or disease. The modern field of biomaterials combines not only chemistry, physics, biology and medicine but also recent influences from materials science, regenerative medicine and tissue engineering (Figure I.1) (4). 24 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Figure I.1. Main disciplines that integrate the biomaterials sciences and engineering. Modified from von Recum and LaBerge (4). Biomaterials science and engineering is an interdisciplinary research field that requires knowledge from different areas, which are mentioned in Table I.1. Table I.1. Disciplines required for the development of biomaterials. Discipline Essential Knowledge Basic Science Physics, chemistry and Biology. Materials Science and Engineering Structure, property, relationship of synthetic and biological materials including metals, ceramics, polymers, composites, body tissues, design and prototype development, applied mechanics, thermodynamics, etc. Medical Science Composition of human body (cells, tissues, organs and systems), anatomy, physiology, pharmacology, immunology and pathology with some biochemistry, microbiology, molecular biology, genetics, etc. Clinical Science Clinical specialties, among which are includes: surgery, orthopedics, maxillofacial, plastic and reconstructive surgery, dentistry, ophthalmology, neurosurgery, obstetrics and gynecology, , cardiovascular surgery, veterinary medicine, etc Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 25 The global market of biomaterials was around 70 billion of USD in 2016 and is expected to reach about 150 Billion of USD by 2021 with a compound annual growth rate (CAGR) of 16.0% (5). According to their applications, the biomaterials market is sectioned into cardiovascular, orthopedic, ophthalmology, dental, plastic surgery, wound healing, tissue engineering, neurological/central nervous system, and other applications (5). The plastic surgery and wound healing segments and the polymeric biomaterials are expected to grow at the highest CAGRs during the forecast period. As well as, polymethylmethacrylate (PMMA) is expected to lead the polymer biomaterials market. Among the main companies that commercialize biomaterials we can find: Royal DSM (Netherlands), BASF SE (Germany), Corbion N.V. (Netherlands), Covestro (Germany), Invibio Ltd. (U.K.), Carpenter Technology Corporation (U.S.), Evonik Industries AG (Germany), Berkeley Advanced Biomaterials, Inc. (U.S.), CAM Bioceramics BV (Netherlands), and Celanese Corporation (U.S.) (5). I.1.1 Biomaterial classification Biomaterials can be classified according to the nature of the material, being the most important groups: ceramics, metals, polymers and composites. They can be defined either as passive biomaterials (generally remain neutral in their biological environment and have no inherent power of action) or active biomaterials (able of interact with their environment and may even become an integral part of the body) with different bioactivity (6). I.1.1.1 Metallic biomaterials Metallurgic industry is able to produce a large number of metals and alloys, but only a few are biocompatible and long-term successful as an implant material. The combination of high mechanical strength and fracture toughness makes metals more suitable for load-bearing applications compared with ceramics or polymeric materials. At the beginning, metallic implants were developed to be used mainly in bone repair (internal fracture fixation of long bones), playing a major role in most orthopedic devices, including temporal and permanent devices (e.g. bone plates, pins, 26 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) screws and total joint replacements) (7). The main methods to obtain them are by casting, forging, pressing, rolling and machining. Afterwards, the metallic implants developed were used not only in orthopedic surgery, but also in dental and orthodontic practice, including tooth fillings and roots. Currently, several researches point to the use of metallic biomaterials in application of nonconventional reconstructive surgery of hard tissues/organs (NiTi shape memory alloys as vascular stents) and to develop new magnesium-based alloys for bone tissue engineering and regeneration (7). Chen and Thouas summarized the current status and clinical applications of the four classes of metallic biomaterial (7): Stainless steels, Co-based and Ti-based alloys are routinely applied in: temporary devices (fracture plates, screws, hip nails, etc.); stem and cup of total hip replacements with cobalt–chromium-molybdenum or ceramic femoral heads; total joint replacements (wrought alloys); dentistry castings and other permanent devices (nails, pacemakers). The last category corresponds to "Miscellaneous" and within there are different alloys such as tantalum alloys, used as a radiographic marker and wire sutures for plastic surgery and neurosurgery, both approved by US Food and Drug Administration (FDA); magnesium alloys, that have been reported in different papers as a biodegradable orthopedic implants and nickeltitanium alloys (known as Nitinol), which have been FDA approved for different applications: orthodontic dental archwires, vascular stents, vena cava filter, intracranial aneurysm clips, catheter guide wires, orthopedic staples and contractile artificial muscles for an artificial heart. Despite all its advantages for bone repair current metallic biomaterials present limitations or disadvantages, among which can be highlighted (8):  possible release of toxic metallic ions and/or particles through corrosion or wear processes, reducing biocompatibility, leading to the beginning of the inflammatory cascades and triggering tissue loss Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 27  the elastic moduli are not well matched with that of natural bone tissue, producing stress shielding effects reducing the remodeling and stimulations of new bone growth, compromising the implant stability  are essentially neutral in vivo, remaining as permanent fixtures, causing a second surgical procedure after the tissue has healed sufficiently to remove them safely, such are the cases of plates, screws and pins used to secure serious fractures. I.1.1.2 Ceramic biomaterials Ceramic materials are composed of inorganic, non-metallic substances crystalline, semi-crystalline or non-crystalline (amorphous compounds), glasses and glass-ceramics (partially crystallized glasses). Ceramic biomaterials (Table I.2) can be categorized according with the interactions or attachment between material and host tissue (bioinert or bioactive). Furthermore, the bioactive ceramics may be divided as resorbable or non-resorbable and they may be manufactured either in granulates, as coatings, porous or dense in bulk form (9). Ceramic biomaterials have been widely used to repair the skeletal system, comprising bone, joints and teeth and to augment both hard and soft tissue (10). Additionally, these materials have been used as carriers for enzymes, antibodies, antigens and as microinjectable delivery system for radioactive isotopes for in situ treatment of tumors. They are also used for eyeglasses, diagnostic instruments, chemical ware, thermometers, tissue culture flasks and fiber optics for endoscopy (9). In dentistry, ceramics are used as restorative materials (including inlays and onlays), in gold-porcelain crowns, glass-filled ionomer cements, multi-unit bridges and dentures (1,9). 28 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Table I.2. Types of Ceramic biomaterials (9). Interaction with host tissue Properties Implant–Tissue Response Ceramic Bioinert Non-toxic material dense or non-porous and biologically inactive (nearly inert) A fibrous tissue of variable thickness is formed and adhered by growth into surface irregularities by cementing the device into the tissues or by pressfitting into a defect Aluminum oxide (single crystal and polycrystalline) Zirconia (10) Non-toxic material porous (pore diameter between 50-150µm) and biologically inactive (nearly inert) An ingrown occurs that mechanically attaches the tissue to the material Aluminum oxide (polycrystalline) Hydroxyapatitecoated porous metals Bioactive Non-toxic material biologically active, dense or non-porous, glasses and glass-ceramics with a reactive surface They connect directly by chemical bonding with the tissue forming an interfacial bond Bioactive glasses Bioactive glassceramics Hydroxyapatite (HA) Non-toxic resorbable material biologically active, dense, non-porous or porous able to dissolve in biological environment The surrounding tissue replaces it slowly Calcium sulfate Tricalcium phosphate Calcium– phosphate salts The main methods of manufacturing ceramic biomaterials are plasma spraying, liquid phase sintering (vitrified), hydrothermal synthesis and solid state sintering (9). In addition, the main advantages of the use of ceramic biomaterials are their high biocompatibility, high resistance to compression, lower wear rates, resistance to corrosion, chemical inertness, release of very low concentrations of ‘inert’ wear particles, low thermal and electrical conductivity, easy obtaining and low production costs (10,11). Its main disadvantages are associated with the low impact resistance and the difficulties in processing specific geometric forms (11). Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 29 I1.1.3 Polymeric biomaterials Polymers are the biomaterials most used in medicine because they show several unique properties which make them useful in a wide range of applications, such as hard and soft tissue replacements, orthopedics, dental or cardiovascular devices (12). The polarity, stiffness and organization of the polymeric chains into very interesting architectures are the main characteristics of polymeric biomaterials (12). These materials are often mechanically weaker than other classes of biomaterials (metals and ceramics), because they are linked by secondary interactions, such as hydrogen bonding, dipole-dipole interactions and London forces. Nonetheless, they can exhibit physical behavior more similar to native tissue (12). Polymers biomaterials can be classified regarding its origin in natural or synthetic. Biodegradability is considered the main characteristic of natural polymers, but synthetic polymers are more available and usually have a more cost-effective fabrication compared with the obtaining of natural polymers (13). Many researchers consider that natural polymers are finding increasing applications in the area of bone replacement and hard tissue augmentation but its availability and cost of obtaining is its main disadvantage (13). Among the most used natural polymers can be found: collagen, chitosan, alginate, starch, cellulose, hyaluronic acid, elastin, keratin, silk, etc. (14,15) . Some of them have a highly desirable to induce rapid bone colonization since they can provide a template for biomimetic apatite formation (15). The main methods to process synthetic polymers are thermally induced phase separation, porogen leaching/solvent casting, electrospinning, gas foaming, rapid prototyping or 3D printing are some of the methods used to process synthetic polymers (16). Synthetic polymers such as, poly( ε -caprolactone) (PCL), poly lactic acid (PLA), poly glycolic acid (PGA) and their copolymers, e.g. poly(lactide-co-glycolic acid) (PLGA) have been widely used to create 3D constructs for bone regeneration due to its biocompatibility and their tailorable biodegradation (16,17). 30 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) In 2016, Teo and collaborators reported a list of synthetic polymers approved by the FDA (Table I.3), commonly used as medical implants (13). Table I.3. Synthetic polymers approved by FDA for biomedical applications. Synthetic Polymer Biomedical Applications Polyethylene Anesthesiology, Cardiovascular, Otolaryngology, Gastroenterology, Urology, Hematology, Pathology, Neurology, Ophthalmic and Orthopedic Polytetrafluorethylene Anesthesiology, Cardiovascular, Gastroenterology, Urology, General Surgery, Plastic Surgery, Hematology, Pathology, Neurology and Ophthalmic Polyamide Anesthesiology, Cardiovascular, Gastroenterology, Urology, Hematology, Pathology, Neurology General Surgery, Plastic Surgery and Ophthalmic Polypropylene Cardiovascular, Gastroenterology, Urology, General Surgery, Plastic Surgery, Obstetric and Gynecologic Polyethylene terephthalate Cardiovascular, General Surgery and Plastic Surgery Polydimethylsiloxane Cardiovascular, Otolaryngology, Gastroenterology, Urology, General Surgery, Plastic Surgery and Neurology Polyhydroxyalkanoates Cardiovascular, Gastroenterology, Urology, Neurology and Orthopedic Polymethylmethacrylate Dental and Ophthalmic Liquid crystal polymer Otolaryngology and Neurology Silicone Otolaryngology, Gastroenterology, Urology, General Surgery, Plastic Surgery, Obstetric and Gynecologic Parylene Otolaryngology, Neurology Polyimides Neurology SU-8 Neurology Polyether Ether Ketone Orthopedic Polyurethane Obstetric and Gynecologic Despite all its advantages and applications, polymeric biomaterials present a low mechanical resistance and several of them are easily biodegradable. These two properties must be taken into account when developing a new biomaterial (11). Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 31 I.1.1.4 Composite biomaterials Composite materials were defined by William in the Dictionary of Biomaterials, as a "structural material made of two or more distinctly different materials, where each component contributes positively to the final properties" (18). In our body, many tissues are considered "composite" with an additional complexity due to their hierarchical structure, such as extracellular matrix (ECM), tendons, ligaments, skin, and bone, among others (19). Generally, composite biomaterials are associated to combinations of two or more components from the basic biomaterials classes, metals, ceramic and polymers, rather than a combination of materials within the same class (18). Components present in a composite should be distinguishable at a microscopic level. Distribution, content and interaction of the constituent materials together with their physical, chemical and mechanical properties, have a strong repercussion over the final properties of the composite biomaterials (19). Composite biomaterials usually consist of one or more discontinuous phases (commonly known as filler or reinforcing material) embedded within a continuous phase (known as matrix). The principal functions of the filler are to increase the mechanical properties and bioactivity and change the physical and chemical properties. Ceramic and glass particles as well as carbon, polymer and glass fibers have been the main reinforcing materials used in composite biomaterials (Table I.4). Most composite biomaterials used in biomedical applications have a polymeric matrix (Table I.5), that can be bio absorbable or not. Ceramic and metallic matrices composites are used mainly in non-biomedical applications, only a few composite biomaterials have a ceramic matrix, such as calcium phosphate bone cements (19). 38 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Normal wound healing physiological process requires skin restoration via reepithelialization and collagen formation (30,31). Several type of dressing are widely used to protect the wound, to increase the speed and quality of wound healing. The main characteristic of a wound dressing is to provide the optimum conditions for wound healing and, at the same time, offer protection to the wound from invasion by pathogenic microorganisms and further trauma (31). It is also essential that the dressing does not integrate into the tissue and can be detached without causing any trauma to the wound surface during dressing changes (32). The obtaining of an ideal dressing is very complicated because not only must take into account the characteristics of each type of wound, but also should mimic the properties and functions of human skin (33):  Protecting the organism from its environment even while maintaining it in uninterrupted communication with the environment;  Preventing body fluid from escaping and external fluids from penetrating;  Thermoregulation;  Synthesis (vitamin D and melanin)  Storage (lipids, melanin and water);  Resistance to mechanical stress;  Absorption;  Excretion; etc.) To select an appropriate wound dressing, it is necessary to take into account multiple factors, based on their functional and performance characteristics (24).  Debridement (wound cleansing): Enhances the migration of leukocytes into the wound bed and supports the accumulation of enzymes. Necrotic tissue, foreign bodies and particles prolong the inflammatory phase and serve as a medium for bacterial growth.  Provide or maintain a moist wound environment: Prevents desiccation and cell death, enhances epidermal migration, promotes angiogenesis and connective tissue synthesis, and supports autolysis via the rehydration of desiccated tissue. Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 39  Absorption. Removal of blood and excess exudate: In chronic wounds, there is an excess of exudate, containing tissue-degrading enzymes, which block the proliferation and activity of cells, and break down extracellular matrix materials and growth factors, thus delaying wound healing. The excess of exudate can also macerate the surrounding skin.  Gaseous exchange (water vapor and air): Permeability to water vapor controls the management of exudate. Low tissue oxygen levels stimulate angiogenesis. Raised tissue oxygen stimulates epithelialization and fibroblasts.  Prevent infection: protect the wound from bacterial invasion: Infection prolongs the inflammatory phase and delays collagen synthesis, inhibits epidermal migration and induces additional tissue damage.  Provision of thermal insulation: Normal tissue temperature improves the blood flow to the wound bed and enhances epidermal migration.  Low adherence. Protects the wound from trauma: Adherent dressings may be painful and difficult to remove, and cause further tissue damage.  Cost-effective low frequency of dressing change: Dressing comparisons based on treatment costs rather than unit or pack costs should be made (cost-benefitratio). Although many dressings are more expensive than traditional materials, the more rapid response to treatment may save considerably on the total cost. Global wound dressings market was estimated to be valued at 6.3 billion of USD in 2016, and is expected to grow at a CAGR of 6.0% from 2016 to 2021, to reach around 8.5 Billion of USD by 2021 (34). Advanced wound dressings segment registered the largest market share in 2015 and it is expected to grow at a CAGR of 6.9%. Increasing knowledge regarding wound care management, the high growth potential of emerging economies, and the increasing number of acquisitions by key companies have opened up a wide range of opportunities for the growth of this market in upcoming years. Among the main companies that market wound dressings are Smith & Nephew plc (U.K.), Acelity L.P., Inc. (U.S.), Integra Lifesciences Corporation (U.S.), Coloplast Corporation (Denmark), 3M Company (U.S.), 40 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Organogenesis Inc. (U.S.), ConvaTec Inc. (U.S.), Hollister Wound Care (U.S.), Mölnlycke Health Care AB (Sweden), and BSN Medical (Germany) (34). In the estimated global prevalence by wound type according to the report from MedMarket Diligence, LLC (Figure I.3), it is expected that the most frequent injuries in the world in 2020 will be surgical wounds (about 450 million of people), followed by diabetic ulcers (more than 60 million of people), venous ulcers (about 40 million of people), pressure ulcers and lacerations (more than 20 million of people) and burn wounds (about 10 million of people) (35). This indicates the importance of developing new wound dressings for the efficient treatment of these types of wounds. Figure I.3. Estimate global prevalence by wound type reported by MarketsandMarkets™ in 2016 (34). I.3.1 Wound dressing classification Traditionally, wound dressings are classified based on its nature or action as: passive products (gauze and tulle dressings, which is the largest market segment); interactive products (polymeric films and foams recommended for low exuding Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 41 wounds, which are mostly transparent, permeable to water vapor and oxygen but impermeable to bacteria); and bioactive products (dressings constructed from material which has endogenous activity or dressings able to act as a drug delivery system of bioactive compounds in wound healing) (36). In 1999, wound dressing categories were reclassified by the FDA as follow (36):  Nonresorbable Gauze/Sponge Dressing for External Use (sterile or nonsterile device, to be placed directly on a patient’s wound to absorb exudates).  Hydrophilic Wound Dressing (non-resorbable materials with hydrophilic properties available in sterile or nonsterile form, to cover a wound and absorb exudates).  Occlusive Wound Dressing (non-resorbable material sterile or nonsterile with or without an adhesive backing, intended to cover a wound, to provide or support a moist wound environment and to allow the exchange of gases such as oxygen and water vapor through the device).  Hydrogel Wound and Burn Dressing (non-resorbable matrix sterile or nonsterile made of hydrophilic polymers or other material in combination with at least a 50 % of water, designed to cover a wound, to absorb wound exudates, to control bleeding or fluid loss, and to protect against abrasion, friction, desiccation and contamination).  Interactive Wound and Burn Dressings (reduces the colonization count, exudates level, improves the wound bed moisture retention, improves the wound collagen matrix, removes cellular products and provides protection for the epithelializing bed). At present, wound dressing market offers more than 3,000 of dressing products registered to suit the needs according to the condition of the wound (37) and according to the base of dressings they can be divided into 8 main categories (Table I.7). 42 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Table I.7. Wound dressing categories and its applications. Dressing categories Main features Applications Disadvantages Gauze and Impregnated Gauze Dressings Highly permeable and relatively non-occlusive, inexpensive, and used as a non-time or shortterm use. Secondary dressings commonly used for both infected and non-infected wounds of any size, shape, depth, or etiology. Require more force to remove and may leave residue or lint in the wound bed. Films Flexible, permeable to water vapor and gas, but impermeable to bacteria and water. Superficial wounds such as lacerations, abrasions, partial thickness wounds, sutured wounds, and graft donor sites with minimal drainage. They may also be used on granular wounds and areas of friction. Not to be used on infected wounds, wounds with moderate to heavy drainage (little absorptive capabilities), or patients with fragile skin. Hydrogels Permeable to gas and water, provide moisture to dry wounds and absorbs a minimal amount of fluid, effective for softening eschars. Available in sheets, gels, or impregnated gauzes are indicated for any thickness wounds with minimal or moderate drainage. They can decrease pain and provide padding to decrease shear forces. Most hydrogels are almost non-adhesive, thus requiring a secondary dressing. Should not be used on infected wounds and on heavily draining wounds. Foams Porous 3 layered polyurethane foams able to quickly and effectively draw the exudates deeply into the absorbent material and reliably hold it there. Hydrophobic outer layer prevents bacteria from penetrating. Middle absorption layer is devised to retain absorbed wound exudates. The inner contact layer is hydrophilic with specifically pores size. Universally used on wounds with minimal to heavy exudates, because they are easy to place and provide thermal insulation. Foams dressing should be able to quickly and effectively draw the exudates deeply into the absorbent material and reliably hold it there. It should also provide soft cover to the wound site, in order to manage the wound environment. Semipermeable foams are not indicated for dry or eschar covered wounds because it can adhere to the wound Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 43 Table I.7. Continued Hydrocolloids Hydrophilic colloidal particles with a strong film or a foam adhesive backing sheet provide thermal insulation. Impermeable to water, gas, and bacteria and therefore can be effective barriers against urine, stool, and pathogenic microorganisms. Indicated for any thickness wounds, may be safely used on both granular and necrotic wounds. They should only be used with skin sealants on patients with good skin integrity. Often leave residues after removal. Not appropriate for bleeding or heavily draining wounds. They are contraindicated for infected wounds and must be used with caution on immunocompromised patients. Alginates and Hydrofibers React with serum and wound exudates to form a hydrophilic gel to provide a moist wound environment. They are highly permeable and nonocclusive. Ideal dressings for moderate to highly draining wounds (they can absorb up to 20 times their weight of exudates). Indicated for partial and full thickness draining wounds and can also be used for infected wounds. They require a secondary dressing and are not indicated for dry or minimally draining wounds and on wounds with exposed tendon, capsule, or bone Biological Dressings Derived from natural tissues, such as skin or amnion. Generally, they provide and maintain moist wound environment that is conducive to regeneration and migration of fibroblasts and epithelial cells and act as a bacterial barrier to protect the wound from infection and fluid loss. Decrease the pain associated with open wounds. Some types may adhere firmly to the wound by vascular connections. Temporary coverage of open and large wounds. Ideal skin substitute for use in acute and chronic wounds (38). Low availability and high produce cost. (38) The most serious potential liability of biologic wound dressings is transmission of infection; however, the actual incidence of such transmission is extremely low. 44 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Table I.7. Continued Composite Dressings and Adjunctives Multiple layers and each layer is physiologically distinct. They usually have 3 layers: inner contact layer (is non adherent, preventing trauma to the wound bed), middle layer (absorbs moisture and wicks it away to prevent maceration yet maintaining a moist wound bed) and outer layer (commonly a semipermeable film, serves as a bacterial barrier). Convenient for both partial and full thickness wounds. It can be used as primary or secondary dressings. Less flexibility and more expensive (30). Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 45 As described above, each dressing variety has unique properties, which makes it suitable for different wound conditions. The disadvantage of a dressing for a specific type of wound may be beneficial for the treatment of another wound. Some of the commercial products are described in Table I.8 (36). However, there is still no superior product that heals chronic wounds like venous leg ulcers, diabetic wound and pressure ulcers which often fail to achieve complete healing. Hence, developing a dressing material that addresses the major interfering factors of normal healing process will help patients and wound care practitioners (30). Table I.8. Some commercial dressing materials (36). Dressing Type Company Material Bioclusive® Film Johnson & Johnson PU Cica-Care Film Smith & Nephew Silicone Mepiform® Film Mölnlycke Health Care Silicone Mepilex® Film/form Mölnlycke Health Care Silicone/PU Mefilm® Film Mölnlycke Health Care PU Mitraflex® Film Mölnlycke Health Care PU Omiderm® Film Iatro Medical PU Opsite® Film Smith & Nephew PU Spyrosorb® Film BritCair PU Tegaderm® Film/pad/foam 3M Health Care Hydrocolloid/acrylic Biatain® Foam/film Coloplast PU PolyMem® Foam/film Ferris Mfg. Corp. PU Contreet® Foam/film Coloplast PU containing silver Cavi-careTM Foam Smith & Nephew Silicone Lyofoam® Foam Seton Healthcare PU Allevyn® Foam Smith & Nephew PU Tielle® Foam Johnson & Johnson PU Actisorb plus® Deodorizing Johnson & Johnson Activated charcoal cloth with silver Carbonet® Deodorizing Smith & Nephew Activated charcoal cloth Alione Bioactive Coloplast Hydrocolloid Biofilm Bioactive Biotrol Hydrocolloid CombiDERM® Bioactive ConvaTec Hydrocolloid Comfeel Contour® Bioactive Coloplast AS Hydrocolloid Cutinova Hydro Bioactive Smith & Nephew Hydrocolloid 46 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Table I.8. Continued Kaltocarb® Deodorizing ConvaTec Activated charcoal cloth with alginate GranuGel® Paste Bioactive ConvaTec Hydrocolloid paste Granuflex® Bioactive ConvaTec Hydrocolloid Tegasorb® Bioactive 3M Health Care Hydrocolloid Duoderm® Bioactive ConvaTec Hydrocolloid Mesitran® Bioactive Theo Manufacturing BV Hydrocolloid with honey Aquacel® Bioactive ConvaTec Hydrofiber Versiva® Bioactive ConvaTec Carboxymethyl cellulose Aquaform® Bioactive Aspen Medical Hydrogel Geliperm® Bioactive Geistlich Sons Ltd. Hydrogel Granugel® Bioactive ConvaTec Hydrogel Intrasite Gel® Bioactive Smith & Nephew Hydrogel Nu-Gel® Bioactive Johnson & Johnson Hydrogel Purilon Gel® Bioactive Coloplast Ltd. Hydrogel Sterigel® Bioactive Seton Hydrogel Vigilon Bioactive Bard Hydrogel Algisite® Bioactive Smith & Nephew Alginates Algosteril® Bioactive Brothier Alginates Kaltostat® Bioactive ConvaTec Alginates Melgisorb Bioactive Mölnlycke Health Care Alginates SeaSorb® Bioactive Coloplast AS Alginates Sorbsan® Bioactive Pharma-Plast Ltd. Alginates Tegagel® Bioactive 3M Health Care Alginates Tegagen Bioactive 3M Health Care Alginates Mepore® Traditional Mölnlycke Health Care Nonwoven polyester fabric Debrisan® Beads Pharmacia and Upjohn Ltd. Polysaccharide Iodosorb Beads Smith & Nephew Polysaccharide Iodoflex Beads Smith & Nephew Polysaccharide Acticoat Silver Smith & Nephew High-density polyethylene mesh and nonwoven fabric of rayon and polyester Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 47 Table I.8. Continued Promogran Form Johnson & Johnson Collagen/regenerated cellulose KytoCel® Form Aspen Medical Chitosan fibers CeloxTM Form/pad/granules Medtrade Products Ltd. Chitosan HemCon® Form HemCon Medical Tech., Inc. Chitosan AquanovaTM Pad MedTrade Products Ltd. Chitosan Beschitin® Sponge/film Unitika Medical Products Chitin I.4 Tissue engineering One of the most devastating, costly and frequent problem in human health care is the loss or failure of an organ or tissue (39). Every year, millions of people suffer tissue loss or end stage organ failure. Tissue engineering is considered a new field dating from late 80's, that benefits/takes advantage of the principles of engineering and biology to the development of functional substitutes for damaged tissue. In 1988, tissue engineering was defined by Y.C. Fung as "an interdisciplinary field that applies the principles of engineering and the life sciences toward the development of biological substitutes that restore, maintain, or improve tissue function" (17). In 1993, three general strategies were proposed for the creation of new tissue (39):  Isolated cells or cell substitutes. It allows the replacement of only the affected cells and permits manipulation of cells before using them. Its limitations include failure of the cells used to maintain their function in the recipient and immunological rejection.  Tissue-inducing substances. This approach not only depends on the large-scale production and purification of appropriate signal molecules, but also on the development of methods to deliver these molecules (growth factors, drugs or any bioactive ingredient) to their targets.  Cells placed on or within matrices. The matrices are produced from natural or synthetic materials (commonly known as scaffolds). Immunological 54 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Table I.9. Continuation Actifuse Silicate substituted calcium phosphate. Baxter International Inc., Deerfield, IL, USA Vitoss Bioactive glass and calcium phosphate. Orthovita Inc. Malvern, PA, USA OP-1® Putty rhBMP-7, Type I bovine collagen matrix and putty additive carboxymethylcellulose sodium. Olympus Biotech, Hopkinton, MA, USA OP-1 Implant Osteoinductive and osteoconductive bone graft material consisting of collagen matrix, 1 g of Type I bovine collagen, 3.3 mg of rhBMP-7 and 2–3 cc of saline. I.4.1.1 Classifications of scaffolds in bone tissue engineering Scaffolds used in bone tissue engineering must provide a stable mechanical environment and adequate degradation rate over a certain period of time in order to achieve optimal bone tissue ingrowth and regeneration. In 2013, Henkel and Hutmacher proposed a novel approach to design scaffolds-based bone tissue engineering (Figure I.6) (69). Scaffolds designed for bone tissue engineering should provide a suitable microenvironment for the tissue regeneration capacity, promote the crucial steps in the organ maturation process and tissue remodeling (69). Scaffolds can be classified according to the nature of the material (metals, ceramic and polymer scaffolds with or without modifications). Polymeric scaffolds with modifications (addition of ceramics) are the most used in this field. Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 55 Figure I.6. Targeted design and application of scaffolds for bone tissue engineering. Images from Servier Medical Art, © 2013 SERVIER. All Rights Reserved (55,69). I.4.1.1.1 Metallic scaffolds in bone tissue engineering. Metallic scaffolds act more like permanent implants than scaffolding. Pure metals like iron, magnesium, titanium, tantalum, zirconium and niobium, and different alloys such as, chrome-cobalt, magnesium-calcium, iron-magnesium, titaniumaluminum-vanadium, nickel-titanium and stainless steel, have been used. Clinical applications as well as their advantages and disadvantages are discussed in I.1.1.1 section I.4.1.1.2 Ceramic scaffolds in bone tissue engineering. Bio-ceramics practically mimic the bone tissue and provide a higher osteoblasts adherence and proliferation because solubility and surface topography have a significant influence in cell behavior compared to other materials. Calcium phosphate ceramics such as hydroxyapatite, tricalcium phosphate, biphasic and amorphous calcium phosphates have been greatly studied for bone tissue repair in the form of scaffolding or in combination with polymers or metals. 56 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) I.4.1.1.3 Polymeric scaffolds in bone tissue engineering. Several physiochemical characteristics (solubility, porosity, pore size, enzymatic reactions, biocompatibility and allergic response) are usually more controllable in polymeric materials used to design scaffolds. Natural polymer scaffolds could be tissue or cell derived that show osteoinductive properties. They are composed mainly by proteins (collagen, elastin, gelatin, keratin, fibrinogen and silk), polysaccharides (cellulose, glycosaminoglycans, amylose, chitin and dextran), polynucleotides (DNA, RNA) and extracellular matrix . Table I.10. Most used polymeric scaffolds in bone repair. Modified from Ghassemi et al (70). Name Mechanical Properties Modifications Advantages Synthetic polymers PLLA* +++ HA incorporation to enhance cell growth. - biocompatible - biodegradable - support cell adhesion PGA* +++ Alkaline hydrolysis for increasing cell replacement and cells biomaterials interaction improvement. - biocompatible - biodegradable - support cell adhesion PCL* ++ - High RGD concentration for increasing osteoblast attachment. - CNT addition for mechanical properties, BMSCs proliferation and differentiation enhancement. - biodegradable PLGA* + - HA incorporation for enhancing compressive strength. - Diamond nanoparticles incorporation for higher mechanical resistance. Incorporation of CNTs for higher rate of cell attachment, proliferation, and differentiation. - biodegradable - support cell adhesion PET* +++ - - highly biocompatible - biodegradable - impact resistance PAG + - - biocompatible PAA* + - - non-biodegradable Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 57 Table I.10. Continued PBT* ++ - - highly biocompatible - biodegradable - impact resistance PEG* + RGD peptides for facilitating cell adhesion and spreading. - biocompatible - steering cells into scaffolds - osmotic effects in body PVA +++ CNT and CNF incorporation for higher concentration of ALP and mineralised matrix. - Non-biodegradable - great resistance against organic solvents PPF* ++ linked RGD peptides for osteoblast migration regulation. - biocompatible - suitable physical properties and decomposition rate PU + - - variable degradability - injectable Natural polymers Collagen (type I, II and III) + - mixing with calcium for mechanical integrity increase. - blending with PCL for mechanical improvement. - biocompatible - degradable Chitosan - Nanocrystalline hydroxyapatite and SWCNT incorporation for mechanically and cytocompatibility enhancement. - Chitin + - - biocompatible - biodegradable Alginate + Addition of HA, calcium phosphate cements, bioglass and other natural and synthetic polymers for upgrading cell adhesion and mechanical properties. - biocompatible - degradable - minimally invasive manner (gel-forming) - ease of chemical modification with adhesion ligands and controlled release of tissue induction factors (e.g., BMP, TGF-β) Mechanical properties: +++ good, ++ average, + poor FDA approved: * 58 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) On the other hand, biocompatible and biodegradable synthetic polymers such as PCL, PLA and PGA have been the most utilized in bone tissue engineering for their excellent mechanical properties and its low cost of obtaining. But some polymers lose their compressive strength due to rapid degradation in vivo (poly(propylene fumarate), generating a local acidic environment which can cause adverse tissue responses. Table I.10 shows some of the most used polymeric scaffolds modified and their advantages. I.5 Techniques to produce materials for biomedical applications Biomaterials may be designed and fabricated using all kind of materials: polymers, ceramics and metals, as well as their combinations (71). In the literature, there are diverse methods to produce biomaterials for medical applications, specifically for wound dressing applications and bone tissue engineering (71–73). Salerno and Netti summarized some of them (Freeze drying, Particles sintering, Phase inversion, Solid freeform fabrication, Reverse templating, Spraying, Textiles, Emulsion, Gas foaming, Bioprinting, Microfluidic, Self-assembly) (71). The design of an appropriate biomaterial depends not only on its specific medical application but also on the characteristics of the particular patient. It is very difficult to describe specific characteristics to take into account in such developments, but any device used should be able to maintain its functionality during all the accurate time without rejection and safely. On the other hand, to prepare 3D porous scaffolds, the most used methods are fiber bonding, compression molding, extrusion, high internal phase emulsion templating, emulsion freeze drying, solvent casting/particulate leaching, high pressure processing, superstructure engineering, supercritical fluid processing, gas foaming/particulate leaching, thermally induced phase separation, electrospinning and rapid prototyping (46,74). Electrospinning is a versatile, simple, scalable and cost-effective spinning technique to produce scaffolds or membranes in form of fibers with a random orientation or with a certain degree of alignment (Figure I.7). Currently, several natural (hyaluronic acid, collagen, alginate, cellulose, chitosan, gelatin, pullulan, zein, Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 59 silk fibroin) or synthetic (polyamide, polycaprolactone, polylactic acid, poly (lactic-coglycolic acid), polyvinyl alcohol, polyurethane, poly (ethylene-co-vinil alcohol), polystyreneetc, etc.) (75,76) polymers are electrospun and the fibers obtained by these technique have been used in various applications such as tissue engineering scaffolds, wound dressing and in numerous biomedical applications (75–77). Figure. I.7. Schematic diagram of a typical vertical set up of electrospinning apparatus. Electrospinning technique is governed by electrostatic forces to produce fibers with different diameters from nanometer to micrometer (76,77). To produce an electrospun material polymer solutions are pumped through a needle under an electric field generated by a voltage source. When the electric field applied to the needle reaches a critical value and the repulsive electrical forces overcome the surface tension forces, a Taylor cone is formed on the needle tip (Figure I.7) (76,78–81). Solvent evaporation of the polymers solution occurs between the needle tip and the collector, leaving a solid polymer in fiber form on the collector(76,77,82–86). Several parameters must be taken into account to produce electrospun material, such as solution properties, process parameters, and environmental conditions 60 CHAPTER I | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) (76,86,87). All of them affect the process in a different way, Table I.11 summarizes these effects. Table I.11. Effects and results of different parameters on electrospun morphology (76,83). Parameters Effect on fiber morphology Results Solution parameters Polymer concentration Directly proportional. Increase in polymer concentration produce an increase in fiber diameter. Molecular weight (Mw) of polymer Directly proportional. Number of beads and droplets are reduced with high Mw. Viscosity Directly proportional. Viscosity, polymer concentration and Mw are strongly related. Optimal viscosity range is needed to produce fibers. Surface tension It is not clear. Instability of jets is obtained when surface tension is high. Conductivity/surface charge density Inversely proportional. Decrease in conductivity induces an increase in fiber diameter. Processing parameters Applied voltage Inversely proportional. Increase in voltage cause an decrease in fiber diameter. Feed rate/Flow rate Directly proportional. A decrease in flow rate generates a decrease in fiber diameter. Particles can be obtained if the flow decreases considerably. Tip to collector distance A minimum distance is necessary to facilitate the solvent evaporation and produce homogeneous fibers. Too large and too small distance can generate beads formation. Ambient parameters Humidity It must controlled A high humidity can result in beads formations and small circular pores can appear on the surface of the fiber. Temperature Inversely proportional. Increase in temperature cause a decrease in fiber diameter. Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER I 61 Scaffolds obtained by electrospinning for bone tissue engineering have received a special attention, due to its high porosity (> 90 % porosity) and its similarities with the structure and morphology of the native ECM (85). A successful delivery of differentiation and proliferation factors (growth factors, such as, BMP-2 and BMP7) (88) with the electrospun structures composed by biodegradable polymer filled or impregnated with calcium phosphate could solve the exigency of bone tissue engineering (89). On the other hand, electrospun membranes have shown a great capability for wound dressing due to the extremely high surface area, allowing adjustment of the wound moisture. The electrospun materials preserve the wound from bacterial infections due to their small pore size. The high porosity allows a good exchange of oxygen and water vapor (75–77,82–88,90–94).In the electrospinning process, it is also possible to load antibiotics and antimicrobial agents to obtain controlled delivery systems of one or multiple bioactive factors (20,76,88,94–99). Furthermore, these systems have the potential to be used as wound dressing materials, bone tissue engineering scaffolds, augmentation devices, and antimicrobial filters (29,30,34,37,75,100,101). Micro and/or nanoparticles can be obtained by electrospraying technique, using the same equipment and governed by the same principle of electrospinning (87,88). Loaded and unloaded particles can be obtained in a single step process by electrospraying without the use of any surfactant (87,93,102,103). For this purpose, the polymer solution should fulfill certain characteristics related to flow rate, polymer concentration, viscosity, and polymer molecular weight Mw. On the other hand, the use of highly volatile solvents together with the appropriate needle-collector distance could help the particles synthesis(93,103–105). Particles obtained by electrospray have a higher encapsulate ion efficiency compared with other particles production technique (88,103). 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Controlled protein release from electrospun biodegradable fiber mesh composed of poly(??-caprolactone) and poly(ethylene oxide). Int J Pharm. 2007;338(1–2):276–83. 134. Biomedical Textiles for Orthopaedic and Surgical Applications. 135. Lin J, Li C, Zhao Y, Hu J, Zhang LM. Co-electrospun nanofibrous membranes of collagen and zein for wound healing. ACS Appl Mater Interfaces. 2012;4(2):1050–7. 136. Allen-Hoffmann BL, Rooney PJ. Skin Tissue Engineering and Regenerative Medicine. In: Albanna MZ, Holmes IV JH, editors. Skin Tissue Engineering and Regenerative Medicine. Academic Press; 2016. p. 265–87. CHAPTER II Composite scaffold obtained by electro-hydrodynamic technique for infection prevention and treatment in bone repair Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER III 107 machining is an efficient technique to prepare patterning structures in electrospun polycaprolactone (PCL) membranes (10,11). It was found that electrospun poly(llactide) scaffolds with laser ablated holes exhibit significant better endothelial cell ingrowth (5). McCullen et al. (12) used laser ablation on electrospun PCL scaffolds to favour the adhesion and growth of human adipose-derived stem cells. Electrospun PCL/β-TCP fibers were also treated by laser irradiation to obtain microsized pores (13). Beside the meaningful cellular activity the material showed enhanced compressive strength. Naturally occurring polymers are well suited for various in vivo applications, promoting cell adhesion and growth. However, scaffolds fabricated from these polymers could exhibit poor mechanical properties or fast degradation rates. Native proteins, such as collagen and fibrinogen, are also more expensive and more difficult to source than synthetic polymers (14). Among the synthetic available polymer used to fabricate three-dimensional (3D) scaffolds aliphatic polyesters undoubtedly represent so far the most extensive studied class, since they combine good physico-chemical and mechanical properties with assessed biocompatibility (15). Particularly PCL is a low cost, biocompatible polymer that has a slow degradation rate and distinct rheological and viscoelastic properties that make it suitable for specific long term implantation (14). Its compatibility with a wide range of drugs enables uniform drug distribution in the matrix and its long term degradation facilitates drug release up to several months (16). Polyvinyl acetate (PVAc) polymer has also been applied in many medical fields because of its strong biocompatibility. The hydrogels containing functional groups such as COOH usually show good biocompatibility with blood, body fluids, and tissues (17). This inert polymer has the advantage that it does not induce a deleterious reaction in living tissue (18). Because of all these characteristics PVAc has been used in many medical fields, including drug and cell carries and in tissue engineering (19). One important limitation in the use of synthetic biodegradable polymers as scaffold materials is the lack of bioactivity, in particular for bone tissue applications (20). The main approach to develop bioresorbable and bioactive scaffolds is the addition of bioactive materials to the polymer matrix. Calcium 108 CHAPTER III | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) phosphate ceramics have been extensively investigated to fabricate highly porous scaffolds to engineer bone due to their near similar composition of bone, including excellent biocompatibility, osteoinductive and osteoconductive properties (3). It has been found that the addition of hydroxyapatite (HA) not only enhances the tensile strength of PCL scaffolds but also acts as a chelating agent to accelerate the mineralization of human fetal osteoblast cells to form bone-like apatite for bone tissue engineering (21). Objective The main objective of this work was to obtain a material with appropriate matrix architecture to favor cells adhesion and proliferation composed of biodegradable polymers, with mechanical resistance, good biocompatibility and high bioactivity. PCL/PVAc and PCL/PCL membranes with core-shell structure loaded with synthetic hydroxyapatite nanoparticles (HAn) to increase the bioactivity of the materials were obtained by electrospinning. In addition, the membranes obtained were treated by laser ablation to create desired microscale topographical features in order to favor cell adhesion and growth. III.2 Synthesis of inorganic nanoparticles and electrospun scaffolds III.2.1 Synthesis of hydroxyapatite nanoparticles Synthesis of HA (Ca10(PO4)6(OH)2) nanoparticles was conducted as previously described (22). CaCO3 was used as calcium sources, maintaining a Ca/P ratio of 1.67 during reaction in solution with 0.3 M H3PO4. III.2.2 Preparation of electrospun scaffolds Electrospun scaffolds were prepared using an Yflow 2.2.D-500 electrospinner (Coaxial Electrospinning Machines/R&D Microencapsulation, Malaga, Spain). PCL pellets were dissolved in DCM/DMF (1:1) and PVAc was dissolved in DMF, these solutions were stirred overnight at room temperature. To prepare PCL-HAn scaffolds, HAn powder was dispersed in DCM/DMF with the help of TWEEN® 80 by stirring overnight at room temperature, then this solution was added to the PCL solution and stirred overnight at room temperature. The polymers solutions were Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER III 109 loaded into 20 mL plastic syringes. The PCL-HAn suspension was fed through the inner needle of the coaxial system and PCL or PVAc solutions were fed through the outer one. This last needle was connected to the positive voltage power supply, at a voltage ranging from 7 to 13 kV. The shell and core flow rates and the spinning distance were fixed to 0.5 mL/h (in both needles) and 19 cm respectively. The spun fibers were collected on a static plate (covered with aluminum foil) connected to negative voltage power supply, at a voltage ranging from −2 to 4 kV for 18 h. To create macropores in the fibrous membrane structures, a pulse Nd:YAG laser (TECHNOLOGY Q-Switch) was used. Different conditions were used for each material: output power: 16 W, wavelength: 1064 nm for the PCL-HAn/PCL fibers and output power: 1 W, wavelength: 532 nm for PCL-HAn/PVAc fibers. The characterization techniques and different evaluation methods are described in Appendix 1. III.3 Results and discussion III.3.1 Hydroxyapatite characterization The XRD pattern for the synthesized HA nanoparticles is shown in Figure III.1a, it could be seen that all the peaks could be indexed to the hexagonal phase hydroxyapatite (JCPDS N° 09-0432). The diffractogram do not show any other peaks corresponding to secondary phases or intermediate compounds suggesting the formation of pure HAn phase. Intense diffraction peaks with broad width are indicative of the crystalline nature of the prepared material and the small crystallite size respectively (23). Inset of Figure III.1a shows the peaks corresponding to the (002) and (211) diffraction planes of hydroxyapatite. According to the JCPDS card 09-0432, the reference intensity ratio for (211), (002) is RI(211)/RI(002) = 2.61. The calculated degree of texture index (24) for the synthesized hydroxyapatite is 0.51 indicating a preferred growth orientation along the c-axis (25). These results are consistent with electronic microscope images (Figure III.1b and III.1c), which show fibrous needle-like particles with mean diameters of 20 nm and length of approximately 150 nm. The rod-like shape of the nanoparticles is more clearly seen in TEM images (Figure III.1c). The FTIR spectrum (Figure III.1.d) of the material presents the characteristic bands for PO43- 110 CHAPTER III | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) appearing at 472, 572 and 603 cm−1 related to the asymmetric bending, and signals assigned to symmetric and asymmetric stretching positioned in the 900-1200 cm−1 range (26). The broad peak around 3455 cm−1 is due to the adsorbed water on HAn structure and the absorption peak at 1635 cm−1 is attributed to the bending mode of OH− groups (27). The absorption peak assigned to apatite hydroxyl bond is observed at 3569 cm−1 (28). The shoulder at 878 cm−1 together with the doublet at 1415 and 1458 cm−1 indicate the existence of CO32− (29) probably coming from the atmosphere carbon dioxide during sample preparation and would have been incorporated into the HAn crystal structure (30). All these results imply that the synthesis of rod-like shaped hydroxyapatite nanoparticles was successful. Figure III.1. Characterization of the synthesized HAn: a) XRD, b) SEM images, c) TEM micrographs, d) FTIR spectrum. Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER III 111 III.3.2 Scaffolds characterization III.3.2.1 As spun scaffolds The materials prepared by electrospinning are listed in Table III.1 together with positive and negative voltages applied. Due to the properties (viscosity, surface tension, conductivity) of the solutions the voltages required to obtain a stable Taylor cone were higher to produce fibers containing hydroxyapatite than the voltages used for the pristine polymer fibers (Table III.1). Table III.1. Preparations conditions and characterization results for the obtained electrospun scaffolds. Sample (Core/Shell) PCL/PCL PCL-HAn/PCL PCL/PVA c PCL-HAn/PVAc Applied voltage -/+ (kv) 3.11/10.25 3.54/12.00 2.87/7.21 3.03/8.57 HAn loada (wt%) - (14.2) (13.8) - (11.9) (12.7) Fiber diameterb (nm) 206±62 210±64 264±56 225±47 Porosity (%) 62.0 63.5 65.5 58.6 Pores diameter (nm) 232 210 209 188 a From TGA results, theoretical value between brackets. b Obtained by measuring at least 100 fibers in different SEM images. The HAn load in the fibers was measured by thermogravimetric analysis (TGA, Figure III.2) as the residue after complete polymers elimination at 600 °C (Table III.1). The experimental values of 13.8 and 12.7 wt% for PCL-HAn/PCL and PCL-HAn/PVAC respectively, were close to the theoretical ones (13.2 and 11.9 wt%). The HAn load used was the maximum possible to obtain a stable Taylor cone; higher amount of nanoparticles in the spinning solution increase excessively the viscosity and also hinder good nanoparticles dispersion. 112 CHAPTER III | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Figure III.2. TGA thermogram of PCL-HAn/PCL and PCL-HAn/PVAc fibers. Figure III.3. XRD patterns of PCL/PCL, PCL/PVAc, PCL-HAn/PCL and PCLHAn/PVAc fibers. Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER III 113 XRD patterns of fibers with HAn are shown in Figure III.3. The XRD diffractgram of PCL-Han/PCL has peaks for both hydroxyapatite and PCL. Both the polymer and ceramic retain their crystalline behavior in the electrospun fibers indicating insignificant change in their crystal structure. This confirms the fact that HAn is finely and uniformly dispersed in the matrix (3). In PCL-HAn/PVAc pattern peaks at 25 and 35° establish the presence of hydroxyapatite incorporated in the scaffold (inset of Figure III.3). Figure III.4. SEM images of a) PCL/PCL fiber, b) PCL-HAn/PCL fibers. FE-SEM images of the electrospun scaffolds showed uniform, beadless and nano-scaled fibrous structures randomly oriented for all the prepared materials under the optimum spinning conditions utilized in each case. As an example Figure III.4a) shows SEM micrographs obtained for fibers with PCL in the core and in the shell. Hydroxyapatite nanoparticles are mostly encased inside the fibers, but some agglomerates are observed on the surface (Figure III.4b). The fibers average diameter was measured from SEM images (Table III.1). The presence of hydroxyapatite nanoparticles in the inner solution does not seem to have any effect on the PCL/PCL fiber diameter. It was previously observed for PCL fiber without defects, such as the obtained in this work, that for HA load higher than 5 wt% the average diameter was similar to the pristine fibers (31). This result for b) a) 50 µm 3 µm 500 nm 50 µm 3 µm 500 nm 114 CHAPTER III | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) high mass fraction was attributed to the increased viscosity of the dispersion during electrospinning. For the PCL/PVAc material the presence of hydroxyapatite leads to a slightly lower fibers mean diameter, probably due to the higher voltage needed to obtain a stable cone in the case of the hydroxyapatite loaded fibers (32). The core-shell structure of the PCL/PVAc fibers was investigated by TEM microscopy as shown in Figure III.5a. The PVAc polymer (the transparent part of the outside structure of the fiber) is the shell wrapping uniformly the PCL fiber (the darker part of the inside structure of the fiber) as the core. A TEM micrograph of a PCL/PCL fiber was added in the inset for comparison purpose, as expected core and shell are indistinguishable in this case. The distribution of the HA nanoparticles inside fibers can be observed in Figure III.5b, as mentioned before most of the particles are encapsulated into the polymers. Figure III.5. TEM images of a) PCL/PVAc fiber, inset PCL/PCL fiber, b) PCLHAn/PVAc fiber. III.3.2.1 Laser treated scaffolds After laser irradiation the scaffolds were analyzed by SEM microscopy (Figure III.6). For both materials the laser energy was not sufficient to machine a hole through the electrospun fibers. The images show that the membrane was not significantly affected outside the laser irradiation area. Figure III.6a shows SEM 200 nm 200 nm 200 nm Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER III 115 images of PCL-HAn/PCL ablated scaffolds, the micropores were well obtained with diameters in the range 70–120 μm. It is important to observe the morphology of the remaining fibers since the aim is to improve the porosity without any additional effect (33). Even when some melting and coalescence of the fibers can be observed around the drilled holes, the surface morphology is only slightly changed. For PCL-HAn/PVAC on the other hand, the pores diameter are in the 50– 90 μm range and the change in the fibers morphology around the holes is more important. Besides, in this material not all the holes of the rectangular pattern were produced under irradiation. As mentioned in the experimental section different laser were used in order to get the pores in each material. Figure III.6. SEM images and pore size distribution of laser treated a) and c) PCLHAn/PCL and b) and d) PCL-HAn/PVAc. 116 CHAPTER III | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) III.3.3 In vitro bioactivity Figure III.7. SEM images of the prepared materials after immersion in SBF solution during different times. Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER III 123 Figure III.13. Confocal microscope images of scaffolds surface (left) and viability (right) of human osteoblasts seeded in PCL-HAn/PVAc scaffolds not treated (above) and treated with laser (below) after 14 days of cell seeding. Scale bars 100 μm. Previous studies have reported the biocompatibility of PCL electrospun 3Dscaffolds in a human osteoblast in vitro model in which the high porosity enabled cells to penetrate into the scaffold (38). The grafting of collagen and chondroitin sulphate on modified surface PCL porous scaffolds synthesized by particulate leaching significantly increased the in vitro proliferation of murine chondrocytes four weeks after seeding though porosity was not altered (39) while chemically cross-linked PCL and HAn nanoparticles used to fabricate nanocomposite scaffolds loaded with the growth factor BMP-2 also showed very good cytocompatibility in a rabbit bone marrow stem cells in vitro model (40), supporting PCL suitability as biomedical material. In addition, the fabrication of PCL scaffolds treated with a 124 CHAPTER III | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) femtosecond laser to create pores, and therefore to modify the scaffold surface and porosity, has shown enhanced cellular activities compared to those scaffolds with the same pore size and not treated with laser (41), indicating that laser treatment may significantly improve the potential of these types of scaffolds in bone regeneration as our study shows. Though it is controversy regarding the “ideal” size pore for biomedical applications, in osteoregeneration most authors have pointed to 100–400 μm as recommended to facilitate cell adhesion and growth (42). However, smaller pores are able to increase scaffolds surface and lead to higher cell attachment while larger pores facilitate cell migration (41). Our results showed that laser pulse on PCL-HAn/PVAc scaffolds implied the formation of micropores of 50–90 μm while on PCL-HAn/PCL were slightly larger (70–120 μm). These data, together with the apparently higher cell density recorded in PVAc containing scaffolds, suggest that smaller micropore sizes made easier cell adhesion and proliferation. The incorporation of HAn in the synthesis process of poly(L/DL)-lactide (43) or polylactic acid (PLA) (44) electrospun scaffolds for biomedical applications has been reported in order to improve cell attachment and proliferation. Since this effect is not clear in Figures III.9-III.11, the cell viability 14 days after seeding NHOst on scaffolds with and without HAn was measured by image quantification (Figure III.12). Results show that cell viability on samples with HAn are in the range of the ones without it. Besides, they support our observations pointing to a higher viability of cells seeded on PVAc containing scaffolds, exerting significant differences among PCL-HAn/PVAc group (treated and not treated with laser) vs PCL/PCL and PCL-HAn/PCL groups. Laser treated scaffolds also displayed higher viability percentages than the not treated ones highlighting the laser treatment as an improvement for cell attachment and viability. To our knowledge, this is the first time that the incorporation of PVAc to PCL electrospun scaffolds and treated with laser to enhance their suitability in biomedical applications has been shown these promising effects regarding human osteoblasts adhesion and proliferation, pointing to its potential application in bone repair approaches. Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER III 125 III.4 Conclusions Rod shaped hydroxyapatite nanoparticles were successfully synthesized and incorporated into core-shell PCL/PCL and PCL/PVAc electrospun nanofibers. Only HAn loaded fibers presented hydroxyapatite precipitation after been soaked in SBF for 14 and 30 days. The presence HAn particles would be necessary for the apatite formation on the nanofiber surface. Fibers with PVAc shell suffer some structural change but they still exhibited a well interconnected pore network structure. Even when human osteoblasts growth was observed on all seeded surface, the laser treatment of the surfaces notably increased the scaffolds biocompatibility observing higher viability and cell density. 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Biomacromolecules. 2014;15(3):1019–30. 41. Kim M, Son J, Lee H, Hwang H, Choi CH, Kim G. Highly porous 3D nanofibrous scaffolds processed with an electrospinning/laser process. Curr Appl Phys. 2014;14(1):1–7. 42. Roosa SMM, Kemppainen JM, Moffitt EN, Krebsbach PH, Hollister SJ. The pore size of polycaprolactone scaffolds has limited influence on bone regeneration in an in vivo model. J Biomed Mater Res - Part A. 2010;92(1):359–68. 43. Rajzer I, Menaszek E, Kwiatkowski R, Chrzanowski W. Bioactive nanocomposite PLDL/nano-hydroxyapatite electrospun membranes for bone tissue engineering. J Mater Sci Mater Med. 2014;25(5):1239–47. 44. Morelli S, Salerno S, Holopainen J, Ritala M, De Bartolo L. Osteogenic and osteoclastogenic differentiation of co-cultured cells in polylactic acidnanohydroxyapatite fiber scaffolds. J Biotechnol. 2015;204:53–62. CHAPTER IV Polymeric electrospun scaffolds for bone morphogenetic protein 2 delivery in bone tissue engineering. The contents of this chapter have been adapted from the following published work: Polymeric electrospun scaffolds for bone morphogenetic protein 2 delivery in bone tissue engineering. Javier Aragó n, Simóna Salernó, Lóredana De Bartóló, Silvia Irusta and Gracia Mendóza. Jóurnal óf Cóllóid and Interface Science, 531 (2018) 126–137. DOI:10.1016/j.jcis.2018.07.029. “The development of novel scaffolds based on biocompatible polymers is of great interest in the field of bone repair for fabrication of biodegradable scaffolds that mimic the extracellular matrix and have osteoconductive and osteoinductive properties for enhanced bone regeneration. Polycaprolactone (PCL) and polycaprolactone/polyvinyl acetate (PCL/PVAc) core– shell fibers were synthesised and decorated with poly(lactic-co-glycolic acid) [PLGA] particles loaded with bone morphogenetic protein 2 (BMP2) by simultaneous electrospinning and electrospraying. Hydroxyapatite nanorods (HAn) were loaded into the core of fibers. The obtained scaffolds were characterized by scanning and transmission electron microscopy, Fourier-transform infrared spectroscopy, and thermogravimetric analysis. The in vitro potential of these materials for bone regeneration was assessed in biodegradation assays, osteoblast viability assays, and analyses of expression of specific bone markers, such as alkaline phosphatase (ALP), osteocalcin (OCN), and osteopontin (OPN). PLGA particles were homogeneously distributed in the entire fiber mat. The growth factor load was 1.2–1.7 µg/g of the scaffold whereas the HAn load was in the 8.8–12.6 wt% range. These scaffolds were able to support and enhance cell growth and proliferation facilitating the expression of osteogenic and osteoconductive markers (OCN and OPN). These observations underline the great importance of the presence of BMP2 in scaffolds for bone remodeling as well as the good potential of the newly developed scaffolds for clinical use in tissue engineering.” CHAPTER IV ................................................................................................................................................................... 130 Polymeric electrospun scaffolds for bone morphogenetic protein 2 delivery in bone tissue engineering. ................................................................................................................................................... 130 IV.1 Introduction ...................................................................................................................................................... 133 Objective ...................................................................................................................................................................... 134 IV.2 Scaffold fabrication ........................................................................................................................................ 135 IV.3 Results and discussion .................................................................................................................................. 137 IV.3.1 Scaffold characterization .................................................................................................................... 137 IV.3.2 In vitro protein release ........................................................................................................................ 141 IV.3.3 In vitro enzymatic degradation ........................................................................................................ 144 IV.3.4 Cell viability and morphology .......................................................................................................... 148 IV.3.5 Osteogenic, osteoinductive, and osteoconductive activities of scaffolds ....................... 150 IV.4 Conclusions ........................................................................................................................................................ 155 References ................................................................................................................................................................... 156 Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER IV 139 was a slightly higher concentration in the shell, the fluorescent compound was almost uniformly distributed inside the microparticles. No significant phase separation of the polymer and/or drug was observed. HAn was also characterized by means of the SEM and TEM (Figure IV.3). These images confirmed our previous results (4), in which fibrous needle-like particles with a rod-like shape were also seen (length and diameter of approximately 150 and 20 nm, respectively). Figure IV.3. SEM (a) and TEM (b) images of synthesized HAn. Once the particle synthesis was optimized, the simultaneous electrosprayelectrospinning process was carried out to generate the final scaffolds of HAn-loaded PCL/PVAc fibers decorated with PLGA particles (Figure IV.4). The optimized electrospray-electrospinning process produced fibers with uniform morphology and a random distribution that led to the creation of a scaffold with porosity of ≈60–64%. Because of the simultaneous synthesis, PLGA particles with a mean diameter of 1.2 ± 0.5 µm were homogeneously distributed in the entire fiber mat. It is noteworthy that the attachment of particles onto fibers did not influence fiber morphology (Figure IV.4). However, some of the particles slightly deviated from the spherical shape (Figures IV.2 and IV.4). The mean diameter of fibers (380 ± 108 nm) forming the scaffolds was in the range of collagen fibers (from 50 to 500 nm), mimicking the nanostructures of the 140 CHAPTER IV | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) natural ECM. SEM images clearly revealed the formation of scaffolds with an interconnected network of large pores with size between 0.9 and 2.5 µm for PCL and PVAc scaffolds and smaller pores ranging from 0.2 to 0.6 µm. The mean pore size was 1.14 µm for PVAc and 1.46 µm for PVAc:PLGA-BMP2 owing to the presence of a nanofibrous structure that confers microporosity on the scaffold (Figure IV.4). This structure provides a large surface area-to-volume ratio for cell attachment as well as sufficient porosity for nutrient diffusion. Indeed, the measured hydraulic permeance was 8.00 ± 0.30 and 6.60 ± 0.03 L/h m2 mbar for PVAc and PCL scaffolds, respectively. An increase in hydraulic permeance by 16% was observed in the PVAc:PLGA-BMP2 scaffolds (9.5 ± 0.4 L/h m2 mbar) with respect to the non-loaded scaffold. Figure IV.4. SEM images of the synthesised electrospun fibres (left) decorated with BMP2-loaded PLGA electrosprayed particles (right). Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER IV 141 TGA analyses (Figure IV.5) confirmed HAn loads of 13.8 and 12.7 wt% for PCL and PVAc fibers, respectively, in agreement with our previous results (4). In PCL:PLGABMP2 and PVAc:PLGA-BMP2 scaffolds, owing to the presence of the PLGA polymer, the HAn load was reduced to 12.6 and 8.8 wt%, respectively. However, BMP2 encapsulation efficiency was similar between the two formulations, reaching percentages of 39% ± 5% for PCL:PLGA-BMP2 and 40% ± 11% for PVAc:PLGA-BMP2. A final BMP2 concentration of 39 and 40 µg/ml was achieved in the scaffolds, in line with the clinical required dose (21). These similar efficiency rates led to growth factor loads of 1.2 and 1.7 µg/g in PCL and PVAc composite fibers, respectively. Figure IV.5. TGA thermogram of the four types of scaffolds synthesized: PCL, PVAc,PCL:PLGA-BMP2 and PVAc:PLGA-BMP2. IV.3.2 In vitro protein release The kinetic profile of the BMP2 release from both kinds of scaffolds was determined in the course of 28 days (Figure IV.6a) and the morphology of PVAc:PLGABMP2 scaffolds after the release for 28 days was also studied under the SEM (Figure IV.7). 142 CHAPTER IV | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Figure IV.6. BMP2 release kinetic curves obtained for both types of scaffolds (a) and their mathematical adjustment (b). The pattern of the growth factor release from both scaffolds showed a strong initial burst within the first 8 h, followed by a slow release until the end of the experiment. This behavior should synergistically enhance bone generation because it was suggested that the ideal BMP2 release strategy includes both an initial burst and a subsequent sustained release. The reason is that the former helps to recruit osteoprogenitor cells to the delivery system and the latter promotes osteogenic differentiation (22). The burst release reached ≈ 68 % of the BMP2 load for PCL:PLGABMP2 and only 40 % for PVAc:PLGA-BMP2. The difference could be due to the absorption capacity of PVAc (350 % water absorption in 24 h (23)) that may retain the protein released from the PLGA particles during the swelling process. It is important to point out that the hydrophilicity of both membranes is determined by the presence of HA, which turns both scaffolds into highly hydrophilic materials. SEM images of scaffolds (Figure IV.7) after the release (28 days) confirmed this pattern: there were no PLGA particles on the surface of the scaffolds or only a few damaged ones. Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER IV 143 Figure IV.7. SEM images of PVAc:PLGA-BMP2 scaffolds before and after release for 28 days. The kinetic mathematical models were fitted to the experimental results of the release (Table IV.2, Figure IV.6b). Preliminary calculations (not shown) indicate that the Higuchi model was not satisfactory for fitting to the experimental release data. Given that some of these models should be used only for the first 60% of the release, only the data representing the first 24 h were fitted because it was the release from PCL:PLGABMP2 achieved at that time (24). The best fit solution was identified by evaluating coefficient of correlation R2. The highest values of the coefficient indicated that both samples were better described by the Peppas–Sahlin release kinetic model (Figure IV.6b). In this model, term k1Ptn represents the Fickian diffusional contribution to the release (F), whereas term k2t2n is the case-II relaxational contribution (R). The negative values obtained for k2 should be interpreted in terms of a relaxation mechanism being insignificant compared to the diffusion process (25). In accordance with the Peppas– Sahlin equation, the value of exponent n for the Fickian release mechanism from a polymeric system with sphere geometry should be ≈0.43 (26). The lower values found for PCL and PVAc samples may be related to the wide particle size distribution mentioned before (27). 144 CHAPTER IV | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) Table IV.2. Release modeling parameters. Model Parameter parameters PCL:PLGABMP2 PVAc:PLGABMP2 Korsmeyer-Peppas k1K 0.42 0.21 n 0.12 0.18 R2 0.77 0.89 Peppas-Sahlin k1P 0.50 0.23 k2 -0.11 -0.04 n 0.30 0.38 R2 0.91 0.98 IV.3.3 In vitro enzymatic degradation An ideal scaffold for bone tissue engineering should be biodegradable and bioresorbable while being able to support the growth of new bone. The degradation behavior of biomaterials in physiological environments plays an important role in the engineering process of a new tissue. The greatest advantage of degradable polymers such as PCL, PVAc, and PLGA is that they are broken down into biologically acceptable molecules that are metabolized and removed from the body via normal metabolic pathways (28). To investigate the effect of enzymes on the degradation behavior of the synthesized scaffolds, they were incubated in phosphate-buffered saline in the presence of lipase or lysozyme for different periods at the same concentrations as those found in human serum. The degradation was followed by gravimetric measurements (Figure IV.8). PVAc and PCL scaffolds manifested similar behavior in the presence of lysozyme, thus reaching degradation of 10 % ± 1 % and 13 % ± 1 %, respectively, after 60 days. This value for PCL-based fibers was higher than the one expected based on the results of Banerjee el at. (29). These authors did not find any significant PCL film degradation in the presence of lysozyme. The greater degradation observed in our PCL scaffolds could be explained by the fibrous structure and high porosity that increase the surface area exposed to the enzyme solution. In the presence of lipase, the degradation profile of all scaffolds was similar to that observed in other studies, showing a significant weight loss with time, which reached values ≈ 90 % after 60 days (29). Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER IV 145 Figure IV.8. Enzymatic degradation of membranes by lipase (a) and lysozyme (b). Table IV.3. Area ratio in lipase and lysozyme degradation media. SAMPLE Area ratio C=O/ PO4-3 C-O-C/PO4-3 PCL 2.9 0.7 PCL_LIPASE 0.5 0.2 PCL_LYSOZYME 2.2 0.6 PCL:PLGA-BMP2 2.1 0.6 PCL:PLGA-BMP2_LIPASE 0.2 0.1 PCL:PLGA-BMP2_LYSOZYME 1.9 0.6 PVAC 2.4 1.8 PVAC_LIPASE 1.7 1.7 PVAC_LYSOZYME 2.2 2.0 PVAC:PLGA-BMP2 2.1 1.4 PVAC:PLGA-BMP2_LIPASE 2.3 2.0 PVAC:PLGA-BMP2_LYSOZYME 2.2 1.6 The effects of lipase and lysozyme on the PCL were corroborated by infrared (IR) spectroscopy (Figure IV.9). Bands at 1725 and 1240 cm1, related to C=O and C-O-C 146 CHAPTER IV | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) bonds, respectively, slightly changed after enzymatic degradation. Besides, the ratios between signals assigned to PO43 (1040 cm1), owing to the presence of HAn, and the above-mentioned polymer bands indicate a minor change in sample polymer content after the enzymatic attack (Table IV.3). Even when it is known that PVAc can be degraded by microorganisms, especially filamentous fungi (30,31), to our knowledge, there is no published study about PVAc degradation by lysozyme. The mass decrease observed in the PVAc scaffold may be related only to the PCL degradation. Figure IV.9. FTIR spectra of PCL (a, b) and PVAc (c, d) scaffolds loaded (b, d) and not loaded (a, c) with PLGA-BMP2 particles after enzymatic degradation with lipase and lysozyme. In the presence of lipase, the fibers with PVAc in the shell showed lower degradation than did the fibers consisting only of PCL (Figure IV.8a). After 40 days, PCL lost almost 75% of the mass, while PVAc scaffolds lost only 29% of their mass, as expected because PCL can be hydrolytically and enzymatically degraded by lipase through hydrolysis of the ester bond (32). On the other hand, each lipase has distinct specificity towards the side chain hydrolysis of PVAc (33). The polymer backbone is not hydrolyzed, and only the side chains containing ester linkages undergo hydrolysis Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER IV 147 catalyzed by lipase. The smaller mass loss of PVAc fibers could be due not only to the smaller amount of PCL in the sample (48 and 86 wt% for PVAc and PCL samples, respectively) but also to the reported reduction in the degradation of PCL with the addition of PVAc in PCL–PVAc blends (34). PCL degradation was also confirmed by IR spectroscopy (Figure IV.5a), which revealed relative enhancement of a band at 1040 cm1 associated with HAn, thereby implying a polymer mass decrease. On the other hand, the addition of PLGA particles to the fibers slightly changed the degradation of PVAc scaffolds in the presence of lysozyme and did not affect the behavior with lipase. PLGA particles synthesized from different polymer formulations may possess different degradation rates because the average molecular weight and lactide:glycolide ratio affect the diffusion rate and permeability of the PLGA polymeric matrix, which consequently affect the degradation rate (35). PLGA (75:25) requires more than 56 days to lose ≈ 8 % of its mass in the presence of lysozyme (36). The polymer used in this work (50:50) is a more degradable formulation because of the preferential degradation of the glycolic acid domain owing to its higher hydrophilicity (37). However, it is frequently used to encapsulate lysozyme, and the release in PBS takes place via diffusion through pre-existing pores and channels in the polymer matrix (38). This fact suggests that the degradation of PLGA 50:50 should be very slow, and therefore the release of the remaining protein (32 % for PCL:PLGA-BMP2 and 60 % for PVAc:PLGA-BMP2) may be sustained long-term while PLGA particle degradation proceeds. This process may enable a BMP2 release in the nanogram range as reported to take place during in vivo normal bone regeneration (39) as well as during successful in vivo bone repair in an experimental model (40). As a consequence, the prepared scaffold may manifest both required release types (an initial burst followed by a sustained release) to enhance bone regeneration. It has been extensively studied and proven that biodegradation of HAn is very limited. However, due to the highly osteoconductive nature of HAn, its use in bone graft substitutes is crucial for improving the cellular response and the mineralization process (41). 148 CHAPTER IV | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) IV.3.4 Cell viability and morphology Human osteoblasts seeded onto the different types of the newly developed scaffolds adhered and spread, thus forming (after 2 weeks) a layer that covered the surface as a result of their proliferation (Figures IV.10, IV.11 and IV.12). The nanofibrous structure of scaffolds mimicking the architecture of the ECM favored adhesion and proliferation of osteoblasts as well as the maintenance of their phenotype. Figure IV.10. Cell viability of osteoblasts seeded in a 2D system and onto PCL and PVAc scaffolds, decorated and not decorated with PLGA-BMP2 particles, at different time points. Mean ± SD (9 samples). Differences between groups were considered significant when p < 0.05. MTT assays (Figure IV.10) showed cell viability and proliferation during the experiments for up to 4 weeks after seeding. Osteoblast seeding in a 2D environment (directly onto a culture plate well) yielded slightly faster growth after 2 and 3 weeks, as compared to the absorbance recorded in the first week, and decreased again after 4 weeks. This decrease is consistent with saturation of the growth surface, preventing cell proliferation and decreasing cell viability. Furthermore, the results obtained clearly show the significant differences in cell viability and proliferation between 2D and 3D Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) | CHAPTER IV 155 IV.4 Conclusions It has been previously reported that electrospun PCL nanofibers loaded with HA promote apatite formation, whereas the presence of PVAc in the fiber shell increases hydrophilicity and favors osteoblast adhesion and proliferation (4). Given that BMP2 could be an alternative way to increase osteogenic activity (8), in this work, those fibers were decórated with PLGA‐BMP2 particles óbtained by electróspraying for osteoinductive and osteoconductive purposes in bone regeneration. Fibers and particles kept their size and structure after the particles were electrosprayed, though HAn loading was slightly lower (≤ 3.9 %) when PLGA was present in the samples. BMP2 was successfully loaded into PLGA particles, and their loading and activity after electrospraying were confirmed, yielding gr ówth factór lóads up tó 1.7 μg/g in PCL:PLGA-BMP2 and PVAc:PLGA-BMP2 scaffolds with a final BMP2 concentration of 39 and 40 µg/mL, respectively; a mid-range concentration can fuse bone defects without adverse effects (53). The BMP2 release was 40–68 %, being lower for PVAc-containing scaffolds; this phenomenon may be attributed to their absorption capacity. Enzymatic degradation of the scaffolds highlighted their good biodegradation profile, supporting their good potential for bone regeneration. In addition, osteoblast viability and proliferation increased when the FDA-approved growth factor BMP2 was present in the scaffolds; the expression of bone formation and maturation markers was also improved by BMP2. Taken together, these results suggest that BMP2 is a key factor for the potential improvement of PCL/PVAc scaffolds. 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Cambridge University Press; 2013 Jan;1502:mrsf12-1502t01-09. 52. Chen D, Harris MA, Rossini G, Dunstan CR, Dallas SL, Feng JQ, et al. Bone morphogenetic protein 2 (BMP-2) enhances BMP-3, BMP-4, and bone cell differentiation marker gene expression during the induction of mineralized bone matrix formation in cultures of fetal rat calvarial osteoblasts. Calcif Tissue Int. 1997 Mar;60(3):283–90. 53. Halling Linder C, Ek-Rylander B, Krumpel M, Norgård M, Narisawa S, Millán JL, et al. Bone Alkaline Phosphatase and Tartrate-Resistant Acid Phosphatase: Potential Co-regulators of Bone Mineralization. Calcif Tissue Int. Springer US; 2017 Jul;101(1):92–101. CHAPTER V Electrospun asymmetric membranes for wound dressing applications The contents of this chapter have been adapted from the following submitted work (under review): Electrospun asymmetric membranes for wound dressing applications. Javier Aragó n, Clarinda Cósta, Isabel Cóelhósó, Gracia Mendóza, Ana Aguiar-Ricardó and Silvia Irusta. Materials Science and Engineering C (MSEC_2018_3013) “To accomplish a rapid wound healing it is necessary to develop an asymmetric membrane with interconnected pores consisting of a top layer that prevents rapid dehydration of the wound and bacteria penetration and a sub-layer with high absorption capacity and bactericidal properties. Polycaprolactone (PCL)/polyvinyl acetate (PVAc) asymmetric membranes loaded with the bactericidal monoterpene carvacrol (CRV) were synthesized and characterized by scanning electron microscopy and Fourier transform infrared spectroscopy. Mechanical properties in dry and wet conditions and fluid handling behavior were also assessed. In addition, biological studies regarding their bactericidal effects, cytocompatibility and wound closure properties were also developed. Loading efficiencies of 40-50 % were achieved in the prepared samples and 85-100 % of the loaded CRV was released in simulated wound pH evolution medium. The significant inhibition of Gram negative (Escherichia coli S17) and Gram positive (Staphylococcus aureus ATCC 25923) bacteria growth clearly showed the suitability of the fabricated membranes for wound healing applications. Furthermore, cytocompatibility of the loaded membranes was demonstrated both in 2D and 3D human dermal fibroblast cultures, as well as cell migration was not impaired by released carvacrol from the membranes. These results highlight the potential of these polymeric electrospun membranes for wound healing.” CHAPTER V ......................................................................................................................................................... 161 Electrospun asymmetric membranes for wound dressing applications ............................. 161 V.1 Introduction ..................................................................................¡Error! Marcador no definido. Objective ................................................................................................¡Error! Marcador no definido. V.2 Membrane preparation ............................................................¡Error! Marcador no definido. V.3 Results and discussion .............................................................¡Error! Marcador no definido. V.3.1 Membranes characterization by SEM and FTIR ....¡Error! Marcador no definido. V.3.2 Mechanical properties .....................................................¡Error! Marcador no definido. V.3.3 Fluids handling properties .............................................¡Error! Marcador no definido. V.3.4 Carvacrol release ................................................................¡Error! Marcador no definido. V.3.5 Antimicrobial properties ................................................¡Error! Marcador no definido. V.3.6 Cytocompatibility...............................................................¡Error! Marcador no definido. V.3.7 Cell scratch assay ...............................................................¡Error! Marcador no definido. V.4 Conclusions ...................................................................................¡Error! Marcador no definido. References .............................................................................................¡Error! Marcador no definido. 190 GENERAL CONCLUSIONS | Erasmus Mundus Doctorate in Membrane Engineering (EUDIME) GENERAL CONCLUSIONS The materials prepared in this work have demonstrated the potential of electrodynamic technique to obtain biomaterials for two different applications, bone scaffolds and wound dressing membranes. The most relevant conclusions of each prepared material are summarized in this chapter. Three polymeric multifunctional scaffolds were developed using electrospinning and electrospray for bone repair and the prevention or treatment of bone infection.  PCL and PCL/PVAc core/shell fibers obtained by electrospinning can be decorated with RFP or BMP2 loaded PLGA microparticles using the electrospraying technique.  The incorporation of HAn particles in the polymer fibers favors the apatite formation on nanofiber surface.  The combination of PCL, PVAc, and PLGA-BMP2 is a novel and promising therapeutic approach for bone repair. Given that BMP2 could be an alternative way to increase osteogenic activity and the PVAc fiber shell increases hydrophilicity and favors osteoblast adhesion and proliferation.  Enzymatic degradation of the scaffolds highlighted their potential for bone regeneration.  Laser treatment of the surfaces notably increases the scaffolds cytocompatibility resulting in higher viability and cell density.  The synthesized scaffold shows a porous network providing a large surface area-to-volume ratio for cell attachment as well as sufficient porosity for nutrient diffusion.  The obtained scaffold presents a homogeneous distribution of RFPPLGA microparticles along the entire scaffolds thereby ensuring a continuous release of the RFP.