Nanotechnology-based approaches for bone tissue engineering

– I – Mar Bonany Mariñosa Nanotechnology-based Approaches for Bone Tissue Engineering Supervisors: Prof. Maria Pau Ginebra Molins Dr. Montserrat Español Pons Doctoral Program of Biomedical Engineering Universitat Politècnica de Catalunya – II – Nanotechnology-based Approaches for Bone Tissue Engineering Mar Bonany Mariñosa ADVERTIMENT La consulta d’aquesta tesi queda condicionada a l’acceptació de les següents condicions d'ús: La difusió d’aquesta tesi per mitjà del repositori institucional UPCommons (http://upcommons.upc.edu/tesis) i el repositori cooperatiu TDX (http://www.tdx.cat) ha estat autoritzada pels titulars dels drets de propietat intel·lectual únicament per a usos privats emmarcats en activitats d’investigació i docència. No s’autoritza la seva reproducció amb finalitats de lucre ni la seva difusió i posada a disposició des d’un lloc aliè al servei UPCommons o TDX. No s’autoritza la presentació del seu contingut en una finestra o marc aliè a UPCommons (framing). 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In the using or citation of parts of the thesis it’s obliged to indicate the name of the author. – III – PhD Thesis NANOTECHNOLOGY-BASED APPROACHES FOR BONE TISSUE ENGINEERING Doctoral Program of Biomedical Engineering Mar Bonany Mariñosa Supervisors: Prof. Maria Pau Ginebra Molins Dr. Montserrat Español Pons Biomaterials, Biomechanics and Tissue Engineering Group Department of Materials Science and Engineering Universitat Politècnica de Catalunya Barcelona, 2021 – IV – Als meus pares, que s’ho mereixen tot. I a la meva àvia, que és bona, forta i entregada. Au milieu de l’hiver, j’apprenais enfin qu’il y avait en moi un été invincible. Albert Camus – V – Table of Contents ABSTRACT ............................................................................................................... VIII RESUM .......................................................................................................................... IX ACKNOWLEDGEMENTS ........................................................................................... X OBJECTIVES OF THE THESIS ............................................................................ XIII ABBREVIATIONS .................................................................................................... XIV CHAPTER 1: INTRODUCTION .................................................................................... 1 1.1. Bone Tissue ............................................................................................................... 2 1.1.1. Function ............................................................................................................... 2 1.1.2. Bone Structure ..................................................................................................... 2 1.1.3. Bone Composition ............................................................................................... 3 1.1.4. Bone Cells ........................................................................................................... 5 1.1.5. Bone Remodelling Process .................................................................................. 5 1.2. Bone Affections ......................................................................................................... 6 1.2.1. Bone Cancer ........................................................................................................ 6 1.2.2. Bone Defects ....................................................................................................... 7 1.3. Bone Cancer Therapy .............................................................................................. 7 1.3.1. Hydroxyapatite Nanoparticles in Cancer Treatment ........................................... 7 1.3.2. Cytotoxicity Mechanisms of Hydroxyapatite Nanoparticles .............................. 8 1.4. Bone Grafting Strategies .......................................................................................... 9 1.4.1. Bone Grafts of Natural Origin ............................................................................. 9 1.4.2. Synthetic Bone Grafts ....................................................................................... 10 1.4.3. Three-Dimensional Printing of Bone Grafts ..................................................... 10 1.5. Bioinks for Bone Regeneration.............................................................................. 12 1.5.1. Bioink Formulations .......................................................................................... 12 1.5.2. Strategies to Tune the Mechanical Properties and Degradability of Alginate Bioinks ......................................................................................................................... 12 1.5.3. Strategies to Confer Bioactivity to Alginate Bioinks ........................................ 14 1.6. References ............................................................................................................... 16 CHAPTER 2: HYDROXYAPATITE NANOPARTICLES-CELL INTERACTION: THE FATE OF MEMBRANE-BOUND AND INTERNALISED NANOPARTICLES ......................................................................................................................................... 30 Scope ............................................................................................................................... 31 2.1. Introduction ............................................................................................................ 32 2.2. Experimental Section ............................................................................................. 33 – VI – 2.2.1. Synthesis and Characterisation of Non-doped and Magnesium-doped Hydroxyapatite Nanoparticles ..................................................................................... 33 2.2.2. Fluorescent Functionalisation of the Nanoparticles .......................................... 34 2.2.3. Cell Culture ....................................................................................................... 34 2.2.4. Nanoparticles – Cell Membrane Interaction ..................................................... 35 2.2.5. Flow Cytometry Assay ...................................................................................... 35 2.2.6. Intracellular Calcium Evaluation....................................................................... 36 2.2.7. Transmission Electron Microscopy ................................................................... 36 2.2.8. Cryo-Soft X-ray Tomography ........................................................................... 36 2.2.9. Statistical Analysis ............................................................................................ 37 2.3. Results and Discussion ........................................................................................... 37 2.3.1. Characterisation of the Nanoparticles ............................................................... 37 2.3.2. Functionalisation of the Nanoparticles .............................................................. 39 2.3.3. Study of the Nanoparticles – Cell Membrane Interaction ................................. 40 2.3.4. Flow Cytometry Assay ...................................................................................... 43 2.3.5. Intracellular Calcium Evaluation....................................................................... 46 2.3.6. Transmission Electron Microscopy and Cryo-Soft X-ray Tomography ........... 49 2.4. Conclusions ............................................................................................................. 54 2.5. References ............................................................................................................... 54 CHAPTER 3: MICROSPHERES WITH ION-DOPED HYDROXYAPATITE NANOPARTICLES FOR ENHANCED BONE REGENERATION ............................. 60 Scope ............................................................................................................................... 61 3.1. Introduction ............................................................................................................ 62 3.2. Experimental Section ............................................................................................. 63 3.2.1. Synthesis of Mg-, Znand Sr-Doped Hydroxyapatite Nanoparticles ............... 63 3.2.2. Synthesis of Nanoparticle-Charged Gelatine Microspheres ............................. 64 3.2.3. Physicochemical Characterisation of the Nanoparticles ................................... 64 3.2.4. Physicochemical Characterisation of the Microspheres .................................... 64 3.2.5. Study of the Ion Release in Cell Culture Media ................................................ 65 3.2.6. Statistical Analysis ............................................................................................ 65 3.3. Results and Discussion ........................................................................................... 65 3.3.1. Physicochemical Characterisation of the Doped HA Nanoparticles ................. 65 3.3.2. Physicochemical Characterisation of the Microspheres .................................... 67 3.3.3. Ion Release in Cell Culture Media .................................................................... 70 3.4. Conclusion ............................................................................................................... 76 3.5. References ............................................................................................................... 77 CHAPTER 4: MICROSPHERE INCORPORATION AS A STRATEGY TO TUNE THE BIOLOGICAL PERFORMANCE OF BIOINKS .................................................. 84 Scope ............................................................................................................................... 85 4.1. Introduction ............................................................................................................ 86 4.2. Experimental Section ............................................................................................. 87 4.2.1. Synthesis of Gelatine Microspheres .................................................................. 87 4.2.2. Synthesis of Hydroxyapatite-Containing Gelatine Microspheres ..................... 87 – VII – 4.2.3. Synthesis of Calcium-Deficient Hydroxyapatite Microspheres ........................ 88 4.2.4. Physicochemical Characterisation of the Microspheres .................................... 88 4.2.5. Preparation of the Bioinks ................................................................................. 89 4.2.6. Rheological Study ............................................................................................. 90 4.2.7. Direct Cell Culture on the Microspheres ........................................................... 90 4.2.8. Cell Viability Analysis ...................................................................................... 90 4.2.9. Cell Proliferation and Cell Morphology Evaluation ......................................... 91 4.2.10.Gene Expression Assessment ........................................................................... 91 4.2.11.Statistical Analysis ............................................................................................ 92 4.3. Results ...................................................................................................................... 92 4.3.1. Physicochemical Characterisation of the Microspheres .................................... 92 4.3.2. Rheology and Shape Fidelity of the Bioinks ..................................................... 95 4.3.3. Cell Cultures on the Microspheres .................................................................... 96 4.3.4. Cell Viability and Cell Migration in the Bioinks .............................................. 98 4.3.5. Cell Proliferation and Cell Morphology in the Bioinks .................................. 101 4.3.6. Gene Expression of the Cells in the Bioinks ................................................... 102 4.4. Discussion .............................................................................................................. 103 4.5. Conclusions ........................................................................................................... 107 4.6. References ............................................................................................................. 107 CHAPTER 5: BIOINK FUNCTIONALISATION WITH INTEGRIN-SELECTIVE PEPTIDOMIMETICS ................................................................................................... 113 Scope ............................................................................................................................. 114 5.1. Introduction .......................................................................................................... 115 5.2. Experimental Section ........................................................................................... 116 5.2.1. Alginate Functionalisation and Characterisation ............................................ 116 5.2.2. Cell Culture ..................................................................................................... 116 5.2.3. Bioink Synthesis .............................................................................................. 117 5.2.4. Cell Viability and Proliferation ....................................................................... 117 5.2.5. Osteogenic Differentiation .............................................................................. 117 5.2.6. Statistical Analysis .......................................................................................... 118 5.3. Results and Discussion ......................................................................................... 119 5.3.1. Alginate Bioink Characterisation .................................................................... 119 5.3.2. Cell Viability and Morphology ....................................................................... 120 5.3.3. Cell Proliferation ............................................................................................. 123 5.3.4. Osteogenic Differentiation .............................................................................. 123 5.4. Conclusions ........................................................................................................... 126 5.5. References ............................................................................................................. 126 GENERAL CONCLUSIONS ..................................................................................... 133 PUBLICATIONS AND CONFERENCES ................................................................ 136 Publications .................................................................................................................. 137 Conference Participation ............................................................................................ 138 – VIII – Abstract Substantial research efforts in the field of tissue engineering have been done in recent years to overcome the bone regeneration limitations of the human body and to treat bonerelated pathologies like osteosarcoma. Nanotechnology offers limitless ways to improve bone regeneration and to fight bone cancer by means of hydroxyapatite (HA) nanoparticles (NPs) and through the design of small organic molecules with the capacity to trigger specific biological reactions. In the present thesis, the implementation of these molecules in cancer therapy and in the bone regeneration field is studied, and their interaction with biological systems is investigated. Chapter 1 offers a general perspective of the topic and introduces the state of the art with the main strategies for bone cancer therapy and bone regeneration reported in the literature. Chapter 2 explores the use of hydroxyapatite nanoparticles for bone cancer therapies, with a view to elucidate their underlying cytotoxic mechanism. Moreover, the techniques commonly used to study the cellular uptake of these materials are compared with more novel characterisation alternatives. Indeed, the implementation of cryo-soft Xray tomography is presented as a key tool for tracking and monitoring the fate of internalised HA NPs, as it allows visualising both solid and liquid calcium deposits resulting from the dissolution of the nanomaterials. Chapter 3 investigates the use of gelatine microspheres (MS) as carriers of HA NPs for bone regeneration. Additionally, the ionic doping of these NPs with Mg2+, Zn2+ and Sr2+, with recognised benefits in the bone remodelling process, is studied and their release kinetics in cell culture medium is further quantified. Although dissolution and precipitation events take place simultaneously, MS demonstrate to be an effective method to deliver the therapeutical loaded ions. Chapter 4 and Chapter 5 focus on the design of bioinks as a strategy for bone tissue engineering. Alginate is chosen as the bioink binder due to its mild and straightforward cross-linking process, as well as for its high tuneability and adequate rheological properties. Two different strategies are proposed to solve the poor bioactivity of this polymer by conferring biological functionality to the cells embedded in the inks. Chapter 4 studies the incorporation of different types of MS (gelatine, gelatine containing HA NPs and calcium-deficient HA) to the formulation of these cell-laden inks. Furthermore, two cross-linking procedures are compared. Both, the addition of MS and the cross-linking using Ca2+-supplemented cell culture medium, resulted in increased stiffness of the bioinks. Importantly, the capacity of cells to migrate to the MS in the bioink formulation affected the cellular performance. In the case of inks loaded with gelatine-containing MS, promotion of cell migration, attachment and proliferation inside the bioink was observed. In addition, despite the superior osteogenic differentiation potential of the fully mineral MS when cells were cultured directly onto them, this behaviour was not evidenced when these particles were added in the bioink formulation. Finally, Chapter 5 explores the use of two peptidomimetics with selective activity towards α5β1 and αvβ3 integrins, which play a crucial role in the osteogenic pathway. Alginate polymer is functionalised with either molecule prior to the preparation of the cell-laden inks. The use of both non-peptidic ligands results in enhanced osteogenic differentiation of the mesenchymal stem cells embedded in the bioinks. Besides, the boost of osteodifferentiation is higher than in the traditional RGD-coupled alginate hydrogels. – IX – Resum En els últims anys s’ha fet un gran esforç de recerca en el camp de l’enginyeria de teixits per tal de superar les limitacions de regeneració òssia del cos humà, així com per tractar patologies relacionades amb l’os com ara l’osteosarcoma. La nanotecnologia ofereix múltiples maneres de millorar la regeneració òssia i de lluitar contra el càncer d’os, gràcies a l’ús de nanopartícules d’hidroxiapatita (NPs HA) i a través del disseny de molècules orgàniques molt petites amb capacitat per desencadenar reaccions biològiques específiques. En aquesta tesi, s’estudia la implementació d’aquestes molècules en la teràpia anticàncer i en el camp de la regeneració òssia, així com la seva interacció amb els sistemes biològics. El capítol 1 ofereix una perspectiva general del tema i introdueix l’estat de l’art amb les estratègies per al tractament del càncer d’os i la regeneració òssia més comunes reportades en la literatura. En el capítol 2 s’explora l’ús de les nanopartícules d’hidroxiapatita com a teràpia per al càncer d’os amb la intenció d’investigar-ne el mecanisme de citotoxicitat subjacent. A més, les tècniques més utilitzades per a l’estudi de la internalització cel·lular d’aquests materials es comparen amb alternatives de caracterització més innovadores. D’aquesta manera, la implementació de la crio-tomografia de raigs X tous es presenta com una eina clau per al seguiment i la monitorització del destí d’aquestes NPs internalitzades, ja que permet la visualització de dipòsits de calci en estat sòlid i líquid, que deriven de la dissolució del nanomaterial. En el capítol 3 s’investiga l’ús de microesferes (MS) de gelatina com a portadores de NPs HA per a la regeneració òssia. S’estudia el dopatge iònic de les NPs amb Mg2+, Zn2+ i Sr2+, que són ions molt implicats en el procés de remodelació òssia, i es quantifica la seva cinètica d’alliberació en medi de cultiu cel·lular. Tot i que hi ha processos de dissolució i precipitació simultanis, es demostra que les MS són un mètode efectiu per alliberar els ions terapèutics carregats. El capítol 4 i el capítol 5 se centren en el disseny de biotintes en el camp de l’enginyeria de teixits ossis. S’escull l’alginat com a matriu, donat el seu procés d’entrecreuament simple i suau per a les cèl·lules, així com per la seva alta capacitat de modificació i les seves propietats reològiques. Es proposen dues estratègies diferents per solucionar la falta de bioactivitat d’aquest polímer, i donar així funcionalitats biològiques a les cèl·lules incorporades a les biotintes. En el capítol 4 s’estudia la incorporació de diferents tipus de MS (gelatina, gelatina amb NPs HA, i HA deficient en calci) per a la formulació d’aquestes biotintes. També es comparen dos protocols d’entrecreuament. Tant la incorporació de MS com l’entrecreuament usant medi de cultiu suplementat amb Ca2+ incrementen la rigidesa de les biotintes. També s’aprecia com la capacitat de migració de les cèl·lules cap a les MS en la formulació de les biotintes afecta el comportament cel·lular. En el cas de les tintes carregades amb MS que contenen gelatina, s’observa una promoció de la migració, adhesió i proliferació cel·lular dins les biotintes. D’altra banda, tot i la major capacitat de diferenciació osteogènica de les MS minerals quan les cèl·lules s’hi sembren a sobre, no s’aprecien aquests altres fenòmens. Finalment, en el capítol 5 s’explora l’ús de dos peptidomimètics amb activitat selectiva per a les integrines α5β1 i αvβ3, que s’han vist molt implicades en la ruta osteogènica. El polímer d’alginat es funcionalitza amb aquestes molècules abans de preparar-ne biotintes. L’aplicació d’aquests dos lligands no peptídics té com a resultat una millor diferenciació osteogènica de les cèl·lules mare mesenquimals incorporades en les biotintes. A part, aquest estímul de l’osteodiferenciació és més important que en els hidrogels d’alginat funcionalitzats amb RGD, que és la solució tradicional. Chapter 1 Introduction Introduction – 2 – Bone is reported to be the second most transplanted tissue after blood [1], and bone injuries account for more than 500000 grafting procedures per year only in the United States [2]. Moreover, age-associated musculoskeletal diseases are the 4th main cause of disability worldwide, and their incidence is expected to increase due to the global ageing of the population. Indeed, it is predicted that the proportion of the world's population over 60 years will duplicate by the end of 2050 [3]. In addition, bone cancer is another major problem of this tissue, being osteosarcoma (OS) the most common type of primary solid tumour originated in bone. This disease mainly affects children and young adults, as well as adults with previous bone pathologies [4]. The following sections provide an overview of the main characteristics of bone tissue, describing its function, composition and structure, as well as its remodelling process. Afterwards, a brief description of the two major bone-related alterations is exposed. The main bone cancer treatments will be covered, as well as the use of hydroxyapatite nanoparticles in bone cancer therapy. Finally, an insight into bone grafts and bioinks will be given, together with different approaches to both tune the mechanical properties and the cell behaviour of alginate cell-laden inks. 1.1. Bone Tissue 1.1.1. Function Bone is a complex organ that exerts important functions in the body. On the one hand, bone provides shape and support to the body through the skeletal system which, in turn, protects and supports other internal organs and soft tissues. In this sense, bone has a clear locomotive and biomechanical function as it acts as attachment points for the muscles [5]. On the other hand, at the metabolic level, bone tissue serves as a storage site for certain ions, especially calcium and phosphorus, regulating their levels throughout the body. Finally, bones harbour marrow which is divided into yellow and red marrow. Whilst yellow bone marrow serves as fat storage, the red marrow produces blood cells (i.e. red blood cells, white blood cells and platelets) through haematopoiesis [6,7]. 1.1.2. Bone Structure The human skeleton is composed of more than 200 different bones, which are usually classified by their shape. The four big families are: (1) long bones, presenting a central shaft called diaphysis and two edges (epiphyses) covered by cartilaginous tissue, such as tibia, femur, ulna and radius; (2) short or cuboidal bones, like carpals and tarsals; (3) flat bones, that usually protect organs or serve for muscle attachment, for example, parietal, scapulae, ribs or sternum; (4) irregular bones with complex shapes, such as vertebrae or maxilla [8]. There are two types of bone tissue, i.e. cortical or compact tissue and trabecular tissue, also known as spongy or cancellous bone. Although all the bones present a similar structure, the distribution and concentration of each type of tissue will depend on the bone’s overall function. In an adult human, cortical bone accounts for approximately 80 wt% of the skeletal system. This tissue exhibits a high calcification and is responsible for providing mechanical strength to the system. It possesses high rigidity and resistance to bending and torsion. Differently, trabecular bone comprises the remaining 20 wt% of the Chapter 1 – 3 – skeleton. This tissue presents an important porosity, that permits the accommodation of the bone marrow [9,10]. 1.1.3. Bone Composition Bone can be described as a composite material, formed by the combination of an extracellular matrix (ECM), different kinds of cells and water. The ECM is composed of a major inorganic phase (representing the 60 wt%) together with an organic phase (that accounts for the remaining 40 wt%) [11]. While the inorganic phase provides stiffness to the structure, the organic phase supplies toughness [12]. The main inorganic components of the bone ECM are calcium and phosphorus ions that precipitate to form non-stoichiometric hydroxyapatite (HA) crystals, with a length of 20 – 50 nm, a width of 15 nm and a thickness of 2 – 5 nm [13]. Moreover, hydroxyapatite has an open crystalline structure that allows several ionic substitutions in its crystal lattice, where different monovalent or divalent cations such as Na+, K+, Mg2+, Zn2+ or Sr2+ can substitute Ca2+, while PO43and OHsites can be substituted by other anions such as CO32- (Figure 1.1). Indeed, the chemical formula for biological HA can be described as Ca10xMx(PO4)6-y(HPO4, CO3)y(OH)2-zNz, where M and N represent cationic and anionic ions, respectively, and x, y and z stand for the molar number of the incorporated ions [14]. Table 1.1 summarises the composition of bone tissue in the human body. As can be observed, bone presents small amounts of certain foreign ions that cause the poorly crystalline structure of the biological apatites, presenting nanosized hexagonal plate-like crystals [15]. Table 1.1. Composition of the mineral phase of bone, adapted from [16]. Element (Symbol) Quantity Calcium (Ca) 36.60 wt% Phosphorus (P) 17.10 wt% Carbonate (CO3) 4.80 wt% Sodium (Na) 1.00 wt% Magnesium (Mg) 0.60 wt% Chlorine (Cl) 0.10 wt% Fluorine (F) 0.10 wt% Potassium (K) 0.07 wt% Strontium (Sr) 0.05 wt% Silicon (Si) 500 ppm Zinc (Zn) 39 ppm Aside from the crystalline HA structure, another ionic environment can be found in form of a hydrated layer on the surface of the apatite (Figure 1.1). This layer permits the rapid and reversible exchange of ions with the surrounding fluids. It has been observed that this Introduction – 4 – hydrated layer decreases with maturity and, thus, the amount of freely exchangeable ions to assure homeostasis drops [17–19]. Figure 1.1. Scheme of the ionic substitutions in hydroxyapatite crystals, taking place either in the crystalline lattice or in the hydrated layer of the surface. Although presented in small levels, all these ions play a major role in the bone properties and are crucial for the proper development of the tissue. For instance, magnesium is crucial in bone metabolism and is involved in the stimulation of osteoblast proliferation in the initial stages of osteogenesis [20,21]. On the other hand, zinc promotes bone growth and further mineralisation by stimulating osteoblasts and collagen secretion and inhibiting osteoclastic differentiation [22–24]. Finally, strontium inhibits bone resorption and boosts osteoblastic differentiation [25,26]. Since the presence of these ions in biological apatite play such important roles in bone formation, ionic substitutions into synthetic HA have been studied to improve the biological performance of these materials in bone tissue engineering [27]. On the other hand, the organic ECM is 90 wt% formed by type I collagen. This protein plays an important role in determining the architecture of bone by the creation of fibrils that interact with other proteins and generate lamellar arrangements. Non-collagenous proteins account for the remaining 10 wt% of the ECM organic phase. Among them, several proteoglycans, osteocalcin and osteonectin can be found, together with other small integrin-binding ligand N-linked glycoproteins (SIBLINGs) including bone sialoprotein (BSP) or osteopontin (OPN) [28]. In this PhD thesis, hydroxyapatite nanoparticles will be doped with bone-related therapeutic ions as a way to enhance the bone regeneration process. Chapter 1 – 5 – 1.1.4. Bone Cells Bone tissue contains five main cell types: osteoprogenitor cells, osteoblasts, bone lining cells, osteocytes and osteoclasts. Each of them has individual functions, the combination of which makes the bone remodelling process possible. Osteoprogenitor cells come from mesenchymal stem cells (MSCs) and are the precursors to more specialised bone cells such as osteoblasts and osteocytes. They are located in the red bone marrow from where they are released during the remodelling process [29]. With age, the body loses the ability to synthesise this kind of cell, contributing to the ageing of the bones. Osteoblast, derived from osteoprogenitor cells, comprise around 5% of the total resident bone cells and are located in the surface of the bones. These cells are responsible for the ECM deposition and its subsequent mineralisation [30]. Bone lining cells are quiescent osteoblasts covering the bone surfaces. These cells present a thin and flat shape and their functions are not yet completely understood [31]. Osteocytes are the most abundant bone cells and comprise more than 90% of the total bone cells [32]. Osteocytes derive from MSCs through osteoblast differentiation, present a dendritic morphology, and are located within lacunae surrounded by mineralised ECM. Due to their location, osteocytes have a mechanosensitive function as they are able to convert mechanical stimuli to biochemical signals through mechanisms that are still unknown [33]. Osteoclasts are derived from the hematopoietic lineage. These multinucleated cells are responsible for the dissolution and absorption of the bone, either during the remodelling process or motivated by the need for ions in the body [6]. 1.1.5. Bone Remodelling Process During lifetime, healthy bone is under constant renewal. The old bone is progressively reabsorbed giving space for new bone formation. The replacement of old bone by new bone is called bone remodelling or bone turnover. This process permits the continuous adaptation of the bone [34]. Precisely, during childhood, the bone remodelling cycle allows the bone growth and, during adulthood, permits modifications in shape depending on the mechanical demands of the body [8]. Moreover, bone remodelling is of major importance during healing after a bone injury. The remodelling process is complex and requires close interaction of the bone cells with each other and with the ECM. The turnover takes place in basic multicellular units (BMUs) located in the bone cavities [8]. These fundamental units are made of small groups of cells organised into a cutting cone with osteoclastic resorption at the apex and osteoblasts creating new osteoid at the base filling the cavity, together with endothelial cells from the blood vessels [35]. In the remodelling process, five phases can be distinguished: resting or quiescence, resorption, reversal, bone formation and mineralisation, as represented in Figure 1.2. The first step is the separation of lining cells to expose the bone surface, followed by the recruitment and guidance of hematopoietic stem cells by osteoclastogenic factors to the site of interest. During the resorption stage, these cells differentiate into mature osteoclasts that attach to the bone surface by polarising their membrane. Osteoclasts secrete acidic enzymes to locally dissolve both organic and inorganic bone phases, which results in small fragments that are digested within cytoplasmic vacuoles of the osteoclasts Introduction – 6 – [36]. The resulting products are calcium and phosphorus ions that are released into the bloodstream. Resorption lacunae are created as a consequence of this process (reversal stage). Afterwards, the bone formation stage takes place. The MSC-derived osteoblasts deposit the organic matrix (also named osteoid) in the lacunae by secreting collagen type I, as well as alkaline phosphatase (ALP), different growth factors and osteocalcin, among others. Finally, during the mineralisation step, osteoblasts calcify the osteoid. Subsequently, they undergo apoptosis or become trapped into the ECM and transform into osteocytes [37]. Figure 1.2. Schematic representation of bone remodelling process, from [38]. 1.2. Bone Affections The two main bone affections encountered in the clinics are bone cancer and bone defects and, therefore, they represent a field of interest in medical research. These affections have multiple origins and strategies to be treated. Down below these two aspects are developed. 1.2.1. Bone Cancer Bone cancer is the presence of a malignant tumour in the bone tissue. Usually, cancer in the bone is caused by metastasis of other primary tumours, such as breast, prostate or lung cancer [39,40]. Among the primary bone cancers, osteosarcoma (also called osteogenic sarcoma) is the most common one. Although this type of cancer can develop in any bone, it is frequently found in quickly growing bones, like the distal femur (30%), proximal tibia (15%) or proximal humerus (15%) [41]. Therefore, it mostly affects children and young adults, whose bone cells experience rapid growth and a higher risk of mutations [4]. Moreover, OS is developed as an osteoid-producing solid tumour. Despite this type of cancer occurs with a low incidence compared to other solid tumours, it is ranked among the most usual cause of cancer-related child death [42,43]. Interestingly, it has been demonstrated that some cancer cells have ionic channels dysregulated in their membrane, compared to normal cells of the same tissue. For example, the TRPM7 magnesium-transporter is overexpressed on MG-63 osteosarcoma Chapter 1 – 7 – cells [44,45], ZIP4 expression is increased in the membrane of pancreatic cancer cells and ZIP1 is downregulated in malignant prostate cells, both of them being zinc transporters [46]. 1.2.2. Bone Defects Bone defects can be originated by different causes, where the most common is trauma, followed by osteoporosis. Osteoporosis is a skeletal disorder characterised by a progressive reduction of the bone mineral density over time, leading to micro-fractures in the bone architecture and, ultimately, bone fragility. Bone mass loss may increase during menopause due to the reduction in estrogenic levels [47]. Moreover, surgical procedures can cause morphological alterations in bone such in the case of osteotomies to remove bone tumours [48], craniotomies performed to access the brain or harvesting of the patients’ bone to use it as an autologous graft. Additionally, the deficiency in mechanical stimuli can lead to bone resorption, which is frequent in dental medicine [49]. Finally, bone defects can also be the consequence of certain congenital diseases like cleft lip and cleft palate [50]. Despite the remarkable self-regenerative capacity of bone, sometimes the help of a bone graft to regenerate is required. This is the case of critical-sized defects, where the bone is not able to completely heal [51]. 1.3. Bone Cancer Therapy Nowadays, surgical resection of the tumour combined with chemotherapy is the first-line treatment for bone cancer, and specifically, for OS [52]. Although this strategy has increased the cure rate of the patients, the survival rate is still insufficient, around 50% [53]. In addition, conventional chemotherapy lacks sensitivity and specificity, presents drug toxicity and severe side effects, as it affects both malignant and healthy cells [54]. Furthermore, the development of drug resistance is a major problem [55]. Although the advances done in cancer therapy are considerable, these are still far from being an optimum treatment. Therefore, new disruptive solutions based on completely different approaches need to be investigated. 1.3.1. Hydroxyapatite Nanoparticles in Cancer Treatment Besides the well-established use of hydroxyapatite as bone graft, this biomaterial has also been explored in form of nanoparticles (NPs) in various new-emerging biomedical fields such as gene and drug delivery applications, bio-imaging and cancer therapy [56–59]. Regarding the latter approach, NPs are a promising solution due to their selective retention into the tumour mass caused by the enhanced retention and permeability (EPR) effect that characterise malignant cells [60,61]. In addition, the increased metabolic activity of cancer cells leads to a higher uptake of NPs on tumour cells than on healthy ones, requiring lower doses than other chemotherapeutic solutions [62]. Among the different methods to synthesise HA NPs, the wet methods are the most popular and, in particular, chemical precipitation by titration of calcium hydroxide slurry with phosphoric acid (Equation 1.1) at the body or room temperature, which usually results in poorly crystalline needle-like NPs [63]. Introduction – 8 – 10 Ca(OH)2 + 6 H3PO4 → Ca10(PO4)6(OH)2 + 18 H2O Equation 1.1 This process presents several advantages compared to other options, such as its simplicity, water being the only by-product and no need for special equipment or additional chemicals to control the reaction. Another additional asset is the facility in tuning several morphological and chemical properties of the final materials, such as the shape, the size and the specific surface area [64]. Furthermore, wet precipitation allows the ion doping of the HA structure by adding the salt of interest during the precipitation, which has been found to improve their bioactivity, as stated previously [65]. Currently, the use of HA NPs in cancer treatment is based on two different strategies: as chemotherapeutic drug carriers and as a drug by themselves. The application of HA NPs as transfection vehicles solves the degradation and poor solubility problems of some chemotherapeutic drugs, as well as their low internalisation rates into cancer cells. Indeed, this approach has been found to be helpful in the transport of cisplatin and doxorubicin [66–68]. On the other hand, it has been observed that HA NPs themselves were cytotoxic for several cancer cell types (i.e. colorectal cancer, hepatoma, osteosarcoma and glioma) whilst causing minimal side effects on healthy cells [69–72]. Moreover, ion-doped HA NPs have been investigated in cancer therapy. On the one hand, iron-doped HA NPs with magnetic properties have been used in hyperthermia therapy, both in vitro causing malignant liver cell death through elevated ROS levels [73], and in vivo showing an important reduction of the colorectal tumour volume after 14 days [74]. On the other hand, magnesium-doped HA NPs have been reported to present a remarkable killing potential in MG-63 osteosarcoma cells, being more selective than non-doped HA [75]. 1.3.2. Cytotoxicity Mechanisms of Hydroxyapatite Nanoparticles HA NPs are internalised into the cells by endocytosis and, subsequently, degraded under the acidic conditions of these vesicles. However, the interaction of HA NPs with cells and their cell death mechanisms are still under debate. First, while the reactivity of the bulk HA with cell culture media and its consequences in cells have been extensively studied, few works have focused on the effects of the same material in the form of NPs and their influence on cell behaviour [76,77]. Contrastingly, the internalisation of HA NPs is commonly believed to be the cause of cell toxicity. The most accepted hypothesis to explain the cytotoxic effects of HA NPs is that their degradation leads to an increase of intracellular calcium and phosphorus ions, inducing apoptosis [78], as schematised in Figure 1.3. These effects are usually demonstrated by techniques such as transmission electron microscopy, flow cytometry or fluorescent microscopy, all of them based on the visualisation of the uptaken NPs. However, as the degradation of HA NPs is needed to achieve a cytotoxic behaviour, these techniques are insufficient to predict and study their toxicity. Recent alternatives to investigate the HA NPs solubilisation are the use of intracellular calcium straining and the application of cryo-soft X-ray tomography [79– 82]. Chapter 1 – 9 – Figure 1.3. Cytotoxic mechanism of hydroxyapatite nanoparticles (HA NPs) in cancer therapy. (1) Internalisation of the HA NPs through endocytosis. (2) Dissolution of the HA NPs due to the acidic environment in the endocytic vesicle. (3) Massive release of calcium and phosphorus ions to the cytoplasm of the cell. This PhD thesis aims to improve our understanding of the cytotoxicity mechanisms of hydroxyapatite nanoparticles for cancer treatment. To do so, conventional techniques to study their cell uptake will be compared to more advanced strategies. 1.4. Bone Grafting Strategies In spite of the self-regeneration capacity of the bones, when a bone defect exceeds a critical size, the tissue is not able to heal spontaneously. In these cases, external help is needed to achieve the desired regeneration and bone grafts are crucial in this process [51]. Bone grafting is a surgical procedure to reconstruct bone in a defect by means of a bone substitute. The use of bone grafts accelerates the formation of new bone tissue by permitting the regeneration in volume instead of the layer-by-layer process that occurs in the natural bone remodelling. Bone grafts can be classified into four different groups depending on their origin: autografts, allografts, xenografts and synthetic bone grafts. 1.4.1. Bone Grafts of Natural Origin Within this family of bone grafts, autografts stand out as the clinical gold standard. These grafts consist of bone tissue harvested from the patient’s own body and retrieved from a second anatomical site [83]. Although their biocompatibility and bioactivity are unbeatable, they require additional surgery to extract the needed bone tissue which implies added morbidity, longer surgical times and aesthetic concerns. Moreover, there is a limitation on their availability [84,85]. Introduction – 16 – a low cost and are stable under pH and temperature changes. However, the use of such short molecules entails several constraints related to their stability against enzymatic degradation in vivo and their lack of specificity for particular integrins. Moreover, their cellular stimulation is limited compared to full-length proteins [168]. To overcome these drawbacks, synthetic non-peptidic ligands (also known as peptidomimetics) have raised interest. These artificial molecules are designed to present a high affinity for a specific integrin subtype mimicking the pharmacophore of their natural ligands and improving their biological performance. In general, peptidomimetics present high stability in serum and almost no immune reaction [169,170]. v3 and 51 integrin subtypes are found to be essential in the osteogenesis process [171–173] and, thus, artificial ligands selective for these two types of integrins have been developed for bone tissue applications. Both peptidomimetics exhibited selective cell adhesion and high specificity when applied as surface coatings [174–177]. Furthermore, other studies reported enhanced osteogenic differentiation and cell proliferation [178,179]. However, integrin-selective peptidomimetics have only been applied in 2D systems and no data on bioinks or other 3D strategies can be found in the literature. In the present PhD thesis, the use of 51and v3-selective peptidomimetics will be studied as a novel strategy to enhance the bioactivity and osteogenic differentiation capabilities of alginate bioinks. 1.6. References [1] H. Shegarfi, O. 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Alloys Compd. 615 (2014) S670–S675. doi:10.1016/j.jallcom.2014.04.078. [140] B. Sarker, W. Li, K. Zheng, R. Detsch, A.R. Boccaccini, Designing Porous Bone Hydroxyapatite Nanoparticles-Cell Interaction – 32 – 2.1. Introduction Over the past few years, the application of hydroxyapatite (HA) in the form of nanoparticles (NPs) has received considerable attention as a transfection vehicle for the transport of genetic material into cells [1]. HA NPs have also been proposed as a useful tool in cancer therapy. It has been observed that they were able to cause cell death of several cancer cell types whilst having minimal side effects on healthy cells [2–6]. Moreover, ion-doped HA NPs have been investigated for cancer treatment showing improved cytotoxic potential towards cancer cells and being more selective than nondoped HA [7]. Actually, ions are particularly interesting in cancer treatment, as it has been demonstrated that some specific cancer cells have ionic channel dysregulations in their membrane, e.g. TRPM7 magnesium transporter is upregulated in breast and pancreatic cancer, as well as in osteosarcoma MG-63 cell line [8–12]. However, the exact mechanisms of cell death are still under debate. Three different scenarios which can potentially drive towards cytotoxicity can be anticipated: (1) reactivity of the NPs with the surrounding fluids, causing the release or depletion of ions and proteins crucial for cells; (2) interaction of NPs with cell membrane receptors that may induce apoptotic signalling cascades; and (3) NP internalisation and subsequent effects inside the cell. The most accepted hypothesis is based on the third assumption and relies on the degradability of the NPs upon internalisation. Indeed, after HA NPs uptake by endocytosis, they degrade under the acidic conditions in the lysosome, yielding an increase of calcium and phosphorus ions. Whereas a slow and sustained dissolution of HA in the lysosomes would be innocuous to cells and would benefit transfection [13], a more efficient and fast internalisation would release a high concentration of calcium ions that would lead to cell death [14]. This has been hypothesised to be caused either by apoptosis triggered by a disruption of cell homeostasis [15] or by necrosis when the rapid dissolution of the NPs causes an imbalance of the osmotic pressure in the lysosomes that results in their massive breaking [16,17]. In order to describe in detail the HA NPs mode of action, it is necessary to have reliable methods to assess not only their internalisation but also their dissolution once inside the cells. Internalised NPs can be detected either by transmission electron microscopy or by tracking fluorescently labelled NPs with different techniques (e.g. flow cytometry, fluorescent microscopy, etc.). However, with some of these techniques, it is not possible to differentiate the cell surface-associated particles from the ones internalised into the cells [18,19]. Moreover, techniques based on the visualisation of pristine NPs are insufficient to correlate the amount of internalised particles to their toxicity, as they cannot determine their solubilisation [20]. Thereby, various studies have used calcium probes to measure intracellular calcium levels, to check HA NPs dissolution [14,16,17,20]. In addition, in this study, we propose the use of cryo-soft X-ray tomography as a novel approach to analyse the ultrastructure of intact and unstained cells, as well as to obtain quantitative data on the Ca elemental composition of each organelle. Compared to calcium imaging, this technique offers the advantage of visualising soluble calcium-rich vesicles without losing the information of solid calcium material. Beyond the effects associated with NP internalisation, little attention has been paid in the literature to the influence of ionic exchanges between HA NPs and the surrounding fluids on cell viability [7,21,22]. This may be relevant, as drastic ionic fluctuations, especially of calcium, phosphorus and magnesium, elicited by nanostructured hydroxyapatite Chapter 2 – 33 – materials, have been reported to result in cytotoxicity [23–25]. On the other hand, to the best of our knowledge, there is no literature studying the interaction of HA NPs with receptors in the membrane that may be crucial for cell viability. In this context, the objectives of the present work are: (1) to unravel the influence of nanoparticle-cell membrane interaction on cell cytotoxicity, leading to a better understanding of the mechanisms behind the use of HA NPs in cancer treatment applications; and (2) to investigate fundamental aspects of HA NPs internalisation and fate within the cells. The conventional techniques will be compared to calcium fluorescence probes and cryo-soft X-ray tomography as alternatives to track calcium-rich vesicles and the degradation state of internalised HA NPs. 2.2. Experimental Section 2.2.1. Synthesis and Characterisation of Non-doped and Magnesium-doped Hydroxyapatite Nanoparticles Non-doped and magnesium-doped hydroxyapatite nanoparticles (HA NPs and MgHA NPs, respectively) were obtained by the neutralisation of calcium hydroxide with orthophosphoric acid, as described elsewhere [7]. Briefly, 200 mM H3PO4 (85 wt% pure, Panreac) was added dropwise into a solution of 333 mM Ca(OH)2 (96 wt% pure, Fluka) at a rate of 1 ml min-1. The reaction, described in Equation 2.1, was performed under constant stirring at 40 oC in thermojacketed vessels and the pH was continuously monitored. 10 Ca(OH)2 + 6 H3PO4 → Ca10(PO4)6(OH)2 + 18 H2O Equation 2.1 At pH 8, the reaction was stopped and the solution was stirred for 30 min at 40 oC. Afterwards, it was left to mature overnight at room temperature. The suspension was then rinsed three times with bi-distilled water, performing centrifugation cycles of 5 min at 800 g (5430 R, Eppendorf). Finally, the product was freeze-dried (Cryodos, Telstar). The synthesis of MgHA NPs was done similarly, with the incorporation of magnesium chloride (MgCl2 · 6 H2O, 99 wt% pure, PanReac) into the calcium hydroxide solution prior to H3PO4 addition. The final Mg2+ content in these NPs was 2.3 wt% [7]. The NPs were suspended in bi-distilled water to a final concentration of 10 mg ml-1 and sonicated with a high-frequency ultrasound probe sonicator (450D, Branson Digital) to improve their deagglomeration. A 3 mm diameter tip was used with 40% amplitude for 2 min in an ice bath. In addition, specular discs were prepared by compacting 200 mg of NPs into 10 mm diameter moulds and applying a uniaxial pressure of 3 tons for 2 min. Characterisation of the NPs comprised analysis of the phase composition by X-ray diffraction (XRD, D8 Advance, Bruker) using Cu Kα radiation at 40 kV and 40 mA. Data were collected with a step size of 0.02o over the 2θ range from 20 to 40o with a counting time of 2 s per step. Phase identification was accomplished by comparing the results with the standard patterns of HA (ICDD PDF 09-0432). In addition, morphological evaluation of the NPs was assessed by transmission electron microscopy (TEM, JEM-1010, JEOL) after soaking a 300-mesh carbon-coated copper grid into the NPs suspensions. Finally, the surface morphology of the discs was studied by scanning electron microscopy (SEM, JSM-7001F, JEOL). Prior to the observation, the samples were coated with a thin goldpalladium layer using vapour deposition (SCD 004, Balzers). Hydroxyapatite Nanoparticles-Cell Interaction – 34 – 2.2.2. Fluorescent Functionalisation of the Nanoparticles In order to track the NPs, a fluorescent biomolecule was attached to HA by means of strong hydrogen bonds via catechol groups [26]. To this end, a custom-made linear peptide containing DOPA and carboxyfluorescein (CF) (Table 2.1) was synthesised by solid-phase peptide synthesis (SPPS) following the Fmoc/tBu strategy and using Rink Amide MBHA resin (243 mg, 0.45 mmol g-1) as a solid support, according to the protocols optimised in our group [27]. The peptide was used with a purity of ≥ 90%, as determined by reversed-phase analytical high-performance liquid chromatography (RPHPLC, Prominence UFLC XR, Shimadzu). Table 2.1. Chemical sequence of the fluorescent peptide and properties Sequencea Purity (%)b tR (min)b Molecular weight (g mol-1) CF-(Ahx)2-βAla-DOPA2-NH2 90.3 6.127 1031.07 a CF: (5)6-carboxyfluorescein; Ahx: 6-aminohexanoic acid; βAla: beta-alanine; DOPA: L-3,4dihydroxyphenylalanine b Characterised by HPLC using a reversed-phase XBridge (Waters) C18 column (4.6 mm x 100 mm, 3.5 µm) and a linear gradient from 20:100 (0.036% TFA in ACN/0.045% TFA in H2O) in 8 min at 25 ºC. The lyophilised peptide was dissolved in distilled water to a series of concentrations from 500 µM down to 1 µM, adjusting the pH to 7. Carboxyfluorescein (CF) alone was also studied at the same concentrations and pH. The functionalisation of the HA NPs was accomplished by mixing under constant stirring 200 µl of 1 wt% NPs suspension and 100 µl of either peptide or CF solutions for 2 h. The resulting suspension was centrifuged and rinsed three times with bi-distilled water and re-suspended to a final concentration of 1 wt%. After functionalisation, supernatants and washing residues were collected and their fluorescent intensity was measured by means of a microplate reader (Synergy HTX, BioTek Instruments). Knowing the difference between the initial amount of peptide added and the quantity that remained in the supernatants, data were correlated to a calibration curve to determine the final concentration of peptide adsorbed on the NPs. The fluorescent NPs (i.e. HA-F and MgHA-F) functionalised using a concentration of 500 µM of the fluorescent peptide were selected for the cell culture studies. In addition, direct fluorescence intensity was measured in 100 µg ml-1 NP suspension by means of a fluorescence spectrophotometer (Cary Eclipse, Agilent). Finally, the zeta potential of the different NPs (i.e. HA and MgHA, before and after functionalisation) was measured in 300 µg ml-1 NPs suspensions using a Zetasizer Nano (Malvern Panalytical). 2.2.3. Cell Culture Human osteosarcoma MG-63 cells (ATCC) were cultured in Dulbecco’s Modified Eagle medium (DMEM) supplemented with 10% foetal bovine serum (FBS), 20 mM 4-(2hydroxyethyl)-1-piperazineethanesulfonic acid buffer (HEPES), 2 mM L-glutamine, 50 Chapter 2 – 35 – U ml-1 penicillin and 50 µg ml-1 streptomycin (all from Gibco), in a 95% humidified atmosphere containing 5% CO2 at 37 oC. Confluent cells were detached using TrypLE Express (Gibco). For the cell culture studies, the NPs suspensions were sonicated in a bath sonicator (JP Selecta) for 5 min before their addition to the cell culture media, in order to avoid agglomerates. The experiments were carried out in serum-containing (i.e. NPs supplemented or cells seeded on discs in 10% FBS-containing media, FBS+) and in serum-free (FBS-) conditions to assess the influence of the protein corona. Moreover, different time points were studied, ranging from 3 to 24 h, depending on the study, and the supplemented dose of NPs was 100 µg ml-1 in all cases. Controls were prepared seeding the cells on sterile coverslips (Ø=10 mm). 2.2.4. Nanoparticles – Cell Membrane Interaction MG-63 cells were seeded on sterile coverslips (Ø=10 mm) in a 48-well plate at a density of 25,000 cells per well and incubated overnight to allow cell adhesion. The following day, culture media containing 100 µg ml-1 of each NPs were prepared and supplemented to the cells. At the same time, previously compacted discs were sterilised with ethanol 70% for 30 min and rinsed with phosphate-buffered saline (PBS, Gibco). Afterwards, cells were seeded on top of them and left interaction for 4 h. The morphology of the cells was assessed by confocal microscopy (LSM 800, Zeiss) using acridine orange staining (AO, Sigma-Aldrich). Cell morphology was further analysed by scanning electron microscopy (SEM, JSM-7001F, JEOL). To do so, samples were fixed with 2.5% glutaraldehyde (Sigma-Aldrich) in PBS for 1 h at 4 oC. Subsequently, fixed samples were rinsed with PBS and dehydrated in an increasing series of ethanol solutions. Dried discs and coverslips were covered with a thin gold-palladium layer using vapour deposition (SCD 004, Balzers). To quantify the results, a cytotoxicity assay was carried out using WST-1 reagent (Roche) and following the manufacturer’s indications. 2.2.5. Flow Cytometry Assay For the flow cytometry (FC) studies, MG-63 cells were seeded on 6-well plates at a density of 300,000 cells per well and incubated overnight. The following day, fresh culture media containing 100 µg ml-1 of functionalised NPs (i.e. HA-F and MgHA-F) were supplemented into the wells. After 4, 6 and 24 h of exposure, cells were detached from the wells using trypsin and the supernatant was removed after centrifugation. The pellet was rinsed and re-dispersed in PBS prior to analysis by flow cytometry (Accuri C6, BD Biosciences) at an emission wavelength of 488 nm. Controls were used to gate living cells and to determine the fluorescence threshold. A total of 10,000 events were evaluated for each condition, except for serum-free samples at 24 h, due to cell death. Data analysis was performed with FlowJo software (FlowJo LLC, version 10). In addition, cells supplemented with functionalised NPs were imaged in a confocal fluorescence microscope (LSM 800, Zeiss) to assess aggregation of NPs during cell culture. Before visualisation, cells were fixed and stained with Alexa Fluor 568-phalloidin (Thermo Fisher) and 4′,6-diamidino-2-phenylindole (DAPI, Sigma-Aldrich) after 4 h of interaction with the materials. Hydroxyapatite Nanoparticles-Cell Interaction – 36 – 2.2.6. Intracellular Calcium Evaluation For the quantification of intracellular calcium, MG-63 cells were seeded on a black 96well plate at a density of 10000 cells per well and incubated overnight, whereas for the imaging 25000 cells were seeded on sterile coverslips in a 48-well plate. The following day, culture media containing either HA or MgHA NPs were supplemented to the cells. An apoptosis control was also tested by adding H2O2 in the medium at a final concentration of 200 µM, instead of NPs. After 3 h of incubation, cells were stained with Fluo-4 AM intracellular calcium indicator (Thermo Fisher) following the manufacturer’s indications. The green intensity was quantified in a fluorescence microplate reader (Synergy HTX, BioTek Instruments) and the imaging was performed in a confocal microscope (LSM 800, Zeiss). 2.2.7. Transmission Electron Microscopy For TEM imaging, 90-mm Petri dishes were coated with poly-L-lysine prior to MG-63 seeding. After overnight cell attachment, FBS-free culture media containing each type of NPs were supplemented to the cells. A control sample without NPs was also studied. The cells were exposed for 3 h and afterwards fixed with 2.5% glutaraldehyde (SigmaAldrich) in 0.1 M phosphate buffer (PB) for 1 h and detached to form a pellet. The pellet was rinsed and put in contact with a solution of 1 wt% OsO4 and 0.8 wt% potassium ferricyanide for 2 h. Subsequently, it was dehydrated in an ascending series of acetone and infiltrated with EPON resin. Finally, blocks were sectioned with an ultramicrotome (Ultracut UCT, Leica) and the sections were stained with 2 wt% uranyl acetate and imaged with an optical microscope (DM2000 LED, Leica) and TEM (Tecnai Spirit Twin, FEI). 2.2.8. Cryo-Soft X-ray Tomography For cryo-soft X-ray tomography (cryo-SXT), gold Quantifoil R2/2 G200F1 finder grids were coated with fibronectin (Sigma). Cells were seeded on the grids and incubated overnight. The following day, culture media without FBS containing NPs was supplemented to the cells. After 3 h of incubation, the grids were rinsed with PBS and immediately frozen in liquid ethane cooled with liquid nitrogen using an automatic plunge freezer for the bare grid technique (EM GP, Leica Microsystems). The samples were initially examined by means of a cryo-visible light microscope (Axioscope, Zeiss) provided with a cryo-stage (CMS196, Linkam Scientific). Samples were stored under liquid nitrogen and subsequently kept under cryogenic conditions at all times during the measurements by X-ray microscopy in the vitrified state. Cryo-SXT imaging was performed at the MISTRAL beamline of the ALBA Synchrotron light source. Tomographic data was collected at 520 eV, irradiating the samples for 2 – 10 s per projection. 520 eV is the energy range in which water is transparent for X-rays and there are mostly absorbed by carbon, allowing the visualisation of biological material. In tomographic setup, images obtained at different sample orientations are computationally combined to produce a three-dimensional (3D) image, permitting the 3D representation of the sub-cellular ultrastructure of whole intact cells. A tilt series was acquired for each cell using an angular step of 1o on a 70o angular range. Each transmission projection image of the tilt series was normalised using flat-field. This process also considers the possibly different exposure time, as well as the slight decrease Chapter 2 – 37 – of the electron beam current during the acquisition. In order to increase the image quality, wiener deconvolution taking into account the experimental impulse response of the optical system [28] was applied to the normalised data. Finally, the Naperian logarithm was used to reconstruct the linear absorption coefficient (LAC). The resulting stacks were then loaded into IMOD software [29] and the individual projections were aligned to the common tilt-axis using the internal cellular structures as markers. Afterwards, the aligned stacks were reconstructed with algebraic reconstruction techniques (ART) [30]. The visualisation, segmentation and quantification of the volumes were carried out using Amira 3D software (Thermo Fisher). Each voxel of the reconstructed tomogram represents the LAC (μl, in cm-1) of the material contained in it, as it is related to the measured transmission signal I through the Beer-Lambert law (Equation 2.2): Voxel = µl = -ln(I I0)=∫μl(z)dz Equation 2.2 Where  corresponds to the mass absorption coefficient of the cell structure (cm2 g-1), I0 to the incoming flux, and z to the thickness of the material. 2.2.9. Statistical Analysis Data distribution was checked with Shapiro-Wilk test. Significant differences between samples were determined using non-parametric Kruskal-Wallis test followed by multiple pairwise comparisons. The significance level was set for p<0.05. Statistical analysis was performed using Minitab 19. All data are reported as mean ± standard error of the mean. 2.3. Results and Discussion 2.3.1. Characterisation of the Nanoparticles Hydroxyapatite (HA) nanoparticles (NPs) were synthesised using a wet precipitation method, obtaining non-doped HA and magnesium-doped HA (MgHA) powders. Figure 2.1A shows the XRD results of the synthesised NPs. All peaks matched those corresponding to HA (ICDD PDF 09-0432) with no secondary crystalline phases detected, demonstrating the precipitation of phase pure HA. Moreover, the broad peaks indicated the poor crystallinity of the powders obtained. As expected, magnesium doping caused a slight shift of the (002) diffraction peak, which correlates with the lattice contraction caused by the incorporation of Mg into the HA structure substituting Ca ions, due to the smaller ionic radius of the Mg cation compared to Ca [31]. Hydroxyapatite Nanoparticles-Cell Interaction – 38 – Figure 2.1. Physicochemical characterisation of the synthesised HA NPs. A X-ray diffraction for the non-doped (HA) and magnesium-doped (MgHA) NPs. B TEM micrographs for both types of nanoparticles. C Images of the discs made of compacted HA and MgHA NPs. Insets with higher magnification of the SEM pictures showing the nanoparticles of the discs. No differences in the morphology of the two types of NPs were observed (Figure 2.1B), both consisting of needle-like crystals with sizes of around 150 nm in length and 20 nm in width. The discs obtained by NPs compaction showed a homogeneous surface microstructure, with specular faces (Figure 2.1C). A close-up view in the insets allowed to observed the individual nanoparticles in the discs by SEM. Immersion of the discs in the culture medium did not alter their surface and no particles were dislodged from the samples, indicating excellent compaction. Chapter 2 – 39 – 2.3.2. Functionalisation of the Nanoparticles L-3,4-dihydroxyphenylalanine (DOPA) is a molecule that presents an excellent affinity for hydroxyapatite through the catechol groups [26,32]. Although the adsorption of DOPA to HA NPs has been explored to enhance the osteoconductive properties of scaffolds and to improve protein adhesion for drug delivery purposes [33,34], to the best of our knowledge this is the first time that it is used to fluorescently label HA NPs, by combining it with a fluorescent molecule such as carboxyfluorescein (CF). The evaluation of the adsorption capacity of either CF or the DOPA-CF peptide on the nanoparticles was done by measuring the difference in the concentration of the staining solution before and after NPs incubation. The direct adsorption of CF on the NPs was almost non-existent, or binding was so weak that the fluorophore was lost during the subsequent rinsing steps (data not reported). Figure 2.2A illustrates the quantity of DOPA-CF peptide adsorbed on the surface of the NPs, for both non-doped (HA-F) and magnesium-doped (MgHA-F) after rinsing. It is observed that in the case of non-doped NPs, the adsorption of DOPA-CF was about half of the concentration added, whereas in the case of MgHA powders, this adsorption was slightly lower, around 40% of the initially added peptide. This reduction of adsorbed DOPA-CF was further confirmed by direct fluorescence intensity data, shown in Figure 2.2B, where the intensity of MgHA-F NPs was 80% with respect to the HA-F sample. This variation might be due to physicochemical differences between both types of nanoparticles. After functionalisation of the NPs, various rinsing steps with bi-distilled water were carried out to eliminate the non-adhered peptide and the fluorescence intensity of the supernatants was measured. It is worth mentioning that the intensity in the first rinse after vortexing of the fluorescently labelled NPs was about 5%, which explained the excellent binding between DOPA-CF and the NPs. The zeta potential of the NPs measured in water is displayed in Figure 2.2C, revealing a slightly negative surface charge for the pristine nanoparticles. When functionalising the NPs with the fluorescent peptide, the surface charge decreased to -13 mV for both HA-F and MgHA-F, which confirmed the success in the functionalisation of the NPs. Hydroxyapatite Nanoparticles-Cell Interaction – 40 – Figure 2.2. Characterisation of the fluorescently functionalised NPs. A Adsorption of DOPA-CF peptide to HA and MgHA NPs at different initial concentrations. B Direct fluorescence intensity of the HA-F and MgHA-F NPs. C Zeta potential of the different materials. 2.3.3. Study of the Nanoparticles – Cell Membrane Interaction Any NPs internalisation process starts with the interaction of NPs on the cell membrane. Since the cell membrane is rich in receptors and some of them have a strong affinity for calcium, NPs could potentially induce changes in cells. From a fundamental point of view, it is essential to disclose if the cytotoxic effect of the HA NPs comes mainly from their interaction with the cell membrane or it is due to their internalisation inside the cells. Indeed, it has been reported that the toxicity of certain NPs (e.g. positively charged NPs) can derive from their reactivity when interfacing with the cellular membrane [35]. For this purpose, cell culture studies were done either by supplementing HA NPs on the cells or by seeding cells on NPs-compacted discs. In these homogeneous discs, uptake of the nanoparticles into the cells is prevented, and the interaction is limited to surface interaction with the cell membrane. In contrast, when the cells are exposed to the NPs suspended in the cell culture medium, both interactions are possible: not only the contact with the cell membrane but also their internalisation. In addition, the assay was done either with or without the supplementation of FBS in the cell culture medium to validate the direct interaction of the NPs with the cell membrane (FBS-) without the interference of adsorbed proteins, which are known to have a strong affinity for HA [36]. In general, Chapter 2 – 41 – these adsorbed proteins block the interaction of cellular receptors with the nanomaterial surface [37] causing a reduction in the cytotoxic effect due to a lower cell uptake [38]. Figure 2.3A shows the morphology of the cells adhered to the samples imaged with acridine orange (AO) staining in the confocal microscope and by SEM. The well-spread morphology of the cells accounts for their viability when seeded on the discs – both in HA and MgHA –, similarly to the control condition (i.e. cells seeded on coverslips, without material), regardless of the removal of FBS. The cells on the discs of the serumfree samples presented minor differences in their morphology compared to the glass coverslip, exhibiting star-like shapes instead of elongated profiles, probably caused by the nature of the material. However, this difference was not relevant since the cells appeared perfectly attached to the discs’ surfaces. The results of the samples with the NPs suspensions demonstrated that the cells supplemented with both types of NPs in serumcontaining medium presented no cytotoxicity. In contrast, in the absence of FBS, cells in contact with the suspension of HA or MgHA mostly died after 4 h, as observed by the drastic drop in the cell population. Figure 2.3B quantifies these findings by WST-1. The data obtained are consistent with the images, as cell death of 70% for HA NPs suspension was revealed, whilst MgHA NPs presented a killing potential for 80% of the cells, with statistically significant differences between the two conditions. In contrast, although a slight decrease in cell number was observed in HA FBS+ and in both types of discs in the FBScondition compared to the control, the differences were not statistically significant. The higher cytotoxicity observed for MgHA NPs, although limited in the present work, aligns with other studies in the literature, where the authors reported that doping HA NPs with small amounts of magnesium was observed to improve cytotoxicity in MG-63, leaving mesenchymal stem cells alive [7]. The fact that here the effect was small could be related to differences in the aggregation state of the HA and MgHA NPs, as will be later discussed. Regarding the differences observed in NPs suspension between FBS+ and FBSconditions, several authors have demonstrated that the adsorption of proteins on the surface of HA NPs can strongly affect the adhesion properties to the cell membrane, reducing the uptake levels [37,38]. The reason for the reduced internalisation was explained by the electrostatic repulsion between the cell membrane and the negative charge that proteins confer to the HA NPs [7,39]. Overall, the results obtained suggest that the direct interaction of the cell surface with the material was not sufficient to cause cell death and it must be ruled out as the main mechanism of cytotoxicity, HA NPs internalisation being a requirement for cell toxicity. Hydroxyapatite Nanoparticles-Cell Interaction – 48 – Figure 2.6. Intracellular calcium evaluation of cells exposed to the NPs for 3 h and stained with Fluo-4 AM. An apoptosis control was also studied. A Intracellular calcium quantified in a fluorescence microplate reader. B Intracellular calcium imaged by fluorescence confocal microscopy. Scale bars represent 20 µm. * indicates statistically significant differences. Various calcium stores are present in the cell, which could explain the selective staining of vesicles in Figure 2.6. It is well established that the main intracellular calcium store is the endoplasmic reticulum (ER), together with the Golgi apparatus and mitochondria. However, there are other small-volume stores, such as acidocalcisomes, endosomes or lysosomes, dispersed throughout the cell that are considered as secondary Ca2+ stores [45,46]. Despite the classification in main and secondary Ca stores, the concentration of calcium can reach comparable values between the lysosome lumen and the lumen of the ER [47]. As previously mentioned, Fluo-4 in all the serum-containing samples, as well as in the control FBSexhibited homogeneous staining of the cytosol, proving that the fluorophore was not able to enter the ER, Golgi apparatus or mitochondria, as expected. Thus, it seems feasible that the small green-stained vesicles ranging from 0.3 to 0.6 µm in size observed could be attributed to intracellular Ca2+ localised in individual vesicular stores (e.g. lysosomes, endosomes, etc.) enriched by the presence of dissolved HA NPs. A 0 1 2 3 4 5 6 FBS+ FBSRel. fluorescentintensity Control Apoptosis HA MgHA * * B FBS+ FBSControl HA MgHAApoptosis i ii iii iv v vi vii viii Chapter 2 – 49 – We expect these vesicles to have a compromised membrane to allow the entrance of the Fluo-4 molecule. It has to be noted that the fluorescence intensity for the serum-free samples loaded with NPs had to be decreased during the analysis to avoid the saturation of the vesicles. This adjustment reduced the overall fluorescence and masked the contribution of cytosolic calcium, which could be observed before this modification. Hence, the selective staining of calcium-rich vesicles does not exclude an increase in cytosolic Ca2+, which would be responsible for the cytotoxicity. Unfortunately, the results reported by other authors on intracellular calcium determination does not discriminate between free cytosolic Ca2+ from Ca2+ retained in vesicles. This is for instance the case of the studies that rely on the quantification of fluorescence intensity by microplate reader or by FC [17,20]. The observation of such stained vesicles has been sometimes attributed to subcellular compartmentalisation [48]. Even though we demonstrate the absence of these highly fluorescent vesicles in the control samples and in all the serum-supplemented conditions, which clearly discarded this possibility, further analysis has to be done in order to confirm the presence and determine the composition of the vesicles. 2.3.6. Transmission Electron Microscopy and Cryo-Soft X-ray Tomography Transmission electron microscopy (TEM) and cryo-soft X-ray tomography (cryo-SXT) were performed in selected specimens (i.e. serum-free samples) to check the uptake of NPs and to assess the distribution of calcium-rich vesicles inside the cells. TEM, as a high-resolution technique, with a resolution below 5 nm, is typically used to evaluate the internalisation of nanomaterials. In contrast, cryo-SXT is a novel technique that uses Xray to image frozen pristine samples in the water window region (i.e. between the carbon K edge at 284 eV and the oxygen K edge at 543 eV) with a resolution of about 25 to 40 nm. Within this region, X-rays are absorbed an order of magnitude more strongly by carbon and nitrogen-containing organic material than by water, allowing direct visualisation of the cellular structures while water remains nearly transparent [49,50]. Thus, comparing both techniques, while TEM has the advantage of resolution, it requires major processing steps such as fixing, dehydration and slicing to observe the samples. All along these steps, electrolytes and low molecular weight organic compounds can potentially be washed out from the cells and any information concerning the presence of ionic soluble content will be missed [45]. In addition, artefacts can be easily introduced during sample processing. Instead, cryo-SXT has a lower resolution than TEM but enables mesoscale imaging of the whole cellular ultrastructure without any processing other than vitrification [51,52]. Thus, the possibility to image cells under near-native conditions is a strong asset of this technique. Moreover, cryo-SXT offers the possibility to quantify the results and distinguish cellular structures based on linear absorption coefficients, following the Beer-Lambert law [49,53]. With respect to the TEM results, illustrated in Figure 2.7, samples of cells supplemented with HA and MgHA NPs in serum-free conditions were examined, together with a control without NPs. The results not only confirmed the uptake of nanoparticles in both supplemented samples, but they also denoted the absence of electrodense vesicles comparable in size to the ones observed by confocal fluorescence microscopy. Instead, other vesicles (either empty or with residual content) were seen, which was consistent with the idea that electrolytes such as Ca2+ and Pi were lost during sample processing. Hydroxyapatite Nanoparticles-Cell Interaction – 50 – However, it was impossible to know the original composition of these organelles using this technique. Figure 2.7. TEM images of cells without the addition of nanomaterials (control) and after the supplementation of HA and MgHA NPs in serum-free conditions. * indicates NPs. ER, L and M stand for endoplasmic reticulum, lysosomes and mitochondria, respectively. White arrows point vesicles (either empty or with residual content). Scale bars denote 2 µm. In contrast, cryo-SXT revealed interesting differences compared to TEM. Figure 2.8 shows the X-ray projection images that provide an overview of the cell of interest (i and iv), together with relevant tomographic sections taken from the tilt series (ii and v) and their segmentation (iii and vi). Videos of the whole aligned tilt series can be found in the Supplementary Information. In the selected sections, small highly absorbing vesicles of around 0.3 to 0.6 µm and reaching up to 1 µm were readily visible. Similar-looking vesicles have been ascribed by other authors as lipid droplets (LDs), due to their small size and their high carbon content that highly absorbs at 520 eV [54]. However, LDs are usually described as fully dense dark vesicles by cryo-SXT or, in some cases, as vesicles containing one lighter vesicle within [49]. Furthermore, three different components of the Chapter 2 – 51 – endocytic pathway can be distinguished. First, early endosomes that can be considered as sorting vesicles and are agreed to appear as membrane-bound organelles with a relatively uniform interior by cryo-SXT. Multivesicular bodies (MVBs), formed from early endosomes by the invagination of the membrane into their own lumen, which, imaged by cryo-SXT, should clearly contain one or more internal lighter vesicles. And lysosomes, the last organelle of the endocytic pathway that break down biomolecules into simple compounds thanks to many hydrolytic enzymes, and typically show an electrodense lumen with internal granular structures [49,54,55]. In this regard, Figure 2.8 v shows that most electrodense vesicles contained more than 2 low-absorbing vesicles, which suggested the presence of endocytic vesicles rather than LDs and, more specifically, of MVBs. Actually, MVBs are organelles formed by the maturation of early endosomes. They are also known as late endosomes and are always identified in the degradation pathway of any endocytic process [56]. In the case of Figure 2.8 ii, a lower density of MVB vesicles was observed along with larger grainy structures compatible with lysosomes or endolysosomes, as well as a large vesicle containing pristine HA NPs. In general, the large number of the darker MVBs would match with the green-stained vesicles observed by fluorescence confocal microscopy (Figure 2.6B). However, distinguishing LDs from MVBs is not so straightforward. Hydroxyapatite Nanoparticles-Cell Interaction – 52 – Figure 2.8. Cryo-SXT of cells after 3 h of exposure to HA and MgHA NPs. i and iv showing mosaic projections, ii and v display relevant tomographic sections, and iii and vi reveal the segmentation of the same stacked tomograms. N denotes nuclei and * stands for internalised NPs. Nuclei are segmented in yellow, lipid droplets in pink, nanoparticles in blue and multivesicular bodies in purple. Scale bar represents 5 µm. Apart from discrimination of the different organelles by visual inspection, cryo-SXT allows quantitative determination of the linear absorption coefficient (LAC) for each Chapter 2 – 53 – cellular component [50,57]. Therefore, in order to better elucidate the nature of these small vesicles, the selected tilt series were segmented and their LACs quantified (Table 2.2). Nuclei (segmented in yellow) resulted in a LAC of 0.29 ± 0.02 µm3, while internalised NPs presented a high LAC value of 1.94 ± 0.02 m3. In addition, two different vesicles were identified: LDs with a LAC of 1.42 ± 0.03 m3, and MVBs with a LAC value of 1.09 ± 0.05 m3. It is worth noting that the standard deviation obtained for these values is very low, this indicates the high similarity of LACs between both sample conditions. The differences in LAC values between LD and MVB proved compositional differences between both types of vesicles. Thus, based on LAC values we could unambiguously map MVBs and LDs from the analysed cells (Figure 2.8 iii and vi). It is known that LD is typically the structure with the highest LAC value due to its increased carbon content compared to the other organelles, which agrees with the results obtained (LAC of LDs was 1.42 m3 vs. LAC of nuclei 0.29 m3). The fact that cells with internalised NPs contained MVBs with LAC values close to the LDs ones, implies that these vesicles should contain highly absorbing elements. In combination with the very high LAC value of the solid NPs, it indicates that the electrodense vesicles may contain degraded NPs. Accordingly, these results suggest that the Ca2+-rich vesicles identified by confocal microscopy, should not be considered as artefacts and they likely correspond to MVBs. Table 2.2. Linear absorption coefficient (LAC) of the different cellular components Nucleus (µm3) Nanoparticles (µm3) Lipid Droplets (µm3) Multivesicular Bodies (µm3) Yellow Blue Pink Purple HA 0.29 ± 0.02 1.94 ± 0.02 1.42 ± 0.03 1.11 ± 0.03 MgHA 1.07 ± 0.04 Moreover, although both samples had Ca-rich vesicles with similar LACs (i.e. 1.11 ± 0.03 m3 and 1.07 ± 0.04 m3 for HA and MgHA, respectively), the segmentation revealed different size and distribution of these vesicles. Specifically, HA presented a lower amount of vesicles but of larger dimensions, whilst the vesicles in MgHA were smaller but present in higher amounts. Analysis of the tomograms obtained for the cell exposed to HA NPs showed the capture of a less advanced degradation state of the NPs, which together with the tomogram of the cell exposed to MgHA allows seeing the evolution of the degradation of these materials. An additional aspect to explore in the future from cryo-SXT is cryospectroscopic imaging by changing the imaging energy from 520 eV. In a recent work by Gal et al. the authors set the energy to 353 eV, which is the absorbing energy for Ca, to investigate the distribution of calcium-containing organelles in calcifying and non-calcifying species in algae [58]. A very distinct Ca2+ map was shown, proving to be a powerful tool to trace calcium stores in pristine conditions. This technique applied to our samples would allow gaining further insights into the quantification and distribution of intracellular calcium. The study of intracellular calcium is of major relevance when trying to elucidate the fate of internalised NPs, as changes in the cytosolic calcium of the cells help to understand Hydroxyapatite Nanoparticles-Cell Interaction – 54 – the cytotoxicity caused by HA NPs. Through the combination of different techniques, both conventional and advanced, we have been able to detect the presence of multiple MVB which potentially carry degraded NPs. However, discrimination of these Ca-rich stores from cytosolic calcium would be needed to better assess toxicity. We have demonstrated the limitations of various techniques in this regard, but also the great potential of intracellular calcium probes combined with cryo-SXT. 2.4. Conclusions The main goal of this work was to understand and to prove the cytotoxic mechanism of hydroxyapatite NPs using the common techniques in the field together with cryo-SXT imaging at the ALBA synchrotron. The studies revealed the internalisation of the HA NPs as the main toxicity mechanism, with minor contribution coming from the material interaction with the outer cell surface. 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The solution was sonicated with a highfrequency ultrasound probe sonicator (450D, Branson Digital) using a 3 mm diameter tip at 40% amplitude for 2 min. Afterwards, the NPs dispersion was mixed 1:1 with 30 wt% gelatine solution (Fluka) at 50 oC and 40 ml were added dropwise into a beaker containing 400 ml of olive oil at 50 oC under stirring for the synthesis of the microspheres. After 10 min, the emulsion was cooled down to 4 oC and stirred for 30 min. Subsequently, 400 ml of cold acetone were added and the mixture was left in stirring for 1 h before being filtered to collect the synthesised microspheres. The final product was rinsed with acetone, dried and further sieved to 40-100 µm. Moreover, a control condition consisting of gelatine MS without the addition of NPs was included in the study. To synthesise these particles, 40 ml of 15 wt% gelatine solution at 50 oC were added into the initial beaker and the previously described protocol was followed. 100 mM N-(3-dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC, SigmaAldrich) and 150 mM N-hydroxysuccinimide (NHS, Sigma-Aldrich) solutions were prepared in water/cold acetone (20:80 v/v). They were mixed 1:1 and used to immerse the sieved MS for 1.5 h. Afterwards, the powder was thoroughly rinsed with ddH2O and dried. 3.2.3. Physicochemical Characterisation of the Nanoparticles Characterisation of the NPs comprised analysis of the phase composition by X-ray diffraction (XRD, D8 Advance, Bruker) using Cu Kα radiation at 40 kV and 40 mA. Data were collected with a step size of 0.02o over the 2θ range from 20 to 60o with a counting time of 2 s per step. Phase identification was accomplished by comparing the results with the standard patterns of HA (ICDD PDF 09-0432). The crystallite size (D) of the NPs was determined following Scherrer’s equation: D =K·λ β·cosθ Equation 3.2 Where K is a constant varying with crystal habit and here is set to 0.9, λ corresponds to the wavelength of the X-ray beam (i.e. 0.15406 nm for Cu K radiation),  is the full width at half maximum for the (002) diffraction peak, and  is the diffraction angle. In addition, Raman spectroscopy (inVia Qontor, Renishaw) was conducted with a 723 nm laser at 100% power. The frequency range analysed was 726-1812 cm-1 with 1 s exposition and 40 accumulations per sample. 3.2.4. Physicochemical Characterisation of the Microspheres The morphological evaluation of the different MS was performed by optical microscopy (AE2000, Motic). Moreover, the particle size distribution was assessed through image analysis (FIJI, ImageJ software, [45]) of the obtained acquisitions. MS composition was examined through Raman spectroscopy, using the previously described parameters. Chapter 3 – 65 – 3.2.5. Study of the Ion Release in Cell Culture Media In order to study the ionic release of the microspheres in cell culture medium, 10 mg of each MS were put in contact with 1 ml of Dulbecco’s modified eagle medium (DMEM) supplemented with 10% foetal bovine serum (FBS), 20 mM 4-(2-hydroxyethyl)-1piperazineethanesulfonic acid buffer (HEPES), 2 mM L-glutamine, 50 U ml-1 penicillin and 50 µg ml-1 streptomycin (all from Gibco), for four days at 37 oC. The medium was changed daily and the supernatants were collected and diluted 20-fold in 2 wt% HNO3 (69 wt% for trace metal analysis, PanReac AppliChem). Ca2+, Pi and Mg2+ concentrations were determined by inductively coupled plasma optical emission spectrometry (ICPOES, 5100, Agilent Technologies), whereas inductively coupled plasma mass spectrometry (ICP-MS, 7800, Agilent Technologies) was used to study Sr2+ and Zn2+ levels in cell culture medium. 44Ca, 31P, 24Mg, 66Zn and 88Sr signals were calibrated against a multi-element standard solution (Inorganic Ventures). This experiment was done in triplicate (n=3) for each MS type. In addition, the MS after the ion release experiments were rinsed in water followed by ethanol and, upon drying, were analysed by Raman spectroscopy using a laser power set to 10%, 2 s of exposition and 100 accumulations. 3.2.6. Statistical Analysis All data were reported as mean values ± standard deviation. A normality test was performed to determine if data were modelled by a normal distribution. Statistically significant differences between groups were analysed by one-way ANOVA test followed by Tukey’s post-hoc pairwise comparison. Statistical analysis was performed using Minitab Statistical Software. 3.3. Results and Discussion The aim of this work was to develop microspheres for bone regeneration applications with enhanced bioactivity and the capacity to deliver therapeutic ions. To this end, composite MS made of gelatine and HA NPs doped with either Mg2+, Zn2+ or Sr2+ were synthesised by a simple water in oil emulsion method. The main consequences of this doping strategy were investigated by studying the ion incorporation to the crystal lattice of HA NPs, the morphological implications in the final MS and their ion release behaviour, which would determine the availability of the therapeutic ions in the surrounding media. This strategy permitted the sustained release of trace metals critical in the bone remodelling process while supplying attaching moieties to the cells. 3.3.1. Physicochemical Characterisation of the Doped Hydroxyapatite Nanoparticles Hydroxyapatite nanoparticles were successfully synthesised using a wet precipitation method, obtaining non-doped hydroxyapatite (HA), magnesium-doped (MgHA), zincdoped (ZnHA) and strontium-doped (SrHA) NPs. Figure 3.1A shows the XRD patterns of each synthesis. All peaks matched those corresponding to HA (ICCD PDF 09-0432), demonstrating the precipitation of phase pure HA with no secondary crystalline phases Microspheres with Ion-Doped Hydroxyapatite Nanoparticles – 66 – obtained, except for the Zn10 specimen, where a small peak between the (102) and (201) diffraction peaks revealed the formation of a new phase consisting on Zn3(PO4)2. In addition, broader peaks were observed for the two ZnHA samples, indicating a reduction in the crystallite size and/or an increase in the crystal strain with the incorporation of Zn2+, as also reported by other authors [46]. It has been extensively reported that ionic substitutions in the crystal lattice of HA can create distortion of the crystal structure due to discrepancies between the radius of the substituting ion and the substituted one. This mismatch usually causes the contraction or expansion of the lattice, resulting in peak shifting in the XRD spectra. Therefore, in the present work, the peak shift of the (002) diffraction peak was investigated because it does not overlap with other reflections and has a sharp shape [47]. As expected, the detailed examination of the (002) diffraction peak revealed slight shifts when introducing the doping ions. The incorporation of Mg2+ and Zn2+ produced a slight shift towards higher 2θ angles, whereas a deviation to lower angles was found for SrHA NPs. Moreover, these shifts were accentuated as more doping ions were incorporated during the synthesis. This phenomenon is caused by the substitution of Ca2+ ions by the different doping ions into the HA crystalline structure. It has been previously reported that the smaller size of the Mg (0.065 nm) and Zn (0.075 nm) radius compared to the Ca ionic radius (0.099 nm) reduces the cell parameters of HA, which explains the (002) peak shift towards higher angles. In contrast, Sr2+ presents a larger ionic radius (0.120 nm) than Ca, and its incorporation on the lattice is reported to cause an expansion of the parameters, producing the peak shift to lower angles [20,48]. An estimation of the crystallite size was calculated using Scherrer’s equation in the (002) reflection, and the results are displayed in Figure 3.1C. The value does not have a standard deviation because it was calculated only in one X-ray diffractogram. A crystallite size of 27.85 nm was found for the non-doped HA NPs and was maintained at 27.97 nm for the Mg5 sample. This parameter decreased to 24.00 nm in the case of Mg10, to 25.73 and 24.68 nm for the Zn5 and Zn10 samples, respectively, and to 26.91 nm for the Sr5. In contrast, a crystallite size of 31.52 nm was obtained for the Sr10 sample. This follows the expected trend, as in general, the incorporation of ions in the crystal lattice should entail a reduction in crystallite size regardless of the type of ion, as has been repeatedly reported in the literature [23,25,46,49]. The low incorporation of Mg2+ in the Mg5 condition may explain the maintenance of the HA crystallite size. On the other hand, the increase of this parameter for the Sr10 sample was unexpected since the literature reported a reduction of crystal size for strontium substitutions of up to 50 at% [50]. The results of Raman spectroscopy are summarised in Figure 3.1B. They revealed the presence of the characteristic vibration bands of HA in all the samples. The band at ≈960 cm-1 corresponds to the v1 symmetric stretching vibration of the P-O bond and the vibration bands in the region between ≈1030-1070 cm-1 to the phosphate v3 antisymmetric stretching vibrations [51]. Additionally, the 960 cm-1 band presented a marked shift to the left in both SrHA samples, which has been reported by other authors to result from the crystal lattice expansion induced by the incorporation of Sr2+ ions [52]. Furthermore, the same band was wider in the Mg10 sample, a phenomenon that has also been reported in the literature. Indeed, the broadening of the 960 cm-1 band is associated with the deformation of PO43structure by surrounding magnesium ions, which demonstrates the Mg2+ substitution in the crystal lattice of HA that causes the formation of a new vibration band at 950 cm-1 [48,53]. Chapter 3 – 67 – Figure 3.1 Characterisation of the synthesised ion-doped hydroxyapatite nanoparticles. A X-ray diffractograms and magnified region (25-26.5o) of the (002) peak. B Raman spectra and amplification of the 960 cm-1 band. Arrows indicate small changes in the spectrum. C Crystallite size (in nm) of the different NPs, calculated in the (002) reflection with Scherrer’s equation. Previous characterisation studies by ICP performed in our group using the same synthesis protocol further confirmed the substitution of Mg2+, Zn2+ and Sr2+ ions for Ca2+ ions in the HA structure, with a decrease of the Ca/P ratios in a dose-dependent manner for all the substituted NPs, while maintaining the (Ca + doping ion)/P ratios at ≈1.7 [43]. The ion doping concentration in the synthesised NPs were 1.23 wt% and 2.17 wt% for the Mg5 and Mg10 samples, 3.98 wt% and 8.04 wt% for the Zn5 and Zn10 NPs, and 3.54 wt% and 5.70 wt% for the Sr5 and Sr10 specimens. The ionic incorporation of Mg was close to the limit reported in the literature (i.e. 2.5 wt% [25]) for the higher concentration sample, while the strontium addition was far away from the expected [26]. Furthermore, the Zn incorporated was more than 80%, indicating a high addition efficiency. 3.3.2. Physicochemical Characterisation of the Microspheres The morphology of the synthesised microspheres both in a dry and wet state was analysed by optical microscopy after sieving (Figure 3.2). The incorporation of NPs reduced the transparency of the gelatine MS, regardless of the type of NPs. The sphericity of the gelatine MS was maintained when introducing HA, 5 wt% of ZnHA, and SrHA NPs at both 5 and 10 wt%. However, the MS tended to lose their shape with the magnesium doping. It was hypothesised that in these syntheses, the gelification of the MS was shielded by the interaction between the NPs and the gelatine chains. This was supported by the work of Xing et al., where they purified a gelatine solution from divalent metal ions and the subsequent chemical cross-linking with EDC formed a stronger gelatine network thanks to the removal of the interactions between divalent cations and the carboxylic groups of the gelatine chains [54]. Thus, we believe that the leaching of certain ions from the NPs interferes with the cross-linking process. Moreover, the swelling Microspheres with Ion-Doped Hydroxyapatite Nanoparticles – 68 – observed for the wet MS despite the cross-linking process was due to the hydration of the gelatine chains. Figure 3.2 Morphology of the different microspheres after the incorporation of Mg2+-, Zn2+- and Sr2+-doped HA nanoparticles, either in a dry state or in water suspension. Microspheres without nanoparticles (Gel) and with non-doped nanoparticles (Gel-HA) were included as controls. Image analysis of the optical acquisitions was performed to determine the size distribution of the different syntheses. Figure 3.3 contains the data of each type of MS once sieved, cross-linked and dried. All the conditions presented a monomodal normal particle size distribution (Figure 3.3A), with an average diameter size of 62 ± 13 µm. The control – gelatine – MS showed a mean size of 72 ± 15 µm, while the addition of HA NPs led to a marked decrease in the diameter size. Precisely, the incorporation of non-doped HA NPs reduced the mean size down to 59 ± 13 µm, whereas the zinc and strontium doping diminished it in a dose-dependent manner. The diameter sizes were 48 ± 14 µm and 55 ± 14 µm for the MS with higher doping levels (i.e. Gel-Zn10 and Gel-Sr10, respectively). In contrast, the addition of magnesium-doped NPs in the MS presented a completely different pattern: whilst Gel-Mg5 exhibited similar results compared to the Gel-HA samples, the MS with the higher Mg2+ content (i.e. Gel-Mg10) showed a marked increase in the mean diameter up to 91 ± 16 µm. This may be attributed to changes in the viscosity of the initial gelatine-NPs solution. As explained above, the leaching of certain divalent cations from the NPs can potentially interact with the carboxylic groups of the gelatine chains, creating hydrogen bonds between them. This has been shown to increase the Chapter 3 – 69 – hydrogel viscosity and to impair the subsequent covalent cross-linking through EDC chemistry [54]. Figure 3.3 Particle size distribution of the microspheres. A Distribution plot representing the frequency of microsphere diameter observations, and B Boxplot indicating mean (x), quartiles, max, min and outliers (o). Each outlier corresponds to a single microsphere. Different letters indicate statistically significant differences between conditions. Furthermore, Raman spectroscopy was conducted to identify the chemical composition of the synthesised MS. Several characteristic bands were studied, the major one of HA and four corresponding to gelatine, as displayed in Figure 3.4A. The presence of HA NPs was demonstrated with the ≈960 cm-1 band, which corresponds to the symmetric stretching mode v1 vibration of the PO43groups [56]. Regarding the gelatine, four different group of bands were identified in the Raman spectra. The first one, from ≈800 to 1040 cm-1, corresponds to the stretching mode vibrations of different C – C bonds, mainly assigned to amino acids such as proline, hydroxyproline, tyrosine and tryptophan [56]. The amide III band of the gelatine was observed from ≈1240 to 1300 cm-1, and is related to C – N stretching vibrations and N – H bending vibrations. The third band (≈1450 cm-1) corresponds to CH2 in-plane bending vibration mode of the gelatine, while the last region (≈1650 to 1700 cm-1) is the amide I band that correlate to C = O stretching bond vibrations [55,57]. The cross-linking of the gelatine of the MS did not result in major changes in the spectra when compared to the non-cross-linked MS. The four characteristic group of bands of the gelatine were observed for all the MS after cross-linking them (Figure 3.4B), which demonstrated that the reticulation of the gelatine was not altering the composition of the final MS. In addition, the band corresponding to the HA was found in all the samples containing HA NPs, which accounts for the proper incorporation of the particles in the MS, without losing them after the cross-linking and subsequent rinsing processes. It was observed that the cross-linking process resulted in a slight increment in the intensity of the amide vibration bands, which was more pronounced in the amide I band. This observation is in agreement with previous works [58,59], which attributed this phenomenon to the formation of new amide groups by the EDC cross-linking between carboxylic acids and primary amines. Microspheres with Ion-Doped Hydroxyapatite Nanoparticles – 70 – Figure 3.4 Raman spectroscopy of the MS (A) before and (B) after the cross-linking process. The grey band at 960 cm-1 is characteristic of HA, while the yellow regions are distinctive of gelatine: carbon bonds (C-C) region, assigned to ν(C-O-C) of proline, hydroxyproline and tyrosine at 813 cm-1, ν(C-C) of proline, hydroxyproline, tyrosine and tryptophan at 853-873 cm-1, ν(C-C) of proline at 919 cm-1, and ν(C-C) of phenylalanine at 1002 cm-1; amide III band at 1240-1300 cm-1; vibrations of CH2 bonds at 1450 cm-1; and amide I band at 1650-1700 cm-1 [56]. 3.3.3. Ion Release in Cell Culture Media HA NPs are reactive when placed in solution, exchanging ions with the environment. This has been precisely the main attraction of these NPs as they can potentially deliver key ions introduced within their structure. The main concern is to have an appropriate dose released, within the therapeutic range. The release behaviour of the ions from the MS into the cell culture media along time was checked by ICP analysis. The major ions (i.e. Ca2+, Pi and Mg2+) were quantified by ICPOES, while the minor ions (i.e. Zn2+ and Sr2+) were analysed by ICP-MS. These results are represented in Figures 3.5, 3.6 and 3.7. Figure 3.5 summarises the concentration of major ions in the cell culture medium (DMEM). It can be observed that the concentration of these elements remained stable in DMEM during the course of the study: 2.03 mM Ca2+, 1.30 mM Pi and 0.85 mM Mg2+ (Figure 3.5A), in agreement with the levels reported by the manufacturer. Although the control (Gel) MS slightly lowered the concentrations of these ions in the medium since the first moment, the decrease was small and the levels remained quite stable throughout the culture (i.e. ≈1.77 mM, ≈1.14 mM and ≈0.75 mM for Ca, P and Mg ions, respectively, Figure 3.5B). The ion uptake by the Gel MS could be attributed to either the electrostatic attraction of free cations by the carboxylate groups of the gelatine or to the precipitation of solid calcium phosphate [54]. On the contrary, when DMEM was put in contact with MS containing any type of NPs, a clear decrease in the concentrations of Ca and P ions from day 1 onwards was found, indicating a depletion of these ions. At 6 h there was a higher content of Ca2+ in the medium for some of the MS (Gel-Mg10 (2.31 mM, Figure 3.5E), Gel-Zn5 (2.19 mM, Figure 3.5F) and Gel-Sr10 (2.88 mM, Figure 3.5I), as well as a higher content of Pi in the case of Mg-containing MS (1.55 mM in both Gel-Mg5 and Chapter 3 – 71 – Gel-Mg10, Figure 3.5D and E). Nonetheless, the concentrations of calcium decreased down to 1.10 - 1.40 mM on day 4 and the phosphorus concentration was around 0.70 and 0.90 mM on the same day for all the NPs-containing samples. Figure 3.6 shows the ICP-MS quantification for trace metal ions in the cell culture medium, namely zinc and strontium. Their ionic levels in the control DMEM were stable during the study at around 18 µM and 2.8 µM, respectively, Figure 3.6A. Very similar concentrations were obtained when incubating Gel MS, indicating no interaction between these ions and the gelatine chains. However, a slight decrease in the zinc content was observed on the 4th day of culture (13.9 µM, Figure 3.6B). This reduction is attributed to the electrostatic attraction of the cation for the carboxylic groups of the gelatine chains, as suggested previously for Ca and Mg ions. Contrastingly, an increase in the initial Zn and Sr concentrations was observed in the NPs-containing MS. As expected, this phenomenon was especially pronounced in Znand Sr-doped MS for Zn2+ and Sr2+ ions, respectively (Figure 3.6F-I). Indeed, the concentrations increased notoriously for the samples containing higher amounts of doping ion (i.e. >1000 µM Zn2+ for Gel-Zn10 and >500 µM Sr2+ for Gel-Sr10). Furthermore, in the conditions with Zn and Sr doping, it was observed that the levels of these ions were always above those found in the cell culture medium, demonstrating a continuous release from the MS. Notwithstanding, this release was progressively waning with time, following an asymptotic trend. In another vein, one striking result was the release of Zn2+ and Sr2+ in the Gel-HA and Gel-Mg10 MS (Figure 3.6C and E). The presence of such ions in these samples may be explained by impurities in the reagents used in the NPs synthesis. For example, the presence of ionic impurities in the Ca(OH)2 is common and can easily show up when measuring ionic concentrations at the micromolar level [60]. Microspheres with Ion-Doped Hydroxyapatite Nanoparticles – 72 – Figure 3.5 Calcium, phosphorus and magnesium ionic concentrations in the cell culture medium after the incubation of the different MS at different times. Results obtained by ICP-OES. Chapter 3 – 73 – Figure 3.6 Zinc and strontium ionic concentrations in the cell culture medium after the incubation of the different MS at different times. Results obtained by ICP-MS. Additionally, the cumulative release from or uptake by the material was calculated and represented in Figure 3.7A and B. Rather than a release, a sustained uptake of Ca and P ions from the medium was noted for all the conditions. Such behaviour could be explained by the precipitation of CaP fostered by an initial ion release originated by the dissolution of the HA NPs. This phenomenon has been previously reported in the literature [61–63]. On the other hand, the ionic exchange of magnesium ions (Figure 3.7C) was less accentuated than the uptake observed for Ca and P. This indicates the limited implication of Mg in the precipitation of CaP, which can be explained by its marked inhibitory effect on HA nucleation and growth [64,65]. Altogether, the early release of ions in the medium Microspheres with Ion-Doped Hydroxyapatite Nanoparticles – 80 – Technol. Appl. 2 (2021) 100056. doi:10.1016/j.carpta.2021.100056. [42] S. Dasgupta, S.S. Banerjee, A. Bandyopadhyay, S. Bose, Znand Mg-Doped Hydroxyapatite Nanoparticles for Controlled Release of Protein, Langmuir. 26 (2010) 4958–4964. doi:10.1021/la903617e. 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Chapter 4 Microsphere Incorporation as a Strategy to Tune the Biological Performance of Bioinks Chapter 4 – 85 – MICROSPHERE INCORPORATION AS A STRATEGY TO TUNE THE BIOLOGICAL PERFORMANCE OF BIOINKS Scope The modulation of the hydrogel properties is required in the development of bioinks in order to achieve a self-supporting ink that is extrudable and maintains its stability after printing while assuring the viability of the cells embedded within. One underexplored strategy is the incorporation of microspheres to alginate-based formulations as a way to provide cell-adhesion moieties while improving the mechanical properties of the final cell-laden inks. Therefore, in Chapter 4, the gelatine and gelatine-hydroxyapatite microspheres developed in the previous chapter, together with calcium-deficient hydroxyapatite microspheres, will be added in the formulation of bioinks, and their rheological and biological implications will be assessed. Will their surface serve as an anchorage point for the osseous cells? Will the mineral charge be sufficient for the stimulation of osteogenic differentiation of the cells embedded? Tunning Bioinks with Microspheres – 86 – 4.1. Introduction In recent years, the tissue engineering field has progressed from two-dimensional (2D) cell culture to 3D models. Indeed, 3D systems are biologically more relevant, as they can reproduce better the physiological environment of the tissues [1,2]. This evolution has been possible thanks to the expansion of additive manufacturing technologies, that permit to develop complex structures using layer by layer deposition. Moreover, they allow to control and tune different features, such as the scaffold architecture or the size and shape of the porosity for the specific application and cell type [3]. The initial approach concerning 3D models was to generate biocompatible scaffolds that could be further seeded with cells. More recently, 3D bioprinting has advanced these techniques by the loading of cells directly in the ink, creating a promising approach to enhance the regenerative process. Although a lot of advances have been done in the development of extrusion-based bioinks, the optimal mechanical and biological requirements are usually antagonistic, which makes the perfect bioink difficult to obtain [4]. Indeed, low polymer content bioinks, which would be desirable for cell survival and cell migration, may compromise the shape fidelity of the printed struts. In this sense, the ideal cell-laden ink must be mechanically stable and provide high resolution when printed. Regarding the cellular needs, it must be highly biocompatible and allow the proliferation and differentiation of the cells [5–8]. Hydrogels are the most used materials for the formulation of bioinks, mainly due to their ability to sustain living cells and their adjustable mechanical properties. In particular, alginate is currently one of the most attractive candidates, considering its easy crosslinking capability, which takes place through divalent cation gelation, high availability, low cost and safety not only for the human body but also for the embedded cells [9]. Nevertheless, its main drawback is the lack of bioactivity, which has led to the implementation of additional processing strategies. One of them is the incorporation of bioactive polymers, such as gelatine, collagen, hyaluronic acid and chitosan among others, to encourage cell adhesion and proliferation in a tissue-specific manner [10,11]. However, due to the complexity of these proteins, it becomes difficult to control the specific signals presented to the cells. The functionalisation of the initial polymer with short cell-adhesive peptides, mainly RGD-based sequences, has been proposed as an alternative way to mimic the extracellular matrix (ECM) and to improve the cell attachment and spreading within the bioinks [12–14]. The main drawback of this approach is the requirement of an additional cross-linking step to bind the bioactive molecule to the alginate [15,16], which not only lengthens the process but may adversely affect the rheological properties of the final bioink and, in turn, their manipulation and in vitro performance. As an alternative strategy, the combination of alginate with different inorganic materials has been explored, which is especially attractive in the field of bone regeneration as it may allow endowing the ink with additional osteogenic cues. The incorporation of bioglass nanoparticles (NPs) capable of delivering biologically active ions such as calcium and silicon in the bioinks has been studied, reporting increased ALP levels and enhanced osteodifferentiation [17,18]. In the same line, the use of hydroxyapatite (HA) particles in gels and bioinks has been proved to increase osteogenic differentiation [19– 21]. Besides the incorporation of nanomaterials, the addition of microspheres (MS) in the formulation of bioinks may be of major interest as they can solve the cell-adhering issues Chapter 4 – 87 – by providing the cells with a surface to attach and improve the mechanical properties of the final constructs. Some studies explored this strategy, using polymeric MS [22–24]. Levato et al. added PLA MS into GelMA-based bioinks reporting increased compressive modulus, as well as enhanced cell adhesion and osteogenic differentiation for the condition with preseeded MS [22]. Likewise, in the work of Tan et al. the compressive properties of agarose-collagen hydrogels were improved by the incorporation of PLGA MS. Moreover, they demonstrated the viability and proliferation of different cell types preseeded on the MS [23]. Abu Awwad et al. explored the delivery of bone therapeutic agents from PLGA MS in a bioprinted system, showing the promotion of osteogenic differentiation of MSC [24]. In spite of the interesting findings, in these works, the authors only explored the incorporation of PLA/PLGA MS to the bioinks. In the present work, we intend to assess the performance of MS with different compositions, including polymeric, hybrid and ceramic MS, exploring their role not only as a support for cell adhesion but also as a way to tune cell behaviour. We hypothesise that the design of simple, robust and tuneable alginate-based bioinks through the incorporation of microspheres with improved affinity for bone, would be a strong asset to the field of bone bioprinting. MS of three different compositions are analysed, including: i) gelatine, ii) hydroxyapatite nanoparticle (HA NPs)- containing gelatine; and iii) calcium-deficient hydroxyapatite. The performance of the different MS is tested by direct cell seeding on the MS and also after bioink extrusion of a formulation combining an alginate/gelatine hydrogel with MS and cells. The ability of the cells to colonise the MS in the bioink formulation is further tested using two different degrees of hydrogel cross-linking. 4.2. Experimental Section 4.2.1. Synthesis of Gelatine Microspheres Gelatine microspheres (Gel MS) were prepared by a water in oil emulsion. Specifically, gelatine (Fluka) solution 15 wt% in bi-distilled water (ddH2O) was preheated to 50 oC. Subsequently, the solution was added dropwise into 50 oC preheated 400 ml olive oil with mechanical stirring at 900 rpm to form an emulsion. After 10 min, the mixture was put in an ice bath and stirred for 30 min, prior to the incorporation of cold acetone. The emulsion was left stirring for 1 h before being filtered to collect the synthesised MS. The final powder was rinsed with acetone and left to dry before sieving to the desired particle size (i.e. 40-100 µm). The MS were further cross-linked following a protocol slightly different from the one applied in Chapter 3. The MS were immersed in a cold aqueous solution of 50 mM N-(3-dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC, SigmaAldrich) and 75 mM N-hydroxysuccinimide (NHS, Sigma-Aldrich) for 1.5 h and thoroughly rinsed with distilled water. 4.2.2. Synthesis of Hydroxyapatite-Containing Gelatine Microspheres For the synthesis of the hydroxyapatite-containing gelatine microspheres (Gel-HA MS), hydroxyapatite nanoparticles (HA NPs) were incorporated into the initial gelatine solution before following the previous protocol. Therefore, HA NPs were produced by a wet precipitation method, as described elsewhere [25]. Briefly, 200 mM H3PO4 (85 wt% pure, PanReac AppliChem) was added dropwise into a solution of 333 mM Ca(OH)2 (96 wt% pure, Fluka) under constant stirring at 40 oC at a rate of 1 ml min-1, according to the Tunning Bioinks with Microspheres – 88 – Equation 4.1. The supplementation of phosphoric acid was stopped when the pH reached 8 and the product was left stirring for 30 min. Afterwards, stirring was stopped and the suspension was left to mature overnight at room temperature before being rinsed three times with ddH2O after centrifugation (5430 R, Eppendorf). The final powder was freezedried (Cryodos, Telstar). 10 Ca(OH)2 + 6 H3PO4 → Ca10(PO4)6(OH)2 + 18 H2O Equation 4.1 A suspension of 10 wt% HA NPs and 10 wt% sodium citrate (Sigma-Aldrich) was prepared and subsequently sonicated in a high-frequency ultrasound probe sonicator (450D, Branson Digital) using a 3 mm diameter tip at 40% amplitude for 2 min. The dispersion was mixed 1:1 with a 30 wt% gelatine solution and added in the oil as previously described. This resulted in microspheres loaded with 25 wt% of HA (dry weight), which were additionally sieved and cross-linked following the same crosslinking procedure as for the Gel MS. 4.2.3. Synthesis of Calcium-Deficient Hydroxyapatite Microspheres A previously established protocol [26] was adapted for the synthesis of the calciumdeficient HA (CDHA) MS. The powder phase of the calcium phosphate cement consisted of α-tricalcium phosphate (α-TCP), obtained by sintering a mixture of calcium hydrogen phosphate (CaHPO4, Sigma-Aldrich) and calcium carbonate (CaCO3, Sigma-Aldrich) to a Ca/P ratio of 1.5 in an oven (CRN-58, Hobersal) at 1400 oC for 15 h and subsequent quenching in air. The resulting α-TCP was milled in a planetary mill (Pulverisette 6, Fritsch GmbB) to obtain a mean powder-particle size of 2.8 μm [27], and 2 wt% of precipitated HA (Merck) was added as a seed. In order to obtain the calcium phosphate cement slurry, the seeded α-TCP powder and 1 wt% of sodium citrate were mixed with a 5 wt% gelatine solution in phosphate-buffered saline (PBS, Gibco) in a liquid-to-powder ratio of 1.2. This resulted in microspheres with 94 wt% load of mineral content (dry weight). To synthesise the MS, 3 ml of the ceramic slurry were added drop by drop into 300 ml of olive oil previously heated to 60 oC. The emulsion was mechanically stirred (BDC 2002, Heidolph) for 2 h. Afterwards, the MS were separated from the emulsion by the addition of 0.01% Triton X-100 (Sigma-Aldrich) in Ringer’s solution (0.9% sodium chloride, PanReac AppliChem). The MS were further washed with acetone and the remaining gelatine cross-linked with EDC/NHS following the previously described protocol. Finally, the powder was immersed in Ringer’s solution for a week, to allow the hydrolysis of α-TCP to CDHA, according to Equation 4.2. Once the setting was done, the MS were dried and sieved from 40 to 100 µm. 3 Ca3(PO4)2 + H2O → Ca9(PO4)5(HPO4)(OH) Equation 4.2 4.2.4. Physicochemical Characterisation of the Microspheres The morphological evaluation of the different MS was performed by field emission scanning electron microscopy (FESEM, JSM-7001F, JEOL) with a previous carbon coating. The size distribution of the MS was assessed through image analysis (FIJI, ImageJ software [28]) of optic microscopy (AE2000, Motic) acquisitions. Chapter 4 – 89 – Fourier-transform infrared spectroscopy (FTIR, Nicolet 6700, Thermo Scientific) in the attenuated total reflectance (ATR) mode was used to check the presence of the functional groups of apatite and gelatine in the different MS. Data were obtained with a spectral resolution of 4 cm-1 and averaging of 64 scans, in the 700 to 2000 cm-1 range. Moreover, phase composition of both Gel-HA MS and CDHA MS was determined by X-ray powder diffraction (XRD, D8 Advance, Bruker) using Cu Kα radiation at 40 kV and 40 mA, collecting the data with a step size of 0.02o. The counting time was 2 s per step and the acquisition was done from 20 to 40o 2θ range. The diffractograms were compared to the standard pattern of HA (ICDD PDF 09-0432) for the indexing of the peaks. In addition, calcium and phosphorus ionic exchanges in cell culture media were measured. 10 mg of MS were put in contact with 1 ml of Dulbecco’s modified eagle medium (DMEM) supplemented with 10% foetal bovine serum (FBS), 20 mM 4-(2hydroxyethyl)-1-piperazineethanesulfonic acid buffer (HEPES), 2 mM L-glutamine, 50 U ml-1 penicillin and 50 µg ml-1 streptomycin (all from Gibco), for four days. The medium was changed daily and the supernatants were collected, diluted 20-fold in 2 wt% HNO3 (69 wt% pure, PanReac AppliChem) and analysed by inductively coupled plasma optical emission spectrometry (ICP-OES, 5100, Agilent Technologies). 44Ca and 31P signals were calibrated against a multi-element standard solution (Inorganic Ventures). 4.2.5. Preparation of the Bioinks MG-63 pre-osteoblast cells (ATCC) were cultured in supplemented DMEM in a 95% humidified atmosphere containing 5% CO2 at 37 oC. Previous to cell incorporation in the bioinks, the materials and the powders were sterilised with 70% ethanol or by lowpressure plasma treatment, respectively. Three formulations were studied, depending on which type of MS was used: Gel, GelHA or CDHA. Moreover, a bioink without the addition of MS was used as a control. Additionally, two different cross-linkings were tested, resulting in a total of eight different bioinks. In all the cases, sieved MS (40 to 100 µm) were used. Sodium alginate (PanReac AppliChem) and gelatine were dissolved in DMEM at concentrations of 6 and 8 w/v %, respectively, and mixed in a 1:1 ratio. Cells (2·106 cells ml-1) and MS (60 mg ml-1) were incorporated in the inks. The components were thoroughly mixed and homogenised by hand. The mixture was immediately extruded at 30 mm s-1 through a 22 Ga nozzle directly to a 150 mM CaCl2 (Sigma-Aldrich) bath, by means of a customised 3D printer (Fundació CIM). Afterwards, the bioinks were rinsed three times with supplemented DMEM and maintained at 37 oC and 5% CO2 atmosphere. As previously mentioned, two different cross-linking protocols were investigated: (1) a strong cross-linking (sCL) consisting of immersion for 15 min in the CaCl2 bath, followed by maintenance in supplemented DMEM enriched with CaCl2 up to 5 mM and (2) a weak cross-linking (wCL) consisting of immersion for 10 min in the initial CaCl2 bath followed with maintenance in the original supplemented DMEM. Moreover, discs were prepared for rheological studies. The samples were prepared by extruding 1 ml of bioink without cells into a ø= 20 mm mould and cross-linking them using either wCL or sCL protocols. Tunning Bioinks with Microspheres – 96 – Table 4.2. Data obtained from the oscillatory amplitude sweep assay G’eq (Pa) G’’eq (Pa) Damping factor Yield stress (Pa) Flow stress (Pa) Control 268 ± 40 151 ± 13 0.5634 281 ± 30 989 ± 53 Gel MS 598 ± 197 308 ± 59 0.5150 310 ± 04 1409 ± 200 Gel-HA MS 2937 ± 645 958 ± 115 0.3261 917 ± 80 2926 ± 166 CDHA MS 332 ± 71 185 ± 21 0.5572 314 ± 10 1025 ± 128 Figure 4.2. Rheological properties of the synthesised bioinks. A Oscillatory amplitude sweep results showing the G’ and G’’ evolution over strain. B G’ recovery of the bioinks after being put under a high strain to mimic the extrusion process. C G’ of each condition after being cross-linked for 15 or 10 min corresponding to the sCL and wCL protocols, compared to the non-cross-linked samples. D G’ loss of the two cross-linkings studied after 7 and 14 days in culture. Different letters indicate statistically significant differences between time points in the same condition, different numbers indicate statistically significant differences between conditions, and different symbols indicate statistically significant differences between cross-linkings. 4.3.3. Cell Cultures on the Microspheres Cell behaviour on the MS was examined in terms of viability, proliferation, morphology and gene expression. The results after 1 day of cell culture showed that the initial cell Chapter 4 – 97 – adhesion to all MS types was similar, as observed in Figure 4.3A and quantified in 4.3C. After 3 days of cell culture, cell proliferation was promoted especially on Gel MS and Gel-HA MS. Nevertheless, cell counts were comparable between conditions at day 14 (≈3-fold of the initially seeded cells), confirming a proper cell growth in all the samples throughout the study, with a similar proliferation rate over the study for all the MS. Regarding the distribution of the cells on the MS, L/D images revealed that cells adhered to all three types of MS and had the ability to proliferate and fully colonise them. However, a non-negligible number of dead cells were detected in CDHA MS samples on days 7 and 14. The actin fibres and nuclei stainings in Figure 4.3B helped to certify the cell adhesion on the MS of all natures, as well as their organisation on the samples. Indeed, in all the conditions cells tended to attach, spread and multiply on the surface and in-between the MS. Furthermore, the expression of osteogenic genes was examined and is summarised in Figure 4.3D. In general, CDHA MS presented the most osteoinductive properties among the different MS. When compared to the control, cells on CDHA MS exhibited a 4.5-fold expression of ALP at day 3, a 2 or 2.5-fold expression of Col. I at all the time points, a 4-fold expression of OPN at day 7, and a general 2-fold expression of OCN. Moreover, Gel-HA MS also upregulated the expression of Col. I at days 3 and 14, of OPN at day 14 and of OCN at day 7 of culture, all of them with a 1.5-fold relative to the control. The contribution of the HA NPs was observed comparing the results of Gel-HA MS and Gel MS in terms of ALP expression at day 3 and the expression of OPN and OCN at 14 and 7 days, respectively, although not all the differences were statistically significant. Interestingly, RUNX2 had an important overexpression of 4-fold and 5-fold at day 14 for Gel-HA MS and CDHA MS, respectively, yet the expression in Gel MS was the highest with a 7-fold change to the control. Tunning Bioinks with Microspheres – 98 – Figure 4.3. Behaviour of MG-63 cells cultured on the different MS. A Live/dead staining images at days 1, 3, 7 and 14. Live cells stained in green and dead cells in red. Scalebar represents 500 µm. B Cell morphology pictures at 3, 7 and 14 days of culture. Cell nuclei in blue and actin fibres in red. Scalebar represents 40 µm. C Cell proliferation measured through resazurin-based method. D Gene expression of osteogenic markers for the cells seeded in the MS at days 3, 7 and 14. Control samples at day 3 were used as a reference to determine the fold changes. Different letters indicate statistically significant differences between time points in the same condition and different numbers indicate statistically significant differences between conditions at each time point. 4.3.4. Cell Viability and Cell Migration in the Bioinks In order to study the performance of the cells in the bioinks, four different formulations were prepared by mixing the alginate, the gelatine and each type of MS with the cells. A control bioink without the incorporation of MS was also included. The extrusion was Chapter 4 – 99 – done through a 22 Ga nozzle directly into a CaCl2 bath, as described in the experimental section. In addition, two different cross-linking protocols were investigated (i.e., wCL and sCL). The sCL allowed to study the final constructs up to 21 days of cell culture, while the wCL bioinks lost stability after 14 days in culture. The visualisation of the cells embedded in the bioinks was done by confocal microscopy. Representative live/dead (L/D) images in Figure 4.4A and B proved that ≥90% of the cells were viable (green fluorescent cells) and only a small number of dead cells (red fluorescent ones) were observed at 3, 7, 14 and 21 days of culture for all the conditions. In contrast, the L/D ratios immediately after the extrusion (i.e. at day 0) were slightly lower, especially for the bioinks containing Gel-HA MS and CDHA MS, independently of the cross-linking protocol applied. Therefore, the viability at day 0 was quantified for the two initial cross-linking times and compared among the four bioinks. Figure 4.4C shows that in general, the differences between the cross-linkings were minimal, demonstrating that their effects on cells at the initial time point were similar. Moreover, it was observed that the control condition presented the highest cell viability (93.0% ± 0.4), followed by the Gel MS bioink with 90.8% ± 0.9 of the cells alive. A significant decrease in the viability of the cells was found in the other two conditions, indicating that the incorporation of apatite-containing MS reduced the cell survival rate at earlier time points. Indeed, Gel-HA MS exhibited cell mortality of 21.3% ± 1.5, similar to that of CDHA MS bioinks with 18.4% ± 1.2 of dead cells. In spite of this, the initial viability rates for all the bioinks were sufficient for the cells to survive and proliferate, indicating a suitable performance for all the conditions. With regards to the evolution of the cells in the bioinks, it was observed that initially, MG-63 cells were embedded in the alginate matrix. After several days of culture, the cells close to the MS started migrating and attaching to the Gel MS and Gel-HA MS, as can be appreciated in the insets of Figure 4.4A and B. The time needed for this process highly depended on the cross-linking applied: in the bioinks with sCL the migration started at day 7 (Figure 4.4A), while the bioinks with wCL exhibited cell migration and attachment to the MS earlier, at day 3 (Figure 4.4B). Although to a lesser extent, the confocal images also proved that these phenomena occurred in the CDHA MS bioinks. It is worth noting that at the latest time points evaluated (i.e. 14 days for wCL and 21 days for sCL), the cells that remained unattached to the MS increased their size in some of the conditions (marked with arrows), especially in the control and Gel MS bioinks in the case of the wCL bioink (Figure 4.4B). This observation suggested the formation of small cell spheroids as a consequence of cell growth, which will be further analysed in the next section. Tunning Bioinks with Microspheres – 100 – Chapter 4 – 101 – Figure 4.4. Evolution of MG-63 cell viability in the different bioinks. Live/dead images of the bioinks with sCL (A) and wCL (B) at 0, 3, 7, 14 and 21 days of culture. Live cells stained in green and dead cells in red. Insets show magnification of cells attached to the different MS. Dashed circles indicate MS, and arrows cell clusters forming spheroids. Scale bars represent 500 µm for the full images and 100 µm for the insets. C Quantification of the cell viability at day 0, immediately after the bioinks extrusion. Different numbers indicate statistically significant differences between conditions for each CL and different symbols indicate statistically significant differences between crosslinkings within the same condition. 4.3.5. Cell Proliferation and Cell Morphology in the Bioinks Cell proliferation in the bioinks was evaluated through Presto Blue assay at 0, 3, 7, 14 and 21 days (Figure 4.5A and B). The results confirmed the proliferation of MG-63 in all conditions. In the case of sCL bioinks, the cell number in all the MS-containing bioinks remained unaltered for the first week and registered a 3-fold increase from day 7 to day 14, reaching a plateau that lasted until day 21. In contrast, the cell number in the control bioink exhibited a constant growth throughout the cell culture, up to 5-fold increase at day 21. Regarding the wCL bioinks, similar proliferation rates were found also for all the conditions at each time point. Cell number rose sharply from day 0 to day 3 and in all the bioinks, and afterwards cell growth essentially stopped at a 4-fold cell number for MScontaining bioinks and 3.5-fold value for the control samples. In general, the results showed that all MS-loaded bioinks supported cell proliferation, reaching a plateau of cell population earlier or later depending on the degree of cross-linking. The stronger the cross-linking the longer it took for the cells to start proliferating and to reach the final plateau. Figure 4.5C and D display the morphology of the cells inside the different bioinks at the last incubation time point, investigated using phalloidin and DAPI to stain the actin filaments and the nuclei, respectively. Overall, proper attachment and spreading were observed for the cells adhered to the MS in the Gel MS and Gel-HA MS samples (top rows of Figure 4.5C and D). In contrast, CDHA MS bioink was not able to promote cell migration and attachment to the MS. In general, the cells that did not adhere to the MS presented rounded morphologies, typical of alginate-based bioinks. It was noticeable their ability to cluster and form cell spheroids for all the conditions, as evidenced in the bottom rows of Figure 4.5C and D. Bigger spheroids were found in the pristine bioink, compared to the MS-containing formulations. Moreover, although cell morphology was comparable in the two CL conditions, it was observed that the wCL allowed the formation of larger cell colonies, as evidenced by the number of nuclei in each spheroid (bottom row of Figure 4.5C vs. bottom row of Figure 4.5D), especially noticeable in the control specimens. All these findings are in agreement with the L/D results of the previous section. Additionally, SEM images were taken for the wCL Gel MS bioink (Figure 4.5E). The prior dehydration of the sample resulted in the exposure of the embedded MS through the alginate network, showing the cells well attached and spread on the MS, in agreement with the confocal microscopy results. Tunning Bioinks with Microspheres – 102 – Figure 4.5. Cell behaviour in the developed bioinks. A Cell proliferation results of the sCL samples and B of the wCL condition along the study. Different letters indicate statistically significant differences between time points in the same condition and different numbers indicate statistically significant differences between conditions at each time point. C Cell morphology in the sCL bioinks at day 21 and D in the wCL bioinks at day 14, observed by fluorescence confocal microscopy. Top rows show cells attached to the MS inside the bioinks, bottom rows exhibit unattached cells suspended within the cell-laden ink. Nuclei in blue and actin fibres in red. Dashed circles indicate CDHA MS. Scale bars denote 50 µm. E SEM images of the wCL Gel MS bioink. Arrows indicate cells attached and spread on the MS. 4.3.6. Gene Expression of the Cells in the Bioinks Since the wCL bioinks promoted a faster cell response compared to the sCL condition, wCL specimens were chosen for the gene expression analysis of osteogenic markers. Overall, the MS-containing bioinks presented an overexpression of all the osteogenic genes compared to the control sample, irrespective of their nature. Analysing the results in more detail, they indicated that Gel MS significantly upregulated the expression of ALP, RUNX2, OSX and OPN at day 7 but most notably at 14 days of culture, as shown in Figure 4.6. In the case of Gel-HA MS and CDHA MS samples at 14 days, ALP, OSX and OPN were also overexpressed compared to the control. However, not all the differences were statistically significant. Moreover, Col. I and OCN genes presented the highest expression in the Gel-HA MS bioink for the two studied time points, with a 4- Chapter 4 – 103 – fold (at day 7) and 6-fold (at day 14) increase of Col. I gene with respect to the control, with statistically significant differences. In contrast, CDHA MS showed low expression of these genes. Figure 4.6. Gene expression of osteogenic markers for the cells in wCL bioinks at day 7 and 14 of culture. Control bioink at day 7 was used as a reference to determine the fold changes. Different numbers indicate statistically significant differences between conditions at each time point. 4.4. Discussion The development of bioinks requires a compromise between the improvement of their biological performance and their mechanical properties. Alginate-based bioinks are typically chemically modified to provide cell adhesive moieties to improve the behaviour of the embedded cells. However, their rheological properties can be detrimentally altered by these modifications. In the search for the optimal bioink, we propose the incorporation of microspheres with different bone-related functionalities to an alginate-based bioink, in order to enhance both cell attachment and differentiation while improving their mechanical performance. The initial hydrogel composition consisted of a mixture of alginate (6 wt%) and gelatine (8 wt%). On the one hand, the alginate supported cell viability, glued the MS due to its ability to gellify in the presence of Ca2+ ions and provided mechanical stability to the construct. On the other hand, the gelatine was incorporated in the formulation in views of enhancing the rheological properties and to provide also a protective environment for the cells during the extrusion and in the initial hours of culture. Subsequently, it was expected to gradually diffuse from the inks during cell culture, since no gelatine cross-linking treatment was applied. In this study, microspheres of three different compositions were successfully synthesised. Gelatine MS (Gel MS) were chosen for its biocompatibility [34], while calcium phosphate containing-MS were selected as a possible strategy to enhance the osteogenic or osteoinductive properties of the bioinks [35,36]. Their microstructure and chemical composition was verified by SEM, FTIR and XRD (Figure 4.1A, C and D). The size of the cross-linked gelatine-containing MS (i.e. Gel MS and Gel-HA MS) sieved from 40 to Tunning Bioinks with Microspheres – 104 – 100 µm experienced substantial swelling in cell culture medium due to hydration of their gelatine phase (Figure 4.1B). The results presented small variations compared to the findings of Chapter 3 due to the different cross-linking protocols applied. Gel-HA MS presented the most pronounced swelling, which can be explained by the partial screening between Ca2+ from the NPs and the carboxylic groups involved in the cross-linking of the gelatine chains. This partial screening could partly block the cross-linking between the hydrogel chains causing their swelling [37]. Importantly, the bioactive nature of the apatite containing-MS (i.e. Gel-HA MS and CDHA MS), explained by the inherent reactivity of the apatitic nanocrystals [38], was indirectly proved by the substantial changes in the calcium and phosphorus concentrations in the cell culture media when the microspheres were directly exposed to it (Figure 4.1E). Understanding the rheological properties of the bioinks is fundamental to predict their printability. In this sense, all the bioinks developed in the present work presented a viscoelastic behaviour, as observed in Figure 4.2A, which is required in direct ink writing techniques, since the material must be extrudable at high shear rates while assuring a perfect deposition and shape fidelity [39]. Moreover, it was found that the incorporation of microspheres increased both the G’eq and the yield stress that are directly related to the extrudability and shape fidelity of the bioinks. Indeed, the incorporation of fillers in hydrogels is a common strategy to improve their mechanical properties, where the resulting composite partially exhibits the mechanical features of their individual components. When a load is applied to these materials, it will be transferred from the polymer to the stiffer filler (in our case the cross-linked Gel MS and the mineralcontaining MS), which will bear part of the load, increasing the global mechanical properties [40]. We observed that the extent of these changes depended highly on the nature of each MS. Precisely, Gel MS doubled the G’eq compared to the control, Gel-HA MS presented a >10-fold increase, while for CDHA MS only a limited gain was observed. Regarding the yield stress, in Gel MS and CDHA MS remained similar to the control, while in Gel-HA MS it was 3 times higher (Table 4.2). These differences can be attributed to the marked volume gain of the Gel-HA MS in aqueous media compared to the other types of MS (Figure 4.1B), that generated a higher MS/alginate ratio for the same initial volume. Moreover, the apatite-containing MS might present higher rigidity than the Gel MS, which may contribute to the final rheological properties of the hydrogels. Although one could hypothesise that, since these MS contain Ca2+, this could be released and contribute to the cross-linking of the alginate, we consider this as highly improbable for two reasons. First, as shown in the ICP-OES results (Figure 4.1E), no Ca2+ release was recorded in DMEM, which is the medium where the alginate was dissolved to prepare the inks. Secondly, no effect was found in this direction with the CDHA MS, that also contain Ca2+. Regarding the 3-ITT results, a fast and almost complete elastic recovery was registered for all the conditions. This is generally linked to the shape fidelity of the filaments and is fundamental in the future development of 3D-printed structures, together with a proper self-supporting capacity [39]. Apart from controlling the shape fidelity during the extrusion, it was fundamental to ensure sufficient mechanical stability of the bioprinted ink when immersed in the cell culture medium. This was explored investigating two different cross-linking protocols for the alginate hydrogel, which consisted in varying the initial CaCl2 bath time and the supplementation or not of the cell culture media throughout the study (i.e. wCL consisting of 10 min CL in 150 mM CaCl2, and sCL consisting of 15 min CL with 150 mM CaCl2 Chapter 4 – 105 – and additional supplementation of up to 5 mM CaCl2 in the cell culture medium during the culture). Increasing the CL from 10 to 15 min did not have any substantial effect. The initial G’ value for both CL conditions was ≈10000 Pa, comparable to the results of crosslinked hydrogels previously reported in the literature [41–43]. This parameter, which accounts for the stiffness of the bioink and is related to the self-supporting capacity of the extruded and cross-linked filaments, did not change between the two CL (Figure 4.2C). In contrast, the Ca2+ supplementation in the culture media was crucial for the bioink stability during the subsequent days of incubation. Indeed, rheology results displayed important differences in the loss of G’ between the two CL conditions at both day 7 and 14, with variations of up to 40% in the Gel-HA MS samples at day 14 (Figure 4.2D). The loss of stiffness is associated with the hydrogel degradation, which can be explained by the exchange of divalent by monovalent cations that revert the gelation of alginate, destabilising the hydrogel network and solubilising the polymer [44,45]. For this reason, the addition of CaCl2 to the culture media permitted better stability in sCL bioinks compared to wCL ones, allowing to study the cell performance up to 21 days of culture instead of the 14 days of the wCL conditions. In order to investigate the performance of the synthesised MS, cell viability, proliferation and morphology studies with MG-63 cells seeded directly onto the MS were carried out. In addition, the expression of several osteoblastic differentiation genes was explored. For the cell culture studies MG-63 cells were used, which present an immature preosteoblastic phenotype. Being an immortalised cell line, they are an unlimited cell source and present a more reliable reproducibility than primary cells [46]. The large number of conditions that were tested in vitro, together with the high cell concentration these studies required, urged us to work with reliable and fast-growing cells. The results summarised in Figure 4.3 demonstrated that MG-63 cells were able to attach and grow on the MS, showing a good proliferation rate in all the tested conditions. They presented a similar morphology among conditions. Regarding gene expression, CDHA MS showed, overall, higher levels of all osteogenic genes at all time points. Hence, it can be stated that CDHA MS had better osteoinductive properties than the other MS. The high capacity of certain calcium phosphate materials to induce osteogenic differentiation of stem cells has been largely described [27,36], mostly related to the triggered ion exchange and specific morphological features such as their porosity, geometry and SSA. Despite Gel-HA MS also upregulated certain genes compared to the control sample, no major differences were observed between Gel MS and Gel-HA MS. A step forward in this investigation was to assess the behaviour of MG-63 cells embedded in bioinks containing the synthesised MS. Prior to the cell culture, two cross-linkings were performed, a weak and a strong cross-link to assess not only the mechanical stability of the constructs but also the capacity of the cells to migrate to the MS and proliferate. The first aspect explored was the effect of shear stress during the extrusion on cell viability. The results showed that the dispensation through the nozzle was one of the most critical steps for the cells in 3D-bioprinted inks. In fact, during extrusion-based bioprinting cells are exposed to different mechanical forces, where shear stress is the main cause of cell death [47]. Gel-HA MS and CDHA MS bioinks presented a statistically significant reduction of 20% in the cell survival rate (Figure 4.4C), revealing that the incorporation of ceramic-based MS was unfavourable for the cells at the extrusion step. Notwithstanding, it was the rheology of Gel-HA MS bioink which clearly differed from the other conditions, presenting flow stress of ≈3000 Pa compared to the other formulations where it was near 1000 Pa (Table 4.2). Since in the Gel MS condition the survival rate of the cells was high, it indicates that the shear forces applied cannot be the Tunning Bioinks with Microspheres – 112 – [54] L.G. Major, A.W. Holle, J.L. Young, M.S. Hepburn, K. Jeong, I.L. Chin, R.W. Sanderson, J.H. Jeong, Z.M. Aman, B.F. Kennedy, Y. Hwang, D.-W. Han, H.W. Park, K.-L. Guan, J.P. Spatz, Y.S. Choi, Volume Adaptation Controls Stem Cell Mechanotransduction, ACS Appl. Mater. Interfaces. 11 (2019) 45520–45530. doi:10.1021/acsami.9b19770. [55] J. Rychly, B.J. Nebe, Cell-material interaction, BioNanoMaterials. 14 (2013) 153–160. doi:10.1515/bnm-2013-0019. [56] S. Hanson, R.N. D’Souza, P. Hematti, Biomaterial–Mesenchymal Stem Cell Constructs for Immunomodulation in Composite Tissue Engineering, Tissue Eng. Part A. 20 (2014) 2162–2168. doi:10.1089/ten.tea.2013.0359. [57] A.R. Akkineni, T. Ahlfeld, A. Lode, M. Gelinsky, A versatile method for combining different biopolymers in a core/shell fashion by 3D plotting to achieve mechanically robust constructs, Biofabrication. 8 (2016) 045001. doi:10.1088/1758-5090/8/4/045001. Chapter 5 Bioink Functionalisation with Integrin-Selective Peptidomimetics Bioink Functionalisation with Integrin-Selective Peptidomimetics – 114 – BIOINK FUNCTIONALISATION WITH INTEGRINSELECTIVE PEPTIDOMIMETICS Scope Looking for alternative ways to provide bioactivity to alginate bioinks, the following chapter focuses on the functionalisation of the polymer with selective peptidomimetics. Carbodiimide chemistry is a widely used approach for the coupling of small RGD peptides to alginate, achieving an improvement of its biological properties. Moreover, the design of peptidomimetics with selective activity towards one specific receptor has been explored in the literature as a strategy to enhance the cellular response of natural peptides. α5β1 and αvβ3 integrins have been found to be crucial in the osteogenic process, and synthetic ligands have been developed to tackle them. In Chapter 5 the functionalisation of alginate with bone-related integrin selective peptidomimetics will be investigated. Furthermore, these functionalised alginates will be used for the synthesis of new cellladen inks. Chapter 5 – 115 – 5.1. Introduction Three-dimensional (3D) systems allow the creation of environments closer to the in vivo situation than traditional 2D cell cultures. In this context, 3D bioprinting has emerged as a promising technology in the bone regeneration field, offering the capacity to design 3D constructs with precise control of the spatial distribution of biomaterials, cells and bioactive molecules [1]. The majority bioprinting techniques are based on the use of cellladen inks, and their development still needs significant research to optimise both printability and cell function. Alginate is one of the most used hydrogels for the embedding of cells, due to its gentle gelation capacity, high versatility and applicability to different fabrication methods and bioprinting technologies, and the ease in the control/tuning of its mechanical and biological properties [2]. Despite its biocompatibility, however, alginate is not bioactive as it lacks cell adhesive moieties, which are crucial for the correct function of adhering cells [3]. In this regard, the coupling of the cell adhesive sequence arginine-glycineaspartic acid (RGD) by covalent attachment is a well-established approach to improve cell adhesion on alginate hydrogels and other materials [4–6]. RGD is described as the minimal cell-binding motif in fibronectin and many other proteins of the extracellular matrix (ECM), and it is recognised by different integrin subtypes that are involved in multiple biological pathways [7]. Therefore, in addition to cell-adhesive events, integrin targeting may also foster other biological processes, including cell survival and proliferation, migration and differentiation [8]. More specifically, in the bone tissue engineering field several works coupling RGD peptides to alginate hydrogels reported not only increased levels of cell adhesion and proliferation [6,9,10] but also enhanced values of differentiation of mesenchymal stem cells (MSCs) towards osteoblastic lineage [11]. These results highlight the potential of such ECM-inspired approach for the development of alginate-based cell-laden inks, as reported in a few recent studies [12,13]. Nonetheless, the use of linear RGD peptides entails a number of constraints, such as poor to moderate specificity and affinity for integrins, as well as low stability against enzymatic degradation [14,15]. To overcome these limitations, the use of synthetic peptidomimetics based on the RGD sequence has drawn increasing attention over the last few years. These molecules are designed to bind with high affinity specific integrin subtypes by mimicking the pharmacophore of their natural ligands and optimizing ligandreceptor interactions [8]. Moreover, they present high stability in serum and do not elicit immune reactions [16]. Of the whole subset of RGD-binding integrins, α5β1 (known as the fibronectin receptor) and αvβ3 (known as the vitronectin receptor) stand out as prominent receptors for bone tissue regeneration. The β1 integrin subfamily is the most widely expressed in osteoblasts and MSCs [17,18], while αvβ3 is the most abundant integrin in osteosarcoma cells [19]. Moreover, the activation of α5β1 is associated with the overexpression of osteogenic markers, the osseointegration of implants and ectopic bone formation [20–24]. In contrast, the lack of this integrin subtype is related to bone loss [25]. On the other hand, the functions of αvβ3 integrin remain controversial, but some studies attribute to this receptor a significant role in matrix mineralisation and in the osteodifferentiation of MSCs through BMP-2 induction [26,27]. Likewise, both integrins have also been associated with mechanotransduction processes [28–30], which have the potential to further regulate cellular behaviour in terms of adhesion or differentiation by providing mechanical cues to the cells. The capacity to discriminate α5β1 and αvβ3 integrin subtypes in vitro and promote integrinspecific cell adhesion with peptidomimetics has been demonstrated [31,32]. In Bioink Functionalisation with Integrin-Selective Peptidomimetics – 116 – subsequent studies, we further showed that functionalization of titanium with these ligands improved osteoblast behaviour [33], the adhesion and osteodifferentiation of MSCs [34,35] and bone formation in vivo [34]. However, these integrin-selective peptidomimetics have been mostly tested as surface modification strategies, commonly in 2D configurations, but their effects in 3D environments within a hydrogel matrix are unknown. 3D models better reproduce the physiological conditions of the tissues, and are thus biologically more relevant [36,37]. Based on these premises, we propose for the first time the covalent immobilisation of α51 and αv3 integrin-selective peptidomimetics on alginate hydrogels. The main goal of this strategy is to develop new cell-laden inks for 3D bioprinting with improved cell performance compared to standard RGD-alginate conjugated bioinks, as a first step towards the development of advanced 3D-printed bone tissue constructs. 5.2. Experimental Section 5.2.1. Alginate Functionalisation and Characterisation The custom-made α5β1and αvβ3-selective peptidomimetics were synthesised by a combination of solution and solid-phase peptide synthesis (SPPS), as previously described in the literature [31,38]. Their design was based on the RGD sequence and their structures were optimised through docking studies into the crystal structure of the receptors, as well as competitive solid-phase integrin binding assays. Moreover, a linear RGD peptide (H-GGGGRGDSP-OH, Genscript) was also included in the study as a control. The chemical structure of the three compounds is summarised in Figure 5.1A. The two peptidomimetics and the RGD peptide were covalently attached to alginate by standard carbodiimide chemistry, as described elsewhere [4,9]. Briefly, 1% (v/v) sodium alginate (PanReac AppliChem) was dissolved in MES buffer, consisting of 0.1 M 2-(Nmorpholino)ethanesulfonic acid (MES) and 0.3 M NaCl (both from Sigma-Aldrich) adjusted at pH 6.5. Subsequently, N-(3-dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC, Sigma-Aldrich) and N-hydroxysuccinimide (NHS, Sigma-Aldrich) were added to the solution to activate the carboxylic acid groups (50 mg EDC and 15 mg NHS per g of alginate). The mixture was stirred for 15 min and followed by the incorporation of either one of the peptidomimetics or the RGD peptide (4 mg per g of alginate). The reaction was allowed to proceed for 24 h at room temperature. The alginate was then dialysed for 4 days using dialysis tubing with MWCO of 3500 Da (Spectra/Por). The purified alginate was freeze-dried and stored at -20 ºC. The immobilisation of the peptidomimetics and the RGD peptide to the alginate was corroborated by proton nuclear magnetic resonance (1H NMR) spectroscopy. To this end, 5 mg of each functionalised alginate were dissolved in 700 µl D2O (99.9%) and analysed on a 400 MHz Bruker spectrometer at 298 K (Bruker Ascend 400). Deuterated dimethylformamide was used as an internal standard 5.2.2. Cell Culture Human bone marrow-derived mesenchymal stem cells (hBM-MSC, ATCC) were cultured in advanced Dulbecco’s modified eagle medium (adv. DMEM, Gibco) supplemented with 10% foetal bovine serum (FBS), 20 mM 4-(2-hydroxyethyl)-1piperazineethanesulfonic acid buffer (HEPES), 50 U ml-1 penicillin, 50 µg ml-1 Chapter 5 – 117 – streptomycin and 2 mM L-glutamine, all from Gibco. Cells were maintained at 37 ºC, in a 95% humidified atmosphere with 5% CO2 and the culture medium was changed twice a week. Cells at passages 3 to 4 were used to carry out the experiments. 5.2.3. Bioink Synthesis Three bioinks were tested, based on alginate functionalised with (1) RGD peptide; (2) α5β1-selective peptidomimetic; and (3) αvβ3-selective peptidomimetic. Pristine alginate was used as control bioink. Prior to the synthesis of the bioinks, all the materials and the alginates were sterilised with ethanol and low-pressure plasma, respectively. The conjugated alginates were dissolved in adv. DMEM at 3 wt% and manually mixed with 106 hBM-MSC. The mixtures were extruded by means of an extrusion-based 3D printer (Pastecaster BCN 3D+, Fundació CIM) at 30 mm s-1 through a 22 Ga nozzle directly to a 150 mM CaCl2 (Sigma-Aldrich) bath. After 10 min, the cell-laden inks were rinsed with adv. DMEM and left in culture for 21 days. 5.2.4. Cell Viability and Proliferation Cell viability in the hydrogels was investigated by staining the cells at different time points (i.e. 0, 3, 7, 14 and 21 days) using 3 µM calcein-AM (Santa Cruz Biotechnology) as live indicator in green fluorescence and 1.5 µM propidium iodide (PI, Sigma-Aldrich) as dead indicator in red fluorescence. The stained cells were imaged under a fluorescence confocal laser scanning microscope (LSM 800, Zeiss). In addition, cell viability was further quantified from the images obtained at each time point, with the help of ImageJ software [39], calculating the ratio between live cells over the total number of cells. Cell proliferation in the bioinks was assessed using a resazurin-based viability reagent (Presto Blue, Invitrogen), following the manufacturer’s instructions and measuring the fluorescence at 540/590 nm in a microplate reader (Synergy HTX, BioTek Instruments). Data are presented as fold change to the control at day 0. The bioinks were rinsed with phosphate-buffered saline (PBS) to remove the staining and fresh adv. DMEM was added. 5.2.5. Osteogenic Differentiation Differentiation of hBM-MSC was assessed by measuring alkaline phosphatase (ALP) activity of the cells embedded in the alginate hydrogels. On days 7, 14 and 21, the cell culture medium was removed and the bioinks were immediately frozen at -80 oC. To extract proteins from the samples, hydrogels were thawed and mammalian protein extraction reagent (M-PER, Thermo Fisher) was added and incubated for 30 min at room temperature. Afterwards, the bioink was removed and the supernatant was used for the next steps. ALP activity was quantified using SensoLyte pNPP Alkaline Phosphatase Assay Kit (Anaspec), following the manufacturer’s instructions. Briefly, the M-PER supernatants were incubated at 37 oC with the reagents of the kit. After 1 h, the reaction was stopped and the absorbance was measured at 405 nm in a microplate reader (Synergy HTX, BioTek Instruments). For each sample, ALP levels were normalised to cell number. To do so, the release of lactate dehydrogenase (LDH) from the lysates was quantified using the Cytotoxicity Detection KitPLUS (Roche). The solution was incubated for 10 min with the reagents of the kit, the reaction stopped and the absorbance read at 492 nm in a microplate reader (Synergy HTX, BioTek Instruments). Bioink Functionalisation with Integrin-Selective Peptidomimetics – 118 – In addition, cell differentiation to the osteoblastic lineage was determined by measuring gene expression of different osteogenic markers by means of reverse transcription quantitative polymerase chain reaction (RT-qPCR). Total RNA was extracted using TRIzol reagent (Invitrogen) following the manufacturer’s protocol. Briefly, bioinks were collected in Eppendorf tubes and incubated for 20 min at room temperature with 1 ml TRIzol. RNA isolation was performed by chloroform/isopropanol method. Afterwards, centrifugation at 12000 g at 4 oC for 15 min was done to precipitate the RNA. For complete isolation, RNA samples were purified using RNeasy Mini Kit columns (Qiagen). The RNA obtained was quantified spectrophotometrically with Take3 microvolume plate (BioTek Instruments) followed by cDNA synthesis using QuantiTect Reverse Transcription Kit (Qiagen). The gene expression was assessed with QuantiFast SYBR Green RT-PCR Kit (Qiagen) in a Mic qPCR Cycler (Biomolecular systems) using the primer sequences specified in Table 5.1. Negative controls were included to check the absence of contamination. Moreover, in order to ensure the specificity of the reaction, melt curve analysis was done in all runs. Relative gene expression levels were calculated using the 2-∆∆Ct method. GAPDH was used as the housekeeping gene and the data were normalised to the control bioink at day 3. Table 5.1. Primers’ sequences used for RT-qPCR Gene Primers’ sequences (Fw= forward; Rv= reverse) Glyceraldehyde 3-phosphate dehydrogenase (GAPDH) Fw: 5’-TTGCCATCAATGACCCCTTCA-3’ Rv: 5’-CGCCCCACTTGATTTTGGA-3’ Collagen type I (Col. I) Fw: 5’-AGGTCCCCCTGGAAAGAA-3’ Rv: 5’-AATCCTCGAGCACCCTGA-3’ Runt-related transcription factor 2 (RUNX2) Fw: 5’-AAATGCCTCCGCTGTTATGAA-3’ Rv: 5’-GCTCCGGCCCACAAATCT-3’ Osteopontin (OPN) Fw: 5’-AGCTGGATGACCAGAGTGCT-3’ Rv: 5’-TGAAATTCATGGCTGTGGAA-3’ Osteocalcin (OCN) Fw: 5’-ATGAGAGCCCTCACACTCCT-3’ Rv: 5’-CTTGGACACAAAGGCTGCAC-3’ 5.2.6. Statistical Analysis All data are reported as means ± standard error of the mean. Data distribution was checked with Shapiro-Wilk test and significant differences between samples were determined using one-way ANOVA with Tukey’s post hoc test. p-values <0.05 were considered statistically significant. Statistical analysis was performed using Minitab 19 software. Chapter 5 – 119 – 5.3. Results and Discussion Alginate-based bioinks represent a promising solution for tissue regeneration applications due to their biocompatibility and tuneable properties. As alginate is not able to support cell adhesion, functionalisation strategies with the cell adhesive sequence RGD have been widely studied to enhance its bioactivity. However, RGD is not integrin-selective and its biological potential is moderate. Thus, the present work aims to immobilise α51 and αv3 integrin-selective peptidomimetics to alginate, in order to improve the biological performance of the polymer. In detail, the purpose of this strategy is to endow the hydrogels with highly specific integrin-binding cues to achieve improved cell adhesion through interaction with integrins, and, eventually, obtain bioinks with enhanced osteogenic potential. 5.3.1. Alginate Bioink Characterisation Sodium alginate was functionalised either with one of the non-peptidic molecules or with the linear RGD peptide via carbodiimide chemistry. 1H NMR analysis was performed to verify the appropriate coupling of the molecules to the polymer. Figure 5.1B displays the spectra obtained of the original and the modified alginates. The signals corresponding to the typical alginate polymer chains consisting of mannuronate and guluronate blocks were detected in the four spectra (3.5 to 5 ppm). It was also possible to observe signals corresponding to coupling by-products (i.e. EDC) at 1.2, 2 and 2.8 ppm for the modified samples [40]. However, no additional signals were obtained for the anchored peptidomimetics and RGD peptide. Although NMR has been used to characterise different alginate modifications, such as the introduction of norbornene groups [6], PEG molecules [41] and certain peptides [42], the results highly depend on the sensitivity of the equipment. Therefore, the detection of small molecules such as peptide sequences can be hindered due to the large differences in concentration between alginate and the immobilised molecules. Indeed, the peptide/alginate ratio used in the present study was almost half of the one used in previous studies that were able to detect its presence by NMR [42]. The functionalised alginates were mixed with hBM-MSC and extruded through a nozzle directly to a CaCl2 bath, as depicted in Figure 5.1C. Although ionic gelation of hydrogels is associated with poor mechanical stability [43], the final cell-laden inks (Figure 5.1D) were stable and maintained their integrity over 21 days of culture. In fact, physical crosslinking is preferable for the in vivo application of bioinks because it is a reversible process that allows the progressive exchange of divalent by monovalent cations that revert the gelation of the hydrogel while eventually releasing the differentiated cells, together with their surrounding matrix [44,45]. As the loss of integrity was not evidently shown during the cell culture of the present work, this phenomenon is expected to occur in a longer term. Bioink Functionalisation with Integrin-Selective Peptidomimetics – 120 – Figure 5.1. A Chemical structure of the three compounds used to functionalise alginate: α5β1-selective and αvβ3-selective peptidomimetics, and the RGD peptide. B 1H NMR spectra of the different alginates. C Images of the printing process of the alginate bioinks. D Pictures of the resulting alginate bioinks. 5.3.2. Cell Viability and Morphology hBM-MSC were encapsulated within the alginates to form four different bioinks, i.e. a negative control with non-modified alginate, a positive control with alginate functionalised with the linear RGD peptide, and two alginate hydrogels functionalised with either the α5β1or αv3-selective non-peptidic mimetics. Cell compatibility is an essential requirement in cell-laden inks [46]. In our case, to study the cytotoxicity, cells were imaged with live/dead staining using a fluorescence confocal microscope at 0, 3, 7, 14 and 21 days. As shown in Figure 5.2A, most cells were found to be alive (in green) in all cell-laden inks at all the time points tested, which highlights the proper biocompatibility of all the hydrogels. Overall, the highest cell death (cells in red) was Chapter 5 – 121 – observed at the shortest time points (i.e. days 0 and 3) in all the conditions, which indicates that the extrusion through the nozzle is a critical step for cell viability, as previously stated in the literature [47,48]. To better quantify the differences observed in the images, cell viability percentages were analysed and graphed in Figure 5.2B. It was observed that at day 0 all the bioinks presented 70% of living cells. The viability increased throughout the study, especially for the cells in the three modified-alginate samples, which reached 90% viability at day 7 of culture. In contrast, the percentage of living cells in the control bioink remained stable at 70% for 7 days and afterwards it achieved the levels of the other bioinks, close to 95%, at days 14 and 21. Furthermore, live/dead staining showed that cells were homogeneously distributed in all cell-laden inks, exhibiting spherical morphologies and comparable sizes within all hydrogel condition and time points (Figure 5.2A). Although functionalisation with integrin binding ligands is generally associated with enhanced cell spreading and focal adhesion formation on 2D metallic and polymeric substrates [49–51], the complexity of the 3D system did not allow to visualise an enhancement in cell adhesion, neither using the two peptidomimetics, nor the linear RGD peptide. In this regard, it should be noted that, unlike 2D environments, 3D hydrogels limit cell spreading due to the dense matrix network and the fact that integrin activation is not restricted to the ventral side of the cells. Hereof, several works highlight the importance of ECM stiffness and stress relaxation as a fundamental property to understand and modify the cell-ECM interactions. 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The following conclusions were reached: - Regarding the use of hydroxyapatite (HA) nanoparticles (NPs) for cancer treatment purposes, the results demonstrated that their main mechanism of toxicity comes from their internalisation into cells, with a minor contribution coming from the interaction of the material with the cell surface. Moreover, as it is necessary to evaluate NPs internalisation and subsequent degradation in the tumour cells, the use of conventional characterisation techniques was explored. It was demonstrated that flow cytometry and transmission electron microscopy, which are frequently used to evaluate the presence of solid NPs, present important drawbacks when working with these particles, such as the overestimation of the internalised NPs due to membrane bounding, and the lack of information about the NPs degradation, respectively. Contrastingly, the use of an intracellular calcium probe revealed the presence of small vesicles filled with soluble calcium, which were further studied through cryo-soft Xray tomography. They resulted to be multivesicular bodies generated by the degradation of the internalised HA NPs. The combination of all these techniques allows a more detailed analysis of the toxicity of internalised NPs. (Chapter 2) - Apart from the interest of HA NPs as anticancer drugs, doping of these NPs with therapeutic ions is drawing great attention in the bone regeneration field as they can help activate specific cellular cascades. Therefore, composite microspheres (MS) consisting of gelatine and ion-doped HA NPs were successfully synthesised for future use in the formulation of novel bioinks. The introduction of therapeutic ions such as Mg2+, Zn2+ or Sr2+ entailed morphological changes in terms of sphericity and size distribution of the MS. Importantly, despite the complex dissolution/precipitation reactions of the MS in the cell culture medium, progressive liberation of the incorporated ions was found in all the conditions. Moreover, it was demonstrated that the extent of delivery could be easily controlled by tuning the doping concentration in the NPs. (Chapter 3) - With regards to the biological performance of the mineral loaded MS, various types of MS with different HA loadings (i.e. gelatine, gelatine containing HA NPs, and calcium-deficient HA) were seeded with MG-63 cells on them. The results proved that cells were able to attach and proliferate in all the materials. In addition, CDHA MS were found to better promote osteogenic differentiation than the other conditions. The subsequent introduction of the MS into alginate-based bioinks was studied as an approach to overcome the lack of bioactivity of this polymer. This strategy allowed tuning the rheological properties of the final constructs and provided adhesion sites to the embedded cells. Furthermore, supplementation of Ca2+ in the media proved to have a pronounced effect on the maintenance of the bioink matrix stiffness, which influenced, as well, the behaviour of the embedded cells. Although all the formulations showed perfect biocompatibility, the stiffness relaxation permitted earlier cell migration to the MS and their proliferation. Gene expression results of the cell-laden inks revealed enhanced osteogenic differentiation in the constructs where it was easier for the cells to migrate on the MS, demonstrating a high influence of direct interaction of the cells with the MS. (Chapter 4) General Conclusions – 135 – - Apart from the use of NPs, functionalisation of alginate bioinks with α5β1 and αvβ3 integrin-selective peptidomimetics was studied as an alternative approach to introduce bioactivity. This system proved to be a good strategy to endow alginate with osteogenic activity, outperforming the potential of the classical RGD sequence, the current gold standard employed in alginate modification. The results demonstrate that the two non-peptidic molecules significantly promoted the expression of both early and late osteodifferentiation markers, by MSC embedded in the constructs and did not compromise the stability and biocompatibility of the ink throughout the study. (Chapter 5) Publications and Conferences Publications and Conferences – 137 – PUBLICATIONS AND CONFERENCES Publications Y. Raymond, M. Bonany, C. Lehmann, E. Thorel, R. Benítez, J. Franch, M. Espanol, X. Solé-Martí, M-C. Manzanares, C. Canal, M-P. Ginebra. Hydrothermal processing of 3Dprinted calcium phosphate scaffolds enhances bone formation in vivo: a comparison with biomimetic treatment. Acta Biomater, (2021). https://doi.org/10.1016/j.actbio.2021.09.001 A. Barba, A. Diez-Escudero, M. Espanol, M. Bonany, J-M. Sadowska, J. Guillem-Marti, C. Öhman-Mägi, C. Persson, M-C. Manzanares, J. Franch, M-P. Ginebra. The impact of biomimicry in the design of osteoinductive bone substitutes: nanoscale matters. ACS Appl. Mater. Interfaces, 11 (9), 8818-30 (2019). https://doi.org/10.1021/acsami.8b20749 A. Diez-Escudero, M. Espanol, M. Bonany, X. Lu, C. Persson, M-P. Ginebra. Heparinization of Beta Tricalcium Phosphate: Osteo-immunomodulatory Effects. Adv. Healthc. Mater., 7 (5), 1700867 (2017). https://doi.org/10.1002/adhm.201700867 A. Barba, A. Diez-Escudero, Y. Maazouz, K. Rappe, M. Espanol, E-B. Montufar, M. Bonany, J-M. Sadowska, J. Guillem-Marti, C. Öhman-Mägi, C. Persson, M-C. Manzanares, J. Franch, M-P. Ginebra. Osteoinduction by Foamed and 3D-Printed Calcium Phosphate Scaffolds: Effect of Nanostructure and Pore Architecture. ACS Appl. Mater. Interfaces, 9 (48), 41722-36 (2017). https://doi.org/10.1021/acsami.7b14175 M. Bonany, M. Espanol, C. Mas-Moruno, M-P. Ginebra. Bioink Functionalisation with Integrin-Selective Peptidomimetics. (under preparation) M. Bonany, M. Espanol, M-P. Ginebra. Microspheres with Ion-Doped Hydroxyapatite Nanoparticles for Enhanced Bone Regeneration. (under preparation) M. Bonany, M. Espanol, L. Del Mazo, M-P. Ginebra. Microspheres Incorporation as a Powerful Strategy to Tune the Biological Performance of Bioinks. (under preparation) M. Bonany, A. J. Pérez-Berná, T. Dučić, E. Pereiro, H. Martin-Gómez, C. Mas-Moruno, S. van Rijt, M. Espanol, M-P. Ginebra. Hydroxyapatite Nanoparticles-Cell Interaction: the Fate of Membrane-Bound and Internalised Nanoparticles. (under preparation)