Tuning the biological performance of calcium phosphates through microstructural and chemical modifications
Abstract
Bone is the most transplanted tissue after blood. As pointed out by the World Health Organization, musculoskeletal diseases can potentially rise as the fourth largest cause of disability within the next years. Unfortunately, despite the natural ability of bone to self-heal it cannot bridge large bone defects without the help of a material. Still today the gold standard to restore bone function remains the use of natural bone grafts. However, they have several limitations that need to be overcome to accommodate the high demands of a global ageing population. Calcium phosphate (CaP) bone grafts have been known since the 1970s and stand as excellent synthetic candidates due to their composition, similar to the mineral phase of bone which consists of approximately 70 wt% of hydroxyapatite (HA). CaPs, and in particular HA, possess outstanding intrinsic properties such as biocompatibility, bioactivity and the ability to support bone growth. However, HA is too stable and once implanted it hardly degrades. Ideal synthetic bone grafts should integrate in the bone remodelling cycle, balancing implant resorption with its progressive replacement by new bone. This can be achieved either by modulating the material¿s physicochemical properties, or by combining the substrate with biological molecules capable of adequately orchestrating the various cells involved in the bone healing process. The present thesis seeks to explore, on the one hand, the feasibility of modulating the physicochemical properties of CaPs towards improving its degradation behavior, and, on the other hand, to investigate the potential CaP functionalization with heparin as a strategy to improve their biological performance at the various stages of bone healing: during the initial phase of inflammation, and during the stages of bone resorption and bone growth. The first part of the present thesis deals with the in vitro degradation of CaPs in a solution mimicking the osteoclastic environment, focusing specifically on the effect of some properties like porosity, specific surface area, microstructure and composition. The interrelation of all these parameters sometimes masks the relative importance of textural over compositional features, making difficult the prediction of their degradation behavior. Additionally, part of the work explores different strategies to incorporate carbonate ions in the crystal structure of HA as a route to obtain materials that more closely mimic natural bone. To further mimic the biological environment of bone, the second part of the thesis is focused on grafting heparin, a highly sulfated glycosaminoglycan present in the bone extracellular matrix, to CaPs. The affinity of heparin for growth factors (GF) makes this molecule an excellent candidate to capture endogenous GF bringing many benefits in the regulation of cell behavior. It is hypothesized that heparin, given its anti-inflammatory role, together with the known involvement in osteoblasts differentiation (bone forming cells) and osteoclastogenesis (bone resorbing cells formation) could enhance the biological performance of synthetic bone grafts. To this aim, CaPs were heparinized and their biological performance was assessed using human immune system cells, bone forming cells and bone resorbing cells, in an attempt to elucidate the synergies of both immune cells and cells of the skeletal system.
Full text
Tuning the biological performance of calcium phosphates through microstructural and chemical modifications Anna Díez Escudero 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). Aquesta reserva de drets afecta tant al resum de presentació de la tesi com als seus continguts. En la utilització o cita de parts de la tesi és obligat indicar el nom de la persona autora. ADVERTENCIA La consulta de esta tesis queda condicionada a la aceptación de las siguientes condiciones de uso: La difusión de esta tesis por medio del repositorio institucional UPCommons (http://upcommons.upc.edu/tesis) y el repositorio cooperativo TDR (http://www.tdx.cat/?localeattribute=es) ha sido autorizada por los titulares de los derechos de propiedad intelectual únicamente para usos privados enmarcados en actividades de investigación y docencia. No se autoriza su reproducción con finalidades de lucro ni su difusión y puesta a disposición desde un sitio ajeno al servicio UPCommons No se autoriza la presentación de su contenido en una ventana o marco ajeno a UPCommons (framing). Esta reserva de derechos afecta tanto al resumen de presentación de la tesis como a sus contenidos. En la utilización o cita de partes de la tesis es obligado indicar el nombre de la persona autora. WARNING On having consulted this thesis you’re accepting the following use conditions: Spreading this thesis by the institutional repository UPCommons (http://upcommons.upc.edu/tesis) and the cooperative repository TDX (http://www.tdx.cat/?localeattribute=en) has been authorized by the titular of the intellectual property rights only for private uses placed in investigation and teaching activities. Reproduction with lucrative aims is not authorized neither its spreading nor availability from a site foreign to the UPCommons service. Introducing its content in a window or frame foreign to the UPCommons service is not authorized (framing). These rights affect to the presentation summary of the thesis as well as to its contents. In the using or citation of parts of the thesis it’s obliged to indicate the name of the author.
Front cover: TRAP-Hoechst staining of osteoclasts on heparinized CDHA Back cover: Osteoclast filopodia detail on carbonated CDHA
PhD Thesis TUNING THE BIOLOGICAL PERFORMANCE OF CALCIUM PHOSPHATES THROUGH MICROSTRUCTURAL AND CHEMICAL MODIFICATIONS Doctoral Program of Material Science and Engineering Anna Díez-Escudero Supervisors: Dr. Montserrat Español Pons Prof. Maria-Pau Ginebra Molins Biomaterials, Biomechanics and Tissue engineering Group Department of Material Science and Metallurgical Engineering Universitat Politècnica de Catalunya Barcelona, 2017
A Blanca y Víctor Mucha gente pequeña, en lugares pequeños, haciendo cosas pequeñas, pueden cambiar el mundo. Eduardo Galeano
vii Table of Contents Abstract ………………………………………………………………………………………….…..... xi Resum ...…………………………………………………………………………….………….……. xiii Acknowledgements ……….……………………………………………………………………..…… xv Scope and aim of the thesis …………………………………………….………..………………..… xix Abbreviations ……………………………………………………………………………..……….… xxi CHAPTER 1 1.INTRODUCTION ............................................................................................................ 3 1.1.Bonecomposition............................................................................................................................3 1.2.Bonestructure.................................................................................................................................3 1.3.Bonebiology....................................................................................................................................4 1.3.1.Boneremodeling..................................................................................................................................5 1.3.1.1.Bonemicroenvironment:Theroleoftheextracellularmatrixinboneremodeling..................6 1.3.1.2.Inflammation..............................................................................................................................7 1.3.1.3.Osteoclastogenesis.....................................................................................................................9 1.3.1.4.Osteogenesis............................................................................................................................10 1.4.Syntheticbonegrafts.....................................................................................................................11 1.4.1.CalciumPhosphates...........................................................................................................................12 1.4.2.HightemperatureCaPs......................................................................................................................13 1.4.3.LowtemperatureCaPs.......................................................................................................................14 1.4.3.1.Calciumphosphatecements....................................................................................................15 1.5.Biologicalperformanceofsyntheticbonegrafts.............................................................................18 1.5.1.Physicalproperties.............................................................................................................................19 1.5.2.Chemicalproperties...........................................................................................................................20 1.5.3.Functionalizationwithbiomolecules.................................................................................................21 1.6.REFERENCES...................................................................................................................................24 CHAPTER 2 2.IN VITRO DEGRADATION OF CALCIUM PHOSPHATES: EFFECT OF MULTISCALE POROSITY, TEXTURAL PROPERTIES AND COMPOSITION ...... 33 2.1.Introduction...................................................................................................................................34 2.2.MaterialsandMethods..................................................................................................................35 2.2.1.Preparationofbiomimeticcalciumdeficienthydroxyapatite...........................................................35 2.2.2.Preparationofbiomimeticcarbonate‐dopedhydroxyapatite...........................................................36 2.2.3.Preparationofbiomimeticfoams......................................................................................................36 2.2.4.Preparationofsinteredhydroxyapatiteandbeta‐tricalciumphosphate..........................................36 2.2.5.PhysicochemicalCharacterization.....................................................................................................37 2.2.5.1.X‐rayDiffraction(XRD).............................................................................................................37 2.2.5.2.FTIRspectroscopy.....................................................................................................................37 2.2.5.3.Carbonatequantification..........................................................................................................38 2.2.5.4.Morphologicalanalysis.............................................................................................................38 2.2.5.5.SpecificSurfaceArea................................................................................................................38 2.2.5.6.Porosity.....................................................................................................................................38 2.2.5.7.Skeletaldensityandcompressiontests...................................................................................38 2.2.6.Accelerateddegradationstudy..........................................................................................................39 2.2.7.Statisticalanalyses.............................................................................................................................39
viii 2.3.Results...........................................................................................................................................40 2.3.1.PhysicochemicalCharacterization.....................................................................................................40 2.3.2.Accelerateddegradationstudy..........................................................................................................46 2.4.Discussion......................................................................................................................................50 2.5.Conclusions....................................................................................................................................54 2.6.References.....................................................................................................................................55 CHAPTER 3 3.CARBONATION OF LOW TEMPERATURE MACROPOROUS CALCIUM PHOSPHATES ....................................................................................................................... 61 3.1.Introduction...................................................................................................................................62 3.2.MaterialsandMethods..................................................................................................................63 3.2.1.Carbonationmethods........................................................................................................................63 3.2.2.Characterization.................................................................................................................................64 3.3.ResultsandDiscussion...................................................................................................................65 3.3.1.Carbonationlevelandtexturalpropertiesofthefoams...................................................................65 3.3.2.Compositionoffoamsaftercarbonation...........................................................................................67 3.4.Conclusions....................................................................................................................................70 3.5.Appendix.Controlledbiphasiccalciumphosphatemixtures............................................................70 3.6.References.....................................................................................................................................72 CHAPTER 4 4.HEPARINIZATION OF BETA TRICALCIUM PHOSPHATE: IMMUNOMODULATORY EFFECTS AND OSTEOGENIC POTENTIAL ................. 77 4.1.Introduction...................................................................................................................................78 4.2.MaterialsandMethods..................................................................................................................79 4.2.1.Preparationofcalciumphosphates...................................................................................................79 4.2.2.Heparinization....................................................................................................................................80 4.2.3.Materialscharacterization.................................................................................................................81 4.2.4.Cellisolation.......................................................................................................................................82 4.2.5.Chemiluminescencestudy.................................................................................................................82 4.2.6.Macrophagecellculture....................................................................................................................83 4.2.7.Mesenchymalstemcellculture.........................................................................................................84 4.2.8.Statistics.............................................................................................................................................84 4.3.Results...........................................................................................................................................84 4.3.1.Materialcharacterization...................................................................................................................84 4.3.2.Heparinimmobilization......................................................................................................................85 4.3.3.Cellculturestudies.............................................................................................................................87 4.3.3.1.Reactiveoxygenspeciesgeneratedbymonocytesandneutrophilsinvitro...........................87 4.3.3.2.Macrophagecellcultureassays................................................................................................87 4.3.3.3.Mesenchymalstemcellproliferationanddifferentiation........................................................90 4.4.Discussion......................................................................................................................................91 4.5.Conclusions....................................................................................................................................94
xv Acknowledgements A les meves directores: Pau, gràcies per aquesta oportunitat professional i personal. He pogut créixer al teu costat i he après infinitament, no només a fer bona recerca, si no també a ser crítica, a cercar més enllà i a exprimir tot el suc d’allò que feia. I tot això m’ho has traslladat també per a créixer personalment, amb el teu exemple i ajudant-me a conèixer-me’n millor. M’has obert un món d’oportunitats i t’estaré sempre agraïda. Montse, aquest viatge no hauria estat possible sense tu. M’has encomanat la teva passió i el teu compromís per la ciència, per les coses ben fetes i sobretot, el plaer pel coneixement. Ets admirable tant professionalment com personalment. Però sobretot vull agrair-te el teu costat més humà; per guiar el meu cap i per recolzar-me quan veia que tot se m’enfosquia, per tot el teu suport i la teva manera incondicional de compartir tot allò que saps. Als membres del BBT, sembla un tòpic però creeu un ambient de feina genial. Crec que ho trobaré molt a faltar sempre. A la gent d’empreses Miquel, Mònica, Sergi per tirar del carro tant amb la vostra feina. Txell, sense tu, la feina de tot el grup dubto que fos possible. També vull també agrair-te el teu exemple de responsabilitat per les coses ben fetes sempre amb un somriure. Noelia, por hacer que todo sea más fácil para las que desgravamos, por echarme un cable siempre y por todos los momentos que hemos compartido. A tots els professors, Xavier, José Maria, Elisa, Daniel, Marta, als postdocs, Cristina, Carles, Clara, Jordi, Giuseppe, pel vostre bon humor, la gran pinya que feu i el vostre compromís per la recerca. A tots aquells que han passat: Kiara, Carol, Maria, Gemma, Sara, Natalia, María Isabel, David, Yassine, Roberta, Cédric per compartir amb bon ambient les hores de feina; muy especialmente a Edgar, mi compadre, por tanta ayuda incondicional y por enseñarme tanto de forma tan sencilla: sol sale, hierba crece; als que hi són: Kanu, Elia, Mireia, Romain; a l’Albert, per haver donat importància i llum a la recerca feta, i poder discutir tan sanament; als que hi són recents: Inés, Quim, Mar, Diego, Yago, Angélica; a las Danielas, por hablarme con esas metáforas que me sacan siempre una sonrisa, i extensivament a tots els doctorands i personal del CMEM. A tots, trobo que teniu moltes qualitats humanes que fan d’aquest grup un lloc fàcil i plaent on treballar. A Mª Ángeles, por cuidarnos a todos como a sus hijos. Una persona que fa que tot funcioni literalment, Kim, per ensenyarme que un altre món és possible, des dels teus invents a les teves idees. A l’Isaac, al Pedro, al Lluís, al Casi, perquè en els minuts de fer un piti us he pogut conèixer i he pogut sempre compartir xerrades amb un somriure. Vull agrair molt especialment a la Judit, Dra. Xula, pel teu suport, per la teva bondat. A Joanna, dzubus, por aguantar mis chapas y por cuidarme tanto. A las dos, muy especialmente, habéis sido mis suplementos feligénicos durante esta tesis, porque me habéis ayudado a reír en momentos en que me parecía imposible.
xvi A la gent multi-escala, Carla, Lluís, Montse, Trifon, gràcies per la vostra disposició a ajudar-me sempre. Especialment al Trifon, por enseñarme tanto SEM con inmensa paciencia e infinitas conversaciones y alegría; i a la Montse, per tantes altres xerrades, riures i memes per afrontar el doctorat. Al equipo Mimetis, gracias, en especial a David y Yassine, por confiar en mí para esta pequeña aventura y porque me habéis ayudado a seguir al pie del cañón en las arduas horas de escritura pudiendo compaginar horas en el lab que me han dado mucha vida. I would like to thank Prof. Cecilia Persson for the opportunity of a very enriching experience at Uppsala University. Thanks for allowing my stay in Sweden, for giving me the opportunity to work in your group and very specially for teaching me how to do science significantly interesting (p<0.01). Thanks to all MiM group for immersing me in such broad intercultural experience. I would like to specially thank Viviana Lopes for her altruistic guidance with the biological testing. This Swedish experience has been really enriching in several facets. I have met old colleagues and made new ones. Thanks to Gemma and Sara, for helping me to feel a bit closer to home. Especialmente a Sara, mi cara, por cuidarme tanto y, con la ayuda de nuestras salidas por Vit D, mantener una vida social increíble. Gracias por mostrarme Uppsala, pero sobretodo, por darme luz durante el oscuro invierno sueco. Caroline and Daniel, you have been part of this Swedish family, a great part, with our dinners, cheering and philosophical talks, you have made my stay so easy, grateful and enlightening. Thanks for showing me the ‘hidden’ Swedish warmth. Natalie, Shirin, Sotiris, part of this family, thanks for keeping up my level of laughter at any time. Vorrei ringraziare in modo particolare la Dr.ssa Gabriela Ciapetti, ti sarò sempre grata per l'opportunità che mi hai dato. Sei un grande esempio, sia a livello professionale che umano, grazie a te ho imparato cosa sia la passione per la biologia. Sei stata tanto generosa con me, e mi hai sempre fatto sentire a casa. Ringrazio il Prof. Baldini e tutto il gruppo di Patofisiologia per avermi dato l'opportunità di imparare cosa significhi fare del bene all’umanità grazie alla ricerca sull’ osso e le sue malattie. Gemma, Elena, è stata una vera goia lavorare con voi, vi ringrazio per aver avuto pazienza con la mia incapacità di capire tanti concetti di biologia e per essere sempre state pronte a darmi una mano. Alle ragazze de Patofisiologia: Francesca, Fra, AnnaMaria, Silvia, Marghe, Sofia, tutte voi avete fatto del mio soggiorno un’esperienza indimenticabile. A tutto il personale del Rizzoli: purtroppo sapete che parlo troppo e, nonostante ciò, mi avete sempre ascoltato e avete riso con me, facendomi così sentire a casa. Vorrei aggiungere a questa esperienza bolognese Luna e Gianni, per essere stati il primo contatto con questa incredibile città e per aver condiviso con me l’amore per la bella Bologna e la bella vita. Bianca, per essere un vero cuore con gambe, per essere stata la mia famiglia assieme a Dana e Dario e per avermi insegnato ad essere una persona migliore. A tutti
xvii un grazie anche per avermi aiutato con l'italiano, e per condividere con me la passione per questa bella lingua. Mi familia, que me acompaña siempre, que me ayuda a sacar el velo del que a veces se cubre la vida, y mostrarme lo que verdaderamente importa con vuestras sonrisas. Porque hay dos palabras que conozco gracias a vosotros, especialmente a mis hermanos, amor incondicional. Parte de mi familia: a mis amigos. No os nombro que seguro olvido a alguien, ya me conocéis. Gracias por estar ahí pese a las intermitencias y por estar siempre dispuestos a pasar buenos momentos. Finally, you know I like to leave the best for the end. Victoria, thanks infinitely for being such a bright light. You have inspired and encouraged me to shape this thesis in my head even when I couldn’t see how. Thanks for making futile things to be noise, for stopping time when I needed, for being always there. A todos, habéis sido fuegos ardiendo con pasión durante este viaje. GRACIAS. MERCI. THANKS. GRAZIE. TACK "Un hombre del pueblo de Neguá, en la costa de Colombia, pudo subir al alto cielo. A la vuelta contó. Dijo que había contemplado desde arriba, la vida humana. Y dijo que somos un mar de fueguitos. -El mundo es eso -revelóun montón de gente, un mar de fueguitos. Cada persona brilla con luz propia entre todas las demás. No hay dos fuegos iguales. Hay fuegos grandes y fuegos chicos y fuegos de todos los colores. Hay gente de fuego sereno, que ni se entera del viento, y gente de fuego loco que llena el aire de chispas. Algunos fuegos, fuegos bobos, no alumbran ni queman; pero otros arden la vida con tanta pasión que no se puede mirarlos sin parpadear, y quien se acerca se enciende". Eduardo Galeano
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xix Scope and aim of the thesis Despite the excellent osseointegration potential of CaP bone grafts, the major drawback for most apatitic formulations is their limited resorption rates. Consequently, CaPs remain at the bony site instead of being gradually replaced by new bone. It is fundamental to develop more interactive biomaterials, capable not only of replacing bone, but also of stimulating its regeneration. The process of bone regeneration is complex and involves many cell types operating simultaneously. Among them, there are cells of the immune and skeletal system. Inflammation, which is the first barrier faced by the implant, is often overlooked. Recently, the emerging field of osteoimmunology has shed light on the importance of the cross-talk between the immune and skeletal system in bone regeneration. The present thesis focuses on boosting the limited regenerative potential of specific CaP formulations through different strategies: i) by modifying its physicochemical properties (by changing composition, microstructure, porosity, specific surface area, etc.); and ii) through grafting on the CaP surface extracellular matrix molecules, such as heparin, known for its anti-inflammatory, osteogenic, and osteoclastogenic effects. The materials are analyzed in terms of cell response, using cells of the immune and skeletal systems, to understand the contribution of these cell types in bone regeneration. The specific objectives to achieve the aforementioned goals are: Objective 1: Development of calcium phosphate bone grafts with different physicochemical properties to control the materials’ degradation behavior and improve the biomimicry of bone. Development and characterization of calcium phosphates with different physicochemical properties in terms of microstructure, porosity, specific surface area, and composition. Assessment of the relative effect of each parameter on degradation under conditions similar to the osteoclastic resorption. Assessment of different routes for the incorporation of carbonate ions into calcium deficient hydroxyapatite through biomimetic and hydrothermal synthesis. Objective 2: Functionalization of CaPs with heparin, present in the natural extracellular bone matrix to mimic cell microenvironment and the evaluation of the in vitro cell response to the modified CaPs, focusing on those cell types involved in the bone regeneration process:
xx Immobilization of heparin on the surface of CaPs by different grafting procedures (physisorption and chemisorption) and characterization of the heparinized surfaces. Study of the immunomodulatory effects of heparin grafted surfaces in contact with inflammatory primary human cells. Study of the adhesion, proliferation and differentiation of rat mesenchymal stem cells on heparin functionalized CaPs. Study the effects of heparin on the osteoclastogenesis process, and on human osteoclast adhesion, proliferation and activity. Evaluation of the resorptive activity of osteoclasts on rough biomimetic CaP surfaces by focus ion beam–scanning electron microscopy.
xxi Abbreviations ACP Amorphous calcium phosphate Mϕ Macrophage ATPase Adenosintriphosphatase MCPA Monocalcium phosphate anhydrous ALP Alkaline phosphatase MCPM Monocalcium phosphate monohydrate α-TCP Alpha-tricalcium phosphate M-CSF Macrophage colony stimulating factor APTES (3-aminopropyl)triethoxysilane MIP Mercury intrusion porosimetry ATR Attenuated total reflectance MNC Mononuclear cells BCP Biphasic calcium phosphate MMP Matrix metalloproteinase BET Brunauer-Emmett-Teller MSC Mesenchymal stem cell BMP Bone morphogenetic protein NOS Nitric oxide synthase BMU Basic multicellular unit OB Osteoblast BSP Bone sialoprotein OC Osteoclast β-TCP Beta-tricalcium phosphate OCN Osteocalcin CDHA Calcium deficient hydroxyapatite OCP Octacalcium phosphate CaP Calcium phosphate ON Osteonectin CathK Cathepsin K OPG Osteoprotegerin COLL Collagen OPN Osteopontin CPC Calcium phosphate cement PBS Phosphate buffered saline CPF Calcium phosphate foam PG Proteoglycan CSF Colony stimulating factor PGE2 Prostaglandin E2 DBM Demineralized bone matrix PMA Phorbol-12-myristate-13-acetate DCPA Dicalcium phosphate anhydrous PMNC Polymorphonuclear cells DCPD Dicalcium phosphate dihydrate PTFE Polytetrafluoroethylene DNA Deoxyribonucleic acid PTH Parathyroid hormone ECM Extracellular matrix RANK Receptor activator of nuclear factor κ B FBGC Foreign body giant cell RANKL RANK ligand FBS Fetal bovine serum RGD Arginylglycylaspartic acid FGF Fibroblast growth factor RNA Ribonucleic acid FIB Focus ion beam ROS Reactive oxygen species FTIR Fourier transform infrared spectroscopy RUNX2 Runt-related transcription factor 2 GAG Glycosaminoglycan TCPS Tissue culture polystyrene GF Growth factor TGF Transforming growth factor HA Hydroxyapatite TNF Tumor necrosis factor IGF Insulin-like growth factor SEM Scanning electron microscope IL Interleukin SSA Specific surface area IL-1ra Interleukin-1 receptor antagonist VEGF Vascular endothelial growth factor LPS Lipopolysaccharide XPS X-ray photoelectron spectroscopy TRAP Tartrate-resistant acid phosphatase XRD X-ray diffraction
1 Chapter 1
Introduction 9 13 or TGF-β restraining inflammation and initiating tissue repair.[39] This tight interrelation between the immune and the skeletal system has led to a new emerging field called osteoimmunology that seeks to understand and benefit from the tight crosstalk between both systems. As depicted, both the immune and skeletal system share several regulatory, signaling molecules and transcription factors which modulate their crosstalk. Leukocytes and macrophages release proteins which can activate for instance osteoclast maturation, such as IL-1, IL-6 and TNF (osteolytic cytokines) or RANKL[40] or promote osteogenesis by the regulation of BMP.[41] Understanding the different pathways established between both cell lineage is crucial to develop more integrating materials. An adequate inflammatory response is a requisite of successful bone repair since it is also coupled with coagulation and angiogenesis, which are also requisite for cells survival. Figure 1-5. Macrophage polarization and their function and cytokines released, adapted.[32,42] 1.3.1.3. Osteoclastogenesis Osteoclasts are responsible for bone resorption. They can degrade both the inorganic phase of bone by releasing acidic species, i.e. protons, and the organic phase through specific enzymes that degrade the organic components. Osteoclasts are multinucleated cells formed by fusion of cells of the monocytemacrophage lineage upon stimulation with proteins. These proteins are often secreted by bone marrow stromal cells or cells of the osteoblastic lineage.[43] Macrophage colony-stimulating factor (M-CSF), receptor activator of nuclear factor kappa B (RANK), RANK ligand (RANKL), and osteoprotegerin (OPG) are secreted by
Chapter 1 10 osteoblasts or stromal cells and regulate osteoclast activity in what is known as the RANK/RANKL/OPG pathway. Specifically, RANK-RANKL interactions lead to osteoclast maturation, while OPG is a decoy receptor inhibiting RANK-RANKL interactions, as shown in Figure 1-6. Other environmental stimuli can foster osteoclast formation such as PTH, prostaglandin E (PGE 2 ), vitamin D, or cytokines such as interleukin 11 (IL-11) or tumor necrosis factor alpha (TNF-α). [43–45] Figure 1-6. Osteoclast differentiation and activation, adapted.[12] As previously mentioned, osteoclasts are derived from the monocyte-macrophage lineage. Monocytes belong to the immune system and can arrive to bone after an injury, where they become macrophages which have the potential to actively phagocytize pathogens. Fusion of macrophages under the stimulation with CSF and RANKL, among others, generates mature osteoclasts. In fact, during maturation, osteoclasts resemble foreign body giant cells (FBGC). Prior to bone resorption, osteoclasts attach to bone by polarizing their membrane forming a ‘ruffled membrane’ typical from this type of cells. Once attached and sealed by the ruffled membrane, osteoclasts transport acidifying vesicles to degrade bone. Thus, bone resorption by osteoclasts takes places under acidic conditions, normally at pH around 4.5. Adenosintriphosphatase (ATPase) and carbonic anhydrase are some of the proteases mediating this acidification which degrades bone mineral, whereas cathepsin K (CathK) and MMPs degrade the organic components of bone. [46,47] 1.3.1.4. Osteogenesis There are two paths for bone formation both involving the transformation of a preexisting connective tissue (mesenchyme) into bone tissue; intramembranous ossification occurs by direct conversion from mesenchyme to bone and takes place specifically in flat bones; endochondral ossification occurs by a cartilage intermediate which is later replaced by bone. [13]
Introduction 11 The main actors during osteogenesis are osteoblasts, which are in charge of the secretion of the organic component of bone ECM and its subsequent mineralization. Osteoblasts differentiate from the mesenchyme; i.e. from mesenchymal stem cells (MSC). The microenvironment of MSC will dictate the differentiation into the bone lineage cells. Osteoblasts differentiation can be activated by different pathways, being the main regulators bone morphogenetic proteins (BMP), firstly identified by Urist in 1967. BMPs are members of the TGF-β superfamily and are known to be able to recruit immature cells and trigger their subsequent differentiation into the bone lineage. [19] Fully differentiated osteoblasts are typically protein producing cells, both collagenous and non-collagenous proteins, and are responsible also for regulating the bone matrix mineralization. During OB maturation, typical genes are expressed which are associated to different events; proliferation, ECM development and maturation, and finally, mineralization. The first stage is related to the secretion of collagen type I (COLL1), alkaline phosphatase (ALP), osteopontin (OPN), bone sialoprotein (BSP), and RUNX2 that regulates osteocalcin (OCN), which is a later marker during the mineralization stage. The osteoblasts trapped in the ECM during mineralization will later become osteocytes, as shown in Figure 1-7, by down-regulation of ALP and COLL expression. Figure 1-7. Differentiation of MSC into osteoblasts and osteocytes and the markers at each stage, adapted.[48] 1.4. Synthetic bone grafts Bone is the most commonly replaced tissue except for blood. Despite its natural selfhealing ability, large bone defects or critically sized defects cannot heal by themselves and the need of bone grafts, either natural or synthetic, to replace or regenerate the damaged tissue is fundamental. Autografts harvested from the patient are known to be the gold standard; however, their low volume availability and associated morbidity limit their application. [49,50] Cadaveric allografts, demineralized bone matrix (DBM) or interspecies xenografts are also alternatives but they have also some limitations, such as the associated immunogenic response, risk of disease transmission, as well as cultural and ethical considerations. [51] For all these reasons, the use of synthetic bone
Chapter 1 12 grafts represents a good alternative, although it is necessary to enhance their performance in order for them to be truly competitive with the autografts. Synthetic bone graft must be biocompatible, and ideally, they should undergo remodeling and support new bone formation. However, it is difficult to have synthetic bone grafts combining all this features. Calcium phosphates (CaPs), owing to their close similarity to the mineral phase are good candidates for this application. However, the in vivo performance is uneven, depending not only on the composition but also on the processing route used for the synthesis, as described in the following sections. 1.4.1. Calcium Phosphates CaPs as already mentioned, have a composition similar to that of natural bone and possesses several of the required properties for optimal bone regeneration since they can be bioactive, resorbable, osteoconductive and osteoinductive*.[52–56] However, the combination of all these properties into one material requires a tight control and a thorough understanding of the material and the biological site. Still now, this control is not easy and efficient enough. CaPs can be considered as salts of orthophosphoric acid which can be obtained through several synthesis methods. The synthesis path usually dictates the final physicochemical properties of the compound. There exist eleven known CaPs with Ca/P molar ratio within 0.5 and 2.0. Four of them are obtained by solid state reactions at high temperature, and the rest through precipitation in aqueous environment by low temperature synthesis routes. Additionally to the Ca/P molar ratio, acidity/basicity and solubility are the parameters influencing mostly their physicochemical properties,[57] as depicted from Table 1-1. In general, the lower the Ca/P molar ratio is, the more acidic and water-soluble the calcium orthophosphate becomes. * Bioactive: the ability to form a bone-like apatite layer on implant surface in the living body. Resorbable: the ability to degrade or be resorbed by surrounding tissues. Osteoconductive: the provision of a scaffold for the growth of new bone cells on implant surface Osteoinductive: the capacity of the implant to stimulate primitive, undifferentiated and pluripotent cells to develop into bone-forming lineage.
Introduction 13 Table 1-1. Existing Calcium phosphates (CaP).[57] Ca/P ratio Compound Chemical formula Solubilitya 0.5 Monocalcium phosphate monohydrate (MCPM) Ca(H2PO4)2·2H2O ~18 0.5 Monocalcium phosphate anhydrous (MCPA) Ca(H2PO4)2 ~17 1.0 Dicalcium phosphate dihydrate (DCPD) CaHPO4·2H2O ~0.088 1.0 Dicalcium phosphate anhydrous (DCPA) CaHPO4 ~0.048 1.33 Octacalcium phosphate (OCP) Ca8(HPO4)2 (PO4)4·5H2O ~0.0081 1.5 α-Tricalcium phosphate (α-TCP) α-Ca3(PO4)2 ~0.0025 1.5 β-Tricalcium phosphate (β-TCP) β-Ca3(PO4)2 ~0.0005 1.2-2.2 Amorphous calcium phosphate (ACP) CaXHY(PO4)Z·nH2O N.A. 1.5-1.67 Calcium-deficient hydroxyapatite (CDHA) Ca10-X(HPO4)X(PO4)6-X(OH)2-X ~0.0094 1.67 Hydroxyapatite (HA) Ca10(PO4)6(OH)2 ~0.0003 2.0 Tetracalcium phosphate (TTCP) Ca4(PO4)2O ~0.0007 aSolubility in g/L in water at 25°C 1.4.2. High temperature CaPs There are four of the eleven CaP salts which are obtained by solid state reaction at high temperatures. One of the most used as synthetic bone graft is tricalcium phosphate (Ca3(PO4)2), which exhibits two allotropic forms (α-TCP and β-TCP, respectively). β-TCP is obtained by solid-state reaction of calcium and phosphate precursors, for instance through the following reaction (Equation 1)[58]: 2 → (Equation 1) At temperature above 1125°C, β-TCP transforms to α-TCP, which is a metastable phase that in contact with water hydrolyses rapidly to calcium deficient hydroxyapatite (CDHA). Tetracalcium phosphate (TTCP) is the most basic salt, which possesses a higher solubility than HA. It can be obtained at temperatures above 1300°C by reaction of DCPA with calcium carbonate (CaCO3). TTCP is not stable in water and hydrolyses rapidly to HA and calcium hydroxide (Ca(OH)2).[59] HA can be obtained both by solid state reaction at high temperature, and by precipitation in aqueous solutions at low temperature. The high temperature HA is obtained by solid-state reaction of other calcium phosphates (e.g., MCPM, DCPA, DCPD, OCP) with CaO, Ca(OH)2, or CaCO3 at temperatures above 1200 ºC.[57,60] The solubility of HA is the lowest among the listed CaPs in Table 1-1. Biphasic CaPs (BCP), traditionally mixtures of HA and β-TCP, combine a stable and poorly soluble phase like HA with the more soluble β-TCP, and can also be prepared by solid state
Chapter 1 14 reaction. The properties of the resulting material lie in between each of the components. Due to the higher degradability of TCP, the reactivity of BCP increases with the increase in the TCP/HA ratio. Thus, in vivo bioresorbability of BCP can be controlled through phase composition. [61,62] The microstructure of the high temperature CaPs is dependent on the sintering temperature and time. For instance, higher sintering temperatures result in lower specific surface area and lower porosity, but improved compression resistance. [61] Figure 1-8 shows the typical sintered microstructure of a BCP scaffold. Upon sintering, atoms diffuse across the boundaries of the particles fusing them together and forming sintering necks which lead to denser and smoother pieces. [63] Biological apatites contain several ionic substitutions within the hydroxyapatite structure. Chemical modifications can be inserted in high temperature CaPs, generating different chemical behaviors, such as higher solubility or reactivity. Pietak et al. successfully incorporated silicon ions and proved its effect as an inhibitor of grain growth. [64] Several authors have focused their efforts in the production and characterization of high temperature doped CaPs with silicon, magnesium or carbonate. [65–70] Figure 1-8. SEM images of a BCP scaffold obtained at different sintering temperatures: A: 1200; B: 1150, and C: 1100ºC. Lower sintering temperature results in an increase of micropores and a decrease in the crystal size.[71] Macroporosity can also be incorporated in high temperature CaPs. Macropores are usually formed due to the release of volatile materials, i.e. the incorporation of porecreating additives, so-called porogens. The ideal porogen should be non-toxic and, in the case of sintered ceramics, should be finally volatilized upon sintering. [60] Porosity has been demonstrated to have a great effect on the resorbability of the material. [52,72] 1.4.3. Low temperature CaPs Monocalcium phosphate monohydrate (MCPM), monocalcium phosphate anhydrous (MCPA), octacalcium phosphate (OCP), dicalcium phosphate dihydrate (DCPD), dicalcium phosphate anhydrous (DCPA), amorphous calcium phosphate (ACP), and calcium deficient hydroxyapatite (CDHA) can all be produced by a precipitation
Introduction 15 reaction in aqueous solutions at low temperature. However, some of them can also be obtained through high temperature methods. The most relevant low temperature CaP for bone regeneration applications are those compounds found in the body (HA, OCP, ACP, and occasionally DCPD). MCPM and its anhydrous form MCPA (formed upon heating MCPM above 100°C), for instance, are the most acidic CaP and are not naturally found in the body. They are rarely employed as substitutes for bone replacement. Similarly, DCPD can be crystallized at low pH, pH <6.5, and transformed to the anhydrous counterpart (DCPA) at temperature above 80°C. Their minerals are called brushite and monetite, respectively. Oppositely to MCPM, DCPD is found biologically in pathological calcifications and is proposed as one of the intermediates in bone mineralization.[73] OCP is thought to be an unstable transient intermediate during precipitation of hydroxyapatite, and it is found as component in human dental and urinary calculi and plays a crucial role in in vivo formation of apatitic biominerals.[74,75] Analogously, ACP can be encountered as a transient phase during precipitation of CaPs in aqueous systems. Usually, it is described as the first phase appearing when precipitation occurs from high supersaturated solutions.[74] HA can also be synthesized at low temperature, mainly by precipitation, hydrothermal treatments or hydrolysis. By far, the most interesting low temperature CaP is HA or its analogous CDHA due to their similarity both in structure and composition to the mineral phase of bone. Low temperature CDHA is not a stoichiometric compound, but a solid solution with the following formula: Ca10−x(PO4)6−x(HPO4)x(OH)2−x, (where x is between 0 and 1). Therefore, it has a Ca/P ratio between 1.67 and 1.5. CDHA is crystal structured like hydroxyapatite, but possesses cation vacancies (Ca2+) and anion (OH–) vacancies. Some of the sites occupied solely by phosphate anions in stoichiometric hydroxyapatite are substituted by hydrogen phosphate (HPO42–) anions. It can be obtained by precipitation, with poor crystallinity and high specific surface area, which enhance their solubility compared to stoichiometric HA. Thermal decomposition of CDHA results in its transformation to β-TCP when the Ca/P ratio is 1.5, or to biphasic β-TCP/HA if the Ca/P lies between 1.5 and 1.67. 1.4.3.1. Calcium phosphate cements Among the low temperature CaPs, calcium phosphate cements (CPC) stand as a very particular type of material that has generated great interest as bone grafts. CPC were discovered by Brown and Chow in the 1980s and LeGeros et al.[77,78] CPC are formed through a mixture of a calcium orthophosphate salts with a liquid phase, normally water or an aqueous solution, and forms a paste able to harden and set through a dissolution and precipitation process in physiological conditions. The most commonly
Chapter 1 16 used CPCs are based on final products of hydroxyapatite or brushite, as shown in Figure 1-9. One of the most exciting properties of CPCs consists of their ability to form an injectable paste that is able to self-set in physiological conditions. This enables them to be handled directly by clinicians and to be injected into bone cavities with minimal invasive surgery.[5] One of the first kinds of CPCs to be commercialized was the Constanz cement,[79] which consisted of a poorly crystalline carbonated apatite (dahllite). Upon implantation, cements have shown to resorb slowly, which causes the problematic of mismatch of the resorption rates accomplished naturally in bone. Great efforts have been put towards the development of these materials in an attempt to achieve better resorption rates, similar to those of bone, coupling bone formation with implant resorption. Figure 1-9. Calcium phosphate cement formulations.[80]
Introduction 17 Since cementitious reactions involve mixing a powder with a liquid phase, this simple reaction offers a wide range of tunable properties depending on the particle size of the reactants, liquid phase composition, liquid to powder ratio (L/P), setting environment, and setting time. CPCs are intrinsically porous due to the formation of an entangled network of crystals during precipitation, resulting in high specific surface areas that enhance bioactivity. [81] The use of starting powders with small particle sizes makes them more reactive and this favors the formation of many nuclei leading to the precipitation of a higher number of smaller crystals. [82] Cements also possess nanoporosity due to the voids in between crystals. Moreover, increasing the high liquid to powder ratio of the cement formulation leads to microstructures with a higher content of micropores caused by the higher distance between particles. [83] Despite the intrinsic porous nature of CPCs –with pores in the nanoand micrometric ranges, they lack macroporosity, which is crucial for bone ingrowth and to spur resorbability. Similarly to the sintered materials, macroporosity can be easily obtained through the addition of non-toxic porogens [84] or by surfactants that allow creating and stabilizing foamed pastes. [82,85] All porosity levels achievable with CPCs are depicted in Figure 1-10. Figure 1-10. Pore entrance size distribution associated to the microstructure in macroporous CDHA CPC; a) crystal entanglement due to setting reaction, b) crystal agglomerated due to L/P ratio and c) macropores due to foaming process, adapted.[83]
Chapter 1 18 As earlier mentioned, CDHA crystal structure allows many ionic substitutions. In order to further mimic biological apatite, CDHA can be synthesized incorporating the ion substitutions present in bone, such as carbonate, magnesium, strontium, etc. Carbonate represents the major natural ionic substitution, and carbonate ions can be placed both at phosphate and hydroxyl positions in the apatite lattice, named as B-type or A-type respectively. The B-type substitution is the most frequently found in biological apatites, together with charge compensation by a calcium vacancy, and with a hydrogen atom which bonds to a neighboring phosphate.[86] Ionic substitutions have been shown not only to affect the physicochemical properties of CaP bone grafts materials, but also their biological performance which will be addressed in the following sections. 1.5. Biological performance of synthetic bone grafts As previously mentioned, bone grafts, and particularly CaPs, have the great advantage of being recognized by the body since they have similar composition to that of natural mineral bone phase. When properly designed, CaPs can actively participate in the bone remodeling process. CaPs can foster new bone formation and are also susceptible to osteoclastic degradation, which makes them excellent candidates for bone regeneration. During bone remodeling, there is a tight synchronization between bone resorption and new bone formation that needs to be taken into account in the design of CaP materials. The balance between formation and resorption is complex, and as described earlier, there are several different actors guiding the process. The modulation of CaPs biological performance has been investigated by different authors who have described different routes to enhance the biomaterials output.[87,88] It is of paramount importance to approach bone regeneration from a multidisciplinary view. It is necessary to understand the interaction of the different material properties (chemical, textural, mechanical, etc) with the surrounding cells, in order to identify the possible directions to enhance biomaterial’s performance. In addition to chemical composition, microstructural properties, grain and crystal size, microporosity, surface roughness and specific surface area have also been proposed as crucial parameters controlling the biological response of biomaterials. An important property required in bone regeneration is osteoinduction. The ability of a material to induce osteogenic cell differentiation represents a clear advantage to integrate grafts into tissue.[89] Several authors have investigated chemical and textural properties of biomaterials, both in vitro and in vivo, in order to understand the biological reactions triggered by each feature.[90-92]
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Chapter 2
2. In vitro degradation of calcium phosphates: effect of multiscale porosity, textural properties and composition Scope The resorption of biomaterials can be achieved actively or passively by means of cellular activity (so-called bioresorption) and/or by the inherent physicochemical properties of materials, respectively. The following chapter aims to describe the physicochemical properties of a wide variety of CaPs and the extent to which each parameter can modulate degradation. Multiscale porosity, SSA and composition characterization during acidic degradation can shed light on the most effective properties to modulate biomaterial degradation and resorption, which is crucial and mandatory for the regenerative potential of bone grafts. An accelerated degradation method using an acidic media resembling osteoclast resorption environment (pH<4) has been used to unravel the extent to which each property affect CaPs degradation.
Chapter 2 34 2.1. Introduction Bone has a remarkable capacity for self-repair. However, this self-healing capacity is not sufficient to bridge critical sized bone defects and moreover it can be impaired in some pathological situations. In these cases, the use of bone grafts, either natural (autografts, allografts or xenografts) or synthetic, is crucial to restore bone function.[1] An ideal bone graft should provide initial strength at the implantation site while actively supporting bone remodeling. Therefore, it is important to have a tight synchronization between graft resorption and new bone deposition to allow a gradual replacement of the bone graft by newly formed bone.[2] Although bone autografts remain the gold standard for such applications, site morbidity and volume availability are major concerns that limit their use.[3] The need for synthetic bone grafts is clear, but their choice and design remain still complex due to many considerations such as material composition, architecture, mechanical stability, degradation products, etc. that can potentially affect the remodeling process.[4] Among synthetic bone grafts,[5-7] calcium phosphates (CaP) are very interesting for remodeling purposes as they possess a close resemblance to the mineral phase of natural bone consisting of ~70 wt% of nanocrystalline hydroxyapatite (HA).[8] Owing to the close compositional resemblance of CaP to bone mineral, they can induce a biological response similar to that taking place during bone remodeling. Indeed, CaP can potentially be resorbed by osteoclastic cells and can also support bone formation under the action of osteoblastic cells.[9] However, despite the potential of CaP, there is presently no material matching the remodeling rates of natural bone. The bottleneck for most CaP formulations, including hydroxyapatite-based formulations, is their poor resorption rate, especially for sintered HA, which is one of the most commonly used materials.[10] In vivo studies have proved that even after 9 months of implantation sintered HA remained at the site with hardly any sign of resorption.[11,12] To circumvent this problem, HA has been combined with more soluble phases such as β-tricalcium phosphate (β-TCP).[13–16] Alternatively, more inherently soluble phases such as brushite/monetite have also demonstrated higher resorbability.[17,18] In addition, doping the crystal structure of HA with e.g. carbonate ions was found to enhance its resorption and to promote osteoclastic activity in vitro.[19–24] However, not only composition plays a role in degradation rates. The modification of crystallinity, grain size, specific surface area,[25,26] or the porosity content[27–29] are relevant physicochemical features which have been shown to help tailor resorption rates, hence improving the in vivo performance of such implants.[30] Despite the various works proving increased degradation in materials with lower crystallinity, high SSA, decreased grain density, and a higher degree
CaPs acidic degradation 41 Figure 2-2. XRD patterns (A) and ATR-FTIR spectra (B) of all materials studied, prepared with a L/P ratio of 0.55mL/g The skeletal density of the biomimetic compositions slightly differed. CDHA C and CDHA F resulted in 2.70±0.02g/cm 3 and 2.67±0.02g/cm 3 respectively, while the value for C-CDHA C decreased to 2.54±0.03g/cm 3 consistent with the incorporation of carbonate in the crystal structure. As expected, sintered β-TCP and HA compositions, showed the highest values of density with 3.08±0.02 and 3.18±0.04g/cm 3 respectively. TC/EA analyses of carbonate-containing samples are summarized in Table 2-1. The carbonate levels were similar for all samples regardless of their porosity, ranging from 11 to 13%wt. Control samples of CDHA C and CDHA F were included and proved no carbonation occurring from ambient carbon dioxide or from precursor chemicals. Table 2-1.Carbonate content of the different carbonated CDHA samples L/P ratio Carbonate level [mL/g] [%wt] C-CDHA C 0.35 12.25 ± 0.78 0.45 13.30 ± 1.27 0.55 12.33 ± 0.25 0.65 13.23 ± 0.04 Foam-C-CDHA C 0.55 11.33 ± 0.39 Foam-C-CDHA F 0.65 12.58 ± 0.32 CDHA C/F 0.55 0.05± 0.07
Chapter 2 42 Scanning electron micrographs of the different compositions are shown in Figure 2-3. The different biomimetic substrates showed microstructures consisting of aggregates of nano-micrometric crystals. For CDHA C , the aggregates consisted of plate-like crystals. CDHA F instead consisted of needle-like crystals due to the smaller particle size of α-TCP used for its synthesis. The introduction of carbonate resulted in a plate-like morphology of the crystals, both for the CCDHA C and C-CDHA F samples, which showed a very similar microstructure. The high temperature CaPs, β-TCP and SHA, showed the typical polyhedral grain structure of sintered samples. Figure 2-3. Scanning electron micrographs showing the microstructures of the different samples prepared with L/P ratio of 0.55mL/g (scale bar: 500nm). Micrographs taken at a lower magnification of the CDHA and β-TCP nonfoamed samples obtained at different L/P ratios and micrographs of the foamed samples revealed significant differences according to the processing methods of the materials (Figure 2-4). The increase in the L/P ratio led to more open
CaPs acidic degradation 43 structures, consistent with an increase in the separation between the starting αTCP particles, due to the higher amount of liquid. Foaming resulted in the formation of larger macropores, with pore interconnections, i.e., openings between adjacent macropores. Figure 2-4. Scanning electron micrographs at low magnification of the non-foamed CDHAC and CDHAF and β-TCPC samples obtained with different L/P ratios (0.35 and 0.65mL/g), as well as of the foamed counterparts. The scale bar in the two first columns corresponds to 5µm, and in the third column (foams) to 100µm. The specific surface area and total open porosity values for all materials are shown in Figure 2-5. Figure 2-5A displays the values for non-foamed materials and Figure 2-5B, for foamed analogues. Materials synthesized at high temperature showed low SSA values (<1m 2 /g) whilst low temperature biomimetic substrates showed higher SSA values. Specifically, CDHA F with needle-like structure possessed the higher SSA values (40m 2 /g) whereas CDHA C , consisting of bigger plate-like crystals, showed lower values (15m 2 /g). Carbonated samples (C-CDHA C ) showed values between those of CDHA C and CDHA F (27m 2 /g). Carbonated CDHA fine (C-CDHA F ) was not analyzed since porosity, microstructure and SSA was equal to the coarse counterpart. The L/P ratio did not have a significant effect on SSA values. Oppositely, L/P ratio had a great influence on porosity values. The increase in L/P ratio resulted
Chapter 2 44 in higher percentages of open porosity, significantly evident after the foaming process, as assessed by MIP. The opening structure correlated to the increase in L/P ratio was also demonstrated by the decrease in mechanical properties. Compression strength showed a linear decrease upon L/P increase, independently from composition. Sintered β-TCP exhibited the highest compression strength, whilst biomimetic substrates showed similar resistance. Figure 2-5. Specific surface area (SSA) and percentage of open porosity values for biomimetic and sintered non-foamed compositions (A) and for the corresponding foamed formulations (B). Note that for the foams two L/P were studied, 0.55 and 0.65mL/g, corresponding to coarse and fine formulations, respectively. Compression strength of the non-foamed samples for each L/P and for each different composition (C). Values corresponding to CDHAC and CDHAF were taken from reference [41]. Compression strength for the different foamed formulations (D). Interestingly, in the non-foamed compositions, the increase of L/P ratio not only resulted in an increase in the total porosity, but also shifted the pore size distributions to larger sizes, as illustrated in Figure 2-6. All biomimetic samples (CDHA C , CDHA F and C-CDHA C ) showed bimodal pore populations. Larger pores account for the pores in between aggregates (micrometric range), and the smaller pores relate to the distance between crystals (nanometric range). In contrast, sintered β-TCP samples exhibited a large peak centered around 1µm,
CaPs acidic degradation 45 and an almost total disappearance of the nanoporosity, as a consequence of nanocrystal coalescence during the sintering process. The effect of the foaming process on the pore size distribution is illustrated for biomimetic and sintered foams in Figure 2-6E and F, respectively. Biomimetic foams (Figure 2-6E) maintained the trend regarding nano-microporosities, with a superimposed macroporosity up to 100µm. Upon sintering, the macroporosity as well as the microporosity around 1µm were maintained, whereas no nanoporosity was observed as a result of the aforementioned coalescence of the nanocrystals (Figure 2-6F). Figure 2-6. Pore entrance size distributions of the different samples, as determined by MIP.
Chapter 2 46 2.3.2. Accelerated degradation study The results of the accelerated degradation study are illustrated in Figure 2-7. Figure 2-7A depicts the degradation of the different non-foamed materials depending on their L/P ratio while Figure 2-7B shows the degradation of the foamed samples (non-foamed counterparts are included as control samples). Overall, when the liquid to powder ratio increased, the degradation increased for all samples but to varying degrees depending on their composition, microstructure, and textural properties. For a given L/P, CDHA F experienced a larger weight loss than the similar CDHA C , and the degradation was more sensitive to the increase in L/P than the coarse counterpart. Surprisingly, the carbonated apatite (C-CDHA) exhibited the lowest degradation values of all samples, whereas β-TCP, despite being the most soluble phase, showed degradation values between CDHA and C-CDHA. Figure 2-7. Accelerated in vitro degradation for biomimetic and sintered samples (n=3). A) Non-foamed samples groups identified by the same superscripts are not statistically different (P > 0.05). Letters indicate differences between different L/P within the same formulation; numbers identify differences between compositions for the same L/P (p<0.05); B) Degradation for foamed samples and non-foamed counterparts. According to subscripts, coarse samples have L/P of 0.55mL/g and fine analogues 0.65mL/g. * denotes statistically significant differences (p<0.05) between samples, dense and foamed, respectively, C) pH values for L/P ratio 0.35 and 0.65mL/g.
CaPs acidic degradation 47 The incorporation of macroporosity in the foamed samples resulted in remarkable increases in degradation, compared to their non-foamed analogues (see Figure 2-7B). However, the extent of the increase was strongly dependent on the substrate. Thus, biomimetic CDHA showed the highest increase (threeand four-fold weight loss for the foamed CDHAC and CDHAF compared to their nonfoamed counterparts). In the carbonated samples, C-CDHA, the impact of the macroporosity was smaller resulting in a three-fold weight loss increase irrespective of the α-TCP particle size. In the sintered β-TCP, the impact of foaming in the degradation was again dependent on the size of the starting αTCP, being two or three-fold for β-TCPC and β-TCPF respectively. SHA foams showed the lowest degradation among all foamed materials, yielding similar resorption as non-foamed materials. Figure 2-7C shows the pH values measured during the degradation experiment. Steady values in pH were observed for all samples throughout the experiment, irrespective of sample composition and porosity. Figure 2-8 combines the degradation results obtained for each material as a function of porosity. Despite that non-foamed formulations exhibited a linear increase with L/P (differences in slope were observed depending on sample type), the introduction of interconnected macropores in the foamed formulations led to marked changes in the degradation behavior. Degradation increased exponentially (exponential growth function) on the biomimetic samples with correlation coefficients (R2) of 0.986, 0.799, and 0.855 for coarse, fine, and carbonated CDHA respectively, but remained linear for the sintered β-TCP (R2= 0.938). The samples were analyzed after degradation, to investigate possible changes in composition and morphology. The results for representative samples of each composition at a fixed L/P (0.55mL/g) are shown in Figure 2-9. XRD (Figure 2-9A) proved that no additional phases were re-precipitated. SEM images (Figure 2-9B) taken of the surface of the materials clearly demonstrated the degradation of the pristine plate and needle like crystals in the biomimetic CDHA samples, whilst images of β-TCP exhibited a less-obvious degradation only visible at the grain boundaries. The microstructure of the non-foamed samples cross-sections remained intact indicating that the acid was not penetrating the sample bulk.
Chapter 2 48 Figure 2-8. Degradation of the different samples as a function of the open porosity depicted in Figure 2-5. Insets for each type of material corresponds to dense formulations (including 4 different L/P ratios), while full graphs combine dense with the foamed specimens. C-CDHAC and β-TCP include the foamed coarse and fine specimens.
CaPs acidic degradation 49 Figure 2-9. XRD (A) and SEM (B) images for different compositions with L/P of 0.55mL/g before (solid line) and after degradation (dashed line). SEM images of the surface of the specimens for the biomimetic CDHA (coarse and fine) and β-TCP, before and after degradation (scale bar: 1µm).
Chapter 2 50 2.4. Discussion The present study analyses the degradation behavior of different calcium phosphate biomaterials. The simple in vitro model used here does not capture all the complex mechanisms occurring during the degradation process of a biomaterial once implanted in the body. Therefore, the results obtained cannot directly be translated to material performance in vivo. However, this model facilitates comparison of the sensitivity to acidic degradation of different synthetic calcium phosphate ceramics, and, most importantly, it makes easier analyzing the relevance of some textural properties, i.e. specific surface area and porosity, during the acidic degradation of synthetic calcium phosphates. The use of cementitious reactions in the formulation of biomimetic samples was vital to control the level of nanoand microporosity of the samples.[40] Indeed, the hydrolysis of α-TCP at body temperature results in the precipitation of an entangled network of crystal aggregates responsible for the presence of nanoporosity in the samples (Figure 2-4 and Figure 2-6). Increasing the L/P ratio during sample preparation gradually introduced an additional level of porosity within the micron range, due to an increase in the distance between crystal aggregates(Figure 2-4 and Figure 2-6).[40] Furthermore, the foaming process allowed introducing an interconnected network of macropores in the material (Figure 2-4 and Figure 2-6). This provided a platform to systematically assess the role of the multiscale level of porosity in the degradation behavior of calcium phosphates. On the other hand, the change in the starting α-TCP powder size from fine (F) to coarse (C) led to a change in the morphology of the nanocrystals from needles to plates -regardless of the L/P ratiowith a consequent change in SSA. The SSA was controlled by the surface area of the crystals which was readily available due to porous nature of cements (crystals precipitate forming an open entangled network) and was not affected by the L/P ratio. Modification of the cements composition was achieved either by incorporation of carbonate ions during precipitation or by a sintering process. The degradation behavior of the different materials can be analyzed in two separate blocs: the degradation of the non-foamed samples, i.e. micro/nanoporous samples, versus the degradation of the foamed formulations, i.e., macroporous samples. Although all samples were inherently porous, nonfoamed specimens possessed a maximum pore size below 3µm, whilst foamed formulations presented a superimposed macroporosity, typically above 10µm. Regarding the non-foamed CDHA materials, despite the similar porosity values of biomimetic coarse and fine samples at a fixed L/P ratio, fine samples yielded higher resorption (Figure 2-7A), thus revealing the effect of SSA (the only
CaPs acidic degradation 57 [46] L. Berzina-Cimdina, N. Borodajenko, Research of Calcium Phosphates Using Fourier Transform Infrared Spectroscopy, in: Infrared Spectrosc. – Mater. Sci. Eng. Technol., 2012: pp. 123– 148. [47] R.M. Wilson, J.C. Elliott, S.E.P. Dowker, L.M. Rodriguez-Lorenzo, Rietveld refinements and spectroscopic studies of the structure of Ca-deficient apatite., Biomaterials. 26 (2005) 1317–1327. [48] C. Rey, B. Collins, T. Goehl, I.R. Dickson, M.J. Glimcher, The carbonate environment in bone mineral: A resolution-enhanced fourier transform infrared spectroscopy study, Calcif. Tissue Int. 45 (1989) 157–164. [49] M.E. Fleet, The carbonate ion in hydroxyapatite: recent X-ray and infrared results., Front. Biosci. (Elite Ed). 5 (2013) 643–652. [50] B. Kundu, D. Basu, J.M.F. Ferreira, Development of porous HAp and b -TCP scaffolds by strach consolidation with foaming methos and drug chitosan bilayered scaffold based drug delivery system, J. Mater. Sci. Mater. Med. 21 (2010) 2955–2969. [51] S.-T. Kuo, H.-W. Wu, W.-H. Tuan, Y.-Y. Tsai, S.-F. Wang, Y. Sakka, Porous calcium sulfate ceramics with tunable degradation rate, J. Mater. Sci. Mater. Med. 23 (2012) 2437–2443. [52] D.G.A. Nelson, J.D.B. Featherstone, J.F. Duncan, T.W. Cutress, Paracrystalline Disorder of Biological and Synthetic Carbonate-substituted Apatites, J. Dent. Res. 61 (1982) 1274–1281. [53] E. Landi, G. Celotti, G. Logroscino, A. Tampieri, Carbonated hydroxyapatite as bone substitute, Ceram. Met. Interfaces 23 (2003) 2931–2937. [54] R.Z. LeGeros, M.S. Tung, Chemical Stability of Carbonateand Fluoride-Containing Apatites, Caries Res. 17 (1983) 419–429. [55] R.Z. LeGeros, Calcium phosphates in oral biology and medicine., Monogr. Oral Sci. 15 (1991) 1–201. [56] M.E. Fleet, Infrared spectra of carbonate apatites: v2-Region bands., Biomaterials 30 (2009) 1473–1481. [57] M.T. Fulmer, I.C. Ison, C.R. Hankermayer, B.R. Constantz, J. Ross, Measurements of the solubilities and dissolution rates of several hydroxyapatites., Biomaterials 23 (2002) 751–755. [58] S. V. Dorozhkin, Calcium Orthophosphates in Nature, Biology and Medicine, Materials 2 (2009) 399–498. [59] R.P. Félix Lanao, K. Sariibrahimoglu, H. Wang, J.G.C. Wolke, J.A. Jansen, S.C.G. Leeuwenburgh, Accelerated calcium phosphate cement degradation due to incorporation of gluconodelta-lactone microparticles., Tissue Eng. Part A. 20 (2014) 378–388. [60] K. Sariibrahimoglu, J. An, B.A.J.A. van Oirschot, A.W.G. Nijhuis, R.M. Eman, J. Alblas, J.G.C. Wolke, J.J.J.P. van den Beucken, S.C.G. Leeuwenburgh, J.A. Jansen, Tuning the degradation rate of calcium phosphate cements by incorporating mixtures of polylactic-co-glycolic acid microspheres and glucono-delta-lactone microparticles., Tissue Eng. Part A. 20 (2014) 2870–2882. [61] J. An, S.C.G. Leeuwenburgh, J.G.C. Wolke, J.A. Jansen, Effects of Stirring and Fluid Perfusion on the In Vitro Degradation of Calcium Phosphate Cement/PLGA Composites., Tissue Eng. Part C. Methods. 21 (2015) 1171–1177.
Chapter 3
61 3. Carbonation of low temperature macroporous calcium phosphates Scope Carbonate is the major substituting ion found in biological apatites. It has been attempted to be incorporated into synthetic bone grafts, as a strategy to further enhance their biological performance. Mimicking biological apatites composition can render several advantages towards the recognition and integration of bone grafts. The present chapter investigates different routes to dope low temperature CDHA with carbonate as a tool to enhance their biological performance. The carbonation routes have been applied to macroporous CDHA, preserving their macrostructure and controlling micro and nanostructure, separately from the carbonation content. .
Chapter 3 62 3.1. Introduction The development of biomimetic macroporous scaffolds, with a chemical composition and nanostructure close to the mineral phase of bone remains a major challenge. Biomaterials for bone regeneration need to adapt to the bone turn-over to help preserving the biological and mechanical characteristics of the bony site. This can be accomplished using macroporous scaffolds made of materials that closely mimic bone. It is widely accepted that biological apatites have a greater solubility than traditionally sintered calcium phosphate ceramics due to the nano-metric size of the crystals, their poor crystallinity and the presence of foreign ions where carbonate accounts for the major substitution (4-8%wt.).[1] Carbonate inclusion has indeed a major role in bone remodeling where bone resorbing cells, i.e. osteoclasts, have been closely associated to the selective resorption of regions of high carbonate inclusion[2] or by stimulating osteoblast activity.[3] This explains why biomimetic precipitation routes which provide nanosized crystals similar to those found in bone have been extensively investigated.[4– 7] Unfortunately, in these studies the consolidation of nanoparticles into bulk samples require a sintering step negating any benefit of the nanoparticles.[8–10] Calcium phosphate (CaPs) cements are bulk bodies formed through biomimetic routes. One typical formulation is that of α-TCP that hydrolyses into a CDHA under physiological conditions, according to equation (1). 3CaPOH O→Ca HPOPOOH (Eq. 1) The setting reaction results in an entangled network of nano/submicron crystals with an intrinsic nano and microporosity and large SSA.[11] Additionally, CaP cements can be further modulated by the addition of macroporosity.[12,13] Variations in the local ionic environment during CDHA precipitation offer a suitable platform to incorporate foreign ions into the hospitable crystal structure of CDHA.[14] Attempts to develop carbonated CPCs have been made by introducing sources of carbonate such as vaterite or calcite accomplishing high carbonation ratios.[15–20] The major side-effect of performing carbonation during the setting reaction was observed in the change of the rheological properties of the cement paste and the concomitant delay of the setting reaction. Unfortunately this requires re-adjusting the cement formulation to have setting times adequate for implantation. Del Real et al. added sodium bicarbonate in a CPC formulation as a method to create macropores due to the decomposition of CO32to CO2.[21] However, while those formulations cannot be made injectable because of the risk of embolism, they have the added drawback that both carbonate and macropore content are coupled and cannot be adjusted separately. The goal of the present chapter is to decouple carbonation from macropore formation by exploring different strategies for the carbonation of macroporous scaffolds based on
Carbonation of macroporous CDHA 63 the hydrolysis of α-TCP. Carbonation was carried out either by introducing the carbonate during the setting reaction, i.e., simultaneously to the precipitation of CDHA, or by a post-treatment of the pre-set scaffolds. In both cases the macroporous structure of the scaffold was preserved. The extent of carbonate incorporation was investigated for the different methods varying the amount of carbonate and studying its textural and chemical implications. 3.2. Materials and Methods 3.2.1. Carbonation methods Calcium phosphate foams (CPF) were prepared by mixing a powder phase consisting of α-TCP and 2wt% of precipitated HA with an aqueous solution containing 1wt% of polysorbate (Tween 80®, Sigma Aldrich, St.Louis, USA) that acted as surfactant.[22] Both phases were foamed at 7000rpm for 30s using a domestic food mixer, and the paste was transferred into 6x12mm2 moulds. Three different carbonation routes were explored (Figure 3-1): i) carbonation during setting: the α-TCP foams were immersed in a NaHCO3 solution (Sigma-Aldrich, St. Louis, USA) as setting medium, and either kept at 37 ºC for 17 days (coded as BC-S, standing for carbonation during biomimetic setting) or ii) subjected to an hydrothermal treatment by autoclaving at 121ºC and 1 bar for 30 minutes (HC-S, hydrothermal carbonation during setting). Prior to immersion in the carbonate solution the samples were left for 8h in a humid atmosphere at 37oC to gain initial cohesion; and iii) post-setting carbonation: the hydrothermal treatment described above was applied to previously set CDHA foams, immersed in a NaHCO3 solution (PS-HC, post-setting hydrothermal carbonation). Three different NaHCO3 concentrations were used (2.5, 5 and 10wt% NaHCO3). A control sample was prepared following method, using water as setting medium. Controls for BC-S and PS-HC were left 10 days for setting prior to carbonation treatment, whereas HC-S controls were autoclaved in water after 8h in humid atmosphere. After the treatment all samples were thoroughly rinsed with double distilled water to remove all dissolved salts. The schematic representation of the different carbonation methods is shown in Figure 3-1.
Chapter 3 64 Figure 3-1. Schematic description of the different carbonation methods: i) BC-S, ii) HC-S and iii) PS-HC. 3.2.2. Characterization The presence of carbonate was analyzed by ATR-FTIR (Nicolet 6700). Data was acquired in 64 scans with a resolution of 4cm -1 from 4000 to 575cm -1 with a Germanium crystal. The content of carbonate for the scaffolds processed using the highest concentration of carbonate (10wt% NaHCO 3 ) was analyzed by bulk combustion using a Thermal combustion element analyzer (Thermo EA 1108). The amount of carbonate was calculated according to the following equation (2). PercentageofCarbonate % (Eq.2) where, and are the molecular weight of carbonate and carbon, respectively, and % is the percentage of carbon given by the measurement. Samples were grinded and dried at 120°C overnight prior to analysis. Phase composition was identified by XRD (D8 Advance diffractometer, Bruker) with a Cu Kα anode (monochromated, λ=1.5406Ǻ), operated at 40kV, and 40mA. Data were acquired in 0.02° steps over the 2θ range of 10°-70° with 2s per step. For multiple phase quantification EVA software was used (Bruker). The textural properties of the foams were analyzed by scanning electron microscopy (SEM, Zeiss Neon40). Samples were coated with carbon to gain conductivity prior to imaging. The specific surface area of the scaffolds was determined by nitrogen adsorption using the BET (Brunauer-Emmett-Teller) method (ASAP Micromeritics).
Carbonation of macroporous CDHA 65 3.3. Results and Discussion 3.3.1. Carbonation level and textural properties of the foams The presence of carbonate as observed by FTIR (Figure 3-2) was highly dependent on the carbonation method. Carbonate bands specific for carbonate incorporation into CDHA were visible at 1471 and 1419cm-1, corresponding to a B-type substituted apatite. The band at 874cm-1can be assigned to both, HPO42-or to a B-type carbonate substituted apatite.[4,23,24] The remaining bands were all attributed to typical vibrational modes of phosphate in apatite (stretching at 1030cm-1, 964 and 1093cm-1 and the O-PO bending modes at 604 and 563cm-1).[25] The PS-HC resulted in lower levels of carbonate compared to the other methods, as observed from the lower intensity of the carbonate bands. This result was expected, as carbonate ions needed to diffuse into the CDHA crystal structure. Increasing the carbonate content in the solution resulted in better defined bands by ATR-FTIR as observed with the increase in intensity of the ν3 CO32vibration mode (Figure 3-2). However, the maximum content of carbonate achieved by this method was 2.5wt%, obtained after immersion in 10wt% NaHCO3 solution (Table 3-1). On the contrary, carbonation during setting, i.e. during precipitation of CDHA, either at 37oC (BC-S) or by the hydrothermal treatment (HC-S), increased the yield of carbonation (Figure 3-2). Moreover, increasing the content of carbonate in the setting medium increased the intensity of the carbonate bands in the precipitated apatite. The maximum levels of carbonate, obtained with the 10wt% NaHCO3 solution, were 12.3% for BC-S method and 13.0% for HC-S method, similar to previous works.[26,27] The following table summarizes the maximum carbonate contents (using 10%wt. NaHCO3) for each method, together with their SSA compared to control samples. Table 3-1. SSA and maximum carbonate content for the different synthesis methods using 10%wt. NaHCO3. Method CO32- [%wt.] SSA [m2/g] Control Carbonated BC-S 12.3 20 30 HC-S 13 17 32 PS-HC 2.5 20 17
Chapter 3 66 Figure 3-2. ATR-FTIR spectra for the samples obtained by the different carbonation methods, using carbonate solutions with different concentrations. One fundamental aspect shared by all carbonation methods was their ability to preserve the structural integrity of the foams as shown in the low magnification SEM images of the foams cross-sections (Figure 3-3). BC-S method had a great impact on the microstructure of the precipitated crystals. When the dissolution of α-TCP and precipitation of CDHA took place in a carbonate-containing medium, a significant decrease in the crystal size was observed. The micrometric plate-like crystals found in the foams set in water (Figure 3-3A and inset) turned into smaller plates (Figure 3-3B and inset) in presence of carbonate, thus increasing the SSA from 20m 2 /g to 30m 2 /g respectively (Table 3-1). Similarly, HC-S method in water resulted in needle-like structures, whereas fine plate-like crystals were formed when the same treatment was done in presence of carbonate (Figure 3-3C inset and Figure 3-3D inset, respectively), the SSA increasing from 17 and 32m 2 /g, respectively. Finally, no changes in the microstructure were observed for PS-HC method, i.e. in CDHA foams (Figure 3-3E inset and Figure 3-3F inset), as in this case carbonate had to diffuse in pre-existing crystals. The SSA also showed no modification, with values of 20 and 17m 2 /g for control and carbonated samples, respectively.
Carbonation of macroporous CDHA 73 [22] E.B. Montufar, T. Traykova, C. Gil, I. Harr, A. Almirall, A. Aguirre, E. Engel, J.A. Planell, M.P. Ginebra, Foamed surfactant solution as a template for self-setting injectable hydroxyapatite scaffolds for bone regeneration, Acta Biomater. 6 (2010) 876–885. [23] M.E. Fleet, Infrared spectra of carbonate apatites: v2-Region bands., Biomaterials. 30 (2009) 1473–81. [24] I.R. Gibson, W. Bonfield, Novel synthesis and characterization of an AB-type carbonatesubstituted hydroxyapatite., J. Biomed. Mater. Res. 59 (2002) 697–708. [25] S. Padilla, I. Izquierdo-Barba, M. Vallet-Regí, High Specific Surface Area in Nanometric Carbonated Hydroxyapatite, Chem. Mater. 20 (2008) 5942–5944. [26] V. Jokanović, D. Izvonar, M.D. Dramićanin, B. Jokanović, V. Zivojinović, D. Marković, B. Dacić, Hydrothermal synthesis and nanostructure of carbonated calcium hydroxyapatite., J. Mater. Sci. Mater. Med. 17 (2006) 539–46. [27] D. Marković, V. Jokanović, B. Petrović, T. Perić, B. Vukomanović, The efficacy of hydrothermally obtained carbonated hydroxyapatite in healing alveolar bone defects in rats with or without corticosteroid treatment, Vojnosanit. Pregl. 71 (n.d.) 462–466. [28] L. Galea, D. Alexeev, M. Bohner, N. Doebelin, A.R. Studart, C.G. Aneziris, T. Graule, Textured and hierarchically structured calcium phosphate ceramic blocks through hydrothermal treatment, Biomaterials. 67 (2015) 93–103. [29] T. Toyama, K. Nakashima, T. Yasue, Hydrothermal Synthesis of β-Tricalcium Phosphate from Amorphous Calcium Phosphate, J. Ceram. Soc. Japan. 110 (2002) 716–721. [30] A.M.C. Barradas, H. Yuan, C.A. Van Blitterswijk, P. Habibovic, T. Medicine, Osteoinductive Biomaterials: Current knowledge of properties, Eur. Cells Mater. 21 (2011) 407–429. [31] R.Z. LeGeros, Calcium phosphate-based osteoinductive materials., Chem. Rev. 108 (2008) 4742–53. [32] H. Yuan, Z. Yang, Y. Li, X. Zhang, J.D. De Bruijn, K. De Groot, Osteoinduction by calcium phosphate biomaterials., J. Mater. Sci. Mater. Med. 9 (1998) 723–6.
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Chapter 4
4. Heparinization of Beta Tricalcium Phosphate: Immunomodulatory effects and osteogenic potential Scope Surface functionalization has become a powerful tool to modulate the biological interactions of biomaterials. Ideally, surfaces mimicking the cellular microenvironment can help to enhance or stimulate specific cell responses of paramount importance during bone healing. Heparin is the most sulfated member of GAGs family. They are present in stem cells niche as part of the ECM and interact with several GF involved in inflammation, osteogenesis and osteoclastogenesis. The following chapter is devoted to the study of the functionalization of CaP substrates with heparin as a tool to modulate the response of inflammatory cells and mesenchymal stem cells.
Chapter 4 78 4.1. Introduction The implantation of synthetic bone grafts in a host triggers a complex cascade of events, which ideally should lead, either to the graft’s osseointegration or to its progressive resorption and replacement by new bone. The interaction of the biomaterial with different cells involved in these events is of paramount importance. However, the attention that has been paid to the different stages of this process has been uneven. Whereas most engineering approaches focus on the osteogenic potential of the material, other crucial events such as their immune response are often overlooked. The vital role of immune cells in regulating bone dynamics has been emphasized in recent years by the emerging field of osteoimmunology, which has identified the relevant role of immune cells during osteogenesis.[1,2] After implantation, biomaterials cause a foreign body reaction that can critically determine the success or failure of the material. Briefly, the immune cascade starts with the formation of a transient protein layer on the surface of the biomaterial, followed by extravasation of cells such as neutrophils, monocytes or mast cells that try to phagocytize the biomaterial. At this stage, neutrophils and monocytes release proteolytic enzymes and reactive ROS to degrade the implant. Within a few days, neutrophils undergo apoptosis while monocytes, which have adhered to the substrate, turn into macrophages. These macrophages may develop distinct dynamic phenotypes, known as M1 or M2, with pro-inflammatory or anti-inflammatory functions respectively.[3–6] M1 releases chemokines and cytokines that enhance inflammation while M2 secrete anti-inflammatory cytokines, and promote bone remodeling and repair, by regulating osteogenesis and osteoclastogenesis.[7] The polarization of macrophages can be influenced by materials properties. Thus, features such as topography,[8–11] surface chemistry,[12–17] mechanical stimuli,[18,19] or porosity[20,21] are known to determine immune cells fate. GAGs are interesting as modulating molecules for both immune and skeletal systems. GAGs are ubiquitous in tissues (e.g., as part of stem cells niche and the extracellular matrix) and are known for their potential to interact with growth factors.[22–24] GAGs assembled into proteins (PG) are highly present within injured sites. After injury, GAGs are released from their PG back-bone, become soluble and initiate the healing cascade.[25,26] GAGs can bind chemoand cytokines that alter leukocyte migration, endothelial extravasation or cytokine expression.[25,27–30] Decades ago, the antiinflammatory effect of the most sulfonated GAG, heparin, was demonstrated by Dandona et al., showing an inhibitory effect on ROS release by leucocytes.[31] In more recent studies, Zhou et al. also found the down-regulation of inflammatory cytokines such as IL-1β in presence of GAG multilayers.[32]
Immunoregulation and osteogenesis of heparinized β-TCP 79 Due to their ubiquitous nature, GAGs not only interact with the immune system, but also with bone cells.[33] Their ability to bind proteins such as GF has been used by several authors to regulate cell behavior. The degree and position of sulfate groups on GAGs has been shown to influence the affinity for BMPs,[34,35] and to foster osteoblast differentiation[36] by blocking sclerostin which is an inhibitor of Wnt signaling for osteoblast differentiation.[37] In a similar manner, GAGs can foster osteoclastogenesis by blocking OPG, which competes with RANKL, necessary for osteoclast maturation. [38–40] Overall, the studies performed using GAGs have mostly focused on their supplementation in cell cultures but little has been done using them as substrates or as surface platforms in combination with biocompatible substrates, besides their application as artificial ECM in combination with collagen or other organic molecules.[23,24,41–45] We hypothesize that the combination of bioactive, osteoconductive and inductive substrates such as calcium phosphate (CaP) materials with the regenerative potential of GAGs can result in an enhancement of their biological performance. β-TCP as a member of CaP family is similar to natural bone’s mineral phase and has been used successfully in the clinic for decades. Nevertheless, the main focus of many investigations has been their osteogenic potential, while their interaction with the immune system has often been overlooked. Chen et al. have recently made a step forward towards unraveling the tight link between osteogenesis and immunology on ceramic substrates,[16,46,47] mainly β-TCP or porous alumina. However, the studies performed were limited to the evaluation of the effect of β-TCP extracts on cells rather than involving direct cell seeding on the material. The importance of osteoimmunology in the design of bone grafts, and the involvement of GAGs both in inflammatory and osteogenic processes has encouraged the present work. Some studies have previously explored the functionalization of CaP with heparin, as a strategy for the controlled delivery of growth factors.[48–50] The focus of this study is different, as it aims at investigating the functionalization of β-TCP with heparin as a method to modulate the response of inflammatory cells (i.e. human neutrophils, monocytes and macrophages), and mesenchymal stem cells. 4.2. Materials and Methods 4.2.1. Preparation of calcium phosphates β-TCP was obtained by sintering CDHA. CDHA discs were obtained by hydrolysis of α-TCP through a self-setting reaction. α-TCP was prepared by solid state reaction of calcium hydrogen phosphate (CaHPO4, Sigma-Aldrich, St. Louis, USA) and calcium carbonate (CaCO3, Sigma-Aldrich, St. Louis, USA) mixed at a 2:1 molar ratio and
Chapter 4 80 heated for 15h at 1400°C followed by quenching in air. Subsequently, the particles obtained were milled to a final median size of 5.2µm[51] and mixed with 2wt% of precipitated hydroxyapatite (PHA, Merck KGaA, Darmstadt, Germany). This powder was then mixed for 1min in a mortar with distilled water at a liquid to powder ratio (L/P) of 0.35mL/g, resulting in a paste that was transferred into two different size moulds: 15x2 and 5x1mm2 (diameter x thickness) PTFE moulds. The discs were immersed in water at 37°C for 10 days, to allow for complete hydrolysis to CDHA. The discs were subsequently subjected to a sintering process at 1100°C for 9h to obtain the final β-TCP. 4.2.2. Heparinization Heparin immobilization on the surface of β-TCP was performed using two different routes, either physical adsorption or chemisorption, i.e. chemical covalent binding (Figure 4-1A and B, respectively). Freshly prepared heparin (BioIberica S.A.U. Barcelona, Spain) solutions in PBS (phosphate buffer saline, Thermo Fisher Scientific) were used. Physical adsorption was carried out by immersing the samples in the heparin solution for 2 hours under gentle agitation (100 rpm). Heparin covalent binding was accomplished via a 2-step functionalisation method: i) First, silanization with aminopropyltrietoxysilane (APTES, Sigma-Aldrich, St. Louis, USA) was performed using a modified procedure from previous work.[52] Discs were immersed in a solution of APTES (2% vol., i.e. 80mM) in absolute Ethanol (Panreac,AppliedChem) in the presence of distilled water (3% vol.) and left for different times (1 hour or overnight) under agitation. Afterwards, the discs were removed and sonicated in an ultrasonic bath (Ultrasons-HD, J.P. Selecta, Barcelona) in absolute ethanol solution for 5min, followed by three ethanol rinses to remove any unbound APTES; ii) The second functionalization step consisted in the covalent immobilization of heparin onto the aminated surfaces of the materials. Carbodiimide (1-(3-dimethylaminoporpyl)-3ethylcarbodiimide, EDC, Sigma-Aldrich, St. Louis, USA) and N-hydrosuccinimide (NHS, Sigma-Aldrich, St. Louis, USA) were used to couple heparin via its carboxyl groups to the amino groups of the silane. Freshly prepared solutions of heparin in PBS at different concentrations were firstly activated for 15min at pH ~6.5 in presence of EDC/NHS (100 and 150mM, respectively). Upon activation, pH was raised to 7.5 and the silanized discs were immersed in the activated heparin solution (1mL) for 2h under agitation (100rpm).
Immunoregulation and osteogenesis of heparinized β-TCP 81 Figure 4-1. Diagram of the two heparin immobilization methods. A) physisorption and B) chemisorption. The heparin saturation curves were determined on the overnight silanized surfaces by immersing the samples in activated heparin solutions with different concentrations (10, 25, 50, 75, 100, 150, and 200µg/mL) for 2h. Supernatants were collected, and the amount of heparin was determined using a colorimetric assay modified from the literature.[53] Briefly, a solution of 0.005wt% Toluidine blue (Sigma-Aldrich, St. Louis, USA) was prepared in 0.01M hydrochloric acid (HCl, Panreac AppliChem) and 0.2wt% sodium chloride (NaCl, Sigma-Aldrich, St. Louis, USA). Heparin supernatants were mixed with the Toluidine solution at a 1:1 ratio using vortex and centrifuged for 5min at 15000 relative centrifugal force (rcf). Afterwards, 200µL of the sample solutions were placed in a 96-wellplate and the absorbance was read at 631nm (Infinite 200 PRO Microplate reader, TECAN). A standard calibration curve was made, measuring solutions with known heparin concentrations up to 25µg/mL. 4.2.3. Materials characterization SSA of the discs was measured by nitrogen adsorption using the BET method (ASAP 2020 Micromeritics). The porosity and pore size distribution were obtained by MIP (Autopore IV Micromeritics). Phase composition was assessed by XRD (D8 Advance, Bruker) using a Cu Kα anode (monochromated, λ=1.5406Ǻ) operated at 40kV and 40mA. Data were collected in 0.02° steps over the 2θ range of 10°-80° with a counting time of 2 per step. Phase identification was accomplished comparing the resulting patterns to that of β-tricalcium phosphate (JCPDS 09-0169). SEM (FIB Zeiss Neon40) was used to study the microstructure of the samples. Prior to imaging, samples were coated with carbon to enhance conductivity. The microstructure of the samples was studied before and after functionalization to investigate possible morphological changes. X-ray photoelectron spectroscopy (XPS) was used to analyze surface composition of the pristine (control) and functionalized samples. Spectra were acquired in ultra-high vacuum (5 10 -9 mbar) with an XR50 Mg anode source operating at 150W and a Phoibos 150 MCD-9 detector (D8 advance, SPECS Surface Nano Analysis GmbH, Germany). Spectra were recorded at pass energy of 25eV with a step size of 0.1eV for C1s, N1s, O1s, Si2p, Ca2p, P2p, and S2p. C1s peak was used as a reference. CasaXPS software (Casa Software Ldt, UK) was
Chapter 4 82 used for the determination of atomic elemental composition applying the manufacturer set of relative sensitivity factors. 4.2.4. Cell isolation The inflammatory response was evaluated using primary human cells isolated from blood obtained from anonymous volunteer donors at the Uppsala University Hospital. Monocytes and neutrophils were isolated from buffy coats by two different separation procedures. Firstly, monocytes were isolated from blood (15mL) which were then diluted in 1xPBS at a 1:1 ratio and mixed gently by inversion. Afterwards, diluted blood (30mL) were gently poured onto Ficoll-Paque Plus (15mL) (GE Healthcare, Chicago USA) followed by centrifugation at 400rcf for 30min. The upper human plasma layer was collected and kept at 4°C for later use. The monocytes layer was pelleted and resuspended in 1xPBS (10mL) and centrifuged at 100rcf for 15min for a total of 3 washes. Next, the monocytes pellet was resuspended in PBS (2 mL) and the cells were counted by Trypan blue exclusion method (1:5) in a cytometer. Finally, cells were diluted in cell culture medium (4PBS/1RPMI-1640/100mM glucose) containing 10% of human plasma. Secondly, to extract the PMNC, the blood pellet layer obtained after centrifugation and removal of plasma and the mononuclear layer was resuspended in 20mL of 3% Dextran (Sigma) in 0.9% saline solution and incubated for 25min. The supernatant was collected and centrifuged at 250rcf for 10min. Afterwards, the supernatant was treated with 20mL of 0.2% saline for 20s, followed by 20mL more of 1.6% saline in order to lyse erythrocytes. The cell suspension was then centrifuged at 250rcf for 10min. The PMNC pellet was resuspended in PBS and cells were counted using acetic acid (6%) in a cytometer. Finally, cells were diluted in cell culture medium (4PBS/1RPMI1640/100mM glucose) with 10% of human plasma. 4.2.5. Chemiluminescence study A luminol amplified chemiluminescence assay was used to quantify the reactive oxygen species (ROS) generated by both monocytes and neutrophils under phorbol12-myristate-13-acetate (PMA, Sigma Aldrich) activation[54,55]. Prior to the analyses, the 5x1mm2 discs were sterilized in 70% ethanol for 3h, followed by three PBS rinses of 15min each. 200µL of cells were seeded on the pristine and heparinized β-TCP discs at a cellular density of 106cells/mL, followed by the addition of 100µL of luminol solution (500µM). Previously, a luminol solution was prepared by adding 1% of luminol stock solution and 0.2% horseradish peroxidase (1mg/mL, Jackson Immuno Research) as signal amplifier. The luminol stock solution was obtained by dissolving luminol (Fisher Scientific) to 50mM in 0.2M NaOH. Both cell types and positive controls (TCPS+) were activated with 1µM PMA, which was added into the cell
Immunoregulation and osteogenesis of heparinized β-TCP 89 Macrophage morphology on the different substrates was evaluated by SEM at each time point (Figure 4-6). All substrates promoted macrophage adhesion over time. Some of the cells retained the original round shape whereas others had a more elongated shape. This is relevant since macrophage morphology can be used to assess phenotype polarization. [56] Cell elongation was observed on both β-TCP and β-TCP-H substrates, whereas TCPS and TCPS+ were prone to form cell clusters. TCPS at 72h (Figure 4-6G) showed flattened cells with no filopodia (Figure 4-6G inset) which suggests apoptosis, whilst TCPS+ exhibited cell clusters (Figure 4-6H) with connecting filopodia (Figure 4-6H inset). Higher cell elongation was found on β-TCPH compared to bare β-TCP substrate, which at 72h showed still rounded cells (Figure 4-6E and F, red colored). Image analyses of macrophage morphology showed a higher elongation ratio for the cells adhered on β-TCP-H (Figure 4-6I), especially at 72h, with a 25% higher elongation rate compared to TCPS. Figure 4-6. Representative SEM images of macrophages on β-TCP at 6h (A) and 72h (E); βTCP-H at 6h (B) and 72h (F); TCPS at 6h (C) and 72h (G); TCPS+ corresponding to LPS stimulated cells at 6h (D) and 72h (H). Scale bar in images A, B, E, F: 10µm; C, D: 2µm; G, H: 20 µm, inset scale bar: 2µm. Red colored cells represent round macrophages and purple ones represent elongated ones. I) elongation ratio of cells at 6 and 72h (* denotes statistically significant differences from TCPS and TCPS+ at 72h, and # statistically significant differences between time-points).
Chapter 4 90 4.3.3.3. Mesenchymal stem cell proliferation and differentiation Cell culture with rMSCs showed a slow but sustained proliferation on all β-TCP substrates over 14days measured by lactate dehydrogenase (LDH), (Figure 4-7A). Cell proliferation studies exhibited no statistically significant differences between β-TCP and β-TCP-H. Lower adhesion on heparinized substrates was found at 6h as well as significantly lower proliferation at 7 days comparing both calcium phosphate substrates. Nevertheless, no differences were found at 14 days, where cell proliferation rates were similar, irrespective of functionalization. Both substrates promoted lower proliferation than TCPS, which increased over 7 days until a plateau was reached. The measurement of ALP activity demonstrated earlier differentiation on the heparinized substrate β-TCP-H (Figure 4-7B). In the absence of exogenous supplementation, βTCP-H induced higher differentiation than TCPS at 6 hours and 3 days. After the addition of osteogenic medium (day 4), TCPS exhibit higher differentiation rate (7 days) compared to calcium phosphate substrates, which decreased at 14 days. Figure 4-7. Cell proliferation and early differentiation of rMSCs cultured on β-TCP, β-TCP-H and TCPS as control. A) cell proliferation over time (14d) measured by LDH; data has been normalized to area and to cell number of TCPS at 6h and expressed as relative fold change. B) ALP activity over culture time. On day 4, the cell culture medium was replaced by osteogenic medium. * indicates statistically significant differences compared to control TCPS for the same time point (p<0.05) and # indicates statistically significant differences with respect to β-TCP for the same time point (p<0.05).
Immunoregulation and osteogenesis of heparinized β-TCP 91 4.4. Discussion Calcium phosphates are widely used in clinical procedures. However, although their close resemblance to the composition of bone and other advantages such as bioactivity, osteoconduction and osteoinduction make them good candidates for bone grafting, they are not able to outperform autologous grafts in highly demanding situations. The design of new strategies to foster the interaction with the host tissue during the whole process of bone remodeling may lead to the development of materials with superior performance. The bone healing cascade is orchestrated by the cells of both the skeletal and the immune system through different shared mediators such as cytokines, receptors, signaling molecules and transcription factors. Recent studies have demonstrated that some ubiquitous biomolecules that are involved in the cross-talk between the immune and skeletal system, can be used to tune osteogenesis and osteoclastogenesis to achieve proper bone regeneration rates.[1,2,16,57] In this context, the present work explores the immune and osteomodulatory effect of heparin immobilized on β-TCP substrates. Previous works have proved the potential of heparin applied on model glass surfaces as an anti-inflammatory molecule.[32] Moreover, heparin is known have great affinity to growth factors, and this has been exploited with collagen, polyacrylamide, chitosan and other organic substrates as a way to stimulate bone regeneration.[24,42,44,54] The high affinity of heparin for growth factors is highly relevant in the case of β-TCP because, even though numerous proteins have high binding affinity to calcium phosphates via amino acids with phosphateor carboxylate-terminated side chains, most growth factors lack a specific mineral-binding domain.[43] The functionalization with heparin can help to circumvent this limitation. However, studies pertaining to the heparinization of calcium phosphates are scarce, and limited to its use as an anchoring molecule for the immobilization of growth factors on the surface of the biomaterial, for its subsequent delivery.[43,49] Heparin can bind to CaP by physical adsorption or by covalent attachment, as demonstrated in other works with HA.[49] Physical adsorption of heparin takes place through electrostatic interaction between the charged sulfate/carboxylate groups in heparin with exposed Ca ions on the crystal surface of the CaP. Covalent binding is typically accomplished through a silanization step to covalently graft amine groups that can react with the carboxylate groups from heparin through amide bond formation.[49,66] Heparin binding through covalent bonds is more stable than physically adsorbed heparin. It has been reported that in complex environments like body fluids, the presence of other charged molecules can compete for binding, thus displacing electrostatically bound heparin from the surface of the CaP.[58] The results obtained in
Chapter 4 92 the present work for β-TCP show that a covalent binding yielded higher amounts of heparin than mere physical adsorption. Figure 4-3A shows increased heparin grafting densities following covalent immobilization than physical adsorption, and this was further improved with increased silanization time. The maximum amount of heparin that was adsorbed on the surface of β-TCP was around 25µg, but this dose could be tuned depending on the concentration of heparin used during binding (Figure 4-3B). To assess the osteogenic and immunomodulation ability of heparin on β-TCP, covalent attachment was preferred for stability reasons as well as to test heparin’s ability to recruit endogenous GF which can help modulate cell function.[59,60] During the early inflammation process, immune cells can signal the following biological cascades. Neutrophils and monocytes interacting with materials release H2O2, which act as antiinflammatory signaling molecule. H2O2 can cause cell dysfunction and tissue injury depending on the concentration.[61] High concentrations of H2O2 and other ROS are deleterious to cells; optimal concentration is required to properly modulate the inflammation cascade by stimulating the balance between pro and anti-inflammatory cytokines. In the present work, primary human cells were used to assess the inflammatory response of the materials as they reflect better the in vivo scenario upon biomaterial implantation. Lower hydrogen peroxide levels were detected after contact with heparinized β-TCP surfaces compared to the non-heparinized ones (Figure 4-4). The release kinetics were different for the two cell types; whereas neutrophils exhibited a clear monomodal pattern, monocytes showed an acute release within the first 5min followed by a second raise after 80min. Nevertheless, β-TCP-H (Figure 4-4, blank squares) still displayed the lowest percentage of release. The role of GAGs on inflammation has been previously studied by local administration.[25,30,62] Campo et al. found a down-regulation of ROS and inflammatory cytokines in LPS-induced mouse chondrocytes by several GAGs.[30] The highest inhibitory effect was exerted by chondroitin sulfate and heparan sulfate supplementation. Nevertheless, the effect of GAGs has been mainly studied as a supplement in cell culture. Although in the present work the reduction of ROS is low, probably due to low heparin concentration, the trend observed for the heparinized β-TCP surfaces confirm the anti-oxidative activity of heparin. Activated macrophages secrete cytokines, which also modulate the inflammatory response depending on the interaction with the biomaterial and the implant environment.[63] Mantovani et al. extensively described the cytokine system derived from macrophage polarization[3,4] and highlighted the dynamism of this polarization. Even though for decades M1 pro-inflammatory macrophages were seen as mediators of fibrous encapsulation, nowadays it is widely accepted that both M1 and M2 subtypes promote degradation of implants and healing.[64] In the present work, proinflammatory (TNF-α, IL-1β), pro-wound healing (IL-1β) and anti-inflammatory
Immunoregulation and osteogenesis of heparinized β-TCP 93 cytokines (IL-10) were investigated using primary human cells, which although entailing a certain degree of variability between donors, mimic more closely the in vivo scenario. The cytokine release profiles (Figure 4-5) showed similar trends to those of ROS release. TCPS+ exhibited the highest levels of cytokine release upon LPS stimulation. Neither β-TCP nor β-TCP-H caused cells to secrete more cytokines than LPS-activated TCPS (TCPS+). Similarly to previous works with β-TCP particles,[65,66] β-TCP substrates induced higher release of TNF-α and IL-1β than TCPS. Functionalization with heparin of these substrates resulted in a reduction of TNF-α and IL-1β release, which was below the levels of TCPS in all three donors for IL-1β, and in two donors for TNF-α. IL-10, an anti-inflammatory cytokine, was not detected by ELISA in any of the substrates. Overall, the reduction in the inflammatory response caused by heparin on the surface of the β-TCP substrate was in agreement with previous results obtained in GAG multilayer substrates by Zhou et al.[32] Besides the cytokines release profile, cell shape can provide information on macrophage phenotype which can also be supported by immunostaining techniques for specific markers on macrophages. Sridharan et al. concluded that macrophage elongation can be associated to a pro-healing phenotype (M2), whereas proinflammatory macrophages (M1) present more rounded shapes.[63] Similarly, McWorther et al. found elongated shapes being compatible with an expression of M2 phenotype on patterned TCPS.[56] SEM images showed early elongation on β-TCP substrates at 6h (Figure 4-6A and B), which continued up to 72h, more pronounced on heparinized β-TCP-H (Figure 4-6E and F). Image analyses of cell morphology revealed significantly higher elongation ratio for β-TCP-H at 72h (Figure 4-6I), which can be related to M2 polarization, in agreement to the lower cytokine release (Figure 4-5). Osteogenic cell proliferation and differentiation are required to continue the healing process. Some studies have determined that the binding affinity between GAGs and osteogenic markers such as osteocalcin, osteonectin or osterix, is necessary for osteoblast maturation.[67,68] However, these studies took a direct stimulation approach by supplementing GAGs into the cell media or by attaching them on TCPS. The use of osteoconductive substrates such as CaPs in combination with GAGs can further enhance their biological response. β-TCP and TCPS showed similar rMSCs adhesion, while a slightly lower cell adhesion was found for heparinized β-TCP-H, even no significant differences were depicted from statistics analyses (Figure 4-7A). Noteworthy, early differentiation was stimulated by the presence of heparin as there was increased ALP activity compared to pristine β-TCP (Figure 4-7B). Although after osteogenic stimulation (7 and 14 days) ALP expression was higher on TCPS, this might be influenced by cell confluence, as denoted by the plateau reached in cell proliferation for TCPS (Figure 4-7B). Additionally, ALP expression occurred at
Chapter 4 94 shorter times on CaPs substrates, but at lower levels even the osteogenic stimulation. Hence the platform proposed here presents a novel step forward in the design of osteoimmunomodulatory biomaterials taking advantage of the intrinsic properties of CaPs together with immobilized GAGs on their surface. 4.5. Conclusions Heparin was covalently attached on β-TCP substrates by a silanization process and achieved significantly higher grafting densities compared to physisorption. The presence of heparin clearly influenced the response of both immune and osteogenic cells. Heparinization decreased the levels of hydrogen peroxide and inflammatory cytokines released from human immune cells in contact with the material. Cells elongation ratio was found higher on heparinized β-TCP-H, which denote a M2 or prohealing phenotype correlated to the down-regulation of pro-inflammatory cytokines. In addition, heparinized substrates resulted in higher ALP expression in rat mesenchymal stem cells after 3 days of culture in the absence of osteogenic medium, compared to non-heparinized substrates. This may indicate that heparinized surfaces enhance osteogenic differentiation, but this would need to be validated in further studies.
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Inflammation and osteoclastogenesis of heparinized CDHA 105 5.2.3. Inflammatory response Monocytes and PMN were isolated in two steps from human buffy coats from volunteer donors at Uppsala University Hospital similar to previously reported. [31] Briefly, for monocytes isolation, 15mL of human blood diluted in PBS (1:1) were layered onto Ficoll-Paque Plus (GE Healthcare) and monocytes were isolated by density gradient centrifugation. The upper plasma layer was collected and kept at 4ºC for later use. Mononuclear ring was collected and washed 3 times with PBS. Cells were resuspended in PBS and counted by exclusion method using Trypan blue (1:5) in a cytometer. PMN underwent a second separation step using 20mL of 3% Dextran in 0.9% saline solution. After 25 minutes of incubation, supernatant was centrifuged and treated with 0.2% saline solution over 20 seconds and 1.6% saline solution to remove erythrocytes. Cell suspension was further centrifuged and resuspended in PBS. A solution containing 6%wt. acetic acid was used to count cells in a cytometer. Luminol amplified chemiluminescence was used to study the inflammatory response of cells in contact with 5x1mm2 discs by monitoring the release of reactive oxygen species over a time of 2 hours. Cells were activated with 1µM phorbol-12-myristate-13-acetate (PMA, Sigma Aldrich) in a luminol solution as signal amplifier. A negative control consisting of TCPS was used without PMA stimulation. Luminol solution was prepared from stock (50mM of luminol in 0.2 M NaOH) with 1% luminol stock and 0.2% horseradish peroxidase (1mg/mL) and activated with 1µM phorbol-12-myristate-13-acetate (PMA, Sigma Aldrich). Prior to luminol kinetic study, 5x1mm2 discs were placed in opaque white 96 well-plates (Perkin Elmer) and sterilized with 70% ethanol for 3 hours, followed by thrice rinsing with PBS for 15 minutes each prior to cell seeding. Both cell types were diluted in cell culture media (4PBS/1RPMI-1640/100Mm glucose) at a density of 1·106cells/mL. For the plasma containing studies, 10% of human plasma was added to the cell medium. 200µL of cells were added to opaque 96 well-plates (Perkin Elmer) followed by the addition of 100µL of Luminol solution (500µM). Luminescence was monitored at 37ºC every 2 min over 2h of total kinetic study as previously described.[32] Triplicates for each material were used (n=3) and the experiment was repeated twice with independent donor buffy coats. 5.2.4. Osteoclast cell cultures Peripheral blood MNC (PBMC), containing the precursors of human osteoclasts, were isolated from buffy coats of healthy voluntary blood donors to the National Blood Transfusion Service. Donation was anonymous, and institutional review board (IRB) approval was not required. Density centrifugation with Ficoll-
Chapter 5 106 Histopaque gradient (Sigma–Aldrich) was used to isolate the mononuclear cells as described in previous work.[33] Cells were resuspended in Dulbecco’s Modified Eagle’s Medium–High Glucose (DMEM, Euroclone, Milan, Italy) supplemented with 10% fetal bovine serum (FBS, Euroclone) and 1% Penicillin/Streptomycin (complete DMEM). CDHA, CDHA-H, β-TCP and TCPS were sterilized using 70% ethanol for 3h, followed by three rinses with PBS for 15min each. Complete medium was added after the sterilization and kept overnight prior to cell seeding to minimize ionic exchange as preconditioning treatment. Mononuclear cells were resuspended in complete DMEM and seeded on discs and TCPS at a density of 6·106 cells/cm2 and 3·106 cells/cm2, respectively. After 2h of incubation, non-adherent cells were gently removed and complete DMEM supplemented with 25% RANKL-containing human osteoblast supernatant (differentiation medium) was added to induce the differentiation of osteoclast precursors (OCP).[34,35] Cell cultures were performed up to 28 days and differentiation medium was refreshed every 3 days. 5.2.4.1. Osteoclast morphology Osteoclast morphology at 14, 21 and 28 days of culture was investigated using Scanning Electron Microscopy (SEM, FIB Zeiss Neon40). At each time point, cells were rinsed with PBS and fixed using 2%paraformaldehyde/2% glutaraldehyde in 0.1M cacodylate buffer overnight at 4ºC. Afterwards, samples were dehydrated in ethanol series and dried. A thin carbon layer was sputtered prior to imaging to impart conductivity. Tartrate-resistant acid phosphatase and Hoechst staining (TRAP-Hoechst) at 14, 21 and 28 days was used to study OC morphology and differentiation. At each time point, cells were fixed using 3% paraformaldehyde-2% sucrose for 30 min. Afterwards, cell membrane was permeabilized with Triton 0.5% in HEPES for 5 min at room temperature, and finally stained using naphthol AS-BI phosphoric acid and tartrate solution (Acid Phosphatase kit, Sigma–Aldrich) for 60 min at 37ºC for the cytoplasm and with 2.25µg/mL of Hoechst 33258 (Sigma–Aldrich) for 10 min in the dark to for the cell nuclei. TRAP-Hoechst stained samples were analyzed by optical fluorescence microscopy at 360nm excitation and 470nm emission. ImageJ software was used to perform a semi-quantitative analysis of the multinucleated cells found by optical microscopy. Duplicates of 5x1mm2 material discs were used for the staining study and duplicates of plastic chamber slide were used as controls.
Inflammation and osteoclastogenesis of heparinized CDHA 107 5.2.4.2. Osteoclast activity In order to assess OC differentiation, tartrate-resistant acid phosphatase (TRAP) activity was quantified both in cells and culture supernatants at the end of the experiment (28 days). Briefly, cells grown on material substrates were rinsed in PBS and lysed with 0.1% Triton X-100/1M NaCl (Sigma-Aldrich). 50μl of the cell lysate were transferred to a new plate and 50μl 4-nitrophenyl phosphate (4.61 mg/ml)/40mM Na-tartrate/50mM Na-acetate (pH 4.8) were added. After 1 hour incubation at 37 ºC, the reaction was stopped with 50μl of sodium hydroxide (NaOH, 0.2M), and the absorbance was measured at 405nm in a spectrophotometer (Infinite F200 Pro Microplate Reader; TECAN, Mannedorf, Switzerland). Bicinchoninic acid colorimetric assay (BCA, Pierce BCA Protein Assay Kit, Rockford, USA) was used to quantify the total protein in lysates according to manufacturer’s instructions. Material substrates without cells (blank) were measured and subtracted from cell-seeded materials. TRAP activity was then normalized by the total protein content after blank subtraction. Tartrate-resistant acid phosphatase isoform 5b (TRAP5b) activity was investigated after 28 days of culture. Cell supernatants were collected, centrifuged at 400g for 5 min and assayed using the BoneTRAP® Assay kit (Pantecs.r.l., Torino, Italy), according to the manufacturer's instructions. The concentration of TRAP5b protein in the supernatants was determined by reading the absorbance at 405nm in a spectrophotometer. Data are expressed as mean concentration (units/L) after subtraction of the corresponding cell-free readout (blank). Calcium release and pH were monitored over the experimental time at 0, 3, 7, 14, 21 and 28 days. At each time point supernatants were collected, centrifuged at 2800rpm for 5 minutes and kept at 4ºC until measurement. Calcium release was measured using a colorimetric method based on ortho-cresolphthalein complexone (OCPC, Sigma-Aldrich) adapted from previous works[36,37] and the absorbance was measured at 570nm using a UV-vis spectrophotometer (Infinite M200 Pro Microplate Reader; TECAN, Manndorf, Switzerland). The values of pH were measured using a pH-meter (MultiMeter MM 41). Triplicates of 15x2mm2 material discs were used for all experiments. 5.2.5. Statistics Data were obtained using triplicates for each experiment. Data are represented as mean value ± standard error of the mean. Data groups for TRAP measurements and multinucleated cells counting were normally distributed (homogeneous
Chapter 5 108 variance, Levene>0.05) and were analyzed using one way ANOVA and post-hoc Tukey test was used to assess significant differences. The differences in the number of nuclei were not normally distributed and non-parametric KruskalWallis was used to assess significant differences. Values of p<0.05 were considered as significant. All statistics were performed using IBM® SPSS® Statistics 24 software. 5.3. Results 5.3.1. Physicochemical properties of CaP substrates Phase composition as measured by XRD (Figure 5-1A) showed that CDHA consisted of a poorly crystalline hydroxyapatite phase (JCPDS 09-0432), as indicated by the broad peaks. A small amount of unreacted α-TCP was detected (JCPDS 09-0348), amounting a 2wt%, as quantified by EVA Software. β-TCP samples were phase-pure and showed the sharp peaks typical of crystalline sintered materials (JCPDS 09-0169). According to MIP analysis (Fig. 1B), CDHA exhibited a broad pore size distribution ranging from 10 to 100nm. In contrast, β-TCP specimens presented larger pores, centered around 1µm, although they exhibited also some nanoporosity. The total porosity was similar for both substrates (Figure 5-1C). The microstructure of the two ceramic substrates was clearly different. Whereas CDHA consisted of an entangled network of plate-like crystals, β-TCP presented a smooth polyhedral grain microstructure (Figure 5-1C,a and b), which resulted in very different values of SSA, i.e. 24m2/g for CDHA and 0.4 m2/g for β-TCP. Finally, the results of the heparin adsorption isotherms are displayed in Figure 5-1D. The maximum covalent attachment yield was achieved with 500µg/mL, which allowed for the immobilization of approximately 200µg of heparin on the material surface.
Inflammation and osteoclastogenesis of heparinized CDHA 109 Figure 5-1. Physicochemical properties of CaP substrates. A: XRD patterns of CDHA and β-TCP; B: pore entrance size distribution; C: Specific surface area (SSA) and porosity values with representative SEM microstructures for CDHA and β-TCP (a and b, respectively, scale bar: 2µm); D: heparin adsorption isotherms on CDHA after 2h incubation in heparin solutions. 5.3.2. Inflammatory response The release of ROS by PMN and monocytes upon contact with the surface of the different materials activated with PMA is depicted in Figure 5-2 and Figure 5-3. The inflammatory response of PMN with and without plasma is shown in Figure 5-2A and B, respectively. CDHA exhibited the lowest ROS release among all substrates either with or without plasma (black and grey curves, respectively), followed by CDHA-H and β-TCP. All PMA-stimulated substrates showed a peak within the first 5 minutes, shorter times than those presented on Chapter 4 (Figure 4-4A), which could be related to the handling of PMA activation and measurement. Plasma containing studies showed a slight increase in ROS signal compared to the absence of plasma. No release of ROS was observed in the negative control (TCPS) without PMA activation.
Chapter 5 110 Figure 5-2. Kinetics of ROS release by neutrophils in contact with substrates over 2h. A: ROS release in presence of 10% human plasma, and B: ROS release without plasma. ROS kinetic release upon contact with monocytes in absence of plasma showed higher values than PMN Figure 5-3A, however ROS peaks were shifted to longer time (20min). TCPS activated (TCPS+) showed a bimodal kinetic, with two ROS peaks at 20 and 50min.Biomimetic substrates CDHA and CDHA-H exhibited the lowest ROS signals, approximately 60% and 30%, in the presence and absence of plasma respectively, showing a mean peak at 20min in all substrates (Figure 5-3A and B). The presence of plasma increased the signal of ROS for both CDHA and CDHA-H, which was though diminished after 1h to half of its value (Figure 5-3A). No effect of plasma was found for β-TCP which yielded approximately 70-80% of ROS for both absence and presence of plasma.
Inflammation and osteoclastogenesis of heparinized CDHA 111 Figure 5-3. ROS release kinetic of monocytes in contact with substrates over 2h. A: ROS release in presence of 10% human plasma, and B: ROS release without plasma. 5.3.3. Osteoclast morphology The evolution of osteoclast morphology was qualitatively evaluated by SEM. Figure 5-4 shows the OC morphology on substrates at 14, 21 and 28 days. On TCPS, round cells were found at 14 days, which became bigger and more spread after 21 and 28 days. At 28 days, protuberances surrounded by a flat halo were visible for all cells, compatible with mature OC morphology. β-TCP substrate showed a high number of round cells at 14 days which spread and showed several connecting filopodia at 21 days and 28 days. Biomimetic substrates, CDHA and CDHA-H, also supported OC adhesion as shown in the results at 14 days. After 21 days, cells appeared more elongated on CDHA-H than on CDHA at 28 days the presence of cell clusters were observed in some regions of the sample particularly for CDHA-H
Chapter 5 112 Figure 5-4. SEM images of osteoclasts cultured on substrates at different time points (14 days, first column; 21 days, second column, and 28 days, third column) Scale bar: 20µm. Higher magnification images of CDHA-H at 28 days showed regions with degraded crystals compared to pristine plate-like microstructure of the material. This was compatible with OC resorption and was especially evident close to the cells (Figure 5-5). Figure 5-5B shows that the plate-like crystals typical from CDHA appeared degraded at the adjacent zones of OC. The differentiation of OC was monitored using TRAP-Hoechst staining at each time points (14, 21 and 28 days). Few adherent cells were observed on the biomimetic substrates compared to sintered β-TCP. At 14 days, multinucleated cells were evident only on TCPS and CDHA-H (Figure 5-6D and J, black arrows). This aspect was maintained over 21 and 28 days for both substrates (Figure 5-6E and F for CDHA-H, and Figure 5-6K and L for TCPS). Only at 28 days few multinucleated cells were found on CDHA and β-TCP, even if some cell clusters were visible already at 21 days for CDHA (Figure 5-6B).
Inflammation and osteoclastogenesis of heparinized CDHA 113 Figure 5-5. SEM images of CDHA-H at 28 days showing distorted crystals from OC activity. A: low magnification image showing a covered surface by OC (scale bar:10µm), and B: higher magnification image showing degraded plate crystals (white arrows) compared to the pristine crystal morphology (*) at the adjacent zone to cells (+), (scale bar:2µm).
Chapter 5 114 Figure 5-6. TRAP-Hoechst staining showing OC differentiation over time on CDHA (A, B and C), on CDHA-H (D, E and F), on β-TCP (G, H and I) and on TCPS (I, K and L).). Scale bar:25µm. Black arrows indicate multinucleated cells (OC). Image analyses based on TRAP-Hoechst optical images disclosed different cell behaviors (Figure 5-7A). A higher number of multinucleated cells was found on TCPS since 14 days, which was increasing with time (21 and 28 days). CaPs substrates displayed lower number of multinucleated cells compared to TCPS. For CDHA, the presence of heparin (CDHA-H) produced higher number of multinucleated cells at all-time points, whereas β-TCP and bare CDHA showed similar number of multinucleated cells but with less nuclei. A more in-depth study on the morphology of the nucleated cells revealed different fusion patterns over time. Figure 5-7B illustrates the comparison of the number of nuclei in the
Inflammation and osteoclastogenesis of heparinized CDHA 121 differences in TRAP and TRAP5b were noticed only when CaP substrates were compared to TCPS. An indirect information on OC activity can be inferred from the extracellular calcium and pH measurements, being both parameters indicative of OC activity.[66] Osteoclastic resorption of the substrate should lead to an increase in Ca2+ concentration in the culture medium. However, this can be masked by the variations in calcium merely caused by cell culture medium in contact with the CaP substrates, even in absence of cells compared to TCPS. The three CaPs analyzed in this study were prone to uptake calcium from cell culture medium, although this uptake was more pronounced for biomimetic CDHA, as recently reported.[67] In presence of OC, calcium concentration increased for all substrates, which can be associated to OC activity. This increase was more marked for β-TCP (Figure 5-9B), probably due to its inherent higher solubility compared to CDHA.[68] Also for CDHA, the presence of cells led to an increase in calcium concentration in solution, which was more sustained for heparinized CDHA-H samples. Additionally, acidification, which is related to OC activity, was demonstrated for all CaP substrates in the presence of cells compared to TCPS (Figure 5-9D). β-TCP showed the highest pH variations, followed by CDHA-H, which were maintained over the time points of the experiment. Overall, we can conclude that heparin plays an important role as osteoclastogenesis modulator due to the high number of multinucleated cells. Additionally, SEM images supported the OC activity by degraded microstructures adjacent to cells in accordance as well with calcium and pH variations. Similarly to recent literature, we hypothesize the heparin effect on OC maturation is due to the interactions in the OPG/RANKL pathway, even these conclusions cannot be extended to all bone PGs.[69] This behavior, together with heparin participation in other stages of wound healing and tissue regeneration,[70,71] support the development of heparinmimicking materials in tissue engineering and treatment of various diseases.[72] 5.5. Conclusions Biomimetic CDHA substrates showed a lower inflammatory response in terms of ROS release when compared to sintered β-TCP substrates; however, no further reduction of the inflammatory reaction was observed when the surface of CDHA was functionalized with heparin. On the contrary, OC were highly stimulated by the presence of heparin. Heparin immobilized on CDHA-H induced an earlier OC precursors fusion and maturation at 14 days, similar to that found on control TCPS, whereas the number of multinucleated cells on the pristine CaPs was
Chapter 5 122 significantly lower at 14 days, and remained smaller also at longer times. Even if TRAP activity and TRAP5b at 28 days did not demonstrate clear differences in the OC activity, morphological assays (SEM and TRAP-Hoechst) showed well spread and mature OC on the surfaces of CDHA-H. Material degradation was found on CDHA-H due to OC activity as depicted from microstructural changes of the typical plate-like crystals. Overall, we can conclude that heparin binding to CDHA plays a important role as osteoclastogenesis modulator due to the high number of multinucleated cells. This could have an impact in the control of the resorption of bone grafts, and could represent a useful tool to integrate the biomaterial in the bone remodeling cycle. 5.6. Appendix. Osteogenic response to heparinized CDHA In vitro cell studies with MSC on CDHA and heparinized CDHA were performed to further investigate the osteogenic potential of the substrates. CDHA discs were cultured in presence of rat MSC (2.85·104cells/mL) up to 3 days. Gene expression in terms of ALP, BMP-2, COLL and osteonectin (ON) was investigated at 6h, 1 and 3 days. TCPS was included as control. After the first time point (6h), osteogenic media (om) consisting of 50µg/mL ascorbic acid, 10mM β-glycerophosphate and 100mM dexamethasone, was supplemented to all substrates, except to TCPS control. RNA was extracted at each time and quantified by spectrophotometer (NanoDrop ND-1000,Thermo Fisher Scientific Inc., Waltham, USA). QuantiTect Reverse Transcription Kit (Qiagen GmbH, Hilden, Germany) was used for the synthesis of 150µg of complementary DNA. DNA templates were amplified using the primers shown in Table 5-1 using QuantiTect SYBR Green RT-qPCR Kit (Qiagen GmbH, Hilden, Germany) in an RT-qPCR StepOnePlus (Applied Biosystems, Thermo Fisher Scientific Inc., Waltham, USA). RT-qPCR runs for specificity of primers was determined by melt curves analyses. Data was normalized by the expression of housekeeping gene (β-actin) and relative fold changes were calculated related to TCPS control at 6h, according to the following equation: Etarget∆Cq target (TCPS 6h – sample) / Ehousekeeping∆Cq housekeeping (TCPS 6 h – sample)(Eq. 1) where Cq represents the median value of the quantification cycle of the triplicate of each sample and E corresponds to the efficiency of amplification and is determined from the slope of the log-linear portion of the calibration curve, as E=10(-1/slope).
Inflammation and osteoclastogenesis of heparinized CDHA 123 Table 5-1. Sequences of the primers used for quantification of gene expression (fw: forward; and rv: reverse). Gene Gene symbol Primer’ sequences (5’ to 3’) β-actin ACTB fw:CCCGCGAGTACAACCTTCT rv:CGTCATCCATGGCGA ACT Bone morphogenetic protein-2 BMP-2 fw:CCCCTATATGCTCGACCTGT rv:AAAGTTCCTCGATGGCTTCTT Alkalinephosphatase ALP fw:GCACAACATCAAGGACATCG rv:TCAGTTCTGTTCTTGGGGTACAT Collagen COLL fw:GCAGCTGACTTCAGGGATGT rv:CATGTTCAGCTTTGTGGACCT Osteonectin ON fw:GTTTGAAGAAGGTGCAGAGGA rv:GGTTCTGGCAGGGGTTTT The presence of heparin on CDHA (CDHA-H) demonstrated to increase the ALP, COLL and ON at earlier time points (6h) compared to pristine CDHA and TCPS, maintaining their expression after 1d and finally decreasing at 3d (Figure 5-10). Surprisingly, BMP-2 was not enhanced on CDHA-H, since pristine CDHA registered higher expression levels. However, whilst BMP-2 expression for CDHA and TCPS followed a peak expression with a subsequent decay, in CDHA-H, the expression of BMP-2 was maintained throughout the experimental time. Overall, these results provide some preliminary data that suggest the presence of heparin surfaces on CDHA may produce a stimulatory osteogenic effect, but this require further study. Figure 5-10. Gene expression of CDHA and heparinized CDHA-H, together with controls TCPS and osteogenic medium-stimulated TCPS, cultured with rMSC for 3 days.
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