Carburos de silicio biomórficos como cerámicas portadoras de fármacos para su aplicación en implantes óseos

Departamento de Farmacia y Tecnología Farmacéutica Facultad de Farmacia Campus Vida 15782 Santiago de Compostela Carburos de silicio biomórficos como cerámicas portadoras de fármacos para su aplicación en implantes óseos Patricia Díaz Rodríguez Santiago de Compostela, 2014 ” Al comienzo de la tesis, éste parece el momento más deseado, escribir los agradecimientos, ya que es señal de que el trabajo ha llegado a buen término. Sin embargo, no resulta fácil resumir en un par de folios a todos aquellos que de una manera u otra han contribuido a la finalización de esta tesis. Espero no olvidarme de nadie y si lo hago, espero que me disculpéis. En primer lugar, me gustaría expresar mi agradecimiento a mi directora de tesis Mariana Landín Pérez por su apoyo incondicional, sus consejos y su dirección desde el momento en el que me inicié en el mundo de la investigación. Del mismo modo, quiero agradecer a todos los profesores del Departamento de Farmacia y Tecnología Farmacéutica, especialmente a Ramón Martínez Pacheco, José Luis Gómez Amoza, Carmen Álvarez Lorenzo, Ángel Concheiro Nine y Francisco Otero Espinar por su disponibilidad siempre que lo he necesitado. Al Departamento de Física Aplicada de la Universidad de Vigo especialmente al profesor Pío González, la profesora Julia Serra, Miriam y Cosme por permitirme usar sus instalaciones y por su completa disponibilidad a lo largo de toda la tesis. Al profesor Henning Madry y la profesora Magali Cucchiarini de la Universidad de Saarland por su amabilidad y por brindarme la oportunidad de realizar una estancia en su centro. También quiero agradecer a Janina, Angelic, Amos, Jagi y Gertrud por ser mi familia en Alemania y hacerme sentir como en casa desde el primer día. A todos los miembros del Instituto de Ortopedia y Banco de Tejidos de la Universidad de Santiago de Compostela, Maite, Silvia, Ana, Mari y Abel por el buen trato y la ayuda recibida durante todos estos años. A mis compañeros del Departamento de Farmacia y Tecnología Farmacéutica Clara, Eva, Fani, Cibrán, Fran, Negin, Fernando Álvarez, Sandra, Rosalía, Julia, Patri, Bea, Elena Cutrín, Elena Raviña, Marga, Magdalena, Rocío, Andrea, Lorena y a aquellos que ya no están en el laboratorio pero que me han ayudado en todos estos años y que sin ninguna duda echo mucho de menos; Ana Puga, Nano, Mariajo, Laura, Manolo, Luis, Álvaro, Abraham y Fernando. Sin vosotros ésto no hubiera sido lo mismo. A Bárbara, por compartir tantos buenos momentos juntas, tantas risas, por ser mi confidente y por iniciarme en el tute. A Manuela, por preocuparse siempre por todos y por tener una familia excepcional. A Luis Díaz, por ser capaz de generar un dialecto propio muy enriquecedor científicamente. A Isa, por conseguir sacarme una sonrisa sin importar si es lunes por la mañana o viernes por la noche. A Alejandro, por mantener la motivación de quien está empezando una tesis y contagiar ese sentimiento a quien le rodea. A Sonia, por regalarnos una sonrisa todos los días. A Catarina, mi compañera de mesa, por darme siempre su opinión más sincera además de enseñarme los conceptos básicos de portugués. A Susana Simoes, por ser además de compañera de laboratorio, compañera de piso, gracias por estar siempre ahí. A Ana Rey, porque sin tí esta tesis no sería lo que es. Muchísimas gracias por compartir nuestra motivación, a veces excesiva, por la ciencia y ayudarme en todo lo que he necesitado. A Lidia, qué decir, simplemente gracias, gracias por aguantarme a tu lado tanto tiempo, gracias por escucharme y gracias por ayudarme siempre que lo he necesitado. A mis amigos farmacéuticos Sole, Luis, Silvia y Belén que a pesar de llamarme friki, aguantan todas las charlas científicas que les doy y me animan en todo lo que me propongo. A Josefa, Juan, Maruja, Moncho, Ana, David, Pablo, Luz, Ana Álvarez, Tamara y Manolo por conseguir hacerme desconectar los fines de semana y por preocuparse siempre por mí. A Juan Carlos, porque caiga lo que caiga siempre has estado a mi lado sin importar las consecuencias y por ser el copiloto ideal en este viaje. A mi familia, especialmente a mis tíos Suso, Teresa, Gloria, Manolo y mis primos Javi, Rubén y Yessica porque de nada sirve llegar a la meta si te sientes solo. Gracias por hacerme sentir siempre arropada. A mis abuelos Irene, Dolores, Manuel y Ángel aunque algunos ausentes, siempre presentes, por todos sus consejos y por haberme hecho la niña más feliz de este mundo. A mi hermano Efrén, muchísimas gracias por apoyarme en todas mis decisiones y por valorarme tal y como soy. A mis padres, Aurea y Manuel por enseñarme a ver las cosas siempre bajo el prisma de la positividad, a relativizar y a valorar las cosas realmente importantes en la vida. Este trabajo es vuestro también. Finalmente, agradezco al Ministerio de Educación Cultura y Deporte por la beca FPU concedida que me ha permitido realizar esta tesis doctoral, a la Fundación Barrié de la Maza por permitirme hacer una estancia de investigación en la Universidad de Saarland y al proyecto POCTEP0330IBEROMARE1P por la financiación facilitada. A TODOS, MUCHÍSIMAS GRACIAS 3 Resumen/Summary Los hidrogeles formados por alginato y poloxamer también fueron capaces de incluir vectores virales recombinantes adeno-asociados (rAAV) promoviendo su liberación sostenida y manteniendo su eficacia de transducción. El sistema polimérico con estructura más compacta ha obtenido un incremento de la expresión trangénica más prolongado. La incorporación de los hidrogeles cargados en los bioSiCs no ha sido capaz de obtener una adecuada eficacia de transducción siendo necesarios estudios adicionales. En su conjunto, los resultados obtenidos confirman el elevado potencial de las cerámicas de carburo de silicio como sistemas biofuncionales con aplicación en implantes óseos. La posibilidad de emplear numerosos precursores naturales y varias técnicas de carga permitiría la selección del sistema más adecuado en función de la necesidad requerida. 5 Resumen/Summary Summary The incorporation of therapeutic molecules into implantable biomaterials for in situ drug release gives a novel generation of biofunctional materials. This research work is focused on the synthesis of ceramics derived from natural precursors by a two-step process, which allows biomorphic silicon carbides (bioSiCs) to be obtained. bioSiCs potential to regenerate bone tissue was established by evaluating their morphological properties, their cellular responses produced and their ability to load and release therapeutic molecules. The bioSiCs synthesis from different natural precursors (pine, oak and sapelli) has led to systems with various morphological properties. Their structural and surface characterization has led to the conclusion that higher porosity ceramics are those resulting from pine wood; the higher pore size is obtained using oak wood as precursor and sapelli wood generates the densest and the roughest ceramic, also with the highest pore interconnectivity. The evaluation of the biocompatibility of the three ceramic structures has been performed by the cell culture with mesenchymal stem cells showing excellent results. Resumen/Summary 6 In addition, systems with larger pore size induced cell differentiation, leading to higher concentrations of osteoblastic differentiation markers. The study of the interactions of silicon carbide ceramics and their carbonaceous precursors with tissue and blood components showed the morphological and structural parameters that modulate these interactions. The three ceramic systems were loaded with vancomycin by unspecific adsorption showing porous structure dependent release profiles. The therapeutic activity of these systems was evaluated by the ability to treat a bacterial biofilm and the ability of sapelli systems to inhibit its formation. Sapelli systems significantly inhibit biofilm formation and all loaded systems were able to decrease the number of colony forming units (CFUs) present in the preformed biofilm. The Vascular Endothelial Growth Factor (VEGF) was incorporated into the bioSiCs by ionic interactions between the protein and the functional groups of the ceramic surfaces. The loaded systems show a sustained release profile dependent on the porous structure which was able to maintain the bioactivity of the protein both in vitro and in vivo. The addition of VEGF into the ceramic has achieved a faster osteoblastic differentiation, indicating a synergistic effect between VEGF and bioSiCs topography. Hydrogel-ceramic composite systems were obtained by mixing alginate and poloxamer. The systems achieved sustained release profiles of poorly soluble drugs, also modulated by the morphology of the materials. The in vitro evaluation of indomethacin loaded systems showed a satisfactory anti-inflammatory activity for all the samples and also a decrease of the catabolic activity of osteoarthritic chondrocytes. The hydrogels formed by alginate and poloxamer were also capable of including recombinant adeno-associated viral vectors (rAAV) promoting their controlled release and maintaining their transduction efficiency. The polymeric system with the most 7 Resumen/Summary compact structure was able to obtain a longer increase in transgene expression. The incorporation of loaded hydrogels into bioSiCs has not been able to obtain adequate transduction efficiencies, additional studies being required. Taken together, the results of this work confirm the high potential of silicon carbide ceramics as biofunctional systems with applications in bone implants. The large number of natural precursors and different loading techniques available, should allow the selection of the right fit in each case according to the required needs. Capítulo 1 Introducción general 1.1 Implantable materials for local drug delivery in bone regeneration 19 Implantable materials for local drug delivery in bone regeneration matrix (28, 34, 38-40) or a ceramic layer (35), adsorbed directly on to these implant coatings (36, 41-44), nonspecifically adsorbed on to the implant surface (45) or bound to the metal surface (30, 46, 47). Titanium is the most used metal for this purpose. 1.1.3.2 Ceramics Ceramics are defined as inorganic non-metallic materials. It is a complex group including calcium phosphate based ceramics, silica based bioactive glasses and inert ceramics, such as zirconia and silicon carbides. Their use as biomedical materials dates back to late 1960s, when they were introduced to improve the applicability of the metals. Ceramics present, as main advantages, their lower wear rates at the articulating surface and the release of lower concentrations of inert wear particles (48, 49). They are hard refractory, polycrystalline compounds, difficult to shear plastically with high melting temperatures, low electric conductivity and corrosion resistance (15, 15, 48). Ceramics could be classified into two major groups: biodegradable and non-absorbable ceramics. 1.1.3.2.1 Biodegradable ceramics Their ability to interact with tissue environment facilitating their surface mineralization and therefore the bone growth makes them the most attractive group in tissue engineering. Biodegradable ceramics are reabsorbed after implantation by dissolution, thus being replaced by endogenous tissues. The degradation of the ceramic matrix leads to the formation of particles which are phagocyted by macrophages and giant cells whereas large material volumes are reabsorbed by osteoclast (33, 48, 50). Calcium phosphate ceramics. As previously mentioned the main mineral component of bone is hydroxyapatite (HAp), a specific type of calcium phosphates. This fact justifies the interest in calcium phosphate ceramics when trying to simulate the natural bone Capítulo 1.1 20 composition (41, 43). It has been shown that these ceramics are able to improve the bone bonding strength without inducing the formation of fibrous tissue showing excellent biocompatibility (49, 51). Despite the hydroxyapatite chemical composition being the closest to the mineral bone, its high stability makes the development of chemical modified derivatives with better aqueous solution properties to improve resorption rate in vivo necessary (50, 52-54). Calcium phosphate ceramics degradation rate could be tailored by combining two or more calcium phosphates with different stability. Tricalcium phosphate (TCP) with Ca/P ratio of 1.5 and four polymorphs is widely used. The combination of β-TCP and hydroxyapatite named biphasic calcium phosphate is a material whose resorption rate can be modified according to the bone tissue regeneration and with higher osteoinduction over pure HAp, xenografts and in some cases autologous bone graft (48, 55, 56). It has been shown that these ceramics have been replaced in vivo by new and functional bone tissue which represents a significant advantage compared with other biomedical materials (56-58). Calcium phosphate ceramics are commonly used for local drug delivery as porous scaffolds, coatings and self-hardening cements and in the latter years as nanoparticles (32). As cements, one or more calcium phosphates can be mixed with an aqueous solution to give injectable pastes. After implantation and an in situ curing process, a final carbonate apatite product is formed which is characterized by an adequate mechanical strength that avoids the migration of the cement to undesirable sites (53, 59). Their ability to mold the defect site and their osteoconduction makes them attractive for orthopedic recovery, being approved by the Food and Drug Administration (FDA) in 1996 for the repairing of craniofacial defects in humans (60). Drugs could also be incorporated by mixing them into one or the two phases of their components (53). 21 Implantable materials for local drug delivery in bone regeneration Phosphate ceramic nanoparticles are high biocompatible and easy to handle. They are biodegradable giving no toxic degradation products. Nanoparticles can be used as drug delivery systems even for vectoring RNA and DNA. These nanoparticles could be obtained using various synthesis methods which lead to the efficient incorporation of antitumoral agents and proteins (32). These nanoparticles could also be coated by a drug-polymer solution forming nanocomposite systems which are able to release antibiotics and peptides (61-63). A common combination is the use of ceramic coatings in metals which could be performed by different approaches; plasma spraying, sol-gel deposition, electrophoretic deposition, simultaneous vapor deposition, pulsed laser or electron beam deposition and biomimetic precipitation (32, 41). The final material is characterized by higher osteointegration than the substrate alone (42). However, most of these coating techniques use high processing temperatures making drug loading difficult (32). This makes biomimetic calcium phosphate-drug coprecipitation the most used technique in order to obtain drug loaded coating (64-66). Calcium phosphate ceramic porous scaffolds could be loaded by specific and unspecific drug adsorption (67-69). However, the combination with polymers giving a final composite system is the most common technique (55, 70-72). Bioactive glasses and glass-ceramics. Silica-based bioactive glasses are noncrystalline compounds made from fine ceramic crystallites in a glassy matrix, which present as the main advantages high surface reactivity, good degradation rate, osteoconduction and osteoinduction, creating an apatite layer precipitate in the presence of simulated body fluid (48, 49, 73-76). SiO2 is the mayor constituent of the glass providing stability to the material by the formation of a covalently bonded network (73). The incorporation of new techniques in the obtaining process of bioactive glasses has lead to the production of mesoporous bioactive glasses which are characterized by well structured pore Capítulo 1.1 22 channels with an average pore size of 5-20 nm and high surface area. Those properties make the loading of several therapeutic molecules or the surface modification to enhance drug affinity possible, and justify the high number of published studies on this subject (49, 73, 77-81). 1.1.3.2.2 Non absorbable ceramics Non absorbable ceramics are materials that maintain their physical and mechanical properties once implanted into the patient. Alumina, zirconia and pyrolitic carbon are examples of this group. Their excellent biocompatibility, strength and toughness are their main advantages. However, they have a lack of long term stability and poor osteointegration properties (33, 48, 82). Non-oxide ceramics, such as silicon carbide or silicon nitride are considered more stable and not as sensitive to the characteristic slow crack growth of ceramics, which should lead to a better reliability. The obtaining of porous silicon carbide using the wood structure as template has lead to biomorphic silicon carbide ceramics (bioSiCs) characterized by consistent biomimetic microstructure, high degree of interconnectivity, high strength and thermal conductivity, good resistance to oxidation, biocompatibility and thermal conductivity whose final properties are controlled by the density and microstructure of the wood precursor (83-88). 1.1.3.3 Polymers Polymers are the product of covalent bonding of small molecules called unimers forming long chain molecules widely used in biomedical devices (27). According to their origin they could be divided in two groups: natural and synthetic polymers. 23 Implantable materials for local drug delivery in bone regeneration 1.1.3.3.1 Natural polymers The most common natural polymers used in biomedical field are collagen, alginate, agarose, fibrin, chitosan and hyaluronic acid (5, 57). They are biodegradable, bioresorbable and versatile (24) and could be used as gels, porous scaffolds, films and nanofibers (24, 47). Collagen is a structural protein of bones widely used in bone regeneration, the crosslinking with different agents to reduce its degradation rate is a commonly used technique to increase its stability (5, 17). A good alternative to collagen is the use of its denatured state, gelatin. The presence of sequences arginine-glycine aspartic acids in its structure improves osteointegration and stimulates the adhesion of osteoblasts (89). Alginate is a polysaccharide present in brown algae, formed by D-mannuronic acid and L-glucuronic acid, this composition allows the binding of divalent cations (Ca2+, Cd2+, Cu2+, Mn2+) producing the gelation of the polymeric solution. Moreover, calcium is the most frequently used ion to promote gelation of alginate systems for biomedical applications due to its presence in the human body especially in bone tissue (89, 90). The use of this polymer has been successful in the controlled release of growth factors (24). Chitosan is a hydrophilic polysaccharide with a similar structure to naturally formed glycosaminoglycans that could be degraded by human enzymes (3, 5). Its cationic nature facilitates the interaction with anionic glycosaminoglycans and proteoglycans which have great affinity to cytokines and growth factors. On the other hand due to its Nacetylglucosamine moiety, chitosan is able to interact directly with growth factors, receptors and adhesion proteins (91). This fact, together with its antibacterial activity and its appropriate physicochemical and biological properties, make it an excellent Capítulo 1.1 24 material for the preparation of biomedical materials and drug delivery systems (34, 92, 93). Agarose is used for many tissue engineering applications such as drug delivery or cell culture beads (57) whereas fibrin has been used as tissue sealant in surgery. For this application fibrin gels are prepared by the combination of fibrinogen and thrombin solutions containing calcium ions forming an adhesive glue adequate for the administration of endogenous plasma rich in growth factors (24, 90, 94). Hyaluronic acid is a linear polysaccharide, essential component in extracellular matrix. It is characterized by high biocompatibility and may mediate in cellular signaling, wound repair and matrix organization. The use of this polymer in extracellular matrix biomimetic hydrogels has shown good results as growth factors delivery systems (95). Other natural polymers have been useful as implantable bone drug delivery systems such as silk (96), chondroitin sulfate (25) and bacterial cellulose (97). Natural polymeric systems are able to load drugs into the polymeric matrix obtaining diffusion/erosion controlled drug release profiles (90, 98) that can be tailored modeling the drug load charge, the polymer weight and crosslinking variables (99). 1.1.3.3.2 Synthetic polymers Although natural polymers present adequate biological properties their immunogenicity, the difficulty in processing and the potential risk of transmitting animal-originated pathogens make it necessary to obtain synthetic alternatives (47). Poly(α-esters) such as poly(caprolactone) (PCL), poly(propylene fumarate) (PPF), poly(lactic acid) (PLA), poly(glycolic acid) (PGA) and their copolymer poly(lacticglycolic acid) (PLGA) and poly(ethylene glycol) (PEG) are broadly used in the medical field with minimal foreign body reaction (2, 3). These polymers have been approved by the FDA for their use in diverse clinical applications such as sutures, spinal fusion cages, 25 Implantable materials for local drug delivery in bone regeneration coronary stents, systemic drug delivery systems, nerve conduits and fixation screws (2, 43). PLGA is the most used as implantable bone material. Its degradation rate and mechanical properties could be tailored by selecting its molecular weight and copolymer ratio which makes this polymer attractive for controlled drug delivery systems applications (3, 24, 100, 101). Great efforts have been carried out in this field in order to obtain the perfect material with the adequate release profile; the use of PLGA microspheres alone or included into different matrices has been one of the most common technological approaches used to achieve the desirable bone local release. In this sense, good results have been obtained in the release of different therapeutic molecules as BMP-2 (102, 103), BMP-7, IGF (insulin-like growth factor), dexamethasone (104, 105), alendronate (106) or gentamicin (107). The main disadvantage of these molecules (PLA, PGA and PLGA) is their acidic degradation products as a consequence of their hydrolytic process that can cause the degradation of therapeutic molecules (91). PCL is a biocompatible, biologically inert polymer whose mechanical properties and degradation rate could be also modified in order to obtain the desirable characteristics. It shows minimal toxicity and immune response (18, 25). PPF is a linear polyester, highly biocompatible which has been shown to support and guide bone formation once crosslinked (108). The ability of PEG to bind proteins or peptides giving stable nanosized complexes for their controlled release is used in the development of protein commercial drug delivery systems (57, 109). This polymer has also been useful for bone local delivery of BMPs (24). The synthesis of new biodegradable polymers such as poly(lactic acid)-p-dioxanepoly(ethylene glycol) (PLA-DX-PEG) has been useful for the controlled release of BMP- Capítulo 1.1 26 2. This polymer is able to undergo a sol-gel transition by the change of the temperature and its degradation rate could be tailored in order to achieve the desirable drug release profile (24, 110-113). Although, the above mentioned polymers with elevate biocompatibility and biodegradability have high interest in the orthopedic field, the gold standard in polymers for orthopedic prostheses are poly(methyl methacrylate) (PMMA) cements and beads (114). They are characterized by their easy fabrication and excellent physical properties including tensile modulus, tensile strength, flexural rigidity and resistance to creep (15, 33). Today, the strategy approved for local antibiotic delivery in chronic osteomyelitis after surgical debridement of dead bone is the use of antibiotic loaded PMMA implants. However, the poor osteointegration and non-biodegradability of the polymer makes it necessary to perform additional surgery to remove the polymer device once the pathology is repaired (115, 116). Furthermore, the high drug doses loaded into the cements are not completely released due to the absence of degradation. For overcoming this limitation, different approaches have been made with better results as the incorporation of loaded PLGA microspheres into the cements (117) or the combination with other polymers (118). 1.1.3.4 Composites As it can be deduced, each material has advantages and disadvantages. Not strange then that studies aimed at the collection and evaluation of composites are attracting increasingly attention. In the last few years new complex biomaterials have been developed in order to have the advantages of various materials simultaneously. A composite could be defined as a continuous phase material, made from two or more ingredients with significantly different physical and/or chemical properties, whose characteristics are far from those of the raw materials (33). Several possible 27 Implantable materials for local drug delivery in bone regeneration combinations pointed out by Habraken and coworkers and Bose and coworkers (32, 119) can be used as implant material and also for local drug delivery. The most used combination is the inclusion of particles or fibers of calcium phosphate ceramics and bioactive glasses into biodegradable polymeric matrices (120). Composites of chitosan and PLA where PLA provides mechanical strength and stiffness whereas the cationic nature of chitosan minimizes the reduction of pH caused by the PLA degradation (91) are also employed or, with the same objective, mixtures of components able to release calcium and silicon ions, as bioactive glasses or phosphate ceramics with polymeric systems as PLGA (121) have been developed. The combination of two types of ceramics in order to couple both mechanical properties and electrical conduction could be also used; an example is the case of carbon nanotubes and silica for obtaining a final material with good in vitro results stimulating cell proliferation of human osteoblasts after electrical stimuli (122). More complex systems can also be synthesized including metals, ceramics and polymers. In many of them the stability of the polymeric component, mostly PLGA, is increased by the addition of the ceramic phase, calcium phosphates or bioactive glasses that neutralize the surrounding acid environment and reduce its autocatalytic effect (43). These ternary systems have been shown to be useful in controlled release of antibiotics (34). 1.1.4 Therapeutic molecules with interest in bone regeneration Going back to Figure 1.1.1, it can be seen that the bone healing is an extremely complex process depending on different factors. Despite the great development of new materials in the few last years, the ideal implant material has not yet been developed. Some potential candidates, despite possessing appropriate porosity and mechanical properties, have a lack of osteoinduction. In this situation, their loading with bone Capítulo 1.1 28 morphogenetic proteins and/or other growth factors should enhance their value by improving cellular attachment and bone formation (23, 123). Although bone induction is important in the prosthesis osteointegration, implant fixation is also crucial in order to avoid its failure. It has been shown that the local administration of bisphosphonates, antiresorptive drugs, improve implant fixation in vivo (124). PMMA cements are characterized by low porosity and interconnectivity making the growth of bone cells into their inner structure difficult. This complication could be overcome by the addition of N-acetyl cysteine, a strong antioxidant that acts as a scavenging agent during the polymerization, increasing porosity and consequently osteoconductivity (125). As can be seen the combination of biomaterials and drugs has lead to the enhancement of the material properties making the therapeutic success easier. Material-drug combinations are not only adequate for the improvement of implant materials but also allow local drug delivery. Local therapeutic treatments increase the time of permanence of the drug at the target site, maintain high local drug concentration and reduce the occurrence of side effects (23, 117). In addition, the incorporation of labile therapeutic molecules as growth factors to biomaterial systems could enhance their stability protecting them from enzymatic or chemical degradation both in polymeric systems (126-128) and in ceramic matrices (13). The use of solid implant materials loaded with BMP-2 able to support enough space for cell growth has been found to be more adequate for bone regeneration than hydrogels. These systems have achieved higher levels of bone formation due to the stimulation of tisular compression, the key in bone induction (67, 129). Moreover, high implant volume of materials has produced a high bone formation (110). This could be explained by the influence of mechanical signals in the bone healing modulating 35 Implantable materials for local drug delivery in bone regeneration combine hyperthermia with local drug deliver capacity. The high temperature achieved after applying magnetic stimulation of the implant causes the tumor cells death (185). 1.1.4.7 Others Simvastatin, a common hypolipidemic drug is also of interest because it has been reported to stimulate bone formation by inducing the expression of BMP-2 in different animal models, and therefore enhances osteogenesis and also induces osteoblastic differentiation of mesenchymal stem cells (186-189). Jeon and coworkers (190) found that its intermittent release from devices promotes osteoblastic stimulation while continuous stimulation reduces cell viability. The proposal of the use of lactoferrin, a protein present in the granules of neutrophils and in breast milk with osteogenic purposes, is based on its ability to modify proliferation and differentiation of osteoblasts, increasing the calcification of extracellular matrix (191). Lidocaine, is a local anesthetic that could be used for decreasing pain after the surgery. Its short half-life in blood serum (1.5-2 h) makes the design of controlled release systems necessary when using this drug (192). Different therapeutic ions have also been studied as possible candidates for local controlled release from implants. Controlled release of silver ions (Ag) from tricalcium phosphate particles has shown an adequate protection against bacterial re-infection (193). Selenium (Se) released from coated titanium has anticancer activity in vitro against various cancer cell lines (182). Whereas strontium (Sr) promotes bone regeneration and inhibits bone resorption enhancing bone volume and microstructure (194, 195). Capítulo 1.1 36 Some polymers also have therapeutic interest such as poloxamines. These triblock synthetic polymers formed by poly(ethylene oxide)-poly(propylene oxide) with an ethylenediamine core not only are adequate for drug solubilization but also (196) have the ability to promote osteoblastic differentiation of mesenchymal stem cells in vitro (197, 198). Other polymers are able to release proline when they are degraded, promoting an anti-inflammatory effect (199). In the development of new complex systems it is also possible to combine two or more drugs from different groups. Combinations of antibiotics and growth factors have been found useful for the osteointegration process by both enhancing osteoblasts function and increasing antibacterial activity (29) while the combination of dexamethasone and alendronate has been shown to be useful for the enhancement of osteoblastic differentiation both in vitro and in vivo (176). 1.1.5 Mechanism for loading drugs into implant materials and release kinetics Therapeutic molecules could be incorporated into the material matrix or adsorbed on the external biomaterial surface. Figure 1.1.2 shows the different mechanisms for loading implant materials with drugs. Drugs could be included during the material synthesis process (physical entrapment loading) or on the surface, after the final biomaterial is obtained. In the latter case, the drug can be nonspecifically adsorbed on the material surface, covalently linked to the surface or attached by physical interactions to the internal or external surface of the matrices or polymeric gel networks. Metallic implants are usually drug loaded through the formation of a polymeric or ceramic coating therefore drug loading strategy depends mainly on the coating technique selected. However, some of titanium surface chemical modifications enhance 37 Implantable materials for local drug delivery in bone regeneration the presence of functional groups useful in the covalent bonding of BMP-2 showing a controlled release (30, 46). Unspecific adsorption is normally used in ceramic systems when there is no affinity between the drug and the ceramic. Ceramic systems can also be loaded by directly mixing them with drug powders (32) or by co-precipitation (64, 65) (e.g. physical entrapment loading of the drug in calcium cements). Figure 1.1.2 Drug loading techniques on biomaterials, critical factors and associated release mechanisms and profiles. The functional groups of biodegradable ceramics confer great affinity for certain drugs such as bisphosphonates, obtaining the loading by physical interactions. Capítulo 1.1 38 Polymeric implants are the most studied group. Those systems can be loaded with drugs using all techniques. Generally, those made with natural polymers make use of the physical interactions, while the synthetic polymers implants load drugs by physical entrapment (200). 1.1.5.1 Unspecific adsorption The immersion of porous materials into high concentration drug solutions allows the drug loading by unspecific adsorption which are mainly dependent on structural parameters of the biomaterials like specific surface, total porosity and pore size distribution (201). Once in contact with the dissolution medium, these loaded systems undergo the drug desorption. Release profiles are characterized by an initial burst effect caused by the fast dissolution of drug molecules adsorbed on the external structure followed by a release highly dependent on drug solubility and material pore structure (68, 139, 202). For similar total porosities, the presence of small size pores promotes drug sustained release due to the enhancement of the surface available for drug adsorption (77). On the contrary, big pores improve the water uptake and quick drug solubilization and release (203). The modulation of pore size distribution has been used as an approach to control and extend zolendronate drug release from mesoporous bioceramics (204). An alternative method could be applying vacuum during the loading process for obtaining slow drug release profiles. This procedure makes it possible for the loaded solution to reach the small size pores of the structure, hampering the wetting and the drug release (205). The incorporation of a polymeric coating as a complement of unspecific adsorption method has been used as a way to reduce the initial burst in antibiotic release profiles by creating a local drug diffusion barrier (206, 207). 39 Implantable materials for local drug delivery in bone regeneration When growth factors are loaded on polymeric electrospining fibers and scaffolds using unspecific adsorption methods fast and uncontrolled release profiles are achieved (99, 208). This makes the improvement of the loading techniques necessary in order to obtain a more controlled drug release. 1.1.5.2 Physical interactions Physical interactions including hydrogen bonding, Van der Waals forces, electrostatic or hydrophobic interactions are widely used to load highly hydrophobic drugs in biomaterials and achieve desirable controlled release profiles (183, 209, 210). With the exception of the hydrophobic interactions, all are highly dependent on pH conditioning for both drug loading and release processes. The modification of the pH medium is critical in protein loading and delivery and can be used as an approach to modulate protein release. At a pH lower than the protein isoelectric point, the molecule is positively charged whereas at pH higher than isoelectric point are negatively charged. This modification can increase or decrease the affinity of proteins for the biomaterial (211). If the proteins are loaded under hydrophobic conditions different pH values do not cause modification in protein release. The stabilization of the protein molecule under these conditions is unable to ionize at any pH (212). The strong ionic interaction between bisphosphonates and divalent metal ions can be used as an approach for loading biomaterials. In the presence of those ions, bisphosphonates give a low soluble calcic salt precipitate that allows biomaterial surface loading by the formation of a drug coating and the consequent controlled release (42, 45). The electrostatic interactions between calcium phosphates and bisphosphonates are thought to be between two phosphate anions of calcium phosphate and two phosphonate groups of the bisphosphonate (183). These drugs have been shown to have high affinity to calcium phosphate and bioactive glasses obtaining good controlled Capítulo 1.1 40 release systems (69, 176, 213, 214). The binding affinities of bisphosphonates for hydroxyapatite are different according to their molecular structure (from highest to lowest zolendronate > alendronate > ibandronate > risendronate > etidronate > clodronate) (215). Also the ratio between calcium and phosphate, the surface charge of the ceramics and the surface hydrophobicity have an important influence on drug-ceramic affinity, being stronger for ceramics with the highest concentration of calcium (183, 215) and making the loading of hydrophobic carbon nanotubes by hydrophobic interactions possible (122). It has been demonstrated that the therapeutic results of bisphosphonates are better when delivered locally, loaded on ceramic coatings, than after their systemic administration, finding an improved implant-bone contact (42, 216). Acidic surfaces show greater capacity of bisphosphonate incorporation than the neutral surfaces, which allows the load with one higher dose of bisphosphonate in matrices that contain phosphorus due to interactions with the functional groups of the bisphosphonate (175). The loading of growth factors as BMP-2 or VEGF in implants can also be performed through electrostatic interactions using heparin as mediator (217, 218). Heparin is a highly sulfated glycosaminoglycan able to bind many growth factors by electrostatic interactions between its negatively charged sulfate groups and the positively charged amino acid groups of proteins (219). The immobilization of the heparin molecule on biomaterial surface can be performed by different techniques like its binding to polymers by ionic interactions (219) or its chemical conjugation to polymeric systems (220) and demineralized bone matrices (221). The interaction of heparin with proteins not only enhances the drug content in the final material (17) with a controlled release 41 Implantable materials for local drug delivery in bone regeneration (217) but also stabilizes proteins, protecting them against proteolytic degradation and increasing their bioactivity (219). Heparin can be also useful for loading other therapeutic molecules containing amine groups in their structure such as gentamicin. It has been shown that the combination of gentamicin and BMP-2 loaded using heparin immobilized on titanium surface, presents a suitable antibacterial activity and promotes pre-osteoblasts differentiation (29). The mechanism for heparin immobilization has an important effect on drug release. When heparin is bound through ionic interactions, the biomaterial-heparin affinity plays a key role in the controlled release of the drug (219). On the contrary, when heparin is covalently linked to the biomaterial surface, it is the drug-heparin affinity, the drug diffusion and the system degradation which are the important variables in drug delivery (221). The drug concentration should not exceed the loading ability of the heparin, otherwise an inefficient loading of the implantable system is produced (37). It is possible to synthesize therapeutic peptides including a mineral binding domain in their structure that allows their ionic interaction with calcium phosphates. Protein release from those loaded ceramics is highly influenced by the ceramic solubility that could be modified by changing the carbonate phase content (153). It has been shown that calcium phosphates have high affinity for some proteins (BMP-2, BSA) increasing the scaffold load capacity and being able to modulate their release (23, 55, 65). As an example, the high affinity of COO-, OH and NH2 groups of BMP-2 and the negatively charged hydroxyapatite gives a strong interaction through water bridged hydrogen bonds that promotes the slow release of the protein (222, 223). Higher surface available for the interaction with proteins increases the loading ability (224). Those interactions could be also used for loading antibacterial peptides that simultaneously inhibit bacterial growth and enhance bone formation (225). Capítulo 1.1 42 Additionally, bioactive glasses could interact with charged ampicillin (34), gentamicin (202) or vancomycin (226) obtaining systems with slow antibiotic release profiles. POH and Si-OH groups of bioactive glasses are able to generate hydrogen bonding with antibiotic hydroxil and amine groups. Those combinations would be sensitive to a pH that can be modulated to achieve the antibiotic controlled release (227, 228). The increase of Si-OH groups of bioactive glasses raises the loading ability of proteins due to the presence of higher number of functional groups. This effect also depends on the molecular weight of the loaded protein. The higher molecular weight of protein, the lower its loading due to the need of larger available surface for the interaction with the scaffold (210). The inclusion of strontium into mesoporous bioactive glasses has been shown to be adequate in order to achieve a controlled release of dexamethasone (229). Polymeric systems can also be loaded through physical interactions. Frequently, chemical modifications of natural polymers such as dextran have been carried out for increasing their affinity for BMP-2 (174). Another approach is to incorporate an additive to modulate the therapeutic molecule-polymer affinity. For example, the incorporation of keratose to collagen hydrogels increases growth factors affinity through ionic interactions giving final hydrogels with good controlled release properties during four weeks (230). On the contrary, the addition of chondroitin sulfate to collagen reduces its affinity for proteins by enhancing the polar groups of the hydrogels and the surface area, this modification allows a higher burst effect to be obtained that increases the osteoinductive activity in vivo possibly by the enhancement of chemotaxis (163). The condroitin sulphate applied as a coating on PCL/β-TCP scaffolds has been also shown useful for promoting controlled release of BMP-2 for 15 days. Additionally, the amount of loaded protein can be controlled by the pH modification (161). 43 Implantable materials for local drug delivery in bone regeneration Chemical interactions also explain the different approaches of developing drug controlled release systems based on other polymers as gelatin or chitosan. Gentamicin interacts with gelatin through a Schiff reaction giving slow drug release profile particles. The drug loaded amount and the release mechanism (erosion, diffusion) can be determined by the selection of appropriated basic or acid gelatin (89, 191, 231). The interactions between growth factors (TGF-β1) and chitosan allow the production of surface-loaded chitosan beads that release TGF-β1 at an adequate rate to promote osteoblastic differentiation (232). Bone morphogenetic proteins could be incorporated into polyelectrolyte films by the pH modification (25, 161). In those systems drug release is determined by the charge, the temperature or the degradation of the polyelectrolyte multilayer (233). 1.1.5.3 Physical entrapment Therapeutic molecules (antibiotics, growth factors and ions) can be physically entrapped into ceramics or polymeric systems during their synthesis or production process (234, 235). Homogeneous drug distribution is an important factor to be considered in the development of those systems in order to avoid an initial burst release. They generally achieve more prolonged release profiles than drug-adsorbed systems, the release controlled being mainly by diffusion and biomaterial erosion depending on the polymer degradation (2, 213). For incorporating drugs into orthopedic cements, they can be physically mixed (181, 236) or dissolved (53) in one of the components of the cement system or in all of them. Both techniques have advantages and disadvantages. The addition of the drug in the solid phase allows more slow-release systems to be obtained, while incorporating the drug in the liquid phase gives rise to more homogeneous systems (170). The Capítulo 1.1 44 incorporation of high amounts of drugs into these systems changes the properties of the final cement (237). The extreme operation variables during cements production or their final characteristics (e.g. pH) can limit the utility of this loading procedure. As an example, the high temperature necessary for the polymerization process during PMMA cement production may accelerate thermo sensitive drug degradation (136). The pH of the cement is crucial to maintain the activity of therapeutic molecules degrading vancomycin at pH 9.5 (238). Finally, PMMA cements have a low porosity and an extremely slow degradation rate that limit the release of drugs physically entrapped into those materials (136). In order to overcome these limitations, several polymers as carboximethycellulose (117) or ceramics as silica nanoparticles (142) can be mixed with PMMA cements. Those additives modify the porosity and/or the water uptake into the PMMA matrices enabling the dissolution and the release of the drug. Drug release can be also improved by increasing the amount of drug into PMMA cements which promotes material fragility and its cracking (239) or through external stimuli like ultrasounds (240, 241). An alternative to the PMMA is the use of calcium phosphate cements in which a hardening process takes place at room or body temperature. They have generally high intrinsic porosity that can even be improved by adding soluble organic additives (60, 242). It has been shown that calcium phosphate cements improve the therapeutic activity of certain proteins (53, 242) showing optimal release profiles for local release of antibiotics and anti-inflammatory drugs (243). Techniques to incorporate therapeutic molecules into ceramics include its biomimetic coprecipitation forming a coating on metallic implants (35), the co-precipitation of 51 Implantable materials for local drug delivery in bone regeneration been described that some of these systems are able to release growth factors through cell-mediated polymer degradation mechanism (280). Drug release profiles from polymeric or ceramic coatings (281) can be modulated by the thickness (282, 283) and the degradation properties of the coating (284, 285). Natural filmogen polymers such as zein or chitosan are an attractive alternative to the controlled release of drugs from implants, being adequate for antitumorals, antibiotics and growth factors with kinetics dependent on drug solubility (40, 254, 286). Additionally, some drug depending factors can be pointed out. The selection of high soluble (e.g. salts) or low soluble derivatives can be used as a technological approach to modulate the release profiles of some drugs, reducing the burst release or improving the dissolution process (192, 252, 287). The inclusion of high concentrations of drugs could facilitate biomaterial degradation by system instability or due to the cellular chemotaxis affecting drug release profiles (111). When calcium phosphate cements are loaded, high amounts of drugs can modify their rheological properties and setting kinetics increasing the porosity of the final material and promoting fast release kinetics. The effect is strong for molecules that interact with calcium and phosphate ions (53, 243). The inclusion of two proteins simultaneously into a hydrogel could affect the release of both proteins due to a competitive mechanism. As an example, the combination of SDF-1 (stromal cell-derived factor 1) and BMP-2 into a gelatin hydrogel increases the burst effect of SDF-1 (156). Capítulo 1.1 52 1.1.6.1 Drug release kinetic analysis Mathematical modeling of drug delivery and predictability of drug release from pharmaceutical formulations is a field of enormous importance that historically has attracted great attention (288). Despite the conditions of the in vitro studies are really heterogeneous and not allowing easy comparisons between them, different authors have modeled the drug release behavior from implants, to make predictions about the therapeutic effect of the device (35), for comparing formulations (64) or for improving the understanding of the drug release mechanism (274). The kinetics most commonly used to evaluate the release of drugs from implants are as follows (289, 290): Zero order kinetics: F = kot , where F is the fraction of released drug at time t, and ko is an apparent release rate. The release rate is independent on the drug concentration (64, 274). BMP-2 and insulin–like growth factor have been delivered at zero order kinetics or pseudo-zero order kinetics from crosslinked gelatin coating systems (168). Similarly, the release of ibuprofen from asymmetric coating titanium allows dependence on osmotic pressure (89). The release of other antibiotics and corticoids electrodeposited onto pure Ti dependent on-demand electrical stimulation can also follow zero order kinetic profiles (291). First order kinetics: ln(1-F) = – k1t, where F represents the fraction of drug released at time t, and k1 is the first-order release rate constant. The release rate is dependent on drug concentration, drug solubility and diffusivity (35, 274). Higuchi model: F = kHt½, where F represents the fraction of drug released at time t, and kH is the Higuchi dissolution constant. This equation perfectly describes release 53 Implantable materials for local drug delivery in bone regeneration processes where drugs are dispersed in monolithic systems with no changes during the release process (constant porosity, no swelling…) and release is purely diffusion controlled with constant diffusion coefficients. This model has been shown to be adequate especially for the initial stages of drug delivery from calcium phosphate ceramics and calcium phosphate cements. After that, drug release is also modulated by matrix degradation (53, 64, 236, 243, 281). Hixon-Crowell model: Mo1/3 – Mt1/3 = kst, where Mo is the initial amount of drug in the biomaterial, Mt is the remaining amount of the drug in the composite at time t, and kS is the constant incorporating the surface volume relation. Dividing the above equation by Mo1/3 and simplifying gives (1-F)1/3 = 1 – kEt, F = 1 – (Mt/Mo), where F represents the drug dissolved fraction at time t, and kE is the release rate constant. This model describes systems where the drug release is dependent on the system erosion (64). Korsmeyer-Peppas model: F = kPtn, where F represents the drug fraction released at time t, kP is the release rate constant and n is the diffusional exponent which indicates the drug release mechanism n = 0.45 Fickian diffusion, 0.45 < n < 0.89 anomalous diffusion, n = 0.89 case II transport, n > 0.89 supercase II transport or typical zero-order release (57, 64). This model has been used to explain antibiotics release mechanisms from titanium coatings of calcium alginate and gelatin (51). The use of this model is also adequate for systems whose release profiles are controlled by various factors especially hydrophilic polymeric systems in where diffusion, dissolution and swelling are key parameters (57, 98, 250). Weibull model: F= Mo [1-e-(t-T)b/a] where F represents the amount of drug dissolved as a function of time t. Mo is the total amount of drug being released, the release profiles are characterized by the relationship with the parameter shape that define the transport mechanism (61). This model, typically used for fractal system characterization, has been successfully applied for characterize drug dissolution from Capítulo 1.1 54 mesoporous composite fibers (202). This model has also been found useful in the analysis of release kinetics from calcium phosphate bone cements with interaction between drug and cement component. In these cases different parameters modulate drug release as drug solubility, interaction between drug and cement composition and the precipitation of bone cement components (150). The suitability of the therapeutic molecule release profile may extremely condition the therapeutic success of the implant. Different authors have analyzed these effects in order to obtain the adequate release profile for the success of growth factor treatments in stimulating an osteogenic response (168). This field can be expected to become an integral part of implant devices development. The wide variety of materials available makes it unlikely that there will be one general theory applicable to any type of implant. It is much more likely that there will be a broad spectrum of diverse mathematical models, applicable to specific types of devices differing in composition, geometry and drug load. In silico prediction and optimization should help in accuracy and easiness of application of drug loaded bone implants. 1.1.7 Translation to the human situation In vitro studies allow drug release profiles with good mathematical fitting to be obtained. However, results are particularly difficult to be extrapolated to the in vivo situation (286). If in vitro-in vivo correlations are always difficult for any administration route, these cases are even more problematic as the physiological environment is variable and dependent on multiple factors. Different authors have described high levels of antibiotic into the bone tissue after implantation of loaded materials despite the release kinetics in vivo were markedly slower than the one obtained in vitro during the release studies in PBS (149). 55 Implantable materials for local drug delivery in bone regeneration Several animal models (rats, mice, sheeps, rabbits, dogs, pigs…) have been used for the in vivo evaluation of drug loaded implanted materials (50, 70, 130, 149, 155, 279), showing that their therapeutic effect is highly dependent on the hematoma size, location and size of injury (292). All these parameters modify the physiological environment and therefore, the drug release (292). A key factor in the elucidation of the in vivo results when studies are carried out using implants loaded with bone regeneration factors is to distinguish between the effect of the therapeutic molecule and the intrinsic bone regeneration characteristics of the animal. Adequate controls are needed. In some studies, not the BMP activity but the regeneration ability of the bone itself is responsible for the good therapeutic results (292). 1.1.8 Future perspectives Successful bone regeneration requires the combination of many events, cells that undergo differentiation to form osteoblasts, biological factors that control growth and cell differentiation and cellular attachment, migration and proliferation (47). To help in this natural process and increasing therapeutic success, biomaterials play an important role. They must be properly tailored to obtain the correct properties and they can be loaded with growth factors to promote bone healing, antibiotics to reduce infections, analgesics and anti-inflammatories to reduce pain and recovery time or antitumoral drugs to avoid metastasis. They can also be used for gene-therapy purposes by incorporating DNA plasmids and small-interfering RNA. The sustained delivery of pDNA and siRNA from mesoporous silica nanoparticles (MSNPs) has been shown to increase the transfection levels of these molecules. All these applications need optimization and well organized strategies for the efficient delivery of drugs at target sites. Capítulo 1.1 56 The delivery of cells with therapeutic activity is the next step in the development of biofunctional materials. It has been shown that the use of autologous bone marrow mononuclear cells and bone marrow mesenchymal stem cells increase angiogenesis and bone regeneration by the sustained release of bFGF (293, 294). Additionally, the combination of osteoblasts and BMP-2 has been found to stimulate alkaline phosphatase activity (295) and human bone marrow mesenchymal stem cells together with VEGF have been useful in the enhancement of the bone regenerative mechanism (296). The bone drug delivery systems are considered combined (orthopedic materials + drug) systems by the FDA which requires a longer registration process than for traditional orthopedic implants hampering their clinical use (297). The development of new local delivery systems for bone regeneration and/or the optimization of the already developed ones need expertise from both, biomaterial engineering and pharmaceutical technology fields which have been traditionally at a distance (41). Multidisciplinary works linking bone substitute production and the “know how” of pharmaceutical companies are crucial to facilitate the clinical use of these new materials. 1.1.9 References 1. 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Drug Discovery Today. 2010;15(13/14):547-52. 1.2 Cerámicas biomórficas de carburo de silicio como nuevos materiales para regeneración ósea 99 Cerámicas bioSiCs como nuevos materiales para regeneración ósea procesos de deposición química por vapor. Sin embargo, estas elaboraciones suelen requerir temperaturas superiores a 2000 ºC y resultan caras, lo que resta utilidad a los materiales para una amplia gama de aplicaciones (36). En los últimos años se ha abordado el desarrollo de nuevos métodos de producción para obtener cerámicas biomórficas de carburo de silicio con menor consumo energético y por lo tanto menor coste (25, 37). Su porosidad abierta, ligereza y excelentes propiedades mecánicas en flexión y compresión son algunas de las ventajas del carburo de silicio biomórfico (bioSiC) frente al carburo de silicio convencional (3841). La Figura 1.2.1 sintetiza las diferentes opciones disponibles para la obtención de cerámicas biomórficas de carburo de silicio; la infiltración con silicio en fase vapor, la reducción carbotérmica y la infiltración con silicio líquido (20). Estos métodos requieren el empleo de altas temperaturas para la conversión de la madera en cerámicas (9, 14) y en todos ellos, la porosidad e interconexión de los poros del precursor determinan el proceso de infiltración (11, 24). Figura 1.2.1 Métodos de síntesis del carburo de silicio biomórfico (bioSiC) a partir de madera y precursores procesados. Capítulo 1.2 100 1.2.3.1 Infiltración con silicio en fase vapor Este procedimiento de síntesis es versátil. Las piezas de madera desecadas y moldeadas de la forma apropiada se pirolizan e infiltran con vapor de silicio (Si7, SiO8 o CH3SiCl3) a elevada temperatura (27, 42). La elección del agente infiltrante permite la obtención de diferentes productos de silicio, como SiC-Si3N4, lo que permite además modular la resistencia a la oxidación del material (43, 44). Sin embargo, su principal inconveniente es la fuente de obtención del silicio, la descomposición de precursores metalorgánicos (45), lo que implica el empleo de altas temperaturas y largos procesos (26, 46). 1.2.3.2 Reducción carbotérmica Cuando se emplea este procedimiento, las piezas de madera se infiltran con dióxido de silicio (SiO2) y posteriormente se pirolizan. El SiO2 puede ser obtenido directamente por una infiltración con silica gel o por el empleo de otros reactivos. Para la obtención del material deseado es necesario llevar a cabo varios ciclos de infiltración-pirólisis. El proceso se finaliza mediante una reducción carbotérmica a temperaturas superiores a 1000 ºC de los precursores infiltrados y pirolizados. Este procedimiento da lugar a materiales principalmente formados por fase β-SiC (47) aunque la obtención de materiales puros SiC frecuentemente requiere tratamientos térmicos adicionales a temperaturas muy superiores debido a la formación de fases intermedias (48). La formación del carburo de silicio está determinada principalmente por las reacciones en fase de vapor, dando lugar a materiales de menor resistencia que los obtenidos por infiltración de silicio líquido (15). Con el fin de mejorar los resultados de este procedimiento se han realizado diferentes estudios centrados en el control del vacío y/o de la presión durante la impregnación sol-gel con SiO2 o en el uso de fluidos supercríticos (12, 45). 101 Cerámicas bioSiCs como nuevos materiales para regeneración ósea La infiltración sol-gel puede realizarse también utilizando otros reactivos, como tetraetilortosilicato (TEOS), poli(metilfenilvinilsilsesquioxano) (PMPVS), poli(metilhidrosiloxano) (PMHS) o poli(carbometilsilano) (PCMS) (27, 49, 50). Este método es sencillo y de bajo coste, lo que constituye sus principales ventajas. 1.2.3.3 Infiltración con silicio líquido (LSI) Según este procedimiento, las cerámicas biomórficas de carburo de silicio (bioSiCs) se fabrican mediante un proceso en dos etapas; una pirólisis controlada de la madera en atmósfera inerte, normalmente argón, seguida por una infiltración reactiva rápida y también controlada con silicio fundido a una temperatura superior al punto de fusión del silicio (1.410 ºC) (36, 51-53). C(s)+ Si(l)→ β-SiC Ecuación 1.2.1 Reacción química durante la síntesis de carburo de silicio mediante la infiltración con silicio líquido (14). El material final está formado principalmente por β-SiC y silicio libre en la superficie de los poros, pero también puede contener algunos elementos traza como Al, S, B, Na (9, 14). El contenido medio de silicio se sitúa entre el 20 y el 30% (14). La mayor parte de la pérdida de peso de la madera durante el proceso de pirólisis se lleva a cabo por encima de 500 ºC. El uso de temperaturas superiores durante esta etapa promueve cambios estructurales, disminuyendo el tamaño de poro de la preforma de carbón y dificultando la posterior infiltración (54). Las características de las maderas precursoras seleccionadas modulan la microestructura de la preforma de carbón y ésta a su vez condiciona el proceso de infiltración de silicio, ya que determina las fuerzas capilares y su cinética. Además, las Capítulo 1.2 102 propiedades de la superficie del carbón influyen en el ángulo de contacto y la humectación con el silicio líquido. Por otra parte, la infiltración del carbón en la dirección axial favorece la conversión a carburo de silicio debido al transporte del fundido a través de los canales naturales del molde (20). Así, se obtiene una buena transmisibilidad de la microestructura entre la original de la madera, la del carbón vegetal y la de la cerámica de SiC final (55). El silicio fundido interacciona con la preforma de carbón dando lugar a una reacción exotérmica espontánea que disuelve el carbón generando grupos Si-C que cristalizan. Si no hay suficiente silicio para disolver el carbón, la capa inicial de carburo de silicio formada inhibe las reacciones adicionales entre Si y C, y la formación de carburo de silicio posterior depende de un proceso de difusión, cuya importancia es alta a temperaturas superiores al punto de fusión del silicio. Este mecanismo explica las características microestructurales observadas en las muestras de carburo de silicio obtenidas (36, 56). Durante el proceso de infiltración, los poros de menor tamaño son eliminados debido a la expansión de volumen del 58% que se genera tras la cristalización del carburo de silicio (54). Las condiciones en las que se produce el proceso de infiltración, además de la naturaleza del precursor, determinan la presencia de carbón residual que afectará negativamente a las propiedades de la cerámica (16, 52, 57). Por ello se han explorado diferentes opciones, como el uso de la radiación ultrasónica, con el fin de mejorar el proceso de infiltración y, por lo tanto, las propiedades finales de los materiales obtenidos (49). La amplia variedad de maderas existente ofrece la posibilidad de producir cerámicas de carburo de silicio biomórfico con diferentes microestructuras tridimensional y propiedades a medida (densidad, porosidad abierta, interconectividad, resistencia a la fractura...) para la aplicación requerida (36, 37). 103 Cerámicas bioSiCs como nuevos materiales para regeneración ósea El método de infiltración con silicio líquido presenta diferentes ventajas para la producción de cerámicas biomórficas de carburo de silicio con respecto a los demás procedimientos (8, 52, 56-58):  Resulta respetuoso con el medio ambiente ya que usa materiales renovables y el proceso es poco contaminante.  Sus requerimientos energéticos y térmicos (reacción exotérmica) son reducidos.  No precisa aditivos adicionales.  El procedimiento es rápido y de bajo coste. El carburo de silicio es químicamente inerte, no reabsorbible y extremadamente resistente a la corrosión y a la erosión. Los carburos de silicio son materiales semiconductores tanto térmica como eléctricamente (2.36 y 3.05 eV para βy α-SiC, respectivamente). En general, las cerámicas de carburo de silicio son duras y resistentes, tanto a temperatura ambiente como a temperaturas superiores, y presentan buenas propiedades tribológicas (19, 59, 60). Su erosión se produce mediante la formación y propagación de grietas laterales y radiales. La variabilidad observada en las propiedades de los carburos de silicio biomórficos se explica en función del material precursor y el proceso de fabricación empleado (temperatura de procesado, tiempo de reacción, relación C/Si empleada), ya que estos determinan su composición final, su microestructura (porosidad, morfología y distribución de tamaño de poro), su anisotropía y su densidad (8, 41). Los efectos de cada variable implicada en su proceso de elaboración sobre sus propiedades se detallan a continuación: Capítulo 1.2 104 1.2.3.3.1 Materia prima La disposición de las células vegetales en la madera, conformando tráqueas o traqueidas, determina la microestructura del carburo de silicio biomórfico (17, 61). La densidad final de los bioSiCs que oscila entre 1,1 y 2,6 g/cm3 (62) es también extremadamente dependiente de la madera precursora, habiéndose encontrado una relación lineal entre ambos parámetros (58): ρSiC=(2,39∓0,01)ρWood R2= 0,99 Ecuación 1.2.2 Relación entre la densidad de la madera precursora y la densidad de bioSiC. El empleo de maderas y procesados de elevada densidad (madera de cedro prensada de alta densidad o tablones de alta densidad) conduce a muestras con una resistencia a la compresión dos veces superior a las obtenidas a partir de moldes de baja densidad (paulonia) (25). El carburo de silicio biomórfico, presenta la anisotropía del material de partida, es decir, sus propiedades son también diferentes en dirección axial respecto a las direcciones radial o tangencial, aunque no de forma tan marcada como en la madera original. Estas diferencias son particularmente importantes en lo que se refiere a las características mecánicas y eléctricas (37, 63). En las cerámicas biomórficas, a diferencia de sus maderas precursoras, la deformación en dirección axial no sólo depende de la compresión axial de las paredes celulares de la madera y en la compresión tangencial, la deformación no es sólo debida a la flexión plástica de las paredes celulares. Los sistemas cerámicos finales son complejos y su comportamiento no solamente está condicionado por las estructuras celulares implicadas (63). 105 Cerámicas bioSiCs como nuevos materiales para regeneración ósea La resistividad eléctrica del bioSiC aumenta con el contenido de silicio residual en dirección axial, mientras que en la dirección tangencial, este fenómeno no se observa (64). El carácter anisotrópico mecánico de las muestras de carburo de silicio se puede modular mediante la infiltración con mezclas fundidas de Al3-Si9-Mg, obteniendo carburos de silicio compuestos, reforzados de aluminio, que presentan una mejor resistencia transversal a la compresión y una mayor rigidez (16). 1.2.3.3.2 Relación Si/C empleada La relación inicial entre Si y C es un parámetro crítico durante el proceso de producción. Cuanto mayor es la cantidad de silicio respecto a la de carbón, más elevada es la proporción de SiC/C en el producto final, y por lo tanto, la densidad, la dureza y la resistencia mecánica del material obtenido (65). Índices Si/C por debajo de 2,33 dan lugar a cerámicas con poros vacíos. La formación de cristales de carburo de silicio obturan los poros de la preforma de carbón y reducen la porosidad. Así, proporciones superiores a 2,33, dan lugar a materiales con poros parcial o completamente obturados con silicio residual (11, 66). Cuando la proporción de Si/C es mayor que tres, las cerámicas biomórficas obtenidas están formadas por SiC y Si residual y no presentan carbón sin reaccionar (26). La cantidad de silicio residual condiciona, entre otras, las propiedades eléctricas del bioSiC, especialmente a temperaturas entre -268 y 228 ºC. En este rango, el Si residual constituye una red interconectada, responsable de su comportamiento metálico. Este efecto no se observa a temperaturas superiores a 228 ºC, en las que el bioSiC se convierte en un material semiconductor como consecuencia de la mayor contribución del carburo de silicio (37). Capítulo 1.2 106 El exceso de silicio residual puede ser eliminado mediante su tratamiento con ácidos, usando una solución de ácido fluorhídrico (HF) y ácido nítrico (HNO3) en una relación molar de 1,66 en agua. Controlada mediante un mecanismo de difusión, se lleva a cabo la reacción estequiométrica siguiente (67): 3Si + 12HF + 4HNO3 3SiF4 + 4NO + 8H2O Ecuación 1.2.3 Reacción estequiométrica de eliminación de silicio mediante tratamiento ácido. El tiempo de tratamiento ácido determina el silicio residual y la aparición de poros que actúan como núcleos de formación de grietas. Tiempos de tratamiento extensos, mejoran la interconectividad de las muestras (68). Esta reacción también depende de la anisotropía y la porosidad de la muestra, siendo más rápida en la dirección axial y obteniéndose diferentes coeficientes de difusión eficaz en función de la porosidad. 1.2.3.3.3 Temperatura de procesado La temperatura de procesado condiciona el proceso de infiltración de silicio y el tipo de producto final obtenido. A temperaturas alrededor de 1.550 ºC se obtiene la forma β-SiC, mientras que entre 2.200 y 2.500 ºC se producen los politipos hexagonales (αSiC) (11, 66). Se ha demostrado que los bioSiCs producidos a altas temperaturas, son muy densos y tienen una elevada resistencia a la deformación (9). 1.2.3.3.4 Tiempo de reacción El tiempo de reacción contribuye, al igual que la proporción Si/C utilizada, a la modulación de la relación entre el SiC y C en los materiales finales. A medida que el tiempo se incrementa, el porcentaje de carbón residual disminuye, y paralelamente, la porosidad del material y su resistencia a la fractura (26). El tiempo de reacción optimo 107 Cerámicas bioSiCs como nuevos materiales para regeneración ósea depende de la estructura porosa de la preforma de carbón, así como del tamaño del material infiltrado y de la temperatura empleada. 1.2.4 Aplicaciones del carburo de silicio biomórfico Las cerámicas de carburo de silicio biomórfico poseen excelentes propiedades mecánicas y de resistencia a la oxidación, la corrosión, la temperatura y el desgaste que las hacen adecuadas para su empleo en aplicaciones de ingeniería avanzada, así como para el desarrollo de nuevos materiales con fines biomédicos (56). 1.2.4.1 Soportes catalíticos Debido a su estructura porosa, los bioSiCs se han mostrado útiles como soportes catalíticos de níquel para la oxidación parcial de metano en la producción de gas de síntesis y también para la oxidación selectiva de H2S, el tratamiento de gases de escape para motores de trabajo pesado y la deshidrogenación de n-butano. Además, su recubrimiento con zeolita permite incrementar la utilidad para esta aplicación, ya que ésta presenta una elevada superficie específica que contribuye positivamente a los procesos de adsorción, separación y catálisis (69). 1.2.4.2 Refuerzo cerámico en hormigones El refuerzo de hormigones con partículas en forma de aguja de carburo de silicio es una forma efectiva para mejorar sus propiedades mecánicas mostrando, desde un punto de vista mecánico, resultados similares al refuerzo con filamentos metálicos pero con una mayor estabilidad química tras el proceso de curado (56, 70). 1.2.4.3 Aplicaciones de alta temperatura Los bioSiCs se pueden emplear en el desarrollo de motores avanzados y como componentes estructurales de intercambiadores de calor (44, 49). Con el fin de Capítulo 1.2 108 mejorar su conductividad térmica se puede obtener compuestos de cobre-carburo silicio, de forma que se combine, en un único material, la elevada conductividad térmica del cobre y el bajo coeficiente de expansión térmica del carburo de silicio, abriendo nuevas posibilidades para las aplicaciones de gestión térmica (54). La capacidad de las cerámicas de silicio covalentes (SiC, Si3N4...) para formar una capa de óxido superficial (SiO2) con una baja permeabilidad al oxígeno hace posible la obtención de materiales cuya resistencia a la oxidación es la más elevada de entre las cerámicas no oxídicas (41, 44, 49). 1.2.4.4 Sistemas de filtración La microestructura de los carburos de silicio biomórfico los hace adecuados como materiales de filtración para la limpieza de gases en caliente en la industria de generación de energía eléctrica, donde se necesitan como principales características, una baja emisión y una elevada eficiencia. Para estas aplicaciones, se han propuesto como los más adecuados los obtenidos a partir de tablones de densidad media, con tiempos cortos de reacción (71). 1.2.4.5 Aplicaciones biomédicas El carburo de silicio se ha propuesto como material de recubrimiento de dispositivos biomédicos (stents coronarios, válvulas) ya que es capaz de incrementar su hemocompatibilidad mediante la reducción de su trombogenicidad y de la respuesta inflamatoria (72-74). En los últimos años, el carburo de silicio se ha probado como material útil para el diseño y desarrollo de nuevos dispositivos biomédicos como membranas, agentes de diagnóstico por imagen y biosensores. Además, es particularmente interesante su potencial aplicación como biomaterial en implantes ortopédicos (59). 115 Cerámicas bioSiCs como nuevos materiales para regeneración ósea 20. Calderon NR, Martinez-Escandell M, Narciso J, Rodriguez-Reinoso F. 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Lopez-Alvarez M, Gonzalez P, Serra J, de Carlos A, Chiussi S, Leon B. Innovative bioinspired SiC ceramics from vegetable resources. En: Khang G editor. Handbook of Intelligent Scaffold for Tissue Engineering and Regenerative Medicine. Singapore: Pan Stanford Publishing Pte. Ltd.;2012. 77. de Arellano-Lopez AR, Martinez-Fernandez J, Varela-Feria FM, Orlova TS, Goretta KC, Gutierrez-Mora F, et al. Erosion and strength degradation of biomorphic SiC. J Eur Ceram Soc. 2003;24(5):861-70. 78. Klenke FM, Liu Y, Yuan H, Hunziker EB, Siebenrock KA, Hofstetter W. Impact of pore size on the vascularization and osseointegration of ceramic bone substitutes in vivo. J Biomed Mater Res, Part A. 2008;85A(3):777-86. 79. Oh SH, Park IK, Kim JM, Lee JH. In vitro and in vivo characteristics of PCL scaffolds with pore size gradient fabricated by a centrifugation method. Biomaterials. 2007;28(9):1664-71. 80. Lopez-Alvarez M, Pereiro I, Serra J, Gonzalez P, de Carlos A. Porous silicon carbide scaffolds with patterned surfaces obtained from the sea rush juncus maritimus for tissue engineering applications. Int J Appl Ceram Technol. 2012;9(3):486-96. 81. de Carlos A, Borrajo JP, Serra J, Gonzalez P, Liste S, Leon B. In vitro cytotoxicity testing of wood-based biomorphic SiC ceramics. Key Eng Mater. 2005;284-286:581-4. 82. Gonzalez P, Borrajo JP, Serra J, Chiussi S, Leon B, Martinez-Fernandez J, et al. A new generation of bio-derived ceramic materials for medical applications. J Biomed Mater Res, Part A. 2009;88A(3):807-13. 83. Borrajo JP, Gonzalez P, Serra J, Liste S, Chiussi S, Leon B, et al. Cytotoxicity study of biomorphic SiC ceramics coated with bioactive glass. Bol Soc Esp Ceram Vidrio. 2006;45(2):109-14. Capítulo 1.2 122 84. Borrajo JP, Gonzalez P, Serra J, Liste S, Chiussi S, Leon B, et al. Biomorphic silicon carbide ceramics coated with bioactive glass for medical applications. 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Bioactivation of biomorphous silicon carbide bone implants. Acta Biomater. 2010;6(12):4488-94. 90. Filardo G, Kon E, Tampieri A, Cabezas-Rodriguez R, Di Martino A, Fini M, et al. New bio-ceramization processes applied to vegetable hierarchical structures for bone regeneration: An experimental model in sheep. Tissue Eng, Part A. 2014;20(3-4):76373. 91. Arcos D, Vallet-Regi M. Bioceramics for drug delivery. Acta Mater 2013;61(3):890911. Capítulo 2 Objetivos 131 Bio-inspired porous SiC ceramics loaded with vancomycin for preventing MRSA infections 3.1 Abstract Implant-related infections are a serious complication in orthopaedic and dental surgery resulting in prolonged hospitalization, high medical costs and patient mortality. The development of porous implants loaded with antibiotics may enable a local drug delivery for preventing surface colonization and biofilm formation. A new generation of bio-derived porous ceramic material that mimics hierarchical structures from nature was evaluated. Silicon carbide ceramics (bioSiCs) derived from sapelli wood were obtained by pyrolysis of Entandrophragma cylindricum wood followed by infiltration with molten silicon. This process renders disks that keep the bimodal pore size distribution (3 and 85 µm) of the original material and are highly cytocompatible (BALB/3T3 cell line). The ability of the bioceramic to load the antimicrobial agent vancomycin was evaluated by immersion of disks in drug solutions covering a wide range of concentrations. The disks released at pH 7.4 an important amount of drug during the first 2 h (up to 11 mg/g bioSiC) followed by a slower release, which is related to the presence of macro and mesopores. Finally, the antibiofilm effect against methicillin resistant Staphylococcus aureus was assessed and a considerable reduction (92%) of the bacterial film was observed. Results highlight the bioSiC potential as component of medicated medical devices. Capítulo 3 132 3.2 Introduction After roughly 100 years of clinical use of ceramics in dentistry and orthopaedics, there is still a need for novel biomaterials. Chevalier and Gremillard (1), in their recent and extensive review on ceramics for medical applications, stated the imperative need for obtaining novel, tough and stable materials (with special mention to non oxide ceramics as silicon carbide or silicon nitride) as orthopedic material candidates. Additionally, these authors pointed out the interest of developing those new materials through a biomimetic approach. Huebsch and Mooney (2) emphasize that there is a considerable body of research on the importance of physical variables, including topological and mechanical properties of biomaterials, in guiding a biological response. Some materials originated from living organisms, as bones, have outstanding properties due to their inorganic nature but also to their complex structural organization. To achieve a synthetic material that matches to bone, one has to take care of both these aspects. Bones are made up of a collection of materials built out of a common basic building block, the mineralized collagen fibril, that can be arranged in different patterns. All forms of bone possess mechanical strength and toughness out of reach from its constituent materials (3). Numerous attempts have been made to mimic the structure of bone (4–6). Highly porous calcium phosphate ceramic scaffolds or organic–inorganic composites are well known examples that have been used successfully (7), but none of them simultaneously combined the microscopic and macroscopic structure of the bone. Mimicking bone structure continues to be a challenging task. Bio-derived silicon carbide based ceramics (bioSiC), obtained by Si-melt infiltration of carbonaceous scaffolds derived from wood templates have been proposed as new engineering ceramic materials with potential use in biomedical applications (8, 9). As a result of the evolution, wood combines a good balance between resistance/weight ratio and fluids circulation. BioSiC are non oxide ceramics that keep the complex natural 133 Bio-inspired porous SiC ceramics loaded with vancomycin for preventing MRSA infections structural organization of wood, resembling to a certain extent that of bones. Biocompatible bioSiC can be produced with low cost, nearnetshape and adequate mechanical properties (8–10). The possibility of varying the wood cellulosic preforms and/or the process variables makes the bioSiC approach extremely versatile and allows to tailor microstructures resulting in materials with different densities, morphologies, pore size distributions, levels of anisotropy, mechanical strength, etc. (8) and, thus, potentially useful as candidates for the development of orthopedic and dental implants. Bone infections are typically caused by bacteria introduced from trauma, surgery, implant use, or by direct colonization from a proximal infection or via systemic circulation. Postoperative osteomyelitis is still an important problem in orthopedic and dental surgery (11). The bacterial biofilm, extremely resistant to both the immune system and antibiotics, is considered the primary cause of implant-associated infection. During the first 6 h after surgery, an implant is particularly susceptible to surface colonization and biofilm formation (11). Efficiency of the systemic treatment of osteomyelitis is limited by the difficult access of the antibiotic to the infection site (12, 13). The use of implants loaded with antimicrobial agents is a promising approach to prevent post-operative infections (14). The effectiveness of the antibiotic/device combinations is strongly dependent on the mechanism and the rate of drug release (15). Sub-inhibitory drug concentrations (i.e., those below the minimal inhibitory concentration) does not prevent the formation of a microbial biofilm and may even exacerbate complications or induce resistance in wound-site bacteria. An initial ‘‘burst’’ release from the device may sufficiently reduce the likelihood of the primary biofilm infection and, as a consequence, improve the prophylaxis against infection and speed-up the patients recovery. Recent studies have demonstrated that microporous materials may be particularly useful as drug-eluting implants. Porous hydroxyapatite enables more absorption and longer antibacterial activity (up to 2 days) in vitro than dense Capítulo 3 134 hydroxyapatite (only 12 h) (16). On the other hand, mesoporous silica with a highly regular nano-porous structure and a vast surface area provided controlled release and an excellent protection for the loaded guest molecules (17). The aim of the present work was to evaluate the potential of a bioSiC from sapelli wood for the loading and the release of vancomycin in order to prevent local bacterial infections, including biofilms formed by methicillin resistant Staphylococcus aureus (MRSA). To the best of our knowledge the suitability of bioSiC as drug delivery system has not been evaluated yet. Vancomycin is a highly soluble (>100 mg/mL) antimicrobial agent, which is usually administered systemically after bone surgery to prevent bacterial infection because of its broad spectrum and particular efficiency against staphylococci (18, 19). First, the porosity and topography of the bioSiC was characterized in detail in order to confirm the mimicking of the smart hierarchical structure of the tree template. Cytocompatibility and cell conductive properties were then tested. Finally, the ability of bioSiC to load vancomycin and to prevent the biofilm formation was evaluated and related to the particular hierarchical structure of the disks. 3.3 Materials and methods 3.3.1 Bio-inspired silicon carbide Pieces of bioSiC were obtained from transversal cuts of sapelli wood (Entandrophragma cylindricum). The bioceramization process consisted of drying the wood at 60 ºC for 24 h, after which pieces were subjected to a pyrolysis step up to 800 ºC in an inert atmosphere with well controlled heating and cooling ramps. Finally, the carbon perform obtained was infiltrated with molten silicon at 1,550 ºC in vacuum for 30 min (9). The final material was cut to obtain disks of Ø 6 mm x 2 mm. 135 Bio-inspired porous SiC ceramics loaded with vancomycin for preventing MRSA infections 3.3.2 BioSiC characterization The microstructure and topography of bioSiC disks were evaluated by Scanning Electron Microscopy (SEM Philips XL 30), Interferometric Profilometry (WYKO NT1100) and Confocal Laser Scanning Microscopy (CLSM Bio-Rad MRC 1024). The material density was determined, by triplicate, using a helium-air pycnometer (Quantacrome Mod. PY2, USA). The pore size distribution was evaluated by mercury intrusion porosimetry using a Micromeritics Autopore IV 9500 (Norcross, GA, USA) fitted with a 3 mL penetrometer for solids. The working pressures covered the range 0.6–2.5 x 104 psi. The specific surface area was evaluated by the Brunauer–Emmett– Teller (BET) method (20) which involved the determination of the amount of the adsorptive gas (N2 in this case) required to cover the external and the accessible internal pore surface of the material with a complete monolayer. The disks were degassed by heating at 60 ºC and 10-3 mmHg. Then, samples were exposed to N2 gas at 77 K and 0.01–0.98 relative pressure on an automatic surface area analyzer (Micromeritics ASAP 2000, USA). The BET surface area (SBET) was calculated from the isotherms according to the BET equation: SBET (m2/g) = 4.37 Vm (cm3/g), where Vm is the volume of nitrogen necessary to form the monolayer. 3.3.3 Cell viability test The in vitro cytocompatibility of bioSiC disks was tested, in triplicate, by using a BALB/3T3 cell line (CCL 163, ATCC, USA), according to the 10993-5 protocol of the International Standardization Organization (ISO). BioSiC disks were placed in 24-well plates. Then a cell suspension of 200,000 cells/well in 2 mL of DMEM (GIBCO®), supplemented with 10% fetal bovine serum (FBS) and 1% gentamicin, was added into the wells and the plate was incubated at 37 ºC for 24 h in 5% of CO2 and 90% of relative humidity environment. A control (cells without bioSiC disk) was treated in the same way. The bioSiC samples were collected and dyed in order to assess live/ dead Capítulo 3 136 populations by means of a calcein/propidium iodide staining using confocal microscopy (Confocal Spectral Microscopy Leica TCS-SP2 LEICA, Wetzlar, Germany). To calculate their ratio (viability), live and dead populations were counted using a light microscope (Optiphot2, Nikon, Japan) with green and red filters and an Image Analysis software (Soft Imaging System GmbH, Version 3.2 Build 0.607). Cells remaining adhered to the tissue culture polystyrene (TCP) well were trypsinized and centrifuged. The pellet was resuspended with cell culture medium, cytospinned onto a glass slide, and dyed with calcein and propidium iodide. 3.3.4 Vancomycin loading The high solubility in water of vancomycin enabled the loading of this antimicrobial agent into bioSiC disks by a simple immersion in 3 mL of drug aqueous solutions with concentrations ranging from 0.05 to 42.5 mg/mL. The disks were left in the drug solution for 24 h with mechanical shaking. Vacuum (75 mmHg) was applied for the first 2 h to eliminate the air from the pores of the pieces and to promote the flux of drug solution into the disks. The amount of vancomycin loaded in each disk was calculated as the difference between the initial and the final concentrations in the surrounding solution, determined by UV spectrophotometry at 280 nm (Agilent 8453, Böblingen, Germany). All the experiments were carried out in duplicate. Drug-loaded disks were desiccated at 40 ºC until constant weight. BioSiC material loaded with vancomycin was evaluated as it was indicated above. 3.3.5 Vancomycin release Dried drug-loaded disks were transferred to vials containing 3 mL of phosphate buffer (PBS) pH 7.4 at 37 ºC and kept under mechanical shaking. Samples of the release medium were withdrawn at regular intervals and returned to the vial immediately after their drug concentration was measured spectrophotometrically at 280 nm. Water 137 Bio-inspired porous SiC ceramics loaded with vancomycin for preventing MRSA infections uptake by the disks was monitored in parallel to the release tests by placing of dried disks in a Gay-Lussac pycnometer (Afora, Spain) and weighing the pycnometer at different times after filling with water. 3.3.6 MRSA biofilm formation BioSiC material and vancomycin-loaded bioSiC disks, prepared by immersion in 42.5 mg/mL of drug solution during 24 h, were subjected to this microbiological study by using a clinical MRSA isolate. This bacterial isolate (recovered from a patient at the Ghent University Hospital, Ghent, Belgium) was grown on Tryptic Soy Agar (TSA) (Oxoid, Drongen, Belgium) at 37 ºC. MRSA biofilms were formed in two different model systems. First, MRSA biofilms were formed in the Modified Robbins Devices (MRD), as described previously (21). In this system, growth medium is continuously replaced. Secondly, biofilms were formed on drug loaded-bioSiC disks using 24-well microtiter plates (MTP, Trasadingen, Switzerland). To this end, disks were placed in 1 mL MRSA suspensions with a density of appr. 106 CFU/mL (in 1:5 diluted Tryptic Soy Broth, TSB) for 1 h. Subsequently, disks were gently rinsed with 0.9% (w/v) NaCl to remove non-adherent cells and were placed in 1 mL diluted TSB for an additional 24 h at 37 ºC. To quantify the biofilm formation in both methods the MRD and MTP, each disk was transferred to test tubes with 10 mL 0.9% (v/w) NaCl and the tubes were subjected three times to 30 s of sonication (Branson 3510, 42 kHz, 100 W, Branson Ultrasonics Corp., Danbury, USA) and 30 s of vortex mixing to detach the biofilm from the disks. Using this procedure all cells were removed from the disks and clumps of cells were broken apart. Sessile S. aureus cells were platted on TSA, and incubated at 37 ºC for 48 h. Finally, the number of colony forming units (CFU) per disk was calculated by counting colonies on the plates. All experiments were carried out on at least 3 disks for each composition. The Student t test was used to evaluate the efficacy in reducing cell colonies and the differences between MRD and MTP methods. Capítulo 3 138 3.4 Results Processing variables during bioSiC production may affect the interconnected microstructure of the wood template. Morphological characterization of bioSiC was carried out by SEM micrographs (Figure 3.1) that evidenced the particular porous microstructure of the sapelli tree. Macropores (≈ 80 µm) in groups of two or three, characteristic of the sapelli tree (A), can be seen at the surface of the cross section of bioSiC. More in detail (B) a second population of mesopores (less than 10 µm) at the surface of the material can be noticed, with the silicon carbide crystals making the structure up (C). The walls of the vessels were maintained after infiltration giving a ceramic material with pores unidirectional connected, as it can be noted at the longitudinal section (Figure 3.2) characterized by interferometric profilometry (A) and confocal laser scanning microscopy (B). Figure 3.1 SEM micrographs (cross-section) of bioSiC ceramics produced from Sapelli wood. Three magnifications are shown: (A; X100), (B; X500) and (C; X4,000). Microstructure was also confirmed from mercury intrusion porosimetry measurements (Figure 3.3). Results evidence a total intrusion volume of 0.243 ± 0.005 mL/g and a total porosity of 41.49 ± 0.05% with a bimodal pore size distribution (Figure 3.3) including macropores (mean diameter ≈ 85 µm) and mesopores (mean diameter ≈ 3 µm). Quantitative differences in the microstructural properties between vancomycin loaded 139 Bio-inspired porous SiC ceramics loaded with vancomycin for preventing MRSA infections and unloaded bioSiC disks could not be established using mercury intrusion porosimetry. Figure 3.2 Topographic characteristics of a bioSiC ceramic (longitudinal section) by interferometric profilometry (A) and confocal laser scanning microscopy (B). The nitrogen adsorption analysis confirmed the absence of microporosity and allowed the estimation of specific surface (1.198 ± 0.005 m2/g). The ability of bioSiC disks to sustain cell attachment and growth was assessed by an in vitro biocompatibility test following ISO 10993-5 procedure. The assays were conducted using Balb cell line. After 24 h in contact, a homogeneous well distributed layer of living cells and a significantly smaller amount of dead cells on the top of the bioSiC disks can be observed (Figure 3.4). The ratio between living and dead cells was used to calculate the viability of cells on the disk. The obtained values were 79.9% (sd 13.1) on bioSiC disks and 98.0% (sd 1.9) on the TCP from the wells. Vancomycin loaded bioSiC were also tested to assure in vitro biocompatibility achieving a percentage of cell viability of 95.8% (sd 6.0). Capítulo 3 140 Figure 3.3 Pore size distribution and percentage of porosity of bioSiC derived from sapelli wood. Figure 3.5 shows the adsorption curve of vancomycin on bioSiC material which relates the concentration of the solute on the adsorbent (mg vancomycin per gram of dried material) to the concentration of the solute in the surrounding solution at the equilibrium. The shape of the adsorption curve can be classified as Class S according to Giles and coworker classification (22). At the highest loading concentration studied (6 mg/mL) a significant amount of vancomycin (nearly 50 mg/g of bioSiC) was loaded. Figure 3.4 Calcein-propidium iodide staining observed by confocal microscopy 24 h after seeding on bioSiC samples. Living cells in green and dead cells in red. Capítulo 7 Alginate-poloxamer-silicon carbide composites for the controlled release of indomethacin decrease in vitro inflammation 245 Alginate-poloxamer-SiC composites for indomethacin controlled release decrease inflammation 7.1 Abstract Composites of biomorphic silicon carbides (bioSiCs) and hydrogels are proposed in order to obtain materials able to load and release poor soluble drugs with application in bone pathologies therapy. Hydrogels composed by alginate and poloxamer were loaded with indomethacin, incorporated into the ceramics and crosslinked. The indomethacin release profile is dependent on the microstructure of the bioSiC selected. The loaded oak and sapelli bioSiCs composites have adequate release profiles to promote the decreasing of the secretion of pro-inflammatory cytokines in LPS stimulated macrophages, showing stronger anti-inflammatory effects than pine bioSiC composites. The released indomethacin is able to modulate the degradation of chondrocytes extracellular matrix and promote the formation of new collagen. Particles derived from mechanical wear of biomorphic silicon carbides do not show high toxicity, being similar to the zirconia particles. Capítulo 7 246 7.2 Introduction Bone is a complex hierarchically organized tissue formed by an organic matrix, mainly collagen which is sequential mineralized with an inorganic component, hydroxyapatite (HAp), by the action of specific cells (1, 2). Due to this complex structure the development of new bone biomimetic materials has led to the production of new porous 3D composite systems formed by the combination of one or more than one organic component (natural or synthetic polymers) with an inorganic component. Composite systems should be able to meet all the physical and biological requirements for bone regeneration, combining the advantages of all components (3). These composite systems have been found to be adequate also for the release of growth factors as BMP-2 (4-6), VEGF (7), platelet derived growth factors (8, 9) or drugs such as dexamethasone (10, 11), vancomycin (12, 13) and gentamicin (14). In the last few years, they have been also proposed for releasing a combination of therapeutic molecules such as BMP-2 and vancomycin (15), amikacin and gentamicin (16) or BMP-2 and VEGF (17). This new approach makes them an attractive alternative for the treatment of several bone pathologies. Biomorphic silicon carbide ceramics (bioSiCs) obtained from natural resources (18) have been shown to maintain the original structure of their precursors being highly porous and biocompatible materials (19) suitable for the regeneration and revascularization of tissues. The incorporation of a polymer component into their structure must increase their therapeutic value. The use of ionic crosslinking polymers, such as alginate makes it possible to easily obtain a three dimensional network polymer through the addition of divalent ions (20). Alginate hydrogels can exert the function of an organic matrix suitable for cellular growth and encapsulation within the ceramic system and also facilitate the incorporation of drugs and growth factors into its three dimensional structure while modulating their release. Therefore, the combination of 247 Alginate-poloxamer-SiC composites for indomethacin controlled release decrease inflammation alginate hydrogels and bioSiCs can be presented as a promising strategy in the development of complex systems for tissue regeneration and controlled drug release. The addition of synthetic block copolymers such as poloxamers or poloxamines to the hydrogel component should be able to increase the solubility of low aqueous solubility drugs. The ability of these polymers to form micelles in an aqueous solution, where hydrophobic drugs can be incorporated, leads to an increase on their solubility and therefore makes their administration possible (21, 22). Furthermore, it was observed than poloxamines were able to stimulate the osteoblastic differentiation of adipose derived mesenchymal stem cells by themselves (23). The use of poloxamers has also been found to inhibit P-glycoprotein function decreasing the resistance of multidrug resistant (MDR) cell lines (24, 25). Currently, the long-term stability of the prosthesis continues to be a challenge in the development of bone substitutes. Material wear debris is one of the main drawbacks associated to the use of artificial materials as implantable systems. The production of material particles may cause aseptic loosening and the activation of the surrounding macrophages leading to bone destruction and periprosthetic osteolysis (26-28). Several parameters modulate the inflammatory reaction caused by material particles such as their size (29) or their chemical composition (30, 31). Despite the fact that biomorphic silicon carbide has been previously described as a highly biocompatible material the inflammatory reactions caused by its potential debris have not been documented. The aim of the present work is to develop composites able to load by entrapment and release an anti-inflammatory drug, indomethacin. The development of composite systems formed by a natural (alginate) and a synthetic polymer (poloxamer) together with the biomorphic ceramic (silicon carbide) should allow us to obtain an implant material suitable to load and release, in a controlled way, the indomethacin with an Capítulo 7 248 amount enough to show an adequate anti-inflammatory effect on osteoarthritic chondrocytes. Additionally, the potential toxic effects of particles obtained by mechanical wear of biomorphic silicon carbides are evaluated. 7.3 Materials and methods 7.3.1 Preparation of composite systems Three different types of biomorphic silicon carbide (bioSiC) samples were obtained as described elsewhere from oak (Quercus robur), pine (Pinus pinnaster) and sapelli (Enthandrophragma cylindricum). Disks of 6 mm in diameter were sterilized by autoclaving at 121 ºC for 20 min (32). Two polymeric components were used; Poloxamer 407 (Pluronic F127® (PF127)) that was kindly donated from BASF (Ludwigshafen, Germany) and sodium alginate (GRINDSTED® Alginate PH 155) was purchased from Danisco (Copenhagen, Denmark). Pluronic was dissolved in phosphate buffer (PBS) to achieve a final concentration of 2.5%. After its complete dissolution, indomethacin (2.38 mg/mL) and alginate (2%) were sequentially added. The final solution was autoclaved at 121 ºC for 20 min. 30 µL of the polymeric solution was added on the bioSiC samples and crosslinked by the immersion of the loaded sample into a sterile solution of calcium chloride (Panreac; Barcelona, Spain) at a concentration of 20% for 10 seconds. After the crosslinking, systems were washed twice with 2 mL of PBS for ten seconds. Crosslinked alginate-poloxamer beads prepared by dropping the polymeric solution into the calcium chloride solution were used as control. 249 Alginate-poloxamer-SiC composites for indomethacin controlled release decrease inflammation 7.3.2 Characterization of the polymeric component The mechanical stability of the crosslinked and uncrosslinked polymeric systems was analyzed before and after the autoclaving process using a controlled stress rheometer (Rheolyst AR-1000N TA instruments, Surrey, UK). Ramps of temperature from 15 °C to 60 °C at 2 °C/min with an oscillatory stress of 0.1 Pa at 5 rad/s were carried out for all the samples. The differential scanning calorimetric (DSC Q200, TA instruments, Surrey, UK) was used to evaluate the potential degradation of the polymeric chains during the autoclaving procedure. Ramps of temperature were carried out first from room temperature to -30 °C at 10 °C/min and then from this temperature to 50 °C at 10 °C/min. 7.3.3 Isolation of human osteoarthritic chondrocytes Human osteoarthritic cartilages were provided by the Instituto de Ortopedia y Banco de Tejidos Musculoesqueléticos of the University of Santiago de Compostela. Pieces of the tissue were cut and placed into sterile tubes with trypsin and cell culture medium. The tissue was maintained in the solution at 37 ºC for 30 min in order to kill the fibroblasts present in the extracts. Then, tissue samples were immersed in collagenase 1.5% in culture medium and kept overnight at 37 ºC under mechanical stirring. The solutions were centrifuged at 1,056 g for 4 min. Cells were resuspended in DMEM supplemented with 10% of fetal bovine serum and 1% penicillin/streptomycin and cultured at 37 ºC with 5% of CO2 and 90% of relative humidity. Capítulo 7 250 7.3.4 In vitro release of indomethacin Loaded composite systems were immersed in 3 mL of phosphate buffer at 37 ºC. At preset times the concentration of indomethacin was quantified by UV-visible spectrophotometry at 320 nm. All the experiments were carried out in triplicate. 7.3.5 Anti-inflammatory effect of indomethacin loaded composites The anti-inflammatory effects of loaded indomethacin composites were evaluated in two cell types, extracted human osteoarthritic chondrocytes and a murine macrophage cell line (Raw 264.7). The macrophage cell line was cultured in DMEM-F12 HAM supplemented with 10% FBS and 1% penicillin/streptomycin and maintained at 37 ºC with 5% of CO2 and 90% relative humidity. Loaded composite systems were placed in 24-well plates, cultured with 100,000 cells per well and stimulated with lipopolysaccharide (LPS) at a concentration of 100 ng/mL. The composites anti-inflammatory effect was evaluated after 24 and 72 hours of culture by the quantification of prostaglandin E2 (PGE2) (Arbor), TNF-α (eBioScience), nitric oxide (Cayman) and IL-1α (eBioScience). The composites cytotoxicity was analyzed by the quantification of lactate dehydrogenase (LDH) (Roche). Indomethacin at 100 µM, unloaded silicon carbide samples, and the polymeric solution were used as controls. Composites were also cultured with osteoarthritic chondrocytes in 24-well plates at a density of 60,000 cells per well with 2 mL of supplemented DMEM. The effect on extracellular matrix synthesis of osteoarthritic chondrocytes was evaluated by the quantification of glycosaminoglycans (GAGs) and collagen (I-V) production after 15 days of culture. 251 Alginate-poloxamer-SiC composites for indomethacin controlled release decrease inflammation Cell culture supernatants were centrifuged at 10,000 g for 10 min. Two colorimetric assays were used, Blyscan (Biocolor) for GAGs and Sircol Collagen Assay kit (Biocolor) for collagen. Single cell culture medium was used as negative control and cells treated with indomethacin at a concentration of 100 µM equivalent to 100% of drug release were used as a positive control. Unloaded bioSiCs and crosslinked loaded and unloaded polymeric components were also used as controls. Additionally, the concentration of inflammatory cytokine (IL-1β) and the secreted PGE2 were measured at 1, 2, 5 and 15 days through an ELISA assay (BenderMedSystem) and a commercial colorimetric assay (Arbor) respectively. 7.3.6 Evaluation of inflammation caused by silicon carbide released particles Particles were obtained by direct mechanical friction of two oak silicon carbide samples. One of the samples was immobilized on the surface of a sterile plastic container whereas the other was fixed on an impeller attached to a rotor (Ika RW20DZM) rotating at 150 rpm for one week. Ultrapure water (milliQ) was added to the container in order to facilitate the recovery of the particles. Particle size was analyzed by Scanning Electron Microscopy (ZEISS EVO LS 15, Germany) and their composition by EDX using the same equipment. Particles size distribution was analyzed using a zetasizer (Zetasizer Nano ZSP, Malvern, United Kingdom). The inflammation caused by the particles was evaluated using a macrophage cell line (Raw 264.7) and cultured as previously described. Cells were seeded in 24-well plates with 1 mL of cell culture medium and particles were 10-fold diluted and added to the wells. Cells stimulated with 1 µg/mL of LPS were used as a positive control of inflammation and commercial Zirconium (IV) oxide particles as reference material (Sigma, USA). Capítulo 7 252 Inflammatory cytokines were quantified after 24 hours of culture, IL-1α and TNF-α (Bender MedSystems). 7.3.7 Statistical Analysis Experimental results were expressed as means and standard deviations. The statistical significant differences between treatments were established by the analysis of variance (ANOVA). When the F-ratio suggests significant differences between groups, the least significant difference (LSD) test was used to compare them in pairs. ANOVA and LSD were performed by Statgraphics Centurion®X64 software (USA). 7.4 Results and discussion 7.4.1 Characterization of the polymeric component A strong interaction could be observed between alginate and poloxamer, this interaction is modified by the autoclaving process. Figure 7.1 shows the dynamic and elastic moduli of both, Alg and Alg-PF127 formulations before (Figure 7.1A) and after (Figure 7.1B) the autoclaving process. The addition of poloxamer to alginate solution caused an increase in both moduli and consequently in the complex viscosity at 37 ºC (from 1.94 to 3.12 Pa.s) (Figure 7.1A). The sterilization process led to a markedly decrease in the complex viscosity values at 37 ºC, being 0.06 Pa.s and 0.04 Pa.s for the Alg and Alg-PF127 formulations respectively. After this process the elastic modulus could not be detected. Interestingly, whereas in unsterilized systems the incorporation of poloxamer increased the complex viscosity, after autoclaving these seem to have the opposite effect. Complex interactions were established between alginate chains and poloxamer micelles in the unsterilized systems. Polysaccharide chains could be placed around the micelles hampering the interaction between poloxamer and water and increasing micelle-micelle interactions and therefore, the viscosity of the system (33). In the same way, Lin and coworkers found that the addition of alginate to poloxamer 259 Alginate-poloxamer-SiC composites for indomethacin controlled release decrease inflammation The controlled release of indomethacin by the composite systems, particularly those produced using oak and sapelli bioSiC, promote the highest anti-inflammatory effect according to all the analyzed cytokines. Figure 7.5 Levels of NO and PGE2 secreted by LPS stimulated macrophages after 24 and 72 of cell culture. (C-) is macrophages not stimulated and (C+) is macrophages stimulated with LPS without any treatment. The homogeneous groups are indicated by an equal number of asterisks (*) above the columns (α < 0.05). Osteoarthritis (OA) is a non-inflammatory joint disease and the most common form of arthritis. It is characterized by the loss of cartilage from the articulating surfaces, ostophyte formation, changes in the synovial membrane, subchondral bone sclerosis and an increased volume of synovial fluid with less viscosity and poor lubrication properties (40, 41). Its current treatment is symptomatic, including intra-articular injections of glucocorticoids and hyaluronic acid formulations or the administration of non-steroidal anti-inflammatory drugs (NSAIDs) (40, 42). It is also well known that an imbalance between its metabolism and degradative signals is present in osteoarthritic cartilage. Osteoarthritic chondrocytes increase the secretion of inflammatory cytokines loaded pine SiC loaded sapelli SiC loaded oak SiC pine SiC sapelli SiC oak SiC PF127+alginate+ind Ind CC+ 0 50 100 150 200 250 300 350 400 ** * * *** PGE2 (pg/mL) 24h 72h * loaded pine SiC loaded sapelli SiC loaded oak SiC pine SiC sapelli SiC oak SiC PF127+alginate PF127+alginate+ind Ind CC+ 0 10 20 30 40 50 60 70 80 90 100 110 120 NO (mM) 24h 72h * A A B Capítulo 7 260 and decrease the collagen synthesis (43). It has been reported that the secretion of prostaglandin E2 is enhanced in osteoarthritic cartilage, synovium and synovial fluid (38). Cell viability and the secretion of proinflammatory cytokines after 15 days of cell culture have been analyzed for osteoarthritic chondrocytes treated with the different formulations. Cell viability results, all higher than 60%, (Figure 7.6), show no statistically significant differences between indomethacin alone and drug loaded systems. The addition of the drug to the cell culture medium, shows a slight toxicity and promotes a decrease in cell viability. loaded pine SiC loaded sapelli SiC loaded oak SiC pine SiC sapelli SiC oak SiC Ind 0 20 40 60 80 100 120 cell viability (%) 15 days Figure 7.6 Cell viability values obtained by MTT assay after 15 days of cell culture corrected by the negative control (culture plates). No statistical significant differences were observed (α < 0.05). The levels of NO and PGE2 secreted by osteoarthritic chondrocytes after 48 hour of treatment with the different formulations (Figure 7.7) point out that loaded composites promote a significant decrease in cytokines production in comparison to untreated chondrocytes (C-). 261 Alginate-poloxamer-SiC composites for indomethacin controlled release decrease inflammation Figure 7.7 Levels of NO and PGE2 secreted by osteoarthritic chondrocytes after 48 hours of cell culture. The negative control (C-) is osteoarthritic chondrocytes without treatment. The homogeneous groups are indicated by an equal number of asterisks (*) above the columns (α < 0.05). The secretion of IL-1β induces the catabolic cascade of osteoarthritic chondrocytes including the activity of cyclooxygenase enzymes and, therefore, the production of PGE2, the main pro-inflammatory factor (38). The Levels of IL-1β were undetectable for all the treatments studied. In order to evaluate the extracellular matrix catabolism and metabolism, the concentration of the main structural macromolecules, glycosaminoglycans (GAGs) and collagen I-V, in the cell culture medium after fifteen days of assay have been quantified. Figure 7.8 shows the not statistically significant differences between materials with regard to the amount of released GAGs and synthetized collagen for both loaded and unloaded composites. The release of indomethacin from loaded composites promotes a significant decrease in GAGs together with a significant increase in collagen production from the osteoarthritic chondrocytes which can be explained by the loaded pine SiC loaded sapelli SiC loaded oak SiC pine SiC sapelli SiC oak SiC PF127+alginate+ind CInd 0 50 100 150 200 250 300 350 400 48h PGE2 (pg/mL) loaded pine SiC loaded sapelli SiC loaded oak SiC pine SiC sapelli SiC oak SiC PF127+alginate+ind CInd 0 10 20 30 40 50 60 * * NO (mM) 48h * A B Capítulo 7 262 decreasing of their anabolic activity, leading to the synthesis of new collagen molecules and to the reduction of the digestion of the extracellular matrix. Figure 7.8 Levels of GAGs (A) and collagen I-V (B) secreted by osteoarthritic chondrocytes after 15 days of cell culture. The negative control (C-) is osteoarthritic chondrocytes without any treatment. 7.4.4 Evaluation of inflammation caused by silicon carbide released particles Despite the excellent tribological properties of silicon carbides and their high corrosion resistance under normal biological conditions (26), it is relevant to evaluate the inflammation and macrophage activation caused by their mechanical wear debris after their hypothetical in vivo implantation. The simulation of mechanical wear between two silicon carbide samples allowed the obtaining of small particles characterized by a mean diameter of 481.33 nm to (Figure 7.9). The analysis of particle composition show that they were mainly composed by silicon and carbon atoms according to EDX measurements. loaded pine SiC loaded sapelli SiC loaded oak SiC pine SiC sapelli SiC oak SiC PF127+alginate+Ind cc+ 0.00 0.05 0.10 0.15 0.20 0.25 0.30 GAGs (mg/mL) 15 days loaded pine SiC loaded sapelli SiC loaded oak SiC pine SiC sapelli SiC oak SiC PF127+alginate+Ind cInd 0.0 0.5 1.0 1.5 2.0 2.5 3.0 3.5 4.0 4.5 5.0 5.5 Collagen I-V (mg/mL) 15 days A B 263 Alginate-poloxamer-SiC composites for indomethacin controlled release decrease inflammation 100 1000 10000 0 2 4 6 8 10 12 14 Mean Intensity (%) Size (nm) Figure 7.9 Particle size distribution obtained using a particle size analyzer at room temperature. Macrophages were able to uptake wear debris from the medium as can be seen in Figure 7.10 where the particles are present inside the cytoplasm of the cells (black points) while the nuclei remain transparent. Figure 7.10 Macrophage cellular uptake of biomorphic silicon carbide particles after 24 hours of cell culture (B) compared to negative control (A) at 20X. The activation of macrophages causes the secretion of numerous pro-inflammatory cytokines that cause the inflammatory cell recruitment as IL-1β, TNF-α, IL-6, and IL-8 A B Capítulo 7 264 (44). Two cytokines were selected to evaluate the activation of macrophages; tumor necrosis factor alpha (TNF-α) that had been found to be crucial mediating inflammatory cell recruitment derived from particulate systems, (45) and interleukin-1 (IL-1) that stimulates bone resorption and the consequent periprosthetic osteolysis (30). The levels of TNF-α and IL-1 were compared with the cytokines secreted after a similar treatment with zirconia particles. Figure 7.11 Levels of tumor necrosis alpha (TNF-α) and interleukin 1 beta (IL-1β) secreted by the murine macrophage cell line compared to positive control (cells stimulated with lipopolisaccharide). The homogeneous groups are indicated by an equal number of asterisks (*) above the columns (α < 0.05). After the addition of particles obtained by mechanical wear similar levels of TNF-α secretion than for zirconia particles (46) were found. Both levels were significantly lower than those promoted by the positive control (Figure 7.11). On the other hand the secretion of IL-1β was quantitatively similar to the one promoted by the positive control and the zirconia particles. It was found that the addition of lipopolisaccharide (control+) was not able to stimulate the production of IL-1β. It has been shown that zirconia particles show similar levels of inflammation to alumina with better results SiC particles Zirc particles C+ 0 1000 2000 3000 4000 5000 24h TNF-(pg/mL) * SiC particles Zirc particles C+ 0 10 20 30 40 50 60 70 IL-1 (pg/mL) 24h A B 265 Alginate-poloxamer-SiC composites for indomethacin controlled release decrease inflammation than polymers like high-density polyethylene (46). Although the presence of zirconia particles may produce some adverse reactions, they are less intense than those promoted by titanium particles, therefore zirconia is considered the most suitable biomaterial in terms of inflammation promoted by material debris (47). Similar results were obtained for silicon carbide and zirconia particles showing the acceptable biocompatibility of our biomorphic silicon carbide particles. 7.5 Conclusions New composite systems including biphasic alginate-poloxamer hydrogels and bioSiCs from different precursors have been developed and analyzed with regard to their capability of loading and control indomethacin release. All the ceramic-hydrogel composites have shown high biocompatibility. The loaded oak and sapelli bioSiCs composites had adequate release profiles able to promote the decreasing of the proinflammatory cytokines secretion in LPS stimulated macrophages, showing stronger anti-inflammatory effects than pine bioSiC composites. Those differences could not be observed when treating osteoarthritic chondrocytes. In those cases the indomethacin released from the composites was also able to modulate the degradation of chondrocyte extracellular matrix and promote the formation of new collagen. Particles derived from wear debris of biomorphic silicon carbide did not show high toxicity, being similar to the zirconia particles. These developed composites present great potential for the local treatment of bone pathologies. 7.6 References 1. Raghavan RN, Muthukumar T, Somanathan N, Sastry TP. Biomimetic mineralization of novel silane crosslinked collagen. Mater Sci Eng, C. 2013;33(4):1983-8. Capítulo 7 266 2. Silverman LD, Lukashova L, Herman OT, Lane JM, Boskey AL. Release of gentamicin from a tricalcium phosphate bone implant. J Orthop Res. 2007;25(1):23-9. 3. Dessi M, Borzacchiello A, Mohamed THA, Abdel-Fattah WI, Ambrosio L. Novel biomimetic thermosensitive β-tricalcium phosphate/chitosan-based hydrogels for bone tissue engineering. J Biomed Mater Res, Part A. 2013;101A(10):2984-93. 4. Hernandez A, Sanchez E, Soriano I, Reyes R, Delgado A, Evora C. Material-related effects of BMP-2 delivery systems on bone regeneration. Acta Biomater. 2012;8(2):78191. 5. 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Prop Appl Silicon Carbide. 2011:297-343. 275 Controlled release of rAAV vectors from alginate-poloxamer-SiC composite systems 8.1 Abstract Viral vectors are commonly used to deliver a gene into a specific cell type in order to restore its normal protein expression. The inclusion of viral vectors into implantable systems is a promising alternative to the conventional therapy. However, little is known about the requirements of release systems to successfully include viral vectors and achieve a suitable transgenic expression. This work is focused on the development of different hydrogel structures, able to include recombinant adeno-associated viral vectors (rAAV) and promote their controlled release in order to elucidate the best conditions to achieve the highest transduction efficiency. The inclusion of rAAV loaded hydrogels into silicon carbide ceramics allow the obtaining of implantable systems for the release of viral vectors for local administration. The transduction efficiencies of the loaded hydrogels systems were in agreement with their in vitro release profiles. The addition of poloxamer to the systems was able to enhance viral vectors transduction efficiencies after one day of study. The incorporation of loaded hydrogels into bioSiCs has not been able to obtain adequate transduction efficiencies, being required additional studies. Capítulo 8 276 8.2 Introduction Gene therapy is an attractive approach for the treatment of numerous diseases due to its ability to stably introduce a functional gene into a target cell, allowing for the sustained production of a therapeutic candidate molecule. This can be performed either via direct injection of DNA, by encapsulation of DNA in cationic lipids or polymers, or through viral gene delivery systems (1, 2). Several non-integrating and integrating viral vectors are available to achieve this goal. Non-integrating vectors such as those derived from adenoviruses or the herpes simplex virus (1, 3) are of relatively limited clinical interest due to the initiation of deleterious, virus-specific immune and/or toxic reactions by the host. The development of adequate delivery systems is thus necessary in order to ensure that safe gene expression may occur at appropriate levels over extended periods of time using such vectors (3, 4). The most commonly used viral vectors are based on retroviral and recombinant adeno-associated viral (rAAV) vectors (1). Administration of rAAV vectors that derive from the human nonpathogenic AAV virus is the goal of human gene therapy. They have been found to be most adequate as they can transduce both dividing and nondividing cells in marked contrast with the retroviral vectors that can only modify dividing cells, carrying the additional risk of insertional mutagenesis. Furthermore, rAAV allow for the direct, effective transduction in animal models with low immunogenicity, being suitable for the treatment of human pathologies, especially those that affect the articular cartilage (traumatic defects, osteoarthritis) (5, 6). Transduction efficiencies up to 80% have been reported in articular chondrocytes in vitro, in situ, and in vivo using reporter but also therapeutic candidate genes including the insulin-like growth factor I (IGF-I), transforming growth factor beta (TGF-β), and fibroblast growth factor 2 (FGF-2), restoring a close to normal metabolic balance in osteoarthritic cartilage (7, 8). rAAV have been also successful to activate the processes of 277 Controlled release of rAAV vectors from alginate-poloxamer-SiC composite systems chondrogenic differentiation in human bone marrow derived mesenchymal stem cells (hMSCs) in vitro upon gene transfer of the cartilage-specific transcription factor SOX9 (9) and to enhance the healing of osteochondral defects following direct administration of the same vector construct (10). Overall, these findings suggest that this vector type might be the most adapted gene vehicle to elaborate future, effective treatments against cartilage injuries in patients. While systemic injection of rAAV vectors can be performed in human subjects, the high amounts of vectors generally needed to achieve a therapeutic effect (~ 5 x 1013 vg/kg) severely restrict their use in the clinics via this route of administration (11). Several approaches have been explored to improve the delivery of such vectors by including them into polymeric systems through different technological approaches as a means to increase their stability while decreasing potential immune responses that may be raised against the viral capsids (3, 12, 13). The ideal system would increase the levels and duration of transgene expression and improve the safety of the gene transfer system through controlled release of the vectors, ensuring longer residence time by reduction of the viral clearance (3, 14). To achieve this goal, several natural and synthetic polymeric systems like fibrin, gelatin, collagen, poly(ethylene glycol) (PEG), and agarose have been tested (15, 16). Parenteral administration or co-administration of cationic polymers such as poly-L-lysine or polyarginine have been also described, allowing to enhance the efficacy of rAAV gene delivery. Such polymers may affect the permissivity of the targets to the vector particle either by modifying the charge of the capsid or by increasing the interactions between the viral particle and its cell membrane receptor, a heparan sulfate proteoglycan (11). Despite recent advances, the development of controlled viral vector release systems remains challenging. The high size of viral particles generally impairs their release from the most conventional pharmaceutical technology approaches. Some studies showed the ability of poloxamers like Pluronic F68® to increase rAAV-mediated transgene Capítulo 8 278 expression in different tissues like adipose tissue (17). Biomorphic silicon carbide (bioSiC) ceramics have been reported for their excellent mechanical properties, with an adequate porosity and pore interconnectivity to promote bone regeneration and a good biocompatibility to induce the osteoblastic differentiation of hMSCs (18). To our best knowledge, such compounds have never been examined for their ability to release rAAV vectors as a means to improve the vector delivery process for cartilage repair. The goal of the present work was therefore to produce porous alginate-poloxamer systems in a suitable network structure that is capable of releasing such promising, clinically relevant gene delivery vectors for optimal transduction efficacy of reporter genes in hMSCs. The inclusion of the optimized hydrogel into a porous bioSiC ceramic may provide adequate biphasic systems for future applications to treat articular cartilage defects in patients. 8.3 Materials and methods 8.3.1 Reagents Sodium alginate (GRINDSTED® AlgPH155) was purchased at Danisco (Copenhagen, Denmark). Poloxamer 407 (Pluronic F127® (PF127)) was kindly provided by BASF (Ludwigshafen, Germany). Biomorphic silicon carbide (bioSiC) samples were obtained from sapelli wood (Enthandrophragma cylindricum) as previously reported (19). The Cell Proliferation Reagent WST-1 and β-gal Staining Set were from Roche Applied Science (Mannheim, Germany). The Beta-Glo® Assay System was from Promega (Mannheim, Germany). 8.3.2 Cells Human bone marrow derived mesenchymal stem cells (hMSCs) were prepared from bone marrow aspirates obtained from the distal femurs of patients undergoing total knee arthroplasty (n = 5). The study was approved by the Ethics Committee of the 279 Controlled release of rAAV vectors from alginate-poloxamer-SiC composite systems Saarland Physicians Council. All patients provided informed consent before inclusion in the study. All procedures were performed in accordance with the Helsinky Declaration. Cells were isolated, expanded in DMEM, 100 U/mL penicillin and 100 µL/mL streptomycin, 10% FBS (growth medium), and characterized for cell surface markers and multilineage potential as previously described (9, 20). Culture medium was replaced every 2-3 days. Cells at passage 1-2 were used for the experiments. 8.3.3 rAAV plasmids and vectors The constructs were derived from pSSV9, an AAV-2 genomic clone (21, 22). rAAVlacZ carries the lacZ gene for E. coli β-galactosidase and rAAV-RFP a Discosoma sp. red fluorescent protein (RFP) cDNA fragment, both under the control of the cytomegalovirus immediate-early (CMV-IE) promoter (5, 8-10, 23). The vectors were packaged as conventional (not self-complementary) vectors using a helper-free, twoplasmid transfection system in the 293 cell line (an adenovirus-transformed human embryonic kidney cell line) with the packaging plasmid pXX2 and the Adenovirus helper plasmid pXX6 as previously described (8). The vector preparations were purified by dialysis and titered by real-time PCR (5, 8-10, 23), avering 1010 transgene copies/mL. 8.3.4 Capsule preparation and characterization In order to achieve adequate vector stability and release profiles, three different conditions were selected from the initial conditions studied. Capsules were prepared with 0.3% alginate (AlgPH155) or with alginate containing 9% PF127 (AlgPH155 + PF127) in a solution of 10% sucrose. In the latter case, the effects of temperature upon viscosity and final structure of the combination containing a thermosensitive polymer were examined by crosslinking the systems at room temperature (AlgPH155 + PF127 [C]) or at 50 °C (AlgPH155 + PF127 [H]). Unloaded and rAAV vector-loaded capsules Capítulo 8 280 were prepared by dropping the dispersion of polymers alone or containing rAAV to a calcium chloride solution (102 mM) in a 2 mL tube using a syringe with a needle of 18 ½ G. The crosslinking procedure was performed at room temperature or at 50 °C for a brief period of time (30 sec) to avoid irreversible crosslinking of all the polymeric bead. Capsules were then kept in culture in 96-well plates. Polymeric dispersions of AlgPH155 and AlgPH155 + PF127 in sucrose 10% were prepared by direct dissolution and used as controls. The rheological properties of the polymeric dispersions were evaluated using a rheometer (Rheolyst AR-1000N TA instruments, UK) equipped with a Peltier plate for temperature control and a cone-plate geometry (60-mm diameter with an angle of 1.58°, gap 59 µm). Ramps of temperature from 15 °C to 60 °C at 2 °C/min with an oscillatory stress of 0.1 Pa at 5 rad/sec were carried out. Gel temperature (Tgel) was estimated from the cross point between the storage moduli (G´) and loss moduli (G´´). 8.3.5 rAAV vector encapsulation efficiency The ability of the capsules to entrap rAAV during crosslinking was evaluated by measuring the number of viral particles in each bead. The capsules were placed in 50 µL of bead dissolution medium (55 mM sodium citrate, 0.15 M sodium chloride, 30 mM EDTA) (24), vortexed for 2 min followed by the addition of 50 µL viral dilution buffer (10% SDS, 1 M Tris pH 7.5, 0.5 M EDTA) and incubated for 10 min at 56 °C. After a quick spin, the vector concentrations were measured spectrophotometrically at 260 nm (VPcapsule). The initial amount of viral particles in the loading solution was also evaluated (VPloading) and used as 100% of rAAV vector encapsulation efficiency (EE) calculated as: EE (%) = [VPcapsule/VPloading] x 100 281 Controlled release of rAAV vectors from alginate-poloxamer-SiC composite systems 8.3.6 Capsule stability and degradation The stability of the vector-loaded and unloaded capsules was evaluated by testing their weight and size upon immersion in cell culture medium in the presence or absence of hMSCs. Each type of bead was placed in a well of 96-well plates containing 150 µL of growth medium. At preset times (0, 1, 3, 5, 10 and 21 days), each capsule was weighted and the mean diameters and total areas were measured under light microscopy (Olympus BX45, Hamburg, Germany) using the analySIS® program (Olympus). Similarly, stability against dilution was analyzed by modifying the medium volume where the capsules were immersed. The results were expressed as the percentage of area loss or weight loss of the initial area and weight, respectively. 8.3.7 rAAV vector release from the capsules The vector release profiles were obtained by placing the capsules in 96-well plates containing 150 µL of growth medium. At preset times (1, 3, 5, 10 and 21 days), the release medium was removed and replaced by fresh medium. The number of viral particles in the removed medium at the selected time points was estimated by spectrophotometry as described above. Unloaded capsules were used as negative controls and loading solutions as positive controls. After 21 days, the capsules were placed in 50 µL of bead dissolution medium and the number of remaining viral particles was estimated in a similar way. 8.3.8 Cell viability hMSC viability was estimated by placing rAAV-lacZ-loaded or unloaded capsules in contact with cells in monolayer culture (3,500 cells/well in 96-well plates) for 1, 3, 5, 10 or 21 days using the Cell Proliferation Reagent WST-1, with OD proportional to the cell numbers as previously described (9). Equivalent vector solutions with or without alginate and poloxamer were used as controls. Capítulo 8 282 8.3.9 Transduction efficiencies from released rAAV vectors In order to analyze the potential effects of alginate and poloxamer on the rAAV transduction efficiencies, hMSCs in monolayer culture (7,500 cells/well in 96-well plates) were treated with rAAV-lacZ solutions with and without equivalent concentrations of polymers used to prepare the beads. After 24 h, the efficiencies were measured from the luminescence produced after 30 min of incubation with the BetaGlo® assay according to the manufacturer’s recommendations. The values obtained were expressed as Relative Luminescence Units (RLUs) normalized to the cell numbers as determined by using the Cell Proliferation Reagent WST-1 (9). The transduction efficiencies of released vectors were quantitatively and qualitatively using hMSCs in monolayer culture (109,400 cells/cm2). For the first approach based on rAAV-lacZ gene transfer, X-Gal staining was performed following fixation of the cells and further processing according to the manufacturer’s recommendations to examine positive staining under light microscopy (Olympus BX45). Quantitative estimation of the transduction efficiencies of rAAV-lacZ-loaded and unloaded capsules was carried out measuring the luminescence produced after 30 min of incubation with the BetaGlo® assay as described above with RLUs normalized to the cell number. The transduction efficiencies were qualitatively estimated by detection of live fluorescence in transduced cells as with polymers alone after 1, 3, 5 and 10 days of contact with polymeric rAAV-RFP-loaded capsules under a fluorescent microscope with a 568 nm filter (Olympus CKX41). 8.3.10 Hydrogel-ceramic composites In order to evaluate the possibility to include the hydrogel systems in a ceramic matrix, known amounts of each optimized polymeric solution were added to bioSiC samples. The composite systems obtained were then crosslinked with calcium chloride by 283 Controlled release of rAAV vectors from alginate-poloxamer-SiC composite systems immersion in calcium solution (102 mM) for 30 sec. The transduction efficiencies of the composite systems using rAAV-lacZ were evaluated by X-Gal staining after 1, 5 and 10 days of monolayer culture with hMSCs (12,500 cells/well in 24-well plates). The effects of the presence of bioSiC on the transduction efficiencies were evaluated similarly. 8.3.11 Statistical analysis All experiments were performed in quintuplicate for each condition and time point. Results are expressed as mean ± standard deviation. Statistical significant differences between treatments were evaluated by analysis of variance (ANOVA) using the Statgraphics Centurion® X64 software (Statpoint Technologies, USA). P values of less than 0.05 were considered statistically significant. 8.4 Results 8.4.1 Hydrogel and bead characterization The potential interactions between AlgPH155 and PF127 were studied by analyzing the rheological properties of the polymeric dispersions. The dispersion of AlgPH155 in water produced a viscous solution characterized by a complex viscosity that decreases as the temperature increases with no gelation phenomena. The low concentration of AlgPH155 selected produced solutions with negative values of elastic modulus at all the temperatures studied (data not shown). PF127 dispersion showed a gelation temperature of 55.50 °C (Figure 8.1A). The combination AlgPH155 + PF127 promoted a reduction in the gel temperature to a lower value (44.55 °C) (Figure 8.1B), probably due to enhanced interactions of poloxamer chains via AlgPH155 (Figure 8.1C). Capítulo 8 284 Figure 8.1 Hydrogel and bead characterization. The rheological properties of the polymeric dispersions were analyzed by using ramps of temperature from 15 °C to 60 °C at 2 °C/min with an oscillatory stress of 0.1 Pa at 5 rad/sec. Values of storage moduli (G´) and loss moduli (G´´) are shown for PF127 (A) and AlgPH155 + PF127 (B). The gel temperatures were obtained by the cross point between both moduli. A schematic distribution of polymeric chains after crosslinking with calcium ions is depicted in (C). The stability parameters as evaluated by measuring the loss of capsule area and weight are presented in Figure 8.2. The differences in porous structures resulted in variations 15 20 25 30 35 40 45 50 55 60 65 1 10 100 G', G'' (Pa) temperature (°C) G' G'' 20 25 30 35 40 45 50 55 60 65 1 10 100 G', G'' (Pa) temperature (°C) G' G'' A B C 291 Controlled release of rAAV vectors from alginate-poloxamer-SiC composite systems Figure 8.6 Live fluorescence emission of transduced hMSCs in monolayer cultures after 1, 3, 5 and 10 days of treatment with rAAV-RFP-loaded capsules. Qualitative analysis was also carried out by X-Gal staining of hMSCs treated with unloaded and rAAV-lacZ-loaded capsules. Figure 8.7 shows lacZ expression after 1, 3, 5 and 10 days. Capsules containing alginate and poloxamer crosslinked at room temperature and at 50 °C showed higher staining after one day of treatment than capsules formed by alginate alone. On the other hand, capsules containing single AlgPH155 or AlgPH155 + PF127 [H] showed higher transduction efficiencies after one day of assay. Different cell morphology was observed in these two types of capsules, cells treated with alginate capsules show fibrous shape whereas cells treated with AlgPH155 + PF127 capsules show round shape. It was clearly observed that while positive control suffered a decrease in lacZ expression after 3 days of culture, Capítulo 8 292 AlgPH155 and AlgPH155 + PF127 [H] maintained staining intensity over the time. Controlled release of rAAV was able to promote a suitable transduction efficiency over time in such systems. Figure 8.7 X-Gal staining of hMSCs in monolayer culture after 1, 3, 5 and 10 days of treatment with capsules including rAAV-lacZ. Estimation of the transduction efficiencies was carried out using rAAV-lacZ-loaded and unloaded capsules as controls. Figure 8.8 shows RLUs/cell number for each one of the studied conditions at different times. According to the experimental results, all systems were able to achieve significant higher values of transgene expression than negative control (Control -). Capsules including PF127 were found to be able of increasing cell transduction in comparison with both, the positive control and single alginate capsules formed after 24 h of study. This effect may be attributed to the initial burst release 293 Controlled release of rAAV vectors from alginate-poloxamer-SiC composite systems and/or the polymer itself in agreement with our previous results (Figures 8.3 and 8.4B). The addition of poloxamer is able to increase the transduction efficiencies of the vectors. AlgPH155 AlgPH155 + PF127 [C] AlgPH155 + PF127 [H] Control - Control + 0 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 * * ** RLU/cell numer corrected 1 day 3 days 5 days 10 days 21 days ** Figure 8.8 Transduction efficiencies of released rAAV-lacZ from capsules containing AlgPH155 with or without PF127. The homogeneous groups are indicated by an equal number of asterisks (*) above the columns (P < 0.05). After the first time point of cell culture, poloxamer-alginate capsules shown similar levels of gene transduction efficiencies than alginate capsules but significantly higher than negative and positive controls. Furthermore, single alginate capsules formed by AlgPH155 were able to achieve stable cell transduction with similar RLU values at all time points evaluated whereas for the other capsules, the transduction efficiencies decreased with time. In spite of the effect of alginate decreasing transduction efficiency found for this polymer alone (Figure 8.4B), the controlled release achieved in these samples was able to promote an adequate in vitro transduction efficiency with statistically significant higher levels to negative control and AlgPH155 + PF127 [C] even Capítulo 8 294 after 21 days of cell culture. Interestingly, capsules formed by the two polymers crosslinked at room temperature were not able to achieve higher levels of transduction efficiencies than negative control after 10 and 21 days of cell culture. According to the experimental data capsules formed by alginate alone have been shown the best release profiles and a more prolonged increase in trangene expression. Poloxamer-alginate capsules crosslinked at 50 °C have also shown good experimental results. These two hydrogel structures were selected for their incorporation into bioSiC ceramics. 8.5.6 Hydrogel-ceramic composites Selected hydrogels including rAAV vector were used to be loaded into ceramic matrices in order to obtain a composite system able of promoting a local controlled release of rAAV useful for bone repair implants. Two compositions were selected for this purpose in light of their better in vitro results: AlgPH155 and AlgPH155 + PF127 [H]. As controls, equivalent concentrations of rAAV were directly loaded in the bioSiC samples, vectors added to cell culture medium were used as positive control and cells cultured without any treatment as negative controls. Figure 8.9 shows the results of XGal staining for the studied conditions. Both positive control and silicon carbide containing rAAV-lacZ showed similar staining intensity at different times studied. The presence of this ceramic thus did not affect the transduction efficiency. On the other hand, low staining was observed on both composite systems. Aginate-bioSiC showed few transduced cells while no stained cells were observed using AlgPH155 + PF127bioSiC. This may be due to the lower hydrogel-cell culture interfacial surface that decreases the release rate of the vectors, thus decreasing transduction. Optimization of the composite process need to be next performed to obtain the desirable therapeutic effect for these composite systems. 295 Controlled release of rAAV vectors from alginate-poloxamer-SiC composite systems Figure 8.9 X-Gal staining of hMSCs in monolayer cultures after 1, 3, 5 and 10 days of treatment with the hydrogel-bioSiC composites including rAAV-lacZ. 8.5 Discussion The combination of two polymers allowed to generate polymeric systems capable of combining the advantages of each polymer. Alginate is a safe natural polymer extracted from brown algae commonly used as tablet binding agent and as a diffusion barrier in controlled release formulations, decreasing drug molecules migration (25). The selection of alginate as part of the polymeric composition allows obtaining crosslinked hydrogels by simple addition of divalent ions (26). Synthetic block polymers such as poloxamers are capable of forming micelles in aqueous solutions with a hydrophilic shell and a hydrophobic core adequate for hydrophobic drug inclusion, increasing their solubility. Moreover, poloxamers can undergo a sol-to-gel transition at a temperature higher than the gel temperature (27). Capítulo 8 296 The addition of alginate to poloxamer dispersions was able to decrease the gel temperature of the systems. This effect could be caused by complex interactions between the chains of both polymers where water could act as crosslinking agent forming hydrogen bonds between the polymers (28). The ionotropic crosslinking of the polymeric dispersions with calcium chloride at two different temperatures was carried out being the obtained capsules of different characteristics. The selection of a crosslinking temperature higher than the gel temperature of the mixture alginate-poloxamer should make possible to obtain hydrogel systems with different porous structures. The aggregation of micelles after the gel temperature (50 °C) could be used as templates in order to obtain systems with higher porous size. The complex interactions between the two polymers where alginate chains should be placed around the polymeric micelles could increase the hydrophobicity of micelles and enhance micelle-micelle interactions and the consequent gelation as is shown in Figure 8.1B. The significant differences in capsule stability parameters suggest that the presence of block polymeric micelles in alginate medium modifies the interaction between alginate chains and calcium ions during the crosslinking process. Micelles could impair these interactions giving final highly porous hydrogels therefore with lower stability. When the crosslinking procedure was carried out at 50 °C, a temperature over the sol-gel transition (Tgel 44.50 °C), the pluronic micelles can undergo aggregation. The increase in viscosity which strongly may hinder the ionic interactions between alginate and calcium together with presence of micelle aggregates, which act as porogen agents, may cause the synthesis of an even more porous hydrogel with larger pore size (Figure 8.2). Good encapsulation efficiencies were obtained for all the systems studied. However, despite the higher viscosity of the polymeric composition, the lowest encapsulation efficiency was obtained for high temperature crosslinked systems. This might be result 297 Controlled release of rAAV vectors from alginate-poloxamer-SiC composite systems of the higher porous structure on their external hydrogel surface layer, leading to a loss of rAAV from the particles when is compared to systems crosslinked at room temperature. The controlled rAAV release was governed by a diffusion mechanism through the hydrogel matrix. The three porous structure obtained were able to achieve different release profiles as a function of the variations in hydrogel structure that are in agreement with the stability results. The evaluation of the effect of the polymer presence on transduction efficiencies has pointed out an increase of trangene expression for PF127 according to a similar phenomenon that has been previously described for a similar block copolymer, PF68 (29). On the contrary the addition of alginate decreases the transduction efficiency. It has been reported that this polymer could decrease the rate of virus transport, thus decreasing the in vivo transfection of cancer cells, but not affecting the bioactivity of the vectors (30). Qualitative evaluation of trangene expression of capsule released vectors through live fluorescence emission and X-Gal staining showed an enhancement on the transduction efficiency for the controlled release vectors when they are compared to positive control at long periods of time. It was previously reported that an adequate rAAV release profile is able to increase transgene expression of cells and reduce macrophage activation, increasing the therapeutic utility of the systems (13). According to the quantitative analysis of transduction efficiencies data it is an evidence that the presence of pluronic in the capsule composition enhance the transduction efficiency of the release vectors after one day of study. This fact could be attributed to a change in the cellular membrane permeability (27) and was previously reported for lentiviral and adenoviral vectors (31). Capítulo 8 298 Biomorphic silicon carbide ceramics have been shown to have highly porous structures in which the optimized hydrogels could be included. The addition of viral vectors to the ceramic structures did not affect the activity of the vectors. 8.6 Conclusions In the present study, we were able to develop three hydrogel networks due to the thermosensitive properties of PF127 and their complex interactions with alginate chains during calcium crosslinking. These systems were stable against dilution and were able to produce three different release profiles of rAAV. The presence of PF127 was able to increase the transduction efficiency of rAAV. Crosslinked capsules showed excellent biocompatibility results and their transduction efficiency evaluation had lead to better results for systems containing alginate alone because of the long time controlled release observed for such systems. Despite the effect of PF127 on the transduction efficiency the faster release on these systems was not able to achieve higher long-term transduction efficiencies than negative control after five days of assay for room temperature crosslinked systems. The inclusion of the hydrogels in a porous ceramic was not able to achieve an adequate transduction efficiency due to the lower hydrogel surface and the consequent lower vector release. Further studies should be done in order to achieve adequate ceramic-hydrogel combinations for the treatment of osteochondral defects. 8.7 References 1. Kootstra NA, Verma IM. Gene therapy with viral vectors. Annu Rev Pharmacol Toxicol. 2003;43:413-39. 2. Ibraheem D, Elaissari A, Fessi H. Gene therapy and DNA delivery systems. Int J Pharm. 2014;459(1-2):70-83. 299 Controlled release of rAAV vectors from alginate-poloxamer-SiC composite systems 3. Wang C, Pham P. Polymers for viral gene delivery. Expert Opin Drug Deliv. 2008;5(4):385-401. 4. Xu X, Yang J, Cheng Y. Pharmacokinetic study of viral vectors for gene therapy: Progress and challenges. Viral Gene Ther. 2011:435-50. 5. Weimer A, Madry H, Venkatesan JK, Schmitt G, Frisch J, Wezel A, et al. Benefits of recombinant adeno-associated virus (rAAV)-mediated insulin-like growth factor I (IGFI) overexpression for the long-term reconstruction of human osteoarthritic cartilage by modulation of the IGF-I axis. Mol Med. 2012;18(3):346-58. 6. Huang S, Kamihira M. Development of hybrid viral vectors for gene therapy. Biotechnol Adv. 2013;31(2):208-23. 7. Cucchiarini M, Thurn T, Weimer A, Kohn D, Terwilliger EF, Madry H. Restoration of the extracellular matrix in human osteoarthritic articular cartilage by overexpression of the transcription factor SOX9. Arthritis Rheum. 2007;56(1):158-67. 8. Venkatesan JK, Rey-Rico A, Schmitt G, Wezel A, Madry H, Cucchiarini M. rAAVmediated overexpression of TGF-beta stably restructures human osteoarthritic articular cartilage in situ. J Transl Med. 2013;11:211-24. 9. Venkatesan JK, Ekici M, Madry H, Schmitt G, Kohn D, Cucchiarini M. SOX9 gene transfer via safe, stable, replication-defective recombinant adeno-associated virus vectors as a novel, powerful tool to enhance the chondrogenic potential of human mesenchymal stem cells. Stem Cell Res Ther. 2012;3(3):22-36. 10. Cucchiarini M, Orth P, Madry H. Direct rAAV SOX9 administration for durable articular cartilage repair with delayed terminal differentiation and hypertrophy in vivo. J Mol Med. 2013;91(5):625-36. 11. Moulay G, Boutin S, Masurier C, Scherman D, Kichler A. Polymers for improving the in vivo transduction efficiency of AAV2 vectors. PLoS One. 2010;5(12): e15576-83. Capítulo 8 300 12. Zeng Y, Tseng S-, Kempson IM, Peng S, Wu W, Liu J. Controlled delivery of recombinant adeno-associated virus serotype 2 using pH-sensitive poly(ethylene glycol)-poly-l-histidine hydrogels. Biomaterials. 2012 12;33(36):9239-45. 13. Lee S, Kim J, Chu HS, Kim G, Won J, Jang J. Electrospun nanofibrous scaffolds for controlled release of adeno-associated viral vectors. Acta Biomaterialia. 2011 11;7(11):3868-76. 14. Kilpatrick LA, Li Q, Yang J, Goddard JC, Fekete DM, Lang H. Adeno-associated virus-mediated gene delivery into the scala media of the normal and deafened adult mouse ear. Gene Ther. 2011;18(6):569-78. 15. Kidd ME, Shin S, Shea LD. Fibrin hydrogels for lentiviral gene delivery in vitro and in vivo. J Control Release. 2012;157(1):80-5. 16. Tseng S-, Kempson IM, Peng S, Ke B, Chen H, Chen P, et al. Environment acidity triggers release of recombinant adeno-associated virus serotype 2 from a tunable matrix. J Control Release. 2013;170(2):252-8. 17. Zhang F-, Jia S-, Zheng S-, Ding W. Celastrol enhances AAV1-mediated gene expression in mice adipose tissues. Gene Ther. 2011;18(2):128-34. 18. Díaz-Rodríguez P, Pérez-Estévez A, Seoane R, González P, Serra J, Landin M. Suitability of biomorphic silicon carbide ceramics as drug delivery systems against bacterial biofilms. ISRN Pharm. 2013:104529-36. 19. Díaz-Rodríguez P, Landin M, Rey-Rico A, Couceiro J, Coenye T, González P, et al. Bio-inspired porous SiC ceramics loaded with vancomycin for preventing MRSA infections. J Mater Sci Mater Med. 2011;22(2):339-47. 20. Elsler S, Schetting S, Schmitt G, Kohn D, Madry H, Cucchiarini M. Effective, safe nonviral gene transfer to preserve the chondrogenic differentiation potential of human mesenchymal stem cells. J Gene Med. 2012;14(7):501-11. 307 Discusión general Figura 9.1 Relación entre la porosidad de los bioSiCs y sus preformas de carbón obtenida mediante porosimetría de intrusión de mercurio. La figura 9.2 representa gráficamente algunos de los parámetros característicos de las cerámicas biomórficas producidas y sus precursores de carbón. Mientras que el pino genera la cerámica de mayor porosidad total (46.97%) y menor superficie específica, con macroporos de pequeño tamaño (50 ± 20 µm), el roble da lugar al bioSiC con menor porosidad total (27.85%) pero con la mayor superficie porosa, una densidad intermedia y macroporos de mayor tamaño 250 ± 20 µm (Capítulo 6). Los bioSiCs obtenidos de la madera de sapelli se caracterizan por presentar: una distribución de tamaños de poro bimodal, una porosidad total similar a la del pino (40.72%), una superficie específica intermedia y la mayor densidad de los tres sistemas estudiados, con macroporos de un tamaño medio de 140 ± 30 µm. 0,01 0,1 110 100 0,00 0,02 0,04 0,06 0,08 0,10 0,12 0,14 0,16 0,18 0,20 bioSiC de Sapelli Carbón de Sapelli Volumen de poro incremental (ml/g) Diámetro medio (m) 0,01 0,1 1 10 100 0,00 0,02 0,04 0,06 0,08 0,10 0,12 0,14 0,16 0,18 0,20 Volumen de poro incremental (ml/g) Diámetro medio (m) bioSiC de roble carbón de roble 0,01 0,1 1 10 100 0,00 0,02 0,04 0,06 0,08 0,10 0,12 0,14 0,16 0,18 0,20 Volumen de poro incremental(ml/g) Diámetro medio (m) bioSiC de pino carbón de pino Capítulo 9 308 0,60,81,01,2 1,4 1,6 1,8 2,0 0 1 2 3 20 30 40 50 60 bioSiC de roble bioSiC de sapelli bioSiC de pino Carbón de roble Carbón de sapelli densidad real (g/cm3) porosidad (%) area superficial (m2/g) Carbón de pino Figura 9.2 Relación entre porosidad, área superficial (obtenida por adsorción de nitrógeno) y densidad real (obtenida mediante picnometría de helio) de los tres sistemas cerámicos bioSiCs. Los tres bioSiCs, a diferencia de sus correspondientes preformas de carbón, se caracterizan por ser materiales hidrofílicos, con un ángulo de contacto medio de 38 ± 7º (Capítulo 5) sin diferencias significativas entre ellos. Sin embargo, el empleo de fluidos de diferente naturaleza permite la estimación de valores de energía superficial variables. Así, los bioSiCs de pino presentaron una energía superficial positiva, mientras que los de roble y sapelli presentaron valores negativos (Capítulo 6). La rugosidad superficial, evaluada mediante perfilometría interferométrica (Capítulo 5), también pone de manifiesto variaciones importantes. El bioSiC de sapelli presenta mayor rugosidad superficial con un Rq de 11.05 µm, muy superior al del pino y el roble, con valores de 6.91 µm y 5.91 µm respectivamente. El análisis de la superficie de los bioSiCs mediante FT-IR mostró en todas ellas, la presencia de grupos funcionales OH y SiO (Capítulo 6). 309 Discusión general La modelización de la estructura porosa de los sistemas cerámicos a partir de las curvas de porosimetría de intrusión de mercurio mediante un software especializado (PoreXpert®) permite valorar la interconectividad de sus poros, su tortuosidad y también realizar predicciones sobre el potencial proceso de penetración de un fluido en la estructura porosa cuando la cerámica entre en contacto con un medio líquido (Capítulo 6). La Figura 9.3 presenta la cinética de captación de agua a 0 MPa simulada para los diferentes sistemas. Como puede observarse, es la microestructura del bioSiC de sapelli la que presumiblemente es capaz de captar más agua y más rápidamente. Este sistema presenta la mayor conectividad y la menor tortuosidad de los tres estudiados. A pesar de que la capacidad de captación de agua es similar en los bioSiCs de pino y de roble a 110 ms, la modelización pone de manifiesto que la conectividad del bioSiC de roble es superior a la del pino. Figura 9.3 Simulación de la captación de agua por parte de los tres sistemas porosos bioSiCs obtenida tras la modelización de la estructura porosa con el software especializado PoreXpert®. Capítulo 9 310 9.2 Evaluación de la respuesta celular e inmunológica a los bioSiCs La captación de fluidos por parte de las estructuras porosa modifica no sólo la potencial capacidad de carga de fármacos, sino también la adhesión celular y la respuesta inflamatoria. Por ello, las diferentes estructuras de carburo de silicio deberían dar lugar a un comportamiento celular diferente y perfiles de carga y cesión variables. De esta manera la evaluación de la respuesta celular y tisular a los sistemas bioSiCs resulta crucial para analizar su potencial aplicación en regeneración ósea. 9.2.1 Evaluación de la biocompatibilidad de los bioSiCs La introducción de nuevos materiales como sustitutos óseos ha de estar presidida por la capacidad del biomaterial para ejercer su función sin producir efectos adversos. El biomaterial ideal útil en regeneración ósea debe ser económico, fiable y seguro, a la vez que biocompatible, osteoinductivo, osteoconductivo y preferentemente biodegradable (3, 4). Si el defecto óseo es grande y el implante ha de fabricarse con material no biodegradable, es necesario que éste sea capaz de estimular la formación de tejido óseo y/o promover su osteointegración. En estos casos las características microestructurales y superficiales del material son críticas. Diversos autores han puesto de manifiesto que el tamaño de los poros del material condiciona la formación de nuevos vasos en el implante y, por tanto, el crecimiento del tejido óseo (5-8). El estudio preliminar de la biocompatibilidad de los bioSiCs fue realizado con una línea celular de fibroblastos (BALB/3T3) usando como material cerámico el bioSiC de sapelli (Capítulo 3). Los resultados experimentales demostraron la excelente biocompatibilidad de los sistemas observándose la formación de una monocapa de células tras 15 días de estudio (Figura 9.4). 311 Discusión general Figura 9.4 Fotografía de microscopía electrónica de barrido (SEM) de sapelli tras 15 días de cultivo con la línea celular fibroblástica BALB/3T3. La biocompatibilidad de los tres sistemas bioSiCs (pino, roble y sapelli) fue posteriormente evaluada usando células madre mesenquimales obtenidas de médula ósea humana (Capítulo 6). Los resultados de viabilidad celular así como las imágenes de microscopía confocal, la cuantificación de citoquinas proinflamatorias (IL-1β) y de marcadores de apoptosis (caspasa-3), muestran que todos los sistemas son altamente biocompatibles con niveles no detectables de IL-1β y concentraciones de caspasa-3 equivalentes al control negativo a todos los tiempos estudiados. De esta manera, todos los carburos de silicio cumplen el requisito de biocompatibilidad. 9.2.2 Efecto de las estructuras porosas de los bioSiCs sobre la interacción con los componentes sanguíneos Las características superficies de los materiales también determinan su interacción con los componentes sanguíneos en el momento de su implantación. La formación de coágulo, la activación de plaquetas y la inflamación condicionan el proceso normal de regeneración ósea y modulan el crecimiento óseo alrededor del implante y la osteoconducción (3, 9-12). También ha sido descrito que la energía superficial, el Capítulo 9 312 ángulo de contacto, la liberación de iones, la resistividad y la rugosidad superficial modifican estas interacciones, y por tanto, la regeneración ósea (13-16). La interacción de las superficies de los carburos de silicio y los componentes sanguíneos fue evaluada en lo que respecta a la hemólisis producida, la adsorción de proteínas, la coagulación, la adhesión de plaquetas y la activación del sistema de complemento con el fin de obtener los parámetros microestructurales clave que condicionan dichas interacciones (Capítulo 5). Así, se ha observado que el tamaño de poro es fundamental a la hora de predecir la hemólisis generada. La utilización de sistemas con menor tamaño de macroporo (pino y sapelli) da lugar a valores de hemolisis más reducidos. La relación entre la adsorción de proteínas hidrofílicas (seroalbúmina bovina) y proteínas hidrofóbicas (fibrinógeno) condicionan la adhesión celular y como consecuencia la posterior osteointegración. Todos los bioSiCs estudiados se caracterizan por poseer una mayor adsorción de albúmina que de fibrinógeno lo que debe facilitar el reconocimiento celular de las superficies. Ello además, se correlaciona con una baja formación de trombo y con una mayor superficie específica de los materiales. De la misma manera, se observó que la activación del sistema de complemento, directamente relacionada con la respuesta inmune, está condicionada por la adsorción de proteínas. La adhesión de plaquetas por parte de los sistemas cerámicos mostró que éstas mantienen una morfología redondeada, lo que indica una adecuada adhesión celular, sin causar su activación. La evaluación de formación de coágulo para los tres sistemas ensayados mostró mayores valores para el bioSiC de sapelli lo que está directamente correlacionado con la mayor rugosidad superficial. Se ha observado que la formación de coágulo es dependiente de la superficie externa mientras que la adsorción de proteínas se ve modulada tanto por la superficie externa como por la interna. Los resultados 313 Discusión general experimentales obtenidos permiten concluir que las cerámicas de carburo de silicio biomórfico presentan una hemocompatibilidad adecuada para su aplicación como sistemas implantables. Además, la formación de coágulo observada debería dar lugar a una adecuada integración de los implantes. La reducida respuesta inflamatoria podría indicar la ausencia de formación de tejido fibroso alrededor del implante favoreciendo su integración (17). De acuerdo con los resultados experimentales de rugosidad superficial y la formación de coágulo observada los carburos de silicio de sapelli se presentan como los más prometedores para una posible aplicación clínica. 9.2.3 Efecto de la topografía de los bioSiCs sobre la diferenciación celular La rugosidad superficial, el ángulo de contacto, la cristalinidad y la composición química modulan la adhesión, la morfología, la funcionalidad y la migración celular en el interior del material lo que condiciona la fijación del implante y su integración (5, 18-27). El efecto de las características topográficas de los carburos de silicio biomórficos así como de su estructura porosa sobre la diferenciación osteoblástica de células madre mesenquimales, se evaluó en función de la secreción de indicadores de diferenciación (osteocalcina, osteopontina y fosfatasa alcalina) (Capítulo 6). De acuerdo con datos bibliográficos previos, se ha observado que las propiedades de los bioSiCs modifican el comportamiento celular. Así, los carburos de silicio que presentan mayor tamaño de poro (roble y sapelli) son capaces de estimular la diferenciación osteoblástica de las células madre mesenquimales tras quince días de cultivo obteniéndose valores de osteocalcina y osteopontina similares a los de las células cultivadas en presencia de medio de diferenciación osteoblástica. Es posible que la diferenciación observada no sólo esté condicionada por la morfología de los sistemas, sino también por la presencia de silicio, ya que se ha descrito que este compuesto es capaz de promover la Capítulo 9 314 diferenciación osteoblástica, e incrementar la proliferación de los osteoblastos y su producción de osteocalcina (28). Los resultados experimentales obtenidos permiten concluir que seleccionando un precursor con una estructura porosa adecuada se podrían obtener cerámicas de carburo de silicio biomórfico con capacidad osteoinductora. 9.2.4 Actividad inflamatoria de las partículas de bioSiC obtenidas por desgaste mecánico Otro aspecto importante de los biomateriales es su potencial para generar productos de degradación tóxicos tras su implante. Por ejemplo, el uso de polímeros derivados del ácido láctico promueve la producción de sustancias de carácter ácido como producto de degradación que generan inflamación en los tejidos adyacentes (29). El desgaste mecánico de los materiales implantados puede dar lugar a la formación de partículas, que causen la activación de los macrófagos circundantes, el desencadenamiento de una respuesta inflamatoria aguda y la consiguiente destrucción de tejido óseo y fallo del implante (30-32). Los carburos de silicio biomórficos son materiales sólidos, poco friables, cuya velocidad de degradación es extremadamente lenta, menor de 30 nm por año en condiciones fisiológicas normales (30). Se estudió la capacidad de las partículas de carburo de silicio biomórfico para estimular macrófagos y provocar una respuesta inflamatoria con vistas a predecir su potencial toxicidad a largo plazo tras su implantación (Capítulo 7). Las partículas de carburo de silicio nanométricas obtenidas por desgaste mecánico a partir de dos piezas de carburo de silicio sometidas a rozamiento fueron internalizadas por los macrófagos, como se muestra en la Figura 9.5. A pesar de ello, la secreción de citoquinas proinflamatorias por los macrófagos (TNF-α e IL-1β) mostró valores similares a los obtenidos con las partículas de zirconio usadas como referencia. De 315 Discusión general acuerdo con los resultados experimentales se puede afirmar que el carburo de silicio causa unos niveles aceptables de inflamación, similares a un material, el zirconio, que está reconocido como el más adecuado desde el punto de vista inflamación de partículas (33). Figura 9.5 Captación de las partículas de carburo de silicio por parte de los macrófagos en comparación con el control negativo (A). 9.3 Desarrollo de sistemas biofuncionales de carburo de silicio La incorporación de moléculas terapéuticas en sistemas implantables permite dotar de valor añadido a los materiales. De acuerdo con lo descrito en el Capítulo 1.1 existen diferentes mecanismos de carga que pueden ser empleados para la obtención de materiales biofuncionales y que condicionan los perfiles de liberación de fármaco obtenidos. En este trabajo se han utilizado tres mecanismos diferentes para la incorporación de moléculas terapéuticas: la adsorción inespecífica, las interacciones iónicas y la inclusión en una matriz. 9.3.1 Adsorción inespecífica de antibióticos A pesar de los tratamientos preventivos con antibióticos a nivel sistémico, la osteomielitis postquirúrgica continúa siendo una seria complicación en la cirugía A B Capítulo 9 316 ortopédica y dental (34). Los materiales implantados en el organismo son a menudo objeto de colonización microbiana, lo que evita la adhesión celular y da lugar al fallo de la prótesis (35). Cuando la contaminación es severa, la osteomielitis se cronifica y la administración sistémica de antibióticos no es capaz de erradicarla debido a las bajas concentraciones locales de fármaco que se alcanzan. El protocolo para el tratamiento de esta problemática supone la retirada de la prótesis y la implantación provisional de sistemas poliméricos de polimetil-metacrilato (PMMA) cargados con antibióticos capaces de alcanzar un adecuado perfil de liberación. El carácter no biodegradable del PMMA hace necesaria su retirada mediante una nueva cirugía (36, 37). La posibilidad de desarrollar sistemas para la regeneración ósea, cargados con antibióticos que prevengan o traten la osteomielitis representa una prometedora alternativa en regeneración ósea. En este estudio se han desarrollados sistemas de carburo de silicio biomórfico cargados con vancomicina mediante un mecanismo de adsorción inespecífica. La necesidad de alcanzar concentraciones elevadas de fármaco a tiempos cortos que funcionen como dosis de ataque a los microorganismos presentes en el medio y eviten la colonización del implante por parte de las bacterias hacen de este mecanismo el más adecuado a priori para el desarrollo de sistemas bioSiC cargados con vancomicina. Como evaluación preliminar de la capacidad de carga y cesión de los bioSiCs se seleccionó el carburo de silicio obtenido a partir de madera de sapelli (Capítulo 3). La isoterma de adsorción del fármaco muestra la formación inicial de una monocapa de moléculas de vancomicina para luego dar lugar a consecutivas capas de fármaco a medida que la concentración de carga se ve incrementada de acuerdo con un perfil característico de adsorción inespecífica. 323 Discusión general (rAAV) han mostrado utilidad para la transducción de condrocitos (60, 61). Sin embargo, el elevado tamaño y la reducida estabilidad de estos virus dificulta su aplicación mediante sistemas de liberación convencionales, lo que abre interesantes perspectivas para el desarrollo de nuevas alternativas para su administración. Como fase final de este trabajo hemos desarrollado combinaciones de estructuras porosas de carburo de silicio biomórfico con polímeros, alginato y poloxamer, con el objetivo de cargar y liberar de forma controlada estos vectores virales para su potencial aplicación en la regeneración de defectos osteocondrales. La presencia de alginato y poloxamer modifica la eficacia de transducción de los vectores virales, de forma similar a lo observado por otros autores (62-64). Así, el poloxamer incrementa la eficacia de transducción, posiblemente debido a la modificación de la permeabilidad de la membrana (65), mientras que la adición de alginato al medio de cultivo reduce la eficacia de transducción de los vectores en solución. La incorporación de los vectores virales en los sistemas binarios alginato-poloxamer reticulados con calcio da lugar a cápsulas con perfiles de liberación condicionados por su composición y condiciones de procesado. De esta manera, las cápsulas formadas por alginato solo poseen una estructura más compacta en su parte externa. Los sistemas formados por alginato y poloxamer reticulados a temperatura ambiente y a 50 ºC poseen una estructura menos compacta. En estos últimos, el realizar la reticulación a una temperatura superior a la de gelificación (44.55 ºC) permitiría que los agregados micelares de poloxamer actuaran como moldes para obtener sistemas de mayor tamaño de poro. Los perfiles de liberación obtenidos de los tres hidrogeles con estructura variable se correlacionan con la eficacia de transducción observada. El perfil de liberación más Capítulo 9 324 controlado, obtenido por los sistemas que contienen solo alginato, fue el que logró una eficacia de transducción más prolongada. Los sistemas que contienen poloxamer mostraron un incremento de la capacidad de transducción de los vectores virales tras el primer día de cultivo. Sin embargo, este efecto no se observó a tiempos más prolongados. La incorporación de los dos sistemas con mejor liberación en el seno de matrices poliméricas de bioSiC no logró un perfil de liberación adecuado capaz de alcanzar unos niveles deseados de transducción (Figura 9.8). La adición de los vectores virales en las cerámicas de carburo de silicio biomórfico no modificó la eficacia de transducción de los vectores. Las condiciones de síntesis de los sistemas complejos deberían de ser modificadas con el fin de alcanzar el nivel óptimo de transducción. Figura 9.8 Tinción X-Gal correspondientes a células madre mesenquimales cultivadas en monocapa tras 1, 5 y 10 días de tratamiento con los sistemas compuestos hidrogelbioSiC conteniendo el vector viral rAAV-lacZ. 325 Discusión general 9.4 Referencias 1. 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Conclusions The work presented and discussed herein led to the following conclusions: 1. The bioceramization process employed for the synthesis of bioSiCs in two steps, pyrolyzation and infiltration with liquid silicon of natural precursors (pine, oak and sapelli wood) allows three ceramic systems with different properties of porosity, pore size distribution, surface properties and density to be obtained, which determine its usefulness as bone substitutes able to load and release therapeutic molecules. 2. All materials showed excellent biocompatibility values when tested on an fibroblast cell line (BALB/3T3) and mesenchymal stem cells derived from human bone marrow. Furthermore, the microstructure of materials promotes the osteoblastic differentiation, particularly those silicon carbides with larger pore size (oak and sapelli) in which, after fifteen days of cell culture, osteocalcin and osteopontin levels were similar to those of cells cultured in the presence of the osteoblastic differentiation medium. Silicon carbide particles produced by mechanical wear showed no toxic effect and an acceptable cellular response. Capítulo 4 340 Significant differences in the interactions of bioSiCs and their respective carbon precursors in contact with blood components were observed according to their microstructure and surface properties. Thus, the hemolytic response is primarily determined by the pore size, while clot formation depends on the surface roughness and the protein adsorption and complement activation on the specific surface. 3. The antibiotic loading of the systems through unspecific interactions lead to multi layers of drug molecules on the surface of the material. These distribution results in a two phase release profile with an initial burst followed by a slower release, dependent on the microstructure of the material selected. The lower porosity of oak bioSiC leads to more sustained release profiles. Loaded systems are able to inhibit bacterial biofilm formation and even to treat an already formed biofilm. 4. The inclusion of VEGF into the ceramics by ionic interactions leads to biofunctional systems which maintain the bioactivity of the protein, increasing the vessel formation in an in vivo animal model. The addition of the growth factor achieves an initial stimulation of the differentiation of mesenchymal stem cells, showing a synergistic effect between the protein and the material topography. 5. The combination of bioSiCs with two polymers, alginate and poloxamer, allows low aqueous solubility drugs to be loaded by the drug inclusion into the synthetic polymer micelles and the hydrogel diffusion controlled release. Furthermore, the bioSiC microstructure determines the medium access and drug release, so that significant differences in the anti-inflammatory effect of indomethacin loaded bioSiCs composites were observed. 341 Conclusiones/Conclusions 6. The mixture of these polymers, alginate and poloxamer, has also allowed the loading and release of viral vectors maintaining their transduction efficiency after 21 days of assay, especially for those systems with closer structures. The incorporation of these hydrogels into the ceramic matrix has not produced systems that achieve adequate gene transduction, thus requires further optimization work. Taken all together, the results of this study confirm the high potential of silicon carbide ceramics as biofunctional systems, with an application as implants for the treatment of various bone pathologies. The large number of natural precursors and different loading techniques available, should allow the selection of the right fit in each case according to the required needs.