Bioceramics: From Bone Regeneration to Cancer Nanomedicine
Research on biomaterials has been growing in the last few years due to the clinical needs in organs and tissues replacement and regeneration. In addi-tion, cancer nanomedicine has recently appeared as an effective means to combine nanotechnology developments towards a clinical application. Ceramic materials are suitable candidates to be used in the manufacturing of bone-like scaffolds. Bioceramic materials may also be designed to deliver biologically active substances aimed at repairing, maintaining, restoring or improving the function of organs and tissues in the organism. Several mate-rials such as calcium phosphates, glasses and glass ceramics able to load and subsequently release in a controlled fashion drugs, hormones, growth factors, peptides or nucleic acids have been developed. In particular, to pre-vent post surgical infections bioceramics may be surface modifi ed and loaded with certain antibiotics, thus preventing the formation of bacterial biofi lms. Remarkably, mesoporous bioactive glasses have shown excellent characteris-tics as drug carrying bone regeneration materials. These bioceramics are not only osteoconductive and osteoproductive, but also osteoinductive, and have therefore been proposed as ideal components for the fabrication of scaffolds for bone tissue engineering. A recent promising development of bioceramic materials is related to the design of magnetic mediators against tumors. Magnetic composites are suit-able thermoseeds for cancer treatment by hyperthermia. Moreover, magnetic nanomaterials offer a wide range of possibilities for diagnosis and therapy. These nanoparticles may be conjugated with therapeutic agents and heat the surrounding tissue under the action of alternating magnetic fi elds, enabling hyperthermia of cancer as an effective adjunct to chemotherapy regimens.
1. Introduction
In the last few years the biomedical research area is going towards materials science aiming applications of materials to health care, the so-called biomaterials. They can be defi ned as implantable materials that must be in contact with living tissues with the fi nal aim of achieving a correct biological
Prof. M. Vallet-Regí , Dr. E. Ruiz-Hernández Departamento de Química Inorgánica y BioinorgánicaFacultad de Farmacia, Universidad Complutense de MadridPlaza Ramón y Cajal s/n, 28040 Madrid, Spain, and Networking Research Center on BioengineeringBiomaterials and Nanomedicine (CIBER-BBN), Spain E-mail: [email protected]
DOI: 10.1002/adma.201101586
interaction between the material and the host. [ 1 ] In the fi rst Consensus Conference of the European Society for Biomaterials (ESB) in 1976, a biomaterial was defi ned as “a nonviable material used in a medical device, intended to interact with biological systems”; however the ESB’s current defi -nition is a “material intended to interface with biological systems to evaluate, treat, augment or replace any tissue, organ or function of the body”. This subtle change in defi nition is indicative of how the fi eld of biomaterials has evolved. Biomaterials have moved from merely interacting with the body to infl uencing biological proc-esses toward the goal of tissue regenera-tion. However, a more recent defi nition has been published: A biomaterial is a sub-stance that has been engineered to take a form which, alone or as part of a complex system, is used to direct, by control of interactions with components of living systems, the course of any therapeutic or diagnostic procedure, in human or veterinary medicine. [ 2 ] Different types of materials can be found depending on the function to perform or the tissue to be replaced.
Biomaterials have been often originated from materials used in diverse research areas, that presented desirable mechan-ical properties but were not specifi cally designed to interact with surrounding tis-sues or with blood. [ 3 ] Among those mate-
rials, the employment of ceramics and glasses for skeletal repair and reconstruction has been increased due to both increases in life expectancy and the social obligations to provide a better quality of life. Depending on the type of ceramics employed and their interaction with the host tissue, they can be categorized as either bioinert or bioactive, and the bioactive ceramics can be resorbable or non-resorbable. [ 4 ] These materials, which may be produced in both porous and dense forms as well as powders, granulates or coatings, are known as bioceramics. [ 4 , 5 ] From the chemical point of view, bioceramics can be prepared from alu-mina, zirconia, carbon, calcium phosphates, silica-containing compounds and some other chemicals. Among them, phos-phates can be used to produce biomaterials since they present high biocompatibility and bone integration, and also present a similar composition to the inorganic fraction of bones. In fact, bioceramics are now used in a number of different applications
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REV
IEW María Vallet-Regí is full pro-
fessor of Inorganic Chemistry at the Department of Inorganic and Bioinorganic Chemistry of the Faculty of Pharmacy at Universidad Complutense de Madrid (Spain). Today she currently leads a research group working on biomaterials, especially in bioceramics, drug delivery systems and smart targeted nanosystems for drug
and gene delivery. Prof. Vallet-Regí has written more than 500 articles and several books and is member of sev-eral Academies. She is the most cited Spanish scientist (regardless of gender), according to ISI Web of Knowledge, in the fi eld of Materials Science since 2000, being awarded with diverse national and international scientifi c prizes and distinctions.
Dr. Eduardo Ruiz-Hernández obtained B.S. degree in Chemical Engineering from Universidad Complutense de Madrid (2005), and received his PhD from the same institu-tion in 2010. His PhD research focused on magnetic biomate-rials for medical applications. He is now a postdoctoral fellow in the Department of Pharmaceutics, Utrecht
University (Netherlands). His current research interests include the design, synthesis and application of targeted nanosystems for triggered drug and gene delivery.
throughout the body, covering all areas of skeleton. There has been a chronological evolution in the research of ceramics as bone substitutes. [ 6 , 7 ]
During the 1950s the fi rst aim was to use inert materials, which had no reaction with living tissues. The lack of toxicity of a specifi c material was enough to consider it as biocompat-ible. [ 8 ] However, this generation of ceramics, such as zirconia or alumina, were not recognized by the organism experimenting foreign body reactions, [ 9 ] that is, the implant was surrounded by an acellular collagen capsule which isolated it from the body.
In the 1980s the approach changed radically to the opposite direction. The goal was to implant ceramics that would react with the surrounding environment to produce newly formed bone, as it is the case of certain calcium phosphates and glasses. [ 10–14 ]
The main purpose in this new century is to obtain porous ceramics that act as scaffolds for cells and inducting molecules able to drive self regeneration of tissues. The initial mate-rial to design those scaffolds should be shaped in the form of pieces with interconnected and hierarchical porosity within the micron range. [ 15–17 ] During the design of porous materials with applicability in the biomedical area, it is very important to take into consideration fi rstly the hierarchical porosity that can be found in Nature. Upon mimicking this porous structure, those materials could perform a similar role to those hierarchically structured porous natural materials. These scaffolds are needed to support cells in tissue engineering applications, as discussed in section 3 of this review.
The scale of porosity of natural materials ranges from mil-limeters down to nanometers, depending on the material and its role within the body. Thus, pore diameters below 1 μ m are responsible for the bioactivity and protein interaction. Pore diameters between 1 and 20 μ m would determine the cellular behavior and the type of attached cells. When porosity ranges between 100 and 1000 μ m, they are responsible for cellular growing, blood fl ux and mechanical resistance. Finally, pore diameters larger than 1000 μ m would determine the shape of the implant and its functionality (See Figure 1 ).
Therefore, the different porosity scales in the materials here targeted would range from 1 and 1000 μ m in tissue engi-neering, co-existing macro-, meso- and micropores, and from 2 to 10 nm for drug delivery purposes. In this review, the importance of the hierarchical structure of porous materials with applications in Biology will be tackled, with special atten-tion to those with possible applications in life science and drug delivery technologies.
Although life sciences in general comprise all fi elds of sci-ence that involve the scientifi c study of living organisms, we will focus in the fi eld of biomaterials and bone tissue engineering including some aspects of cancer nanomedicine. This research area has been growing in the last few years due to the clinical needs in bone replacement and regeneration. In fact, there are more and more cases of skeletal defects that require a clinical solution due to the ageing of population. And cancer nano-medicine has recently appeared as an effective means to combine nanotechnology developments towards a clinical application.
Nanodevices allow a close interaction with biological mol-ecules and processes. Molecular ligand-targeted nanomedi-cines are considered a hot research topic given the possibility
to design highly effective therapies for some diseases in which side effects prevent successful outcomes. [ 18–25 ] In the case of cancer, there is a promise to transform protocols of diagnosis and treatment. For instance, metal and semiconductor nano-particles have been functionalized to improve the sensitivity of cancer sensors and therapeutics selectivity. [ 26–28 ] Highly specifi c early diagnosis could overcome current sensitivity limitations associated to detection methodologies. [ 29 ]
The idea of encapsulating severely cytotoxic chemothera-peutics has been long considered. [ 30 ] Also, biological targeting approaches via specifi c molecular ligands for selective recogni-tion of cancerous tissue have provided invaluable means for new treatment alternatives. Rosenholm et al. have identifi ed the steps to be followed when assessing the targeting potential of novel nanoceramics in vitro . [ 31 ] A quantitative comparison between targeted and non-targeted systems is required to deter-mine ligand selectivity. This fact should be further confi rmed by selective recognition of target receptor expressing cells. Additionally, competition experiments with specifi c ligands
Figure 1 . Size scale of natural and technological materials. The scale below is referred to bone regeneration applications.
may contribute to elucidate uptake mechanisms of these engi-neered nanocarriers. [ 32 ]
The incorporation of nanotechnology products has the poten-tial to combine several recently developed tools for the discovery of cancer biomarkers and the design of engineered material mediators for immunotherapy and hyperthermic-based novel therapies. [ 33–35 ] Moreover, cancer progress monitoring and fur-ther removal may be effi ciently carried out by novel imaging agents based on nanotechnology. [ 36 , 37 ] Advances on this ena-bling technology are expected to complement current medical protocols in fi ghting against such a devastating disease.
2. Bioceramic Materials
Research on bioceramics has evolved from the use of inert materials for mere substitution of living tissues towards the development of third generation bioceramics aimed at inducing bone tissue regeneration. We may fi nd the fi rst bioceramic in our own bones. It is a nanometric size, calcium defi cient and carbonated ( Figure 2 ) biological apatite.
Bone structure exhibits a hierarchical porosity ranging from a micron to hundreds of microns. Regarding the application of biomaterials, we need to attend to the factors that rule the formation of new bone. Among them, some belong to biology and others depend on mechanical stability. Three elements are critical in bone repair and regeneration: osteogenesis cells, growth factors for osteoinduction and scaffold required for osteoconduction.
Biomaterials have experienced a clear evolution: the fi rst generation aimed at replacing the damaged tissue, the purpose
of the second generation was to repair the tissue, whereas with the third generation, the purpose is to regenerate. The mate-rials chosen for replacement purposes had to be bioinert, in order to avoid any reaction with the human body. These mate-rials are identifi ed as foreign by the body, which defended itself by encapsulation. In the 1980s, the implantation of materials which could react with the human body is no longer something to fear, as long as the product of said reaction is benefi cial to the host. Hence, ceramics with two main features are used: bio-active and biodegradable.
Bioactivity is considered as the ability to directly attach to bone, without an interposed fi brous capsule.Therefore, there are several elements to consider, with very different dimen-sions: bones with micrometer porosity, bioapatites of nanom-eter size, cells, which shall in turn form bone tissue, with micrometer sizes and proteins, peptides and bone growth factors, with sizes in the nanometer range. The arrival of the third generation opened the doors to Biology. Second genera-tion materials are used as a starting point, although conformed and shaped as porous pieces (scaffolds). And moreover, it is intended to include an added value, by adequate functionaliza-tion: smart features and the ability to load them with biologi-cally active molecules. Therefore, human body repair may be tackled from two different approaches: the bionic approach, using fi rst and second generation materials, and the regenera-tive clinical practice approach, which includes cell therapy and tissue engineering, using third generation materials.
Second generation biomaterials are good candidates to be used in the manufacturing of scaffolds. Bioceramic scaffolds must comply with certain requirements of porosity, which must be equivalent to bone. Therefore, the fi rst step would be to fi nd
methods of conformation that yield pieces with interconnected porosity, with certain values of porosity in the range of microns, while preserving the nanometric dimensions of precursor materials.
2.1. Calcium Phosphates
The chemical similarity to the mineral component of mamma-lian bones and teeth has fueled the use of calcium phosphates as bone substitute materials. In fact, they can be employed with different shapes and functionalities within the clinical area. The reason for this popularity is that they are non-toxic, bio-compatible and exhibit a bioactive behavior, which leads to an intimate physicochemical bond between the implant and bone, i.e., osteointegration. Additionally, calcium phosphates are also recognized to be osteoconductive, able to provide a scaffold or template for new bone formation, and support osteoblasts adhesion and proliferation. [ 38 ] Several works dealing with cal-cium phosphate based implants have also demonstrated their capability for bone regeneration purposes. [ 39–41 ]
Understanding the solutions applied by Nature to different issues, and how said solutions may be a source of inspiration to solve technological problems in the mineral fi eld ( Figure 3 ) is of vital importance. Biomimetics can be considered as a tech-nology transfer from Nature to the mineral fi eld. A suitable working approach consists in understanding how nature works and try to copy its mechanisms. For this purpose, getting into the nano world is required, since many of the biological species we have to handle present nanometric dimensions. Chem-istry completely penetrates into the biotechnology and medi-cine fi elds. New techniques, new diagnosis routes, accurately
tracking patient health status, nanoelectronic devices, dosage and drug delivery and manufacturing more resistant prostheses are some examples of fi elds where the nano-scale chemical work is necessary.
If we look at how Nature solves the task of fabricating hard tissue, we will fi rstly fi nd that biomineralization processes mainly use calcium and silicon combined with carbonates, phosphates and oxides. [ 42 ] Thus, bone is formed by biominer-alization processes, natural sequences of physicochemical reac-tions that yield the formation of hard tissues in vertebrates or protective tissues in invertebrates and inferior zoological spe-cies. As a result, natural composites are formed. In that way, materials with exceptional mechanical properties that are impossible to obtain with pure materials are reached.
However, before dealing with the production of some ceramics in the laboratory, [ 43–47 ] we should recall that the inorganic phase of our bones is an apatite-like phase. Its structure has the spe-cial ability to accommodate several different ions in its three sublattices. Bone apatites can be considered as a basic calcium phosphate. As indicated above, bones of vertebrate animals are organic–inorganic composite materials whose structure can be described in short as follows: the inorganic component is car-bonated and calcium defi cient non stoichiometric hydroxyapatite. These biological apatite crystals exhibit nanometric size, ranging from 25 to 50 nanometers. [ 48 , 49 ] Such crystals grow at the min-eralization sites of the collagen molecules, which are grouped together forming collagen fi bres (see Figure 2 ). Furthermore, certain hierarchical bone porosity is necessary for several physi-ological functions performed by bone. [ 50 ] The order of magnitude of biological apatites and bone pores is shown in Figure 1 .
However, one of the main limitations of calcium phosphates is their poor mechanical properties, especially their brittleness
Figure 3 . Nature represents the best inspiration to solve technological problems associated to biomaterials. Biomimetics can be considered as a technology transfer from Nature to the mineral fi eld.
and poor fatigue resistance. This is even more evident for highly porous bioceramics and scaffolds, where porosity should be greater than 100 μ m. Consequently, for biomedical appli-cations, calcium phosphates are primarily used as fi llers and coatings. [ 51 , 52 ]
2.2. Bone Cements Based on Calcium Salts
Cements based on calcium phosphates, calcium carbonates or calcium sulphates, attracted much attention as biomaterials due to their excellent biocompatibility and bone repair proper-ties. [ 53–56 ] These cements do not have to be delivered in a prefab-ricated form, and this is a remarkable advantage of this type of material over conventional bioceramics. Most of the injectable calcium phosphates used evolve to an apatitic calcium phos-phate during the setting reaction. Physicochemical properties of these materials, such as setting time, porosity and mechanical behaviour, depend on the cement formulation and the presence of additives. [ 57–60 ] These cements are able to harden in situ , are biocompatible and may be slowly resorbed.
During this gradual process, the newly formed bone grows and replaces the cement. Some aspects that must be still improved are related to their mechanical toughness, setting time, application techniques on the osseous defect and the fi nal biological properties. Research is under way to get shorter set-ting times, even in contact with blood, and to improve mechan-ical toughness. [ 61 ]
2.3. Calcium Phosphate Coatings
As stated above, the poor mechanical properties of calcium phosphate ceramics impede their use in load bearing applica-tions. To overcome this drawback, coatings of hydroxyapatite
(HA) and other calcium phosphates over metallic substrates were synthesized. More common substrates are titanium alloys, Ti 6 Al 4 V and commercially pure Ti, because of their low density, good mechanical strength and resistance to cyclic loads.
Commercial implants coated with HA or other calcium phos-phates [ 62–64 ] are produced by plasma spray. The advantages of the method include rapid deposition rates and a relatively low cost. However, some problems need to be solved including the presence of resorbable amorphous calcium phosphate in the coatings, producing resorption problems. Some other issues to be faced are the adherence to substrate and the instability of HA at high temperatures or the phase transition of titanium at 1163 K. Thus, the effect on coatings of many synthesis param-eters or of the post heat treatments to crystallize the amorphous phases has been investigated.
On the other hand, other techniques have been used to obtain the coatings including physical vapor deposition (PVD), [ 65 ] chemical vapor deposition (CVD), [ 66 ] magnetron sput-tering, [ 67 ] electrophoretic deposition, [ 68 ] pulsed laser deposition (PLD), [ 69 , 70 ] and sol-gel based dip coating. [ 71 , 72 ] Some of these processes allow higher control of coating thickness and crystal-linity of phases. Other processes take place at lower tempera-tures than plasma spray, which involve interesting advantages.
Figure 4 shows the different ways in which calcium phos-phates can be used: as coatings for metal prostheses, dense or porous materials, powdered or in granules, as injectable cements. In all cases, the material may be loaded with drugs for local delivery in the area where the implant is placed, as dis-cussed in Section 4.1
2.4. Bioglasses
Another family of bioceramics, bioactive glasses, are com-posed of a network of silica modifi ed by the addition of certain
Figure 4 . Different ways in which calcium phosphates may be used.
network modifi ers, such as Ca, Na, and P, which are bonded to the network via non-bridging oxygen bonds. Bioactive glasses were fi rst developed by Hench when producing the bone graft known as Bioglass45S5®, with a composition of SiO 2 (45%), NaO 2 (24.5%), CaO (24.5%), and P 2 O 5 (6%). [ 73 ] Since then, these materials have solved different clinical problems, such as bone defects. The success of these materials has been based on their high biocompatibility and the positive biological effects after implantation. [ 74 , 75 ] In fact, these bioglasses bond to and integrate with living bone in the body without the formation of any fi brous capsule or promoting infl ammation or toxicity. [ 76 ] This bone bonding is promoted by the formation on the glass surface of a calcium phosphate layer that evolves to hydroxycar-bonate apatite when in contact with physiological fl uids. Thus, the formation of a biologically active hydroxycarbonate apatite on the glass surface has been proposed as the prerequisite for bonding to living tissues.
Bone tissue regeneration and integration have been widely studied in the fi eld of bioactive glasses. [ 77 , 78 ] Sol-gel glasses favor bone tissue growth from the periphery to the implant. The newly formed bone penetrates the implant, while the glass is slowly but progressively resorbed.
Porous bioglasses started to be produced when the sol-gel technique was applied to the synthesis of these materials, [ 79 ] which up to that moment had been produced following con-ventional melting methods. The employment of this technique allowed relatively low synthetic temperatures, [ 80–89 ] which facili-tated the incorporation of biologically active molecules [ 90 ] and living cells [ 91 ] during the synthesis. Additionally, the prepara-tion of bioactive glasses by the sol-gel method expanded the bioactive composition range of these materials.
The nanostructure of bioactive materials, glassy and hybrids, was recently studied for the fi rst time by high resolution trans-mission electron microscopy. [ 92 ] The nanostructural characteri-zation indicates that the addition of P 2 O 5 to the glass leads to the crystallization of a silicon-doped calcium phosphate, while in the materials without any phosphorous content–binary glass and hybrid–calcium is located in an amorphous silica network. The different rates of positive bioactive response of both glasses (with and without phosphorous) is strongly correlated with their nanostructure, since the distances between [SiO 4 4 − ] tet-rahedra decrease when calcium is not present in the vitreous network, and phosphorous bonds to calcium to form a silicon-doped calcium phosphate. [ 92 ] Figure 5 shows high resolution transmission electron microscopy images of glassy and hybrid bioactive materials.
However, the poor mechanical properties of bioglasses have limited the straight applicability in the biomedical industry. [ 93–95 ] Nowadays, the most promising application of these bioglasses in bone tissue engineering is the association with osteogenic agents to form three-dimensional scaffolds, as below com-mented in section 4 of this review. The matrix would present a bioactive behavior, which involves not only osteoconduction and osteoproduction, but also osteoinduction when implanted in living tissue. In addition, the graft would locally supply osteor-egenerative agents to the place where they are actually needed.
Apart from bioglasses, some other CaO- and SiO 2 -containing bioceramics have been reported for bone tissue engineering purposes. [ 96–98 ] In vivo bone regeneration and resorption rates of calcium silicate bioactive ceramics were higher when com-pared with β -tricalcium phosphate bioceramics. [ 99 , 100 ] Wu and colleagues studied the mechanical properties, in vitro bioactivity
Figure 5 . High resolution transmission electron microscopy of bioactive glassy (G) and hybrid (H B ) materials.
and biocompatibility of bredigite (Ca 7 MgSi 4 O 16 ) bioceramics prepared by ceramic sintering. Interestingly, this material was able to develop an apatite-like layer in simulated body fl uid after one week, and the dissolution products remarkably pro-moted cell growth in concentrations up to 50 mg/ml. [ 101 ] Diop-side (CaMgSi 2 O 6 ) has also been reported to show a bioactive behavior in vitro , improved cell proliferation in comparison with other Mg-containing ceramics, [ 102 ] and its biological eval-uation in rabbits has demonstrated a good integration with newly formed bone. In addition, mechanical strength and resistance to acids of this biomaterial were much higher than sintered hydroxyapatite. [ 103 ] Non-bioactive glass-ceramics with diopside, althausite (Mg 2 PO 4 OH) and hydroxyapatite as crys-talline phases in the composition have also been synthesized, in which a bioactive response was induced after acid chemical etching. [ 89 ] More recently, bioactive diopside scaffolds with varying porosity were shown to support human osteoblast like cell proliferation and differentiation. These scaffolds displayed a high interconnectivity and mechanical strength comparable to human spongy bone. [ 104 ] Porous akermanite (Ca 2 MgSi 2 O 7 ) scaffolds have also been reported to be bioactive, degradable and cytocompatible with bone marrow stromal cells [ 105 ] . Of note, this bioceramic showed similar proliferation behavior and enhanced differentiation of human adipose-derived stem cells, even in non-osteogenic media, after around 10 days when com-pared to β -tricalcium phosphate. [ 106 ]
Zr, Ti, Zn and Sr ions have been incorporated into calcium silicate bioceramics as well. In this respect, Ramaswamy and collaborators found improved proliferation and differentiation of human osteoblast like cells with Ca 3 ZrSi 2 O 9 (baghdadite) when compared to classic pseudowollastonite biomaterials. The attachment, growth and differentiation of osteoclasts
and endothelial cells were also supported in this material. [ 107 ] Remarkably, the osteogenic differentiation of bone marrow-derived mesenchymal stem cells in Zn-containing calcium sili-cate, hardystonite (Ca 2 ZnSi 2 O 7 ), was found to be higher than in β -tricalcium phosphate biomaterials. [ 108 ] The integration of Sr in this kind of bioceramics has led to the preparation of Sr-har-dystonite scaffolds with highly interconnected networks. These materials were reported to induce osteoconductivity in rat bone defects. [ 109 ] With regards to the incorporation of Ti ions into silicate bioceramics, Ca-Si-Ti based plasma sprayed coatings on Ti-6Al-4V alloys were shown to enhance the performance of equivalent hydroxyapatite coatings in terms of mechanical properties and cytocompatibility. [ 110 ]
With the aim to improve the performance of bioactive glass and calcium silicate bone scaffolds, polymer coatings were applied over these materials. Polymer layers on bioceramics have been shown to improve the mechanical stability by a bridging cracks mechanism. [ 111 ] As a result, porous CaSiO 3 scaffolds modifi ed with poly(D,L-lactic acid) displayed higher mechanical strength and lower dissolution rates. Moreover, the bioactive behavior in the material was preserved, and an improved attachment and viability of human bone-derived cells was reported. [ 112 ] From another perspective, Li et al . showed that the incorporation of wollastonite into bioceramic/polymer composites enhanced the hydrophilicity of polyhydroxybutyrate-polyhydroxyvalerate polymer. It was also concluded that the acidic degradation by-products of the polymer were effectively neutralized by a gradual release of Ca and Si ions from the wol-lastonite phase within a 3-week period. [ 113 ] An osteoinductive behavior has also been demonstrated in polymer containing bioactive glass composite fi lms. Tsigkou and collaborators evi-denced higher levels of alkaline phosphatase enzymatic activity
Figure 6 . 3-D reconstruction microtomography images of SBA 15 and osteostatin loaded SBA 15.
and osteocalcin protein synthesis when bioglass particles were incorporated into poly-D,L-lactide matrices. In these materials, extracellular matrix mineralization was found in the absence of osteogenic supplements. [ 114 ]
2.5. Silica Mesoporous Materials
In the last few years, silica-based ordered mesoporous mate-rials have attracted the interest of biomedical researchers due to their applicability in the context of drug delivery, [ 35 , 115 , 116 ] as dis-cussed later in section 4, and tissue engineering [ 117 ] , especially in bone tissue regeneration. [ 118 ] Ordered mesoporous materials, which were reported for the fi rst time in the 1990s, [ 119 , 120 ] are synthesized using surfactants as template of the mesostructure for the assembly and subsequent condensation of inorganic precursors. Upon surfactant removal, the materials present a network of cavities within the silica framework that determines their physicochemical properties. Among other characteris-tics, these materials present high surface area ( ca. 1000 m 2 /g), large pore volume ( ca. 1 cm 3 /g), regular and tunable mesopore diameter (2-50 nm) and pore channel systems homogeneously organized in 2-D or 3-D mesostructures. From the biomedical point of view, the surface of these materials, that is, the chem-ical groups located at their surface, would govern the interac-tion with the living body. Silica-based ordered mesoporous materials are composed of a silica network covered with silanol groups on both the inner surface of the pores and the external surface of the particles. Therefore, and taking into account that they present similar chemical surface to bioglasses, ordered mesoporous silicas should behave similarly to bioglasses. As a result, when immersed in a physiological fl uid they should develop nanoapatite coatings with a similar crystallinity to bio-logical apatites. [ 121 ] Thus, the implanted material would bond to living bone in the body, which makes these materials very attractive in bone regeneration technologies. [ 122 , 123 ]
An added value of these ordered mesoporous silicas is the possibility of hosting osteoinductive agents to promote and accelerate bone formation. This is the case of osteostatin, the 107-111 domain of parathyroid hormone-related protein (PTHrP). The whole protein is an important regulator of bone remodeling, and recent studies have shown that the native C-terminal PTHrP (107–139) fragment can inhibit osteoclastic bone resorption by affecting osteoclastic growth and/or dif-ferentiation; this effect has been ascribed to its N-terminal sequence 107–111 (Thr-Arg-Ser-Ala-Trp) called osteostatin. [ 124 ] When loading this pentapeptide into SBA 15 materials, a 2-D hexagonal silica-based ordered mesoporous material, different osteogenic features were observed in vitro .[ 123 ] In fact, when investigating PTHrP-SBA 15 materials in osteoblastic cell cul-tures, it was observed that the presence of the peptide increased cell growth and the expression of several osteoblastic products, such as alkaline phosphatase, osteocalcin, collagen, osteoprote-gerin, receptor activator of nuclear factor- κ B ligand and vascular endothelial growth factor.
However, in vitro studies only provide a basic estima-tion of the ability of certain cell types to survive in the pres-ence of a biomaterial. Therefore, in vivo studies are essential for an understanding of biocompatibility, with the frequently
discussed issue of in vitro - in vivo correlation (IVIVC) [ 125 ] . When implanting those previously mentioned osteostatin loaded ordered silica materials in a cavitary defect located in the epi-physis of rabbit femurs, the histological examination revealed the absence of signifi cant infl ammation or bone resorption within the time of study (4 and 8 weeks) after implantation. [ 122 ] After 8 weeks of implantation, μ CT analysis showed neoformed bone tissue around the biomaterials ( Figure 6 ). This animal model, that has been previously proven suitable for testing bio-materials, and is mainly based on trabecular bone and a minor cortical bone component at the edge of the defect, was used to highlight the bone regeneration action of these peptide-loaded bioceramic assessed by histological and μ CT analysis and immunohistochemistry.
PTHrP-loaded biomaterials were found to be highly osteo-conductive and osteoinductive by promoting infi ltration of proliferating osteogenic precursor cells and connective tissue formation. In addition, peptide-loaded bioceramics induced revascularization of the defect, related to an increased immu-nostaining for VEGF in the healing bone tissue. All these fi ndings demonstrated that the addition of osteostatin to the ordered silica material signifi cantly improved local bone induc-tion, which could be of great interest for future clinical applica-tions in bone tissue engineering.
Regarding the bioactivity of silica-based mesoporous mate-rials, a signifi cant development was carried out in 2004 by Yan et al., who synthesized for the fi rst time mesoporous bioactive glasses. [ 126 ] These materials show much better bioactive kinetics than pure silica mesoporous materials and even than sol-gel glasses with similar chemical composition (SiO 2 -CaO-P 2 O 5 ). [ 127 ] This is a consequence of their enhanced textural properties; the large surface area results in a higher chemical reactivity, and thus the bioactive process is highly favoured.
Different research groups have focused their attention on mesoporous bioactive glasses during the last decade. [ 128–133 ] They have been also proposed as constituents of cements [ 134 ] and synthesized in the form of microspheres in order to improve the control over drug release. [ 135 ] Nevertheless, most of
recent works in the fi eld points them as ideal components for the fabrication of scaffolds for bone tissue engineering [ 136–139 ] . This fact, together with the outstanding bioactive kinetics, the possibility of releasing active molecules in a controlled fashion and the recently shown biocompatibility of mesoporous bioac-tive glasses [ 140 ] explain the interest of using them for the pro-duction of scaffolds. Several techniques have been used with this purpose: sintering, macroporous templates, 3D rapid pro-totyping techniques or silk modifi ed systems. [ 141 ]
The investigation on these materials has advanced very quickly in only eight years. Thus, it can be concluded that mes-oporous bioactive glasses are a promising research line in the fi eld of bioceramics for bone regeneration.
2.5.1. Antifouling Surfaces
Every surface is a possible place for several different species to attach. If we manage to avoid unspecifi c attachment and selec-tively control molecules deposition, the application is obvious. In materials with high porosity and surface area, this effect is highly important. In this section, two examples of highly bioac-tive ordered mesoporous glasses are presented. [ 132 ]
Bacteria and proteins are two candidate species to colonize biomaterial surfaces. Protein attachment may be a benefi cial process when it takes place specifi cally. [ 142 ] However, random adhesion of bacteria and proteins constitute a serious problem for many applications in the medical fi eld.
Zwitterion polymers are a promising tool to solve these prob-lems. [ 143 , 144 ] The formation of a hydration layer via electrostatic interaction and hydrogen bonding generates repulsive forces that lead to antifouling surfaces. In the case of mesoporous materials, these surfaces may have a high resistance to bac-terial adherence, and therefore to the formation of biofi lms, which causes serious bone infection. If successful, it would be a suitable way for preventing implant infection. Indiscriminate
Figure 7 . Antifouling properties of functionalized SBA-15 ordered mesopor
adsorption of proteins is another problem which might fi nd a solution with this strategy. And once again we are copying nature mechanisms.
To prepare these surfaces, ordered mesoporous materials walls have been functionalized with hydroxyl, amino and carboxyl [ 145 ] groups ( Figure 7 ). Highly ordered zwitterionic SBA-15 type nanostructured materials bearing–NH 3 + /–COO − or–SiO − /–NH 3 + groups and exhibiting 2D-hexagonal structures have been prepared by a co-condensation route. The biocompat-ibility of these mesoporous materials and their high resistance to bacterial adherence have been demonstrated by Izquierdo and coworkers. [ 146 ]
The capability of these materials to inhibit bacterial adhesion has been demonstrated under in vitro conditions simulating severe infl ammation/infection, which is usually associated to a decrease in normal pH values. Moreover, the in vitro biocom-patibility of these zwitterionic mesoporous materials has been evaluated obtaining successful results. Thus, it has been evi-denced that Saos-2 osteoblasts adhere and proliferate on and around these biomaterials maintaining their characteristic mor-phology. Measurements of reactive oxygen species content and lactate dehydrogenase release to the culture medium of Saos-2 osteoblasts indicate that SBA-15 type mesoporous materials do not induce oxidative stress in this cell type and confi rm the integrity of osteoblast plasma membrane.
In vitro bacteria adhesion assays using Escherichia coli onto dif-ferent SBA-15 nanostructured ceramics have been performed. For this purpose, pure silica, -NH2 or –COOH monofunctionalized, and –NH2/–COOH bifunctionalized SBA-15 mesoporous mate-rials have been used. In this way, surface amino groups and acid groups are displayed at the same time. Comparing the results obtained in Escherichia coli adhesion on SBA-15 with hydroxyl, hydroxyl and carboxyl, hydroxyl and amine, and carboxyl and amino groups, some differences are found. When only hydroxyl groups are present, bacteria colonize the surface. This situation
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is not enhanced in the case of carboxyl groups. Amine functionalization represents a slight enhancement, but the amino-carboxyl group confi guration is clearly antifouling. However, further improvement is still needed.
Scanning micrographs collected from these materials surfaces are presented in Figure 8 . As shown, only in the case of SBA15 ATPES CES, bacteria are diffi cult to fi nd on the sur-face of the material. In contrast, the other three types of materials are totally colonized.
Hence, materials characterization reveals that both, –NH 2 /–COOH and –NH 2 func-tionalized SBA-15 materials exhibit a zwitteri-onic character due to the presence of –NH 3 + /–COO − or–NH 3 + /–SiO − moieties, respectively. In vitro adhesion assays have been carried out at the pH in which the zwitterionic nature of both these samples is preserved, i.e. pH 5.5. Results show that the presence of both positive and negative moieties with an overall neutral charge leads to a reduced E. coli adhesiveness (Figure 7 ). In vitro tests with cultured human Saos-2 osteoblasts at the physiological pH of
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Figure 8 . Scanning electron microscopy images assessing the antifouling properties of functionalized SBA-15 ordered mesoporous silica in the presence of bacteria.
7.4 demonstrate that all materials exhibit good biocompatibility, where Saos-2 osteoblasts adhere and proliferate maintaining their morphologic and functional characteristics. This novel family of zwitterionic mesoporous materials opens up promising expecta-tions in diverse biomedical applications such as preventing some side effects associated to bone implant infection. [ 146 ]
Furthermore, the ordered mesoporous arrangement is an added value for this new family of zwitterionic ceramic mate-rials because they could load and subsequently release in a sustained way large amounts of biologically active molecules, such as certain drugs, peptides and small proteins. Therefore, this novel family of bioceramics represents a suitable candi-date in the fi eld of implantable medical devices. The attach-ment of proteins may be managed in a similar way, when a surface must be protected from indiscriminate adsorption. In this case, only amine-carboxyl combinations achieve good results (Figure 7 ). Similarly, the co-condensation route has been employed to bifunctionalize SBA-15 with amine and carboxylic acid groups. [ 145 ] The functionalization process following a one-step route does not affect the mesostructural order of SBA-15, as confi rmed by XRD and TEM, originating mesoporous matrices with outstanding features suitable for purposes that require host matrices with relatively large mesopores, surface areas and volumes. The zwitterionic nature of this material has been evidenced by XPS, FTIR and ζ -potential. Moreover the ultralow-fouling behaviour of this zwitterionic ceramic towards the adsorption a model protein has been confi rmed. This novel generation of zwitterionic ceramics has a great potential of application in catalysis, sensing, biotechnology and biomedi-cine. It should also be remarked that these materials display reactive–COOH and–NH 2 groups able to act as coupling sites to covalently immobilize recognition elements for specifi c uses.
3. Bone Tissue Engineering
According to the defi nition of tissue engineering given by Langer and Vacanti, [ 147 ] bone tissue engineering can be defi ned
as an emerging interdisciplinary fi eld that seeks to address the needs by applying the principles of biology and engineering to the development of viable substitutes that restore and maintain the function of human bone tissues. The term “tissue engi-neering” was offi cially coined at a National Science Founda-tion workshop in 1988 to mean “the application or principles and methods of engineering and life sciences toward the fun-damental understanding of structure-function relationships in normal and pathological mammalian tissues and the develop-ment of biological substitutes to restore, maintain or improve tissue function”.
The challenge of tissue engineering is to mimic what hap-pens in Nature. [ 148 , 149 ] Attempts are being made to engineer in vitro practically every tissue and organ in the body. Work is proceding in creating tissue-engineered liver, nerve, kidney, intestine, pancreas and even heart muscle and valves. In the area of connective tissues, work has been ongoing worldwide for many years in the engineering of tendon, ligament, bone and cartilage. One of the most employed approaches aimed at the creation of new tissue involves the initiation of the regen-eration process in vitro by soaking the scaffold in an appropriate cell culture and in the presence of tissue-inducting substances such as certain peptides, hormones and growth factors. Sub-sequently, the scaffold is implanted into the patient. On the other hand, another approach could be to chemically graft such tissue inductive substances, i.e. peptides or growth factors, into the 3D scaffold to be directly implanted into the patient. These osteogenic agents would act as signals that would induce cells to regenerate bone ( Figure 9 ). This feature is essential to stimu-late natural healing processes aimed at repairing living tissues, which makes tissue engineering a revolutionary approach for end-stage diseases.
In the design of bioceramics, and biomaterials in gen-eral, it is necessary to identify the biological environment that they will fi nd once implanted. In the case of bone, to be able to regenerate new bone, it is important to understand its hierarchical structure. Bone is a natural composite material made of collagen (organic polymer) and carbonate hydroxya-patite (inorganic ceramic). Collagen is a triple helix of protein chains that has high tensile and fl exural strength and provides a framework for bone. Bone mineral is a crystalline calcium phosphate ceramic (carbonate hydroxyapatite) that provides the stiffness and high compressive strength of bone. [ 150 ] There are mainly two types of bone: cortical and cancellous bone. Cor-tical bone, also called compact bone, is a dense structure with high mechanical strength. Cancellous or trabecular bone, also called spongy bone, is less dense and weaker compared to cor-tical bone, due to its porous structure. It is highly vascular and frequently contains red bone marrow, where the production of blood cells takes place.
The diffi culty to provide a suitable vascularization to 3D scaffolds for oxygenation of newly formed tissues is a main drawback for regeneration. The amount of oxygen required for cell survival is limited to a diffusion distance between 150 and 200 μ m from the supplying blood vessels, [ 151 ] so the success of 3D constructed tissues strongly depends on angiogenesis.
Last trends in bone tissue engineering are focused on tissue regeneration, rather than tissue substitution, so it is important to understand bone formation mechanism, called osteogenesis.
Figure 9 . The three pillars of Tissue Engineering.
Bone tissue formation is a complex process, which can be sim-plifi ed as the process where the extracellular matrix of mineral-izable collagen is laid down by osteoblasts. They secrete type I collagen which then mineralizes to form a carbonated hydroxya-patite-collagen structure. Bone changes its structure and shape continuously in response to its local loading environment within the body. This process is called bone remodeling, and is carried out by two types of cells: osteoblasts and osteoclasts. Osteob-lasts are responsible for the synthesis of bone matrix and bone formation, while osteoclasts are able to degrade the mineralized matrix, that is, old bone and bone that is not required (bone resorption). Ideally, there should be a balance between the rates of bone gain and bone loss. However, disruption of this equilib-rium occurs in metabolic bone diseases such as osteoporosis, which is characterized by increased bone remodeling related to a high bone resorption, leading to an increase in fracture risk. Additionally, when minor damage occurs in bone tissue, it can repair itself by the biochemical activity of osteoblasts. However, when the defect exceeds a critical size, which might be produced from trauma or from the removal of diseased tissue, bone is not able to repair itself. To solve this problem, graft implants and synthetic bone fi ller materials are currently being used. In a more modern approach, a tissue engineered scaffold could guide and stimulate bone growth in the so-called approach of regenerative biomedicine. This scaffold should present proper-ties either similar to trabecular bone or the capability of stimu-lating new bone growth and create a biocomposite with similar structure and properties to trabecular bone. In general, an ideal scaffold for bone tissue engineering should fulfi ll some gen-eral criteria. [ 152 ] It should act as a template for in vitro and in vivo bone growth in 3 dimensions, the scaffold should resorb at the same rate as new bone tissue is formed, the scaffold mate-rial should be biocompatible and promote cell adhesion and activity, the scaffold should bond to the host bone avoiding the formation of scar tissue, the scaffold should exhibit mechanical
properties matching that of the host bone after in vitro tissue culture, the processing technique of the scaffolds should be able to produce irregular shapes to match that of the bone defect and the scaffold should fulfi ll the international standard requirements for clin-ical use, including sterilization.
Among all the criteria mentioned, the need of a porous structure seems quite straightfor-ward. In fact, the scaffold must present an open porous structure to allow cell penetra-tion, tissue ingrowth and eventually vascu-larization on implantation. The minimum interconnected pore aperture diameter is recognized to be 100 μ m, to allow the vas-cularization required for complete bone regeneration.
In addition, the material to build the scaf-fold should be carefully considered. In fact, bioactive materials stimulate a biological response from the body such as bonding to tissue and stimulating of bone growth along the material surface. This kind of materials is required in order to attract cells and to stimu-
late their migration and growth. On the other hand, biodegrad-able materials and the chemical cues are required in order to promote the regeneration process. [ 153 ] Examples of materials that may be both bioactive and biocompatible are porous cal-cium phosphates, porous bioglasses and silica mesoporous materials.
3.1. 3D Scaffolds
Synthetic scaffolds are aimed to provide temporary mechanical stability until new bone becomes synthesized, organized, and consolidated into a stable structure. Nowadays, not only a wide range of biomaterials but also several processing techniques for the elaboration of scaffolds with tailored architectures and porosities have become available. Bone tissue engineering scaf-folds will succeed depending on both the intrinsic properties of the biomaterial and the features imposed by the fabrication method.
In the particular case of ordered mesoporous materials, pore dimensions are within the range of 2–50 nm, which are far from those of living cells, which are within 10–200 μ m. This fact makes it impossible to cells to penetrate into the mesopores. Hence, taking into consideration that bone cells rule the bone regeneration process, ordered mesoporous materials should be somehow processed to acquire macroporosity. Attending to the hierarchical structure of Nature, bone porosity ranges between 20 and 400 μ m, [ 50 ] and it is necessary for several physiological functions played by bone ( Figure 10 ). It is for this reason that ordered mesoporous materials should be employed to build scaffolds with porosity similar to natural bone, so it would allow bone cell penetration, adherence, growth and prolifera-tion that would lead to bone in-growth and therefore vasculari-sation after implantation. In addition, the scaffold processing method should preserve the original mesoporosity, so it would
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Figure 10 . Porosity features involved in the design of scaffolds for bone tissue engineering.
be possible to combine macroporosity for bone oxygenation and mesoporosity to allow drug delivery of certain biomolecules. By this way, hierarchical macroporous bioceramics with intercon-nected porosity and sizes in the order of microns would be suit-able as scaffolds for bone tissue engineering.
Nowadays, there are several conformation methods which allow to obtain pieces at room temperature. [ 154 , 155 ] Besides,
Figure 11 . Previous preparation and rapid prototyping 3D printing method for tissue engineerin
working at room temperature allows including biomolecules of interest in many cases to treat different diseases, or to improve the treatment of various bone pathologies. A very appropriate and useful method to pro-duce scaffolds is the rapid prototyping 3D printing ( Figure 11 ). Having prior tomog-raphy information of the tissue to replace, and with a computer assisted design, suitable scaffolds for each bone repair situation may be obtained. Biocompatible and biodegrad-able bioceramics are used for this purpose; a bioceramic paste is prepared to load the device and to mould the scaffold with a pre-viously designed porosity. A combination of different methods, such as sol-gel chemistry, double polymer templating and rapid proto-typing, has lead to 3D mesoporous bioglass-based scaffolds exhibiting hierarchical pore networks, with giant (30–1000 μ m), macro- (10–30 μ m) and meso- (5 nm) porosity. [ 136 ] This fabrication of scaffolds via rapid pro-totyping allows the necessary variations in shapes, and it may be adapted to the require-ments of different tissues and organs.
Considering the fi nal application of this 3D scaffolds for bone tissue engineering, they should be capable of driving cell in-growth. An interesting approach is the covalent grafting of oste-oinductive agents, such as certain peptides and growth factors, which would act as signals to induce cells to regenerate bone. Thus, when ordered mesoporous materials are used as starting materials for the fabrication of scaffolds, the osteoconductive
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Figure 12 . Bone regeneration and local drug delivery by mesoporous-based scaffolds.
agents can be easily grafted to the surface of the silica. In this manner, the scaffold would be decorated with potent osteoin-ductive signals able to promote the appropriate bone cellular functions in the place where they are needed. [ 156 ] In addition to this, the possibility of using mesoporous materials as starting materials to build 3D scaffolds confers the resulting biocer-amics a very important added value: they would be capable of hosting drugs for locally treating different bone pathologies, such as bone infection, osteoporosis and cancer.
Within the application of mesoporous matrices for bone tissue engineering, special emphasis has been given to the development of organic-inorganic hybrid bioceramics resulting from the strong interaction via covalent linkages of osteoinduc-tive agents (peptides, hormones, growth factors, etc) to the scaf-fold surface. These osteoinductive agents would act as attracting signals for bone cells, and the appropriate hierarchical porosity of the scaffold, similar to that of natural bone, would permit cell adherence, proliferation, bone in-growth and afterwards vascularisation on implantation ( Figure 12 ). Novel drug delivery systems for the local treatment of bone diseases, such as infec-tion, osteoporosis, cancer, etc. represents an additional option. Controlled release systems involve the development of hybrid bioceramic matrices exhibiting weak host-guest matrix-drug interactions. Within bioceramics, silica-based ordered mes-oporous materials have been widely reported as good candi-dates to design controlled drug delivery devices. The infl uence of the textural and structural properties of such matrices on drug load and release behaviours is of paramount importance, as discussed in section 4. Chemical strategies for functional-izing mesoporous silica walls with different organic groups able to interact with the chemical groups of the targeted drug
molecule have been also tackled. Following these approaches, it is possible to achieve higher control over drug load and release processes.
3.2. Hydroxyapatite Foams
Novel 3D-macroporous biopolymer-coated hydroxyapatite (HA) foams are prepared as potential devices for tissue engineering scaffolds. [ 157 ] The synthesis is performed by the sol-gel method using pluronic F127 as porosity induction agent. [ 158 ] In this way, pieces with a hierarchical porosity equivalent to bone, where pore size ranges from 10–15 nm to 1–400 microns, are obtained ( Figure 13 ). These foams are coated with biocompatible poly-mers by a dip-coating method. Glutaraldehyde cross-linked gelatin [ 159 ] and polycaprolactone are employed for this purpose. In all cases, the porosity values of the hydroxyapatite foams are preserved. Additionally, the coating improves the mechanical behavior of these foams, which undergo swelling effects in contact with aqueous solutions or in contact with blood. This feature makes them more fl exible and resistant. As a result, the foams are perfectly manageable by surgeons, who may easily implant them (see Figure 14 ). Thus, conformation-free one step hydroxyapatite scaffolds can be obtained, presenting a bone-like hierarchical porosity. These HA foams have applications in tissue engineering, as well as in other health-related fi elds.
Indeed, as is well known, HA is biocompatible and its crystal structure is tolerant to many ionic substitutions and complete replacement of Ca 2 + by Ba 2 + , Sr 2 + , Cd 2 + , Pb 2 + , Zn 2 + and Cu 2 + is possible. [ 160 , 161 ] Although HA foams are commonly used in tissue engineering in combination with polymers as effective
Figure 13 . Hydroxyapatite foam piece and HRTEM-ED analyses.
enhancers of their mechanical properties, [ 162 , 163 ] ceramic HA macroporous foams have been designed to improve and increase ionic exchange and heavy metal immobilization due to their higher diffusion and mass transport. [ 161 ] In this case, they may
Figure 14 . Biopolymer-coated hydroxyapatite foams may fi nd application ei
be biopolymer-coated for decreasing their solubility (normally very high in acidic media as the one in the stomach pH = 1.2) to avoid large toxic amounts of calcium in the digestive tract, dis-aggregation of HA pieces and to improve their handling. [ 149 ]
ther as bone implants or as heavy-metal immobilization agents.
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Therefore, these novel 3D-macroporous biopolymer-coated hydroxyapatite foams are potential devices for the treatment of heavy-metal intoxication by ingest. These foams are designed to exhibit a fast and effi cient metal ion immobilization into the HA structure in acidic media. The capture process of metal ions is stable, and no metal ions are released when the foams are soaked in clean basic media afterwards. These two steps mimic a digestion process. In fact, biopolymer-coated HA foams have shown a fast and high effi ciency on cap-turing Pb 2 + , Cd 2 + and Cu 2 + ions in acidic media, keeping these metal ions in clean basic media. All the process is carried out without any disintegration of the foam. These systems have been proposed as very effi cient antidotes against metal ions intoxication before intestinal absorption and hospital treat-ment (Figure 14 ).
Similarly, 3D-macroporous biopolymer-coated hydroxya-patite foams may also be applied for water treatment. Hence, these materials have been developed as potential devices for the treatment of lead, cadmium and copper contamination of con-sumable waters. These foams have exhibited a fast and effec-tive ion metal immobilization into the HA structure after an in vitro treatment mimicking a serious water contamination case. For this application, polycaprolactone and gelatine cross-linked with glutaraldehyde biopolymer coatings are aimed to improve HA foam stability at contaminated aqueous solutions pH, as well as its handling and shape integrity. Metal ion immobiliza-tion tests have shown high and fast heavy metals capture as a function of the biopolymer hydrophilic behaviour. After an in vitro treatment, foam morphology integrity is guaranteed and the uptake of heavy metal ions rises up to 405 μ mol/g in the case of Pb 2 + , 378 μ mol/g of Cu 2 + and 316 μ mol/g of Cd 2 + . These novel materials promise a feasible advance in devel-opment of new, easy to handle and low cost water purifying method [ 164 ] .
4. Ceramic Materials As Potential Drug Delivery Systems
In the last few years, research on drug delivery systems has been intensifi ed on both academia and biomedical industry due to the features and possibilities that these systems offer to bio-medicine. Among others, drug delivery systems allow several drugs to be administrated using new therapies improving effi -cacy and safety, they facilitate delivering new complex drugs that otherwise would not be possible, they promote improvement of therapeutic responses with continuous drug release patterns rather than pulsatile, and they also offer recent advances in materials science and biotechnology to develop new methods of delivery. [ 165 ] Nevertheless, the development of targeted delivery systems, in which the drugs are directed where they are specifi -cally needed, has supposed a real advance in this type of tech-nologies. Concretely, in the context of bone tissue engineering, implantable drug delivery systems for local drug release in bone tissue are promising tools for orthopedic surgery. Thus, the production of multifunctional materials with the ability of repairing bone tissue, while locally delivering a biologically active molecule, should be an important breakthrough in bone disease treatments.
The research on controlled drug delivery systems using bioce-ramics as host matrices presents two distinct sides: one route aims at embedding pharmaceuticals in biomaterials designed for the reconstruction or regeneration of living tissues, in order to counteract infl ammatory responses, infections, bone carci-nomas and so forth, while the other route deals with the more traditional drug introduction systems, i.e. oral administration.
The incorporation of pharmaceuticals into bioceramic matrices could be very interesting in clinical practice. [ 166 ] Nowa-days, it is rather common for an orthopedic surgeon working in bone reconstruction to use bioceramics. An added value to the production of these ceramics would be the optional addi-tion of pharmaceuticals such as antibiotics, anti-infl ammato-ries, anti-carcinogens, etc. In this sense, if we take into account the infections statistics at hip joint prostheses, the incidence varies between 2 and 4%, reaching up to a 45% in bolts used as external fi xation. One of the main problems in these situations is the access to the infected area of the bone, in order to deliver the adequate antibiotic. If the pharmaceutical could be included within the implant itself, the added value would be straightfor-ward. In search of a method for obtaining bioactive implants, which in turn operate temporarily as delivery systems of antibi-otics, two good candidates are bioactive glasses and gentamicin. Bioglasses have the ability to regenerate bone, and gentamicin is a broad-spectrum antibiotic employed in traumatology. The stabilized bioactive glass should be mixed with gentamicin to form a homogeneous mixture, which may then be formed into pieces. Normally, these pieces are stabilized by heat treatment up to several hundred degrees. This process is not possible in this case due to gentamicin. Alternatively, uniaxial compacting and isostatic pressing can be combined, both at room tempera-ture ( Figure 15 ). [ 167 ] The pieces obtained by this procedure may be designed for use in orthopedic surgery as bone defects fi llers preventing or fi ghting against bone infections.
These pieces have been tested in vivo in New Zealand rabbit femurs for 1, 4, 8 and 12 weeks. The biological response of bone to implant is perfect osteointegration, cortical and spongy bone growth, eventually allowing defect restoration. Besides, bone tissue is in contact with the implant surface, without interposi-tion of fi brous tissue, and the implant is partially resorbed in the medium term. [ 167 ]
Bone tissue regeneration and integration have been widely studied in the fi eld of bioactive glasses. [ 77 , 78 ] Sol-gel glasses favor bone tissue growth from the periphery to the implant. The newly formed bone penetrates the implant, while the glass is slowly but progressively resorbed (Figure 15 ).
4.2. Silica Mesoporous Materials
Among all available drug delivery technologies, silica-based ordered mesoporous materials fulfi ll the necessary require-ments to be employed in bone tissue engineering: their porosity and textural properties have promoted their use as drug delivery systems, [ 116 ] while their chemical composition, similar to bioglasses so the material would bond to living bone, allows their use in osseous regeneration technologies. [ 121 ] In
Figure 15 . Processing of antibiotic-loaded bioglasses to produce bone implant for local drug delivery.
this sense, ordered mesoporous silicas can be employed to manufacture scaffolds for bone regeneration that can locally deliver drugs for the treatment of bone pathologies and com-plex diseases [ 168 ] ( Figure 16 ).
In recent years, research on organic-inorganic hybrid mate-rials has become an important subject of study for materials and biomaterials sciences. The concept of organic-inorganic hybrid
Figure 16 . Mesoporous silica matrices allow drug loading and controlled r
materials appeared in the 1980s with the expansion of soft inor-ganic chemical processes. [ 169 ] Organic-inorganic hybrid materials can be generally defi ned as materials with closely mixed organic and inorganic components. [ 170 ] Furthermore, hybrids are either homogeneous systems derived from monomers and miscible organic and inorganic components, or heterogeneous systems (nanocomposites) where at least one of the domains (inorganic
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or organic) has a dimension ranging from a few angstroms to a few tens of nanometers. It is worth mentioning that the properties of hybrid materials are not only the sum of the individual contributions of both phases, but also a large synergy is expected from the inti-mate coexistence of the two phases through size domain effects and nature of the inter-faces. [ 171 ] Generally, the behaviour of hybrid materials is dependent on the nature and rel-ative content of the constitutive inorganic and organic components, with a close dependence on the experimental conditions.
The clinical need for developing bioac-tive materials for bone regeneration, i.e. those materials capable of forming a bioac-tive bond with natural bone, has motivated much research effort to develop bioactive hybrids. [ 172 ] One of the main advantages of these materials aimed at bone implant tech-nologies is that they have the unique feature of combining the properties of traditional materials such as ceramics and organic poly-mers on the nanometric scale [ 173 ] .
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During the last years, the application of organic-inorganic hybrid materials has notably spread out within the bioceramics world. Thus, hybrid materials can be classifi ed as a function of the type of interaction that takes place between the inorganic and the organic constituents of such materials. One of the sim-plest examples is the use of porous inorganic matrices able to host organic molecules, such as drugs, to subsequently act as sustained release systems. In this case, there is a weak interac-tion between the host inorganic matrix and the guest drug, as organic component.
Within the fi eld of bioceramics, silica-based ordered mesopo-rous materials are receiving a growing interest by the bioma-terials scientifi c community due to their capability to host dif-ferent guest molecules. [ 174 ]
When silica-based ordered mesoporous materials are intended as drug delivery systems, the host-guest interaction takes place between the silanol groups covering the surface of the mesoporous channels and the functional groups of the drug. The parameters that govern drug adsorption and release processes mainly depend on the textural and structural proper-ties of the host-matrix.
The complexity of these hybrid systems increases to adapt their properties to specifi c clinical needs. In many cases, it is necessary to organically modify the mesoporous silica walls through the covalent attachment of functional groups. This functionalisation process leads to hybrid mesoporous mate-rials [ 175 ] that can act as host matrices of a wide range of drugs via weak interactions.
The functionalisation of the silica walls may be necessary for several reasons. In some cases, there are certain drugs with remarkable hydrophobic nature that do not exhibit any trend to penetrate into the hydrophilic mesoporous silica. The function-alisation with hydrophobic functional groups is a good alterna-tive to promote the load of different hydrophobic drugs. This
Figure 17 . Possible interactions of mesoporous silica with therapeutic delivery and bone tissue engineering applications.
vstrategy is also employed to delay the release kinetics of cer-tain drugs from mesoporous channels to the aqueous release medium due to the decrease in the wettability degree of the material surface. On the other hand, there are other situations in which the pharmaceutical molecule can be confi ned into the mesoporous channels. However, higher loads and slower release kinetics can be achieved if the mesoporous silica wall is functionalized with different functional groups. Among the existing organic groups, the functionalisation with amino moie-ties has been widely reported.
In other situations, and above all considering the possible employ of these mesoporous materials as starting materials for the manufacture of 3D scaffolds suitable for bone tissue engi-neering, a strong interaction such as covalent linkages between the host matrix (inorganic or hybrid) and the osteoinductive agent (peptides, hormones and growth factors) is necessary. This feature leads to hybrid scaffolds where these osteoinduc-tive agents act as signals to attract cells responsible for new bone formation. On the other hand, the interconnected macro-porosity of the scaffold allow the cellular adherence, growth, bone in-growth and vascularisation, which are the essential stages for the formation of new bone.
Therefore, organic-inorganic hybrids have remarkable rel-evance within the bioceramics fi eld and they are of great utility in many medical applications. All these aspects are schemati-cally depicted in Figure 17 .
Weak interaction forces between drug molecules and the bioceramic inner surface are established in the case of drug delivery systems. The possibility of locally delivering phar-maceutical molecules from implantable bioceramics can be achieved with the use of silica-based ordered mesoporous materials acting as controlled delivery systems. [ 115 , 176 , 177 ] This property of mesoporous silicas emerged in 2001, when it was fi rst shown that ordered mesoporous silica of the M41S family
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molecules in drug
could be used as a drug delivery system with improved drug adsorption features and predictable release kinetics. [ 174 ] Since then, other mesoporous structures such as SBA-15, FDU-5, MCF, etc. have also been proved to act as controlled delivery systems. [ 178 , 179 ] Therefore, it has been shown that both small and large molecular drugs can be loaded within the mesopores by adsorption proc-esses, using an impregnation method, and subsequently being released via diffusion controlled mechanisms.
Mesoporous silica provides an excellent matrix for guest molecules, due to several attractive features among their textural and chemical properties, which play an impor-tant role in the loading and release of mole-cules, i.e. governs the host-guest interactions between bioceramic matrix and organic guest (Figure 17 ). [ 116 ]
Pore diameter acts as a size selector for the intake of guest molecules within the mes-opore channels. Moreover, this parameter is a limiting factor for the diffusion of the mole-cules to the delivery medium, thus regulating
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the release rate. The adsorption of guest molecules is a sur-
face phenomenon and consequently the amount of molecules loaded will be determined by the surface area of the matrix, i.e. matrices with higher surface areas will be able to incorporate a higher amount of drug due to their higher contact surface. Larger fi lling of the mesopores can be attributed to increased drug-drug intermolecular interactions within the pore voids. In this case, the pore volume may raise drug loading.
Regarding chemical properties, these materials are thermally and chemically stable. For the aimed application as local drug delivery systems in bone implants, a prerequisite fulfi lled by silica based bioceramics is their non toxicity and biocompati-bility. [ 15 , 180 ] Moreover, their in vitro bioactive behaviour has been demonstrated since they are able to develop an apatite-like layer similar to the bone mineral phase onto their surfaces when in contact with physiological fl uids. [ 115 , 181 ] Additionally, tradi-tional sol-gel glasses can be obtained with ordered mesoporous structure. Hence, they are called ordered mesoporous bioac-tive glasses and their preparation combine the sol-gel synthesis of glasses with the use of a surfactant as structure directing agent. With this strategy ordered mesoporous glasses exhibit enhanced bioactive behaviour compared to conventional ones, being their bioactive response time of ca. 1h, meanwhile it is 3 days for traditional glasses. [ 177 ]
The chemistry of silica surface is mainly determined by the presence of abundant silanol groups. [ 115 ] Therefore, these matrices interact with guest species through weak interactions such as Van der Waals forces or hydrogen bonds. Noteworthy, rich silanol containing surface may be organically functionalized through sol-gel chemistry via two main different routes, one-pot synthesis (co-condensation) and post-synthesis (silylation). [ 175 ] The former, also known as one-pot route, involves the addition of the modifying agent during the mesoporous synthesis and all the functionalisation process is carried out in a single step. In the post-synthesis method, the organic modifi cation process is performed under anhydrous conditions once the pure inor-ganic silica matrix has been already formed. The main differ-ence between hybrid mesoporous materials resulting from both functionalisation methods is the organic modifi cation degree. When using the co-condensation method, the organic groups are grafted to the outer and to the inner part of the silica walls. Consequently, the organic functionalities cannot exceed 40 mol% to avoid disordering of the mesoporous structure. On the other hand, after modifi cation using post-synthesis method, the functionalisation agent is located in the outer surface of mes-opore silica walls, leading to a higher functionalisation degree.
Chemical modifi cation of the ceramic pore walls with appropriate functional groups that undergo attracting inter-actions with biologically active molecules allows an effective control of the guest molecule adsorption and subsequent sus-tained release. [ 116 ] Hence, the organic modifi cation provides specifi c host-guest interactions of the bioceramic matrix with drug molecules. In the resulting organic-inorganic hybrid mesoporous matrices, these interactions with the loaded mol-ecules are usually established through electrostatic attracting interactions, hydrophilic-hydrophobic forces or electronic interactions. [ 116 , 182–184 ]
Mesoporous materials are of interest as drug delivery sys-tems and would be suitable candidates as starting materials for
the manufacture of 3D scaffolds. [ 185 ] The porosity of the scaf-fold implies a certain sacrifi ce of their mechanical properties. Hierarchical macroporous materials, with interconnected pore sizes in the order of microns, are suitable as scaffolds for bone tissue engineering. [ 186 ] As explained above, the challenge is to obtain scaffolds combining macroporosity for bone oxygena-tion and vascularisation and mesoporosity to allow the load and release of drugs. Then, the main goal is to design materials that can assist the human body to improve its regeneration proper-ties, not only recovering the structure of the damaged tissue, but also its function.
The manufacture of 3D scaffolds capable to drive cell in-growth is an important task in bone tissue engineering, as stated in section 3. The aim is to prepare porous ceramics that act as support for the newly formed tissue able to drive self-regeneration of bone tissue starting from the most suitable material. Moreover, the appropriate functionalisation of the scaffold would allow the covalent grafting of osteoinductive agents, such as certain peptides and growth factors that would act as signals to induce cells to regenerate new bone.
When mesoporous materials are used as starting materials for the fabrication of scaffolds, such osteoinductive agents should be covalently grafted to the external surface of silica. This approach would allow to ‘decorate’ the scaffold with potent osteoinductive signals able to promote the appropriate bone cellular functions in the place where needed. [ 151 , 187 ] So the fi rst step would consist in optimizing the functionalisation method to specifi cally modify the external surface of mesopo-rous silica and subsequently graft the osteoinductive agents. It is well known that in the post-synthesis functionalisation processes, the external surface is more accessible to silylation than the internal surface of channels. In this way, the reaction of the calcinated material with a low reactive functionalisation agent would mainly yield to the selective functionalisation of the external surface that would allow the grafting of the oste-oinductive signals. Later, the functionalisation of the internal surface of mesoporous channels could be achieved by using more reactive functionalisation agents that would be suitable to optimize the host matrix for drug delivery applications. There is another alternative to achieve a selective functionalisation, such as starting from the material still containing the surfactant into the mesoporous cavities, which would allow functionalizing the external mesoporous silica surface. Subsequently, after the surfactant removal, the inner part of the mesoporous channels could be also functionalized. [ 188 , 189 ]
Regarding the attachment of peptides, hormones or growth factors to bioceramic matrices, the possibility of being incor-porated into the mesoporous channels should not be over-ruled. This pathway implies a non-covalent weak interaction between the host matrix and the osteoinductive agent. Hence, the osteoactive factor could be released to the surroundings just at the time when needed, helping bone regeneration proc-esses. However, in this case it could be serendipity rather than design.
Furthermore, the possibility of using mesoporous matrices as starting materials for the manufacture of 3D scaffolds pro-vides an added value to the resulting bioceramics; their capa-bility to host drugs for locally treating bone pathologies, such as bone infection, osteoporosis, cancer, etc.
Different ordered silica mesoporous systems have been suc-cessfully employed as matrices to host several pharmaceutical molecules providing that their size suits to pore dimensions. However, the adsorbed drug should present active centers that could interact with the pore walls leading to drug retention. Therefore, pore size vs. drug size ratio has been revealed as an important factor to control the amount of drug to be adsorbed, but it is not as signifi cant in the drug delivery rates ( Figure 18 ).
It has been observed that a better control over the drug release kinetics of these delivery systems may be achieved by functionalizing the pore walls of the mesoporous matrices due to the attracting interaction between functional groups from modifying agents and drugs. Thus, the drug dosage can be modulated depending on the patient requirements, ranging from a range scale of hours (unmodifi ed matrices) up to a scale of days (amine-modifi ed matrices). [ 190 ]
Silica-based ordered mesoporous materials have attracted the attention of many research groups as drug delivery systems due to their unique porosity and textural properties. [ 116 , 174 ] How-ever, the presence of drugs into the inner part of the mesopores was not evidenced until 2010 by direct methods. [ 191 ] Previous research works have been based on the sum of indirect char-acterization techniques, such as N 2 adsorption [ 183 ] or thermo-gravimetry. [ 192 ] In a recent work, the direct evidence of drug adsorption into the inner part of the hollow channels of ordered mesoporous silicas has been reported.
Scanning Transmission Electron Microscopy (STEM) employing a spherical aberration (Cs) corrector [ 197 ] is based on scanning a sample by an electron probe, which is focused
down to 1 nm or less on the specimen. Then, the STEM images are formed with the collected scattered electrons in each probe position by the high-angle annular dark-fi eld (HAADF) detector at the bottom of the sample in synchronism with the scanning probe. [ 193 ] Thus, an atomic resolution analytical microscope equipped with a STEM Cs corrector enables at present to per-form outstanding microscopy and analysis with enough resolu-tion in practical use capable of atomic level analysis. [ 194 ] This type of analytical STEM is now being employed to the study of not only perfectly crystalline materials, but also grain bounda-ries and interfaces of crystals, because this technique allows determining the position of the atomic columns, [ 195 ] However, STEM with Cs corrector may also be used to characterize amor-phous silica materials, such as SBA 15 matrices, but taking the advantage of the highest level resolution. Thus, considering that it is possible now to illuminate an individual atom with the electron beam, we should be able to distinguish between the pore wall (silica amorphous material) and the pore space, where the drug molecules are supposed to be confi ned ( Figure 19 ).
In this case, SBA 15 ordered mesoporous materials were pro-duced according to the procedure described by Zhao et al .[ 196 ] Functionalisation was carried out by refl uxing mesoporous silica in dry toluene containing aminopropyl triethoxysilane. Thus, the surface of the matrix was covered with amine groups to improve the zoledronate adsorption capacity and retard the further release kinetics of the drug adsorbed. These effects can be achieved through a chemical interaction between amine groups at the surface and phosphonate groups of the guest mol-ecule. The procedure to confi ne the zoledronate molecules into the mesopores of SBA 15 and SBA-NH 2 was based on impreg-nation at room temperature of the matrices in a concentrated
Figure 19 . HAADF-STEM analysis of pore wall and drug-loaded pore channel of SBA-15 materials.
aqueous solution of zoledronate (40 mg/mL at pH 6 under magnetic stirring).
The small angle X-ray diffraction (XRD) pattern of SBA 15 material was well resolved with a prominent maximum at 0.9º, and two other weak maxima at 1.6º and 1.8º 2 θ . This pattern agrees with the previously reported for SBA 15 material and
Figure 20 . Nitrogen adsorption porosimetry study of SBA 15 materials aftwith aminopropyl groups and zoledronate loading.
S B A 1 5 - N H 2
S B A 1 5 - N H 2 - z o l
ed (ar
b. u
nits
)
SBA 15-NH2-zol
S B A 1 5
S B A 1 5 - z o l
0 0.2 0.4 0.6 0.8 1Relative Pressure (P/P0)
Volume
adso
rb
SampleSBET
(m2/g)Vp
(cm3/g)Vµp
(cm3/
SBA 15 903 1.21 0.04
SBA 15 Zol 239 0.56 0.00SBA 15-NH
2333 0.51 -
SBA 15-NH2Zol 161 0.34 -
SBA 15
SBA 15-zol
SBA 15-NH2
can be indexed according to a hexagonal lattice with a d(100)-spacing of 93.9Å, corresponding to a large unit cell parameter a 0 = 108Å (a0 = 2d10/
√3). [ 197 ] The mesostructure of the host
matrices survived the functionalisation and loading process, which was confi rmed by XRD. The prevalence of the ordered mesostructure was confi rmed by the presence of the same dif-
GmbH & Co. KGaA, We
er functionalization
g)Dp
(nm)twall
(nm)a0
(nm)
6 7.4 3.4 10.8
4 7.4 2.7 10.45.7 5.6 11.35.2 5.9 11.1
fraction maxima of the XRD patterns. Transmission electron microscopy (TEM)
has also been employed to confi rm the non-collapsing of the mesostructure after the loading process. Moreover, Fourier-trans-formed infrared (FTIR) spectroscopy was carried out to confi rm that zoledronate was loaded into the mesoporous matrices. Typical vibration bands at ca. 2900 and 2800 cm − 1 from C-H bonds confi rmed the presence of the alkyl chains from the drug ([1-Hydroxy-2-(1H-imidazol-1-yl)-ethylidene]bisphosphonic acid). The amount of zoledronate adsorbed by the mesoporous matrices was quantifi ed using X-ray Fluorescence (XRF). While SBA 15 materials loaded ca. 4% (wt), SBA 15-NH 2 matrices adsorbed ca. 10%, that is, the appro-priate functionalisation of the mesoporous walls leaded to duplicate, or even more, the amount of loaded drug molecules.
Nitrogen adsorption analysis is commonly employed to fi nd out if the adsorbed mole-cules are inside the porous channels. Typical N 2 adsorption type IV isotherms for these types of materials were obtained. Figure 20
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shows the variation of the textural parameters as a consequence of the amine functionalisation and zoledronate uptake.
The surface area of SBA 15 and SBA 15-NH 2 matrices was reduced as a consequence of the zoledronate loading, which indicates that the drug molecules are being retained on the walls surface through the host-guest interaction (molecular adsorption into mesoporous silicas is a surface phenomenon that is governed mainly by the chemical interactions between the hosts, SiOH in SBA 15 and NH 2 in SBA 15-NH 2 , and the functional groups of the guest molecules, phosphonates in the case of zoledronate). But even more importantly, the available pore volume drastically decreased after the loading process, both in SBA 15 (from 1.21 to 0.56 cm 3 /g) and SBA 15-NH 2 (from 0.51 to 0.34 cm 3 /g). This volume reduction is caused by the drug molecules partially fi lling the mesopores. This obser-vation was conventionally used to justify that the drug mole-cules were being confi ned inside the pores (Figure 18 ).
However, N 2 adsorption analysis is an indirect way of fi nding out if the drug molecules are confi ned inside the pores. The difference of the available pore volume before and after loading process denotes that the drug molecules are occupying this space, but there is not a direct evidence on whether the mol-ecules into the pores are zoledronate molecules or not.
Consequently, to determine the distribution of zoledronate molecules within the SBA 15 materials, a microscope equipped with a STEM Cs spherical aberration corrector and Electron Energy Loss Spectroscopy (EELS) seems to be a powerful tech-nique able to show the distribution of silicon, oxygen, nitrogen and carbon throughout a mesoporous silica network. Previous studies by other authors tried to establish the distribution of an organic functionalisation into the mesopores of SBA 15. [ 198 ] However, the microscopic and analytical techniques used in that case did not present enough resolution to establish such strong conclusions. Traditionally, the employment of characterization
Figure 21 . EELS intensity profi les of Si, O, C and N elements demonstradrug inside the mesopores.
[001]
x 10
^3
nm
230
240
250
260
270
280
290
300
310
320
330
340
350
360
x 10
^3
0 10 20 30 40 50nmnm0 10 20 30 40 50
nmMaximum Si-OMinimum C-N
Maximum C-NMinimum Si-O
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10 nm
techniques such as electron microscopy and their associated spectroscopic techniques to evaluate ordered mesoporous materials presented, basically, two drawbacks: (1) there was not enough resolution to illustrate what was going on into the mes-oporous channels, and (2) the time needed to do the micros-copy studies usually provoked radiation damage and, in some occasions, the decomposition of the sample. On the other hand, the improved mechanical and electrical stabilities of the very recent generation of STEM Cs corrected microscopes enable us to perform outstanding microscopy and analysis at the highest-level resolution. Thus, it is even possible to illuminate an indi-vidual atom with the electron beam to identify an unknown substance. High resolution Cs corrected STEM microscopy has been employed to determine the chemical composition of both matrix walls and pore area, confi rming that drug molecules are confi ned in the inner part of the mesopores of SBA 15 host matrices.
Despite the lack of order from both the silica walls and drug molecules, the employed resolution permits analyzing the chemical elements from the walls of the silica material and from the drug molecules located in the internal area of the mes-opores. This is an important observation because it allows the qualitative determination of the presence of organic molecules into the inner area of mesoporous systems that are applicable as drug delivery systems (Figure 19 ).
In this case, the sample is a mesoporous system with a drug inside the pores. That means there is pore ordering while the material is amorphous. In the images parallel to the c axis, the 8 nm pore diameter is observed, since the resolution of the tech-nique is below 1 nm. With EELS analysis along the pore wall, only Si and O can be detected (see Figure 21 ). However, turning 90 degrees and carrying out the analysis inside the pore, only light elements such as C and N are observed (see Figure 21 ). Moreover, when the analysis is performed in the pore wall, a
bH & Co. KGaA, Weinh
te the presence of
60 70 80 9060 70 9080
maximum of Si and O and a minimum in the intensity of C and N is obtained. Conversely, a maximum in C and N content coincides with a minimum of Si and O. Therefore, there is direct experimental evidence of the presence of drugs (zoledronate in this case) within the mesoporous material.
Since the fi rst publication on the employ-ment of ordered mesoporous silica, MCM 41, as ibuprofen release system, [ 174 ] many other mesoporous structures have been proved to act as controlled delivery systems of many different drugs. [ 199–203 ] When aiming to use these materials for drug delivery technolo-gies, the fi rst step should be the selection of the adequate matrix with a mesopore diam-eter high enough to host each type of drug, since that diameter would determine the size of the molecule that can be hosted. If the molecule to be hosted is smaller than the mesopore diameter, it would be confi ned in the inner part of the mesopores. On the other hand, if the molecule is larger than the mes-opore cavity, the adsorption would take place only at the external surface of the matrix. [ 204 ]
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Figure 22 . Parameters involved in the adsorption capacity of mesoporous silica matrices.
Thus, the pore diameter acts as a size-selective adsorption parameter. [ 205 ]
The surface area is another key parameter in drug adsorption to take into consideration, especially when the drug molecule is far smaller than the pore diameter. In this situation, only some of the molecules can interact with the pore walls, while the rest would not be retained, and the higher the surface area, the higher the amount of drug loaded. [ 183 ] When loading molecules with a large volume, such as proteins, the pore volume is also an important factor. Sometimes it is necessary to modify the synthetic conditions of the ordered mesoporous silica to obtain large enough mesopore volumes to host proteins. [ 177 ] However, the most determinant parameter in the adsorption capacity of a mesoporous matrix, and also in the release kinetics of the adsorbed molecules, is the functionalisation of the material sur-face ( Figure 22 ).
In this sense, grafting alkoxysilanes with different organic groups allows adjusting the chemical properties of the sur-faces. [ 175 ] From the variety of possibilities, the organic modi-fi cation should be selected depending on the biomolecule to be adsorbed, that is, depending on the functional groups of the guest molecule. [ 206 ] An interesting approach that has been recently presented is the one-step synthesis of zwitterionic SBA 15 type mesoporous material containing both amine and car-boxylic acid groups exhibiting ultralow-fouling capability. [ 133 ] This strategy may be implemented in clinical applications since the surface of these materials would hinder the adherence of any kind of bacteria (see section 2.5.1).
4.2.2. Controlled Delivery Concept for Bone Implant Technologies
Ordered mesoporous structures have been revealed as great matrices for drug delivery technologies, as it has been previ-ously shown. Additionally, sol-gel glasses were observed to bond to living bone, so the combination of the sol-gel synthesis of glasses with the supramolecular chemistry of surfactants has
promoted a new generation of highly ordered mesoporous materials for biomedical appli-cations, both for bone tissue engineering and drug delivery technology. These materials, called ordered mesoporous bioactive glasses or templated glasses, exhibit enhanced bio-active behavior compared to conventional glasses. [ 194–207 ]
Although the preparation of highly ordered mesoporous structures with multiple compo-nents might be complicated, the employment of non-ionic surfactants has opened new pos-sibilities. These structure directing agents combined with the well-known method of evaporation-induced self-assembly process (EISA) permitted the successful prepara-tion of a new generation of compounds with diverse chemical composition. [ 208 , 209 ] In fact, different amounts of CaO could be reached, which leaded to different mesostructures, [ 210 ] and which also infl uenced the bioactive behavior of the mesoporous material.
Along the last few years of investigation
in our research group we have observed that silica mesoporous materials play a double function: bioactivity and releasing of biologically active molecules. The examples discussed so far open new applications for silica based ordered mesoporous materials. Also, it would be possible to combine both applica-tions, arriving at the design of new bioceramics with a large added value. However, all these fi ndings should be taken into consideration carefully, because if these materials are used in tissue engineering, the main role has to be played by living cells. Mesopore dimensions of ordered mesoporous materials, between 2 and 50 nm, are too small for cells to be uptaken, needing dimensions in the order of μ m. Bone porosity, ranging between 1 and 3500 μ m [ 211 ] is necessary for several physi-ological functions performed by bone. To produce scaffolds for tissue engineering, the required pieces must exhibit similar porosity to the natural bone, hence the need to apply confor-mation methods that preserve the mesoporosity of these silica materials while providing interconnected macroporosity, within the 20 to 400 μ m range.
Those materials with mesoporosity between 2 and 50 nm are of interest for applications where drugs or biologically active molecules are loaded, and later released. Macroporous mate-rials, where pore sizes are in the order of microns, are adequate as scaffolds for tissue engineering. The challenge then is to obtain scaffolds combining macroporosity for bone oxygenation and vascularisation, and mesopores able to be loaded with bio-logically active molecules. The goal is to design materials that can help the human body to improve its regeneration features, not only recovering the structure of the damaged tissue, but also its function. The present target in biomaterials is to pro-duce three dimensional scaffolds with interconnected porosity, so that cells can proliferate and form tissue in a similar way to the process in human tissues.
The design of scaffolds able to guide cell growth is an impor-tant challenge in tissue engineering. The aim is to fabricate pieces that support and structure the newly formed tissue,
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and said pieces have to be made of the most suitable materials for these tasks. Living cells must be cultured onto this sup-port and subsequently give rise to the growth of new tissue. In order to work in potential hard tissue replacement solutions, it is required to know and bear in mind the bone regeneration process. Wolf’s law dictates that the bone remodels itself as a function of those forces acting on it, hence preserving its shape and density. Mechanical loads of stress, compression, fl exion and torsion in bones and the interstitial fl uid contained in them generate stresses and deformations at the microscopic level, which in turn stimulate cells.
The use of fi nite element calculation methods contributes to the study of interactions between materials, mechanical stim-ulations and biological responses. This theoretical approach allows simulating the conditions generated by bioreactors in the scaffolds and the effect and deformation at each point of the scaffold can be quantifi ed. Thus, the relationship between mechanical loads and cell differentiation can be quantifi ed. [ 212 ]
The fabrication of scaffolds for tissue engineering requires choosing a conformation method that yields pieces with inter-connected porosity. The main purpose now is to obtain porous ceramics that act as scaffolds for cells and inducting molecules, able to drive self regeneration of tissues. At present, the aim is to fi nd bioceramics which induce the regeneration of hard tissues stimulating the response of the cells involved. The requirements for these ceramics are to act as a scaffold and also to be porous so that the cells can do their job. Also, in some cases, porosity allows loading biologically active molecules onto such ceramics, if necessary, such as anticancer drugs or other drugs loaded into the implantable materials aimed at regener-ating osteoporotic bone. Therefore, drugs could be loaded into the mesoporous matrix that would act as a controlled delivery system. Moreover, the scaffold would allow, with an appropriate functionalization, grafting proteins, peptides or growth factors to induce cells to regenerate bone ( Figure 23 ).
As it has been shown so far, mesoporous materials allow designing controlled release systems, with the purpose of loading drugs to be subsequently released in a localized and controlled way. Among other applications, these systems could
Figure 23 . Mesoporous matrices may be used for drug loading and oste-oinductive factors grafting for bone regeneration applications.
be useful in bone infection pathologies. However, they would not be useful if the pores are loaded with peptides or growth factors; the role of these types of molecules is to act as attracting signals for bone formation cells, and these cells could never fi t inside the mesopores. Therefore, although peptides and growth factors do fi t inside the pores, they would not be playing any useful role in the bone regeneration process.
However, it is possible to apply all acquired knowledge with these systems to the fi xation of the mentioned peptides or growth factors onto the external surface, in order to accel-erate the bone formation rate. Even more, they can be used as starting material in scaffold fabrication, with porosity values in the hundred microns range, similar to the porosity of natural bone. These scaffolds would exhibit the added value of hosting drugs on the mesopores which could act against specifi c pathol-ogies and be tailor-made for a specifi c patient. And, of course, there is an added contribution of peptides and growth factors loaded on their surface.
An important feature of these templated glasses, especially taking into account their biomedical applications, is that their biodegradation products are biocompatible. In fact, in vitro compatibility tests have demonstrated the good behaviour of osteoblasts, fi broblasts and lymphocytes in the presence of these mesoporous templated bioglasses. [ 140 ] Therefore, these bioactive glasses can be employed in association with osteo-genic agents to produce three-dimensional scaffolds for bone tissue engineering. The bioactive behaviour of these templated bioglasses would promote osteoconduction, osteoproduction and osteoinduction processes when implanted in living tissue. Additionally, the mesoporosity provides an added value to the possible implant, because the graft not only fi lls and repairs the defect but also acts as a drug delivery system, which would locally supply osteoregenerative agents.
The design of new materials applicable in bone implant tech-nologies has received great attention during the last years. [ 213–218 ] Several bioceramics have been reported as good candidates in bone tissue regeneration acting as local controlled delivery sys-tems of drugs and other biologically active molecules that pro-mote new bone formation. [ 115 ] Therefore, several biologically active molecules such as antibiotics and/or anti-infl amatories have been confi ned into ceramic carriers to fi ght against adverse phenomena, infection and/or infl ammation, which normally take place after the implantation process. [ 116 , 219–223 ] Also the confi nement of drugs employed for osteoporosis treatments, such as bisphosphonates, has been described. [ 183 ] Moreover, the confi nement and release of biologically active molecules, such as amino acids, peptides, proteins and growth factors into ceramic matrices have been reported as well. [ 168 , 177 ] Molecules are commonly loaded into the ceramic matrices by adsorption mechanisms. This represents a great limitation because these drug delivery systems frequently exhibit a burst release effect, where most of the drug loaded is rapidly released to the delivery medium during the fi rst hours of assay. Furthermore, when dealing with bone implants a time-delayed delivery with an ini-tial zero-release or lag time period would be desirable to allow the surgeon to carry out the surgical procedure.
Several strategies have been developed trying to minimize the fast delivery of molecules to the medium, such as organi-cally modifying the ceramic carrier with functions that undergo
attracting interactions with the functional groups of the tar-
geted molecule, then minimizing the initial burst effect. [ 116 , 224 ] Thus, for instance, ca . 55% alendronate loaded into unmodi-fi ed SBA-15 mesoporous material is released to the medium during the fi rst 24 hours of assay. This burst effect drops to ca . 25% of alendronate delivered when the silica walls of SBA-15 are organically modifi ed using amino groups. [ 116 ] However, this problem is still a matter of discussion and requires the develop-ment of new advanced materials that allow solving the clinical situation.
The burst effect drawback could be overridden by using one-step procedures, in which the synthesis of the ceramic material would take place at the same time that the drug is physically entrapped. Therefore, the drug release would be controlled by the slow degradation of the matrix and/or slow diffusion of the drug through the matrix. The one-step incorporation of biologi-cally active molecules into ceramic carriers requires synthesis procedures that employ mild conditions, such as room temper-ature and aqueous media. This can be easily achieved by means of the sol-gel chemistry. [ 225 , 226 ] One of the main advantages of sol-gel technology is that allows not only the entrapment of sev-eral kind of molecules within the resulting xerogel, but also pre-serves their functionality when dealing with biologically active molecules, in the so-called bioencapsulation process. [ 227 , 228 ] This is especially important when the confi nement of drugs with a limited shelf-life is aimed. Thus, molecules bioencapsu-lated into sol-gel matrices are protected from biological degra-dation and are often considerably stabilized against chemical and thermal activation. This would result in improved storage stability and therefore drug release would last over a clinically relevant period of time in an active form.
The release of entrapped molecules from sol-gel matrices can be attained by the design of hybrids containing amino groups,
Figure 24 . Organic-inorganic hybrid materials may be tuned to avoid hydroreach time-delayed drug release.
4 0
Biologically active molecules (H2O)
Sol-GelPrecursors
Bio-dopedhybrid material
Bioencapsulation
Si(OEt)3
DrugsProteinsGrowth faCells
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Lag-TimeGlassy to Rubber Transition
due to the high solubility of these materials in aqueous media. However, the extremely fast dissolution rate of amino-polysi-loxane matrices would lead to a rapid release of the entrapped molecules. This is an effective strategy for the storage and fast release of actives in applications such as fabric care [ 229 ] , but limits their use in the clinical practice in which a time-delayed release followed by a sustained delivery of molecules over a pro-longed period of time at the targeted site is needed.
Previous studies have demonstrated that the problem of aqueous instability of amino-polysiloxane matrices can be suc-cessfully overcome by using an aminosilane precursor together with another alkoxysilane containing a more hydrophobic group. [ 230 ] Organic-inorganic hybrid materials were prepared from different molar ratios of [N-(2-aminoethyl)-3-aminopropyl] trimethoxysilane (DAMO) and 3-methacryloxypropyl)trimeth-oxysilane (MPS) and the results clearly established that the MPS/DAMO molar ratio governed the degradation rate of the resulting hybrid matrices ( Figure 24 ). Moreover, with the addi-tion of small amounts of calcium salts the materials exhibited in vitro bioactive behavior
The development of new controlled delivery systems by entrapping biologically active molecules into amino-polysi-loxane matrices starting from MPS and DAMO and using one-step sol-gel process at room temperature is a promising approach. Vancomycin, the most effective antibiotic against Gram-positive bacteria used clinically to treat osteomyelitis infections, has been used as model antibiotic. In addition, bac-terial inhibition assays previously reported revealed that the bactericidal effi cacy of released vancomycin from silica xerogels is retained. The MPS/DAMO molar ratio was selected to get a slow dissolution of the matrix in simulated body fl uid, in order to obtain a lag time in the drug release and a subsequent sus-tained delivery of the entrapped drug. [ 231 ] The evolution of the
mbH & Co. KGaA, Wei
lytic instability and
ctors
6 0 7 0
release rates and drug delivery profi les were studied as a function of the amount of vanco-mycin incorporated into the polysiloxane net-work. Moreover, with the aim of testing the behaviour of these systems in bone infection situations, in vitro drug delivery assays were performed at two different values, pH 7.4, i.e . physiological pH in healthy bone tissues, and pH 6.5, since a decrease of pH occurs in the surroundings of infected bone tissue [ 232 , 233 ] ( Figure 25 ).
Indeed, vancomycin has been success-fully entrapped into amino-polysiloxane matrices via one-step room temperature sol-gel process, resulting in bio-doped mono-lithic hybrid materials. [ 231 ] The vancomycin-containing matrices have been characterized by means of FTIR and 29 Si MAS NMR solid state spectroscopies, evidencing the effective encapsulation of the drug. In vitro swelling behaviour and delivery tests carried out under physiological conditions (37 ° C, pH 7.4) showed not only the absence of burst effect but also, and remarkably, a zero-release period, or lag time, where no vancomycin was released to the medium during the fi rst
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Figure 25 . Vancomycin may be encapsulated in hybrid bioceramics for the treatment of infections in bone tissue replacement.
hours of assay. Subsequently, there was a sustained release of vancomycin over a prolonged period of days. This lag time is an essential requirement for implantable bioceramics that would allow the surgeon to perform the surgical procedure with zero drug release. The delivery behaviour has been also tested mimicking bone infection conditions (37 ° C, pH 6.5), and the lag time and subsequent sustained release are comparable to those obtained at physiological conditions, in spite of the pH decrease. This research work opens new possibilities for the design of novel time-delayed controlled delivery systems for bioencapsulates, useful in bone implant technologies.
4.2.3. In vitro Stability of Silica Mesoporous Materials
Silica based ordered mesoporous materials have been widely proposed for different biomedical applications [ 15 ] . Their out-standing structural and textural properties together with their capability to be functionalized with different organic groups make them attractive candidates for controlled drug delivery, [ 116 , 234 ] bone tissue regeneration, [ 121 , 168 ] gene transfec-tion, [ 235 , 236 ] cell tracking [ 91 , 237 ] and immobilization of proteins or enzymes. [ 238 , 239 ] Silica mesoporous materials are of particular interest for bone tissue regeneration because they exhibit the added value of being able to act as host matrices for a wide range of biologically active guest molecules to locally promote new bone formation or bone healing.
Within the world of biomaterials technology, an adequate biocompatibility for the designed applications is a mandatory requirement. [ 240 ] Amorphous silica has been generally consid-ered as non-toxic, and it has been reported as biocompatible and degradable in living tissue, [ 241 ] which is required for bone,
cartilage and connective tissue formation. [ 242 , 243 ] Up to date, biocompatibility and stability tests of different silica materials have been reported. [ 244 , 245 ] However, much research effort has to be committed to evaluate the inherent degradability, cytotox-icity and biocompatibility of mesoporous materials. Previous research works demonstrated that the concentration and par-ticle size of mesoporous silicas play a key role in their in vitro toxicity, showing low toxicity at low concentrations. [ 235 , 246 ]
Hudson et al. have performed in vivo studies using different mesoporous materials, concretely, MCM-41, MCM-48 and MCF. [ 247 ] Histological studies performed after subcutaneous injection of such mesoporous materials at the sciatic nerve in rats evidenced a good biocompatibility at all time points. However, intra-peritoneal and intra-venous injections in mice resulted in death or euthanasia. The role of silica in the etiology of this pathogenesis is still unclear. Although local tissue reac-tion of mesoporous silicas was benign, they caused somewhat systemic toxicity. Authors suggest that toxicity could be mitigated by modifi cation of the materials, by means of surface adsorption of proteins, functional groups or polymers and so forth.
The above mentioned results clearly evidence that toxicity of mesoporous silicas highly depends on the dosage and the administration route to the living body. This is tightly related to the amount of silica that the mesoporous matrix releases to the physiological fl uids, where the normal concentration of SiO 2 (in human plasma) is ca. 1 mg/L. [ 248 ] Considering the viability of mesoporous matrices as biomaterials for bone tissue regen-eration, the fi rst step should be the in vitro tests in aqueous medium mimicking physiological conditions. Such studies would allow knowing the materials reactivity, including the amount of SiO 2 leached to the aqueous medium.
Figure 26 . Degradation rates study of unmodifi ed and functionalized SBA 15 mesoporous silica by TEM. Organic functionalization of the material is shown to be a suitable strategy to decrease SiO 2 lixiviated.
It has been demonstrated that the incorporation of several heteroatoms into the silica network led to decreases in the silica lixiviation rate by increasing the stability of materials. [ 249 , 250 ] Another proposed strategy consisted in incorporating osteo-genic agents, such as parathyroid hormone related-protein (PTHrP) to promote bone tissue regeneration and accelerate the bone integration process. [ 123 ]
The organic functionalization of mesoporous materials is a straightforward approach to decrease the silica lixiviation. It is possible to evaluate the degradation in different aqueous media of SBA-15 mesoporous material. [ 251 ] Among those media we can fi nd: saline solution (0.9% NaCl saline solution buffered at pH 7.4), a simulated body fl uid (SBF) (an acellular aqueous solu-tion with inorganic ion composition similar to that of human plasma) [ 252 ] and completed Dulbecco’s modifi ed Eagle medium (DMEM). All of them are aqueous solutions widely employed as release media in drug delivery, bioactivity and cell culture assays. Izquierdo-Barba et al. have monitored the silica degra-dation kinetics from SBA-15 matrix as a function of the degree of ionic complexity of the media and the presence or absence of serum proteins.
The functionalization of ordered mesoporous silica is a well known procedure employed to chemically modify the surface of mesoporous matrices. In this approach, SBA-15 has been modifi ed with two different alkyl groups, methyl (C1) and octyl (C8) chains, and aminopropyl groups in an attempt to make the surface more hydrophobic and hence more stable in aqueous media. After verifying the proper functionalization of SBA-15 with the different organic moieties, the stability of the resulting samples has been investigated in the same way as unmodi-fi ed SBA-15. The differences observed in the silica degradation rates have been determined by colorimetrically measuring the
species of soluble silica leached to the tested media. Moreover, the textural and structural features of the different mesoporous materials after the in vitro tests have been also carried out. In vitro stability tests of SBA-15 in solutions mimicking physiolog-ical fl uids using three aqueous media (saline solution, SBF and DMEM) at 37 ºC reveal that SBA-15 is partially soluble in all tested media. The complexity of the aqueous medium regarding the ionic composition and the presence of proteins does not signifi cantly affect the silica dissolution from SBA-15. The deg-radation of SBA-15 produces some loss of its textural and struc-tural properties. However, the degradation rate is relatively low, since TEM studies indicate that after 60 days of assay there is a partial loss of the ordered mesoporous arrangement in SBA-15, but there are still some domains with 2D-hexagonal structure ( Figure 26 ). This is in agreement with the results derived from XRD measurements.
The reduction in the silica lixiviation rate has been success-fully achieved by organically modifying SBA-15 using different organic groups: alkyl chains (methyl and octyl) and aminopropyl groups. The in vitro degradation tests indicate that functionali-zation of SBA-15 decreases the SiO 2 leaching to the three media tested comparing with unmodifi ed SBA-15. These results open promising expectations for the use of these materials in many biomedical applications where low silica release to the physi-ological medium is essential.
The decrease in the degradation rate of functionalized sam-ples compared to SBA-15 is accompanied by a preservation of the mesostructure arrangement, as confi rmed by TEM and XRD measurements after 60 days of assay. This is undoubt-edly an added value when long-term applications, such as bone tissue regeneration and other biotechnological applications, are targeted, which require mesoporous materials with stable
mesostructural order. This study, which has been performed for biotechnological purposes, could be suitable for other applica-tions that require and evaluation of the stability of mesoporous matrices.
5. Bioceramics For Cancer Therapy
As previously addressed, bioceramic materials can be chemi-cally designed to meet target medical needs. Moreover, exten-sive research on stimuli-responsive systems has triggered a renovated hope for more controlled and selective therapies. In bioceramics for bone substitution, a broad range of mate-rial properties are liable to be tailored to replace damaged bone tissue. In particular cases where a cancerous disease arises, magnetic bioceramics can be synthesized to provide a local treatment of tumor cells. [ 241–255 ]
In the last decades, implantable magnetic bioceramics have been developed to treat localized tumours by hyperthermia. [ 244 ] This technique is especially useful when dealing with deep-seated malignancies in the body. Hyperthermic treatment aims to heat organs or tissues up to temperatures ranging from 43 ºC to 47 ºC. Within this interval, an almost irreversible selective destruction of cancer cells is reached. [ 257–260 ]
Up to now, hyperthermia has shown a high effi cacy when combined with conventional treatments such as chemotherapy or radiotherapy. [ 261–264 ] The main drawback of hyperthermic therapy is the diffi culty to reach target temperatures and fur-ther control the localized tumor heating. Overheating of sur-rounding healthy tissues is particularly diffi cult to avoid under hyperthermic temperatures. [ 265 ] For this reason, implantable magnetic mediators may be appropriate for hardly accessible tumours like bone-related malignancies.
In other respects, bioceramic materials constitute robust tuneable matrices with the ability to locally deliver bioactive substances, such as chemotherapy drugs in the case of cancer diseases. [ 266 ] Pore volume and size, accessible surface area and chemical affi nity are some of the parameters to be adjusted to suitably host each possible therapeutic molecule.
5.1. Selective Mediators for Cancer Targeting
Several processing techniques of bioceramic materials allow the synthesis of sub-micrometric matrices to act as multifunctional devices for theranostic purposes. [ 267 , 268 ] Many of the functions of living organisms occur at the nanoscale. The human body uses natural nanoscale materials, such as proteins and other molecules, to control its systems and processes. Furthermore, nanostructures seem for example to improve drug delivery and especially avoid unwanted side effects. Some new drugs and diagnostic agents based on nanotechnology are now approved or in clinical trials. [ 269–271 ] On the one hand, nanoparticles are being used to release highly toxic drugs directly into tumours, while reducing the amount of chemotherapy which damages healthy tissue. Other devices fi nd applications in imaging techniques such as computed tomography or magnetic resonance imaging.
In the case of drug delivery, it is estimated that only a very small portion of the antibodies and contrast agents for imaging
techniques that are administered intravenously reach their des-tination inside the body. [ 272 , 273 ] By carrier injection into an organ or tissue, it may not only increase the ratio but also achieve the desired effect only in a cell type without affecting other tissue cells. Concerning cancerous diseases, the cytotoxic effect of chemotherapy may be directed only to tumor cells through the use of various bioceramic carriers of antibody-based molecules. Tsai and co-workers have demonstrated selective targeting of mesoporous silica nanoparticles to breast cancer cells by spe-cifi c monoclonal antibodies. [ 274 ]
For this reason, it is not surprising that the development of nano and microparticles for biomedical applications has been declared one of the most promising fi elds of research in the last decade. Advances in the preparation of nanosystems in the medical fi eld make it possible to face new challenges in the design of smart materials capable of meeting the clinical demands. [ 275–279 ] One of the major concerns in medicine is the administration of drugs and diagnostic agents to the patient in a more physiologically acceptable way. In many cases, it is necessary to apply large doses of the drug as a result of non-specifi c release and the limited absorption by the target tissue. This problem is compounded in oncological diseases, in which the risk-benefi t ratio associated with chemotherapy is often intolerable. [ 280–282 ]
It is generally accepted that the absorption of the drug by the organism is favoured by a smaller size of the drug and the coating material that is used in its package. Nanostructures, as in the case of polymer nanofi bers and bioceramic-polymer hybrid materials, have advantages in controlling the release of therapeutic substances when compared with the carriers cur-rently used. [ 283 , 284 ] The dissolution rate of a particular drug can be modulated by the surface area of the polymeric structures or the composition of the bioceramic component. [ 285 − 287 ]
5.2. Multifunctional Nanomaterials
Moreover, the release of drugs is not the only nano-biotech-nological application on target. Molecular recognition, encap-sulation, production of materials and biocompatible coatings, molecular analysis of DNA, bio-inorganic hybrids and diag-nostic techniques [ 288–290 ] are some of the aspects that can be approached from a new perspective with nanomaterials. In the fi eld of tissue engineering it is expected that, in the future, some elements of nanoscale construction can be used for the repair of cartilage, bone or skin. [ 291 ] As for biological testing, the design of nanometric indicators may increase the speed and sensitivity in measuring the presence or activity of certain molecules. [ 292 ] The association of quantum dots of different sizes with each type of molecule will signifi cantly improve cur-rent methods based on excitation with different wavelengths of organic dyes used in staining. In this respect, bioceramic matrices can act as stable protective hosts encapsulating spe-cifi c markers. Hu et al. have reported a simple method to effi ciently encapsulate quantum dots into mesoporous silica providing excellent lumininescence, stability and size mono-dispersity. [ 293 ] This kind of mesoporous silica/quantum dot conjugates has also been coated with PEGylated liposomes to improve biocompatibility. [ 294 ]
Materials scientists are developing an arsenal of structures for
a wide range of biomedical applications: polymer-based nano-carriers, [ 295 ] liposomes, [ 296 ] polymersomes, [ 297 ] dendrimers [ 298 ] and nanoparticles. [ 299 , 300 ] Some of them have already demon-strated their advantages for research or clinical practice. Among others, some families of polymers of polylactic-polyglycolic acid or polyanhydrides, [ 301 , 302 ] are able to maintain a sustained release of therapeutic substances at safe levels with a duration of two weeks to several months. Also, some dendrimers and liposomes are routinely used in gene transfection assays in the laboratory. [ 303 ]
On the use of nanoparticles, they have become widespread in diagnostic techniques over the past 40 years. [ 304 , 305 ] The incorporation of these agents leads to several advantages such as the possibility of greater accessibility to tissues as a result of a high ratio of surface area to volume. In the last decade, there have been many efforts in research and development of magnetic nanoparticles. [ 306 ] These materials offer several pos-sibilities among which are: improving the quality of magnetic resonance imaging, the treatment of cancer cells by hyper-thermia therapy, the targeted drug delivery to an affected area, the manipulation of cell membranes and magnetic separation of cells or other biological entities. One of the biggest obstacles for the therapeutic application of nanoparticles is the diffi culty of driving them to a specifi c area of the body. With magnetic nanoparticles, the potential to attract the particles by magnetic fi eld gradients to a specifi c region within the organism repre-sents a promising approach. In principle, they could be kept in the target place until the end of therapy and subsequently be removed. The properties sought in these materials are the ability to be injected, a high degree of accumulation in the organ or tissue, and especially their biocompatibility.
5.3. Bioceramic Implants for Bone Cancer Treatment
Currently, the established chemotherapy regimens for bone tumors suffer from toxic effects and low levels of tolerance and effi cacy. Because of this, treatments often involve complicated surgical operations in which they come to remove large por-tions of bone tissue. The presence of residual malignant cells in the affected area is an added problem, which can lead to recur-rence of the tumor, signifi cantly worsening disease prognosis.
In this scenario, looking for tissue engineering constructs capable of combining the replacement of damaged parts with local cancer treatment constitutes a medical demand. [ 307 ] Implantable magnetic thermoseeds can be designed to prevent metastases after tumor resection. These materials can be part of magnetic scaffolds that allow adhesion, proliferation and differentiation of healthy cells on their surface and generate increases in local temperature up to values of hyperthermia.
Remarkably, tailor-made magnetic bioceramics combine the potential to replace the damaged bone, with the ability to keep possible remaining metastases under control. [ 308 , 309 ] The asso-ciation of bioactive sol-gel glasses and magnetic glass-ceramics has been suggested as a suitable biocompatible system for the replacement of cancerous bone. [ 310 , 311 ] On the one hand, porous glasses produced by sol-gel techniques display an enhanced capability to bond to the living bone, as discussed above. This
kind of materials lacks suitable mechanical properties for hard tissue replacement, and the combination with glass-ceramics or with a fl exible polymer, such as polydimethylsiloxane or polytetramethyloxide is desirable to obtain bone-like structure and mechanical properties. [ 312 ] On the other hand, as a result of magnetic hysteresis losses, magnetic glass-ceramics can be designed to absorb energy under external alternating magnetic fi elds of certain frequencies, and release that energy in the form of heat. Moreover, this kind of materials shows a biocompatible behaviour and suitable mechanical properties.
The presence of a magnetic component provides the ability to locally generate thermal energy after the composite has been implanted, and every time the biomaterial is subjected to alter-nating magnetic fi elds. It has been shown that the effect of heat in the environment of solid tumours leads to a global involve-ment of cell proliferation accompanied by an inhibition of reg-ulatory signals. [ 313 ] The selectivity of the damage, with limited impact on healthy tissue, makes the magnetic hyperthermia treatment an adjuvant of great interest in anticancer therapies. In fact, the results obtained so far with the addition of hyper-thermia in the protocols against different tumor types, confi rm increases in effi ciency without a signifi cant contribution to the levels of risk. [ 314 , 315 ]
5.3.1. Magnetic Glass-Glass Ceramic Composites
An adaptation of material properties to the needs of such clin-ical situations to occur after surgical resection of bone tumours would encourage its use as implantable thermoseeds. The role of the biomaterial in this scenario would provide for the replace-ment or regeneration of bone tissue removed, and the thermal attack against the proliferation of tumor remains. Several dif-ferent approaches have been proposed so far. Luderer et al . evalu-ated the response of a murine mammary breast carcinoma with subcutaneously transplanted lithium ferrite in a matrix of hem-atite and a SiO 2 -P 2 O 5 glassy phase. This thermoseed was able to provoke a local hyperthermic effect causing approximately 50% tumor regrowth delay and a 12% permanent control. [ 256 ]
Finely dispersed ferrimagnetic, ferromagnetic and super-paramagnetic particles have been incorporated into bioactive matrices as hyperthermic treatment mediators. The bioactivity in this kind of biomaterial is attributed to the formation of bone minerals such as apatite or wollastonite in a physiological envi-ronment. Magnetite and α -Fe fi ne particles have been employed in combination with CaO-SiO 2 -based glassy matrices. [ 309 , 316 , 317 ] The bioactive behaviour of these composite materials was enhanced by modifi cation with Na 2 O or B 2 O 3 added in com-bination with P 2 O 5 . A similar composition was assayed by Kawashita and coworkers, who demonstrated the potential to adjust the magnetic properties and hence the maximum heat generation by controlling the magnetite crystallization process. [ 318 ] Higher values of heat specifi c absorption rates were found by the same authors when assessing ferrimagnetic zinc-iron oxide composites. [ 319 ] The magnetic properties of glass-ceramics with zinc-iron ferrite crystallites can be tuned from partially paramagnetic to fully ferrimagnetic with superior mag-netization values by changing the material composition. [ 320 ]
The infl uence of the amount of crystallized particles on the magnetic properties of glass-ceramic materials has
been assessed in magnetite-based bioactive composites. [ 253 ] Depending on the iron content, glass-ceramic microstructural features vary, consequently affecting the magnetic hysteresis characteristics. Saturation magnetization will then be prevalent when generating heat at intensive magnetic fi elds. In contrast, coercive force values will determine the heating capability under low magnetic fi elds. By controlling the chemical composition of the glass-ceramic, heat generation may fi t with hyperthermic performance requirements.
5.3.2. Magnetite-Wollastonite Based Thermoseeds
Given this background, our research efforts in this topic have focused on analyzing the behaviour of bioactive magnetic ther-moseeds depending on their composition and microstructure. Magnetite-wollastonite based glass-ceramic materials have been evaluated, with the addition of sol-gel glass in the system SiO 2 -P 2 O 5 -CaO at different proportions. Attending to the structural, microstructural and physico-chemical properties of the mix-tures, the effects of the composition on the bioactive behaviour of the implants and their ability to produce heat under alter-nating magnetic fi elds have been addressed. [ 321 ] Bioactivity is associated with the textural properties of the biomaterial, so that high values of specifi c surface area and porosity might increase the reactivity of the composites. The high porosity levels of glasses obtained by the sol-gel method have been linked to a higher carbonate hydroxyapatite growth rate in the bioactive process. [ 322 ]
In this sense, the regenerative capacity of bone tissue can be modulated through the sol-gel derived porous glass content, which is also indirectly involved in the potential for producing hyperthermia. On the one hand, this component is respon-sible for porosity and reactivity in the surface of the materials, leading to a higher bioactivity. On the other hand, the magnetic phase provided by the crystallization of iron oxide glass precur-sors rules the magnetic properties in the composites. However, the hyperthermic performance is affected by both components, since a higher amount of the magnetic phase results in higher magnetization values whereas a higher presence of bioactive glass means a higher coercivity and heating rates. [ 321 ] These fea-tures allow a rational design of magnetic implants to meet dif-ferent clinical requirements.
In addition, cell adhesion and proliferation on these mag-netic thermoseeds have been monitored. Previous studies dem-onstrated the cytocompatible properties of sol-gel glasses. [ 165 , 323 ] Remarkably, a higher content of magnetic phase in the biocer-amic implants gives rise to an improved biocompatible behav-iour. Cell culture tests have shown that the presence of sol-gel glass in the material is responsible for changes in the mito-chondrial activity as well as transitory alterations in intracel-lular reactive oxygen species content. Nevertheless, there were no signifi cant disturbances in the cell cycle, indicating that the interaction between cells and the thermoseeds did not induce signifi cant apoptosis. [ 324 ]
Furthermore, bacterial infection on this kind of biocer-amics has been assessed. Microorganisms have a strong ten-dency to adhere to the surface of biomaterials, and are then able to modify their structure and metabolism towards the formation of a biofi lm. [ 325 ] In this organized microbial system,
polysaccharide-based extracellular matrix is generated, and the bacteria become resistant to the attack of antibiotics. Bacte-rial infection is one of the most common reasons for failure of a biomaterial. The fi rst stage of this process is the adhesion of bacteria on the surface, and depends on many parameters such as chemical composition, electric charge, hydrophobicity, roughness and porosity. [ 326 ] Bacterial adhesion on magnetic thermoseeds has been compared to other newly developed multifunctional bioceramics which are capable of combining osseointegration with controlled release of drugs, such as mes-oporous silica materials. The tests performed with two strains of staphylococci, S. aureus and S. epidermidis , which cause most infections in biomaterials, show a signifi cantly lower bacterial adhesion in the case of magnetic bioceramics. [ 327 ]
5.4. Magnetic Micro- and Nanostructures
Magnetic microspheres have been proposed as suitable media-tors for inducing hyperthermia in cancerous tissues. When implanted through blood vessels, they would get entrapped in the capillary bed of tumours, and locally heat cancers by hys-teresis losses under an alternating magnetic fi eld. [ 328 ] This type of bioceramics can be synthesized following different routes. Magnetic microspheres obtained by melting powders in high-frequency induction thermal plasma and subsequent heating, are composed of 1 μ m size magnetite crystals. However, smaller crystal sizes can be attained by techniques based on precipita-tion from aqueous solution and subsequent heat treatment. In this case, higher heat generation rates are achieved.
The concept of miniaturization of bioceramics provides an added value in the biomedical fi eld. Although hyperthermia is one of the most attractive applications of biomaterials in med-icine, it is not the only possibility for magnetic systems. The intense research on magnetic nanoparticles in recent years has opened up many possibilities in areas of great impact such as targeted drug release, magnetic resonance imaging or gene transfection. Multiple methods have been proposed for the synthesis of these materials, resulting in suspensions with dif-ferent characteristics depending on their intended use. [ 304 ] The potential offered by nanotechnology is leading to profound changes in the design strategies of biomedical devices. The transition from bioceramic-based macroscopic implants to sub-micron particles opens the gate to developing stable magnetic materials suitably dispersed at physiological pH, with the ability to directly interact with protein and nucleic acid structures and provide therapeutic effects at the cellular level. [ 329 ]
Magnetic nanostructures have great potential for applica-tion in areas such as electronics, opto-electronics, magnetic memories, and various aspects of biomedicine. [ 330 ] At present, there is a need to prepare magnetic nanoparticles of controlled sizes and narrow particle size distributions, with the intention of obtaining materials with new properties. [ 331 ] The magnetic behaviour of these particles is dominated by their size and sur-face effects. However, the relationship between shape, surface structure, composition and magnetic properties of nanoparti-cles has not yet been clearly defi ned.
Magnetic particles may form suspensions with a superpara-magnetic behaviour, in which the magnetization ceases upon
removal of the applied fi eld. Among the metals and oxide sus-
pensions of particles with a size ranging from 1 to 100 nm, magnetite (Fe 3 O 4 ) and maghemite ( γ -Fe 2 O 3 ) iron oxides are most widely used for biomedical research purposes. [ 332–334 ] Whereas pure metals display higher values of magnetic suscep-tibility, they are highly toxic and extremely sensitive to oxida-tion. Under ambient conditions, Ni, Co and Fe are respectively oxidized to NiO, CoO and FeO, which are antiferromagnetic. Currently, there is no procedure able to avoid this oxida-tion mechanism, which is a major problem for nanoparticles because of their high surface area. For this reason, even with lower values of magnetization, iron oxides offer better guaran-tees for its implementation.
In all biomedical applications of magnetic nanoparticles, the coating with different molecules is of vital importance. These components serve to isolate the magnetic core from tissues, while ensuring the stability of the suspension in aqueous bio-logical fl uids. Both aspects work together to minimize the long-term toxicity of magnetic particles injected into the body. [ 335 , 336 ] Magnetic fl uids with nanoparticles suspended by the action of surface stabilizers are known as ferrofl uids. Aggregation due to attractive forces associated with the magnetic particles can be avoided by following different strategies: the creation of a bilayer causing electrostatic repulsion between particles, the use of surfactants that act as steric stabilizers, or modifying the iso-electric point with citrate groups or silica coatings. [ 337 , 338 ] There are also various methods for preparing composites of magnetic nanoparticles dispersed in organic or inorganic matrices. The advantage of incorporating the magnetic particles in submicron size diamagnetic matrices is the possibility to functionalize the surface of the material, as well as making it biocompatible.
On the one hand, magnetic nanoparticle suspensions can be stabilized with polymeric surfactants. Dextran with different chain lengths and degrees of branching is a polysaccharide widely used in the stabilization of nanoparticles. The most com-monly used method for coating nanoparticles of maghemite or magnetite with dextran is to include it as a surfactant during the formation of the particles. [ 339 ] Once the particles are sta-bilized, it is possible to functionalize dextran through oxida-tion, creating a larger number of aldehyde groups, capable of binding to protein amino groups in order to carry out magnetic separation. [ 340 ] On the other hand, magnetic nanoparticles have been encapsulated into inorganic matrices of gold, gadolinium, carbon, and mainly silica. [ 341 , 342 ] The presence of silanol groups on the surface of silica facilitates the reaction with alcohols and organosilanes, able to stabilize the suspensions in non-aqueous media, as well as providing groups for the attachment of specifi c ligands. Gold species also create a surface suitable for functionalization, [ 343 ] but the coating is not thick enough to prevent particle agglomeration.
5.5. Magnetic Drug and Gene Delivery Systems
One of the most promising applications of magnetic nano-particle suspensions is in the preparation of devices for drug delivery. An active targeting of drug carriers, which might be concentrated and retained in a certain place by means of an external magnetic fi eld, would allow a selective distribution
within the organs or tissues where the therapeutic agent is required. [ 344 ] Storm and coworkers determined that the mag-netic particles needed both high values of magnetization and a size below 100 nm for targeting purposes, so that they could prevent the action of the reticulo-endothelial system and pen-etrate the defective blood vessel network of a tumor. [ 345 ] The long circulation times that have been observed in liposomal carriers of antitumor agents or antibiotics have confi rmed this point, [ 346 , 347 ] which constitutes a drawback for magnetic sys-tems because the magnetic force is proportional to the volume of the particle.
The fi rst designs of magnetic drug carriers incorporated both anticancer agents and magnetic particles within albumin micro-spheres. [ 348 ] The effi cacy of magnetic albumin microspheres has been demonstrated when releasing doxorubicin through the arteries of rats, reaching peak levels of the drug up to 16 times higher than free doxorubicin and with a reduced accumulation in heart and liver. [ 349 ] However, these albumin-based carriers lack the necessary stability, and immune reactions against them have been reported.
Subsequently, magnetic colloids have been combined with polymer networks and inorganic matrices to produce nano-composites capable of holding the drug and protect it against mechanical and enzymatic degradation. In addition, these materials prevent the release of metals and metal ions from the magnetic component to the outer media, thereby inhibiting a potential cytotoxic effect. The surface of these magnetic struc-tures can be modifi ed with multiple ligands, so that therapeutic agents are retained through various types of chemical bonds that can be later broken in the target tissue environment as a result of changes in pH, hydrolytic processes or externally applied stimuli [ 350–352 ] ( Figure 27 ).
Gene transport and delivery by magnetic conjugates repre-sents a particular challenge. DNA fragments have to overcome several biological barriers when reaching the cell nucleus from the extracellular environment. [ 353 , 354 ] The covalent bonding of dendrimer molecules to magnetic nanoparticles has been revealed as a promising strategy to perform magnetic force-assisted transfection in vitro . In this kind of systems, the mag-netic component is aimed at facilitating an intimate contact between cell culture and gene vectors, thus providing a reduc-tion of transfection times [ 355 ] (Figure 27 ).
Many of the systems proposed in the literature as selective drug carriers are based on core-shell settings in which various species (silica, polymeric surfactants, non-polymeric organic stabilizers) coat magnetic nanocrystals. In 1994, the fi rst core-shell magnetic silica dispersions were obtained, [ 356 ] following a synthetic procedure known as Stöber method in which silica species grow in alkaline mixtures of ethanol and water. [ 357 ] The silica coating on a magnetic core leads to stable and biocompat-ible ferrofl uids, which have the ability to incorporate therapeutic substances or fl uorophores in the silica shell for drug delivery applications or cell labeling, respectively [ 358 ] (Figure 27 ). An alternative to this design is the encapsulation of hydrophobic drugs between magnetic nanoparticles and the silica layer, [ 359 ] so as to obtain a controlled release profi le under the infl uence of an external magnetic fi eld. This stimuli responsive effect can be applied for on-demand release of the drug. Hu and colleagues have proposed a bioceramic material capable of modifying the
Figure 27 . Magnetic multifunctional hybrid bioceramics may be engineered for several biomedical applications such as smart drug delivery, tumor cells targeting and imaging, and magnetic-mediated transfection.
ibuprofen release levels in the medium in response to a high frequency magnetic fi eld. [ 360 ] Some of these devices have also demonstrated an effective internalization in human cells. [ 32 , 361 ]
Inorganic silica matrices are aimed at mechanically and chemically protecting the magnetic cores and the rest of incor-porated species. In addition to core-shell structures, some other systems in which silica microspheres encapsulate a sig-nifi cant amount of magnetic nanoparticles have been investi-gated. [ 362–364 ] The synthesis of mesoporous materials via sur-factant structure directing agents results in excellent character-istics for the adsorption of drugs and their controlled release. A large amount of magnetic cores can also be introduced in a single step, so the fi nal material can reach temperatures in the range of hyperthermia under the external action of alter-nating magnetic fi elds. [ 365 ] This set of attributes allows com-bining hyperthermia treatment and chemotherapy against solid tumors. Both the mesoporous ordering and the magnetic prop-erties of the biomaterial can be adjusted by varying the ratio of surfactant and silica, and the number of encapsulated magnetic nanoparticles, respectively.
As explained in section 4.2, the use of mesoporous silica matrices for adsorption and controlled release of drugs in the body, has emerged as a promising strategy for those treat-ments that require a continuous supply of medications. Several physicochemical features, such as size, surface chemistry, mor-phology, aggregation tendency, homogeneity of dispersions or turbidity, will determine the suitability of these materials for a safe medical use. [ 366 , 367 ] Additionally, dosage levels have been shown critical for determining the toxicity profi le of mesopo-rous silica nanoparticles. [ 247 ] In vitro experiments with various
cell lines, demonstrated that this material is biocompatible at concentrations below 100 μ g mL − 1, [ 368–370 ] while some inhibi-tion of cell proliferation was reported with concentrations over 200 μ g mL − 1 . In vivo experimentation with animal models evi-denced that concentrations around 50 mg kg − 1 were safely tol-erated by mice. [ 247 , 371 ] Tamanoi and coworkers demonstrated that this dosage was suitable for pharmacological application for cancer therapy with 100–130 nm mesoporous silica par-ticles. This material was shown to effi ciently accumulate in tumors, as well as to suppress the growth of established human breast cancer xenografts after intraperitoneal injections of camptothecin loaded nanoparticles in mice. [ 371 ] Kohane et al . reported mice death or euthanasia of mice after intraperitoneal and intravenous injections of high doses of mesoporous sili-cates of particles sizes 150–4000 nm. [ 247 ]
In the case of magnetic mesoporous silica carriers, different synthetic approaches have been followed, such as the prepara-tion of microspheres by the sol-gel method and the tempera-ture-controlled pyrolysis of an aerosol including silica precur-sors and magnetic nanoparticle ferrofl uid at a suitable pH. Since the pyrolysis route is based on self-assembly of hybrid silica-surfactant induced by solvent evaporation (EISA method, Evaporation Induced Self-Assembly), [ 372 , 373 ] the temperature of the process is a critical parameter for obtaining structures with higher degrees of order.
5.5.1. Magnetic Carriers for Gene Delivery
Gene transfection is an alternative or complementary therapy in the treatment of cancer and other diseases with a genetic
Figure 28 . Covalently bonded dendrimer-iron oxide magnetic nanoparticles represent promising non-viral vectors for magnetically enhanced nucleic acid delivery.
+
SiEtO
EtOEtO
R=
γ-Fe2O3
Sol-Gel
MNPs
γ-Fe2O3(maghemite)
Covalent bonds
basis. The two possible strategies for the transport of genes into the cell nucleus are viral and nonviral vectors .[ 374–376 ] The fundamental requirement of these systems is the ability to con-jugate with and transport DNA fragments, protecting it from enzymatic degradation. In the case of viral vectors, transfection effi ciency is very high, but there is a high risk of provoking an immune response or other undesirable reaction. On the other hand, nonviral vectors are not associated with a signifi cant response by the body, and begin to be claimed as a safe alterna-tive to viral vectors. As an example, some polymer families with dendritic structures, such as polyamidoamines and polypropyl-eneimines [ 377–379 ] ( Figure 28 ), have shown promising levels of gene transfection.
The incorporation of magnetic nanoparticles into these devices provides the ability for magnetic guidance, as well as a potential adjuvant in anticancer therapy via magnetic hyper-thermia. There are some precedents for the combination of negatively charged DNA chains with iron oxide superparamag-netic nanoparticles coated with a polycationic molecule. [ 379 ] Thus, the application of an external magnetic fi eld facilitates the transport of genetic material into the cell and reduces process time. [ 380–382 ]
For this purpose, plasmid DNA may be combined with mag-netic nanoparticles to transfect cell cultures. When a perma-nent magnet is then placed below the culture dish, transfection rates may be enhanced. This procedure, known as magnetofec-tion, [ 383–386 ] might be suitable for hard to transfect cell types, and provides a fast alternative to current protocols).
González-Ortiz et al . have combined iron oxide magnetic nanoparticles with different dendrimer generations, and studied the physicochemical properties of the resulting mate-rials. [ 355 ] Their ability to transfect human tumor cells in vitro has been evaluated by conjugating a plasmid coding for the green fl uorescent protein, whose expression can be used to quantify the effectiveness of the process. In order to attach DNA mol-ecules to the magnetic component, iron oxide nanoparticles are previously functionalized with amine-terminated dendrimers,
which electrostatically interact with negatively charged DNA phosphate groups ( Figure 29 ). Besides, these dendrimers have a protecting function of DNA throughout its path towards the cell nucleus.
5.5.2. Magnetic Mesoporous Silica Carriers
Aerosol-assisted methods provide microparticles of spherical homogeneous morphology, useful for inclusion in preparations administered orally or parenterally. [ 387 , 388 ] The stability of these microparticle suspensions can be examined by electrophoretic mobility measurements as a function of pH. The isoelectric point determination of bare iron oxide magnetic nanoparticles revealed that they lose their electrostatic repulsion around phys-iological pH, [ 338 , 389 ] while magnetic nuclei-encapsulated mes-oporous matrices lead to stable suspensions of particles with a high density of negative charges on the surface.
Magnetic mesoporous microspheres have been designed by several groups by incorporating magnetic nanoparticles, and having a network of pores with diameters suitable for loading with drugs. [ 390–392 ] Pore size distribution in the carriers can be tuned by the use of different surfactant molecules in the self-assembly process. When non-ionic Pluronic P123 was employed as structure directing agent, 5.6 nm pore size matrices were produced by an aerosol-assisted method ( Figure 30 ). In this case, encapsulated magnetic nanoparticles kept crystallinity and superparamagnetic behaviour after the silica coating. This fact confi rms that the original properties of magnetic nanoparticles are retained after the pyrolysis process, so that the material preserves the ability to be guided by the action of an external magnetic fi eld. The silica microparticles presented an ordered majority hexagonal array of pores, with high values of specifi c surface area and pore volume. Moreover, the ability to load and subsequently release a model drug was evidenced. In this par-ticular case, the release profi le presented two different regions: an initial burst effect responsible for around 55% drug delivery in the fi rst hours of the assay, followed by a more controlled
Figure 29 . Amine-terminated dendrimer-iron oxide nanosystems are able to interact with negatively charged DNA via electrostatic interactions.
kinetics. [ 393 ] This pattern appears to be suitable in those cases requiring high initial doses, subsequently keeping medication levels in the organism.
Furthermore, the infl uence of the concentration of P123 sur-factant on the evolution of structural and textural parameters
Figure 30 . Aerosol-assisted synthesis stages of mesoporous silica micros
heatindryingN2
Ultrasound generator
Precursorsolution
Si(OH)4
Si(OH)4
Si(OH)4
Mesopha
γ-Fe2O3
of the hybrid composite has been investigated. [ 394 ] The highest pore ordering degree was found when 25% wt of P123 was employed. Similarly, the effect of adding different proportions of the magnetic component in the precursor solution was ana-lyzed, taking into account the structural, textural and magnetic
properties. In this respect, the maximum concentration of mag-
netic nanoparticles tolerated by the mesoporous structure was determined at iron oxide-silica weight ratios of 25%.
A modifi ed Stöber method has also been suggested to design porous structured microparticles suitable for drug carriage. [ 395 ] This synthetic route has some advantages for preparing spher-ical particles with a homogeneous size distribution and tailor-made pore ordering. [ 396 ] For this purpose, cationic surfactants are employed as structure-directing agents, and silica condensa-tion is carried out in alkaline conditions. Moreover, this proce-dure has been combined with aerosol-assisted methods, which allow using several surfactant types (cationic, anionic and even non-ionic). The integration of both processes leads to highly versatile materials with a great variety of arrangements and pore sizes.
López-Noriega et al. reported the combination of magnetic and non-magnetic P123-templated microparticles synthesized by aerosol pyrolisis, with an ordered mesoporous silica coating achieved through the modifi ed Stöber method. [ 397 ] In this mate-rial, two different hexagonal pore arrangements are obtained in each mesoporous phase. In addition, it was shown that the two-step methodology did not affect the magnetic nature of ferrofl uid nanoparticles incorporated in the synthetic precur-sors. These materials with pore hierarchy and high adsorption capacity have several advantages for the control of drug release kinetics.
Magnetic mesoporous silica beads can be internalized without causing adverse nonspecifi c effects in the host cell. [ 398 , 399 ] This feature promotes a selective intracellular release and represents a safe option for the management of toxic substances such as antitumor drugs. Additionally, these systems may perform
Figure 31 . The combination of mesoporous silica with different stimuli-re
magnetic hyperthermia treatments, which in combination with chemotherapy may represent a more aggressive therapy against tumours or lead to a reduction in the needed drug dose.
Highly magnetic mesoporous silica microparticles have shown a biocompatible behaviour when assessed in in vitro assays with several human cancer cell lines. Confocal micro-scopy studies revealed that the material did not interfere with cell extension or morphology. The microparticles were localized outside the cell nucleus, and there were no signs of nuclear condensation or fragmentation, which are characteristic of an apoptotic cell. Furthermore, the metabolic activity was not sig-nifi cantly altered. [ 400 ]
5.5.3. Smart Controlled Mesoporous Carriers
The previously mentioned special characteristics of mesoporous silica materials, such as a stable mesoporous structure, adjust-able pore sizes and high specifi c surface area and pore volume, make them suitable for application in the controlled release of therapeutic substances. The use of mesoporous silica spheres offers the additional advantage of allowing surface modifi cation with many functionalities. [ 401 ] They also offer the opportunity to retain their loading by the combination with several moieties placed in the pore outlets, so that release occurs in response to external stimuli: pH, temperature, redox reactions, reducing agents, enzymes or radiation [ 402 ] ( Figure 31 ).
The so-called stimuli responsive release systems are becoming very important within the framework of ordered mes-oporous materials as drug delivery systems. In this approach, the mesopore channels are designed as drug reservoirs that can be opened and closed on demand, responding to the external
sponsive moieties gives rise to smart materials for controlled drug delivery.
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stimuli. In this way, drugs can be confi ned into the mesoporous cavities where they may be protected against any possible deg-radation, and then locally delivered when needed.
Among smart materials with biomedical applications, it is possible to distinguish between those that respond to environmental stimuli, like intracellular pH or glucose con-centration, and those that respond to generated stimuli, like sonication or electromagnetic fi elds. [ 403–405 ] Several triggers have been reported to activate the release of guest molecules such as chemicals, where a reactant can remove the channels caps that were previously preventing drug molecules from leaching out. [ 406 ] In temperature responsive systems, thermo-sensitive polymers are employed to open or close the channels depending on the surrounding temperature. [ 407–409 ] pH is also used as a trigger, based on the fact that certain body tissues present a slightly more acidic pH level than blood or normal tissue, such as tumor or infl ammatory tissues, allowing a site-selective controlled release mechanism. [ 410–415 ] Pulsatile release of certain drugs has been reported by ultrasound exci-tation. [ 416 ] Moreover, redox systems allow designing nanovalves able to be turned on and off by external reducing/oxidising agents, [ 417–419 ] thereby acting as mesopore gates. Photorespon-sive molecules have been suggested as well, since they allow capping the channels until the exposure to light releases guest molecules. [ 420–423 ] Finally, magnetic nanoparticles are included among the release systems. The application of a magnetic stimulus is proposed not only as targeting mechanism, but also as a means for purposefully deliver the drug payload to a specifi c target. [ 424 ]
Moreover, magnetic nanosystems may integrate various stimuli on a single device. Giri et al . reported on the synthesis of mesoporous silica nanoparticles forming a covalent bond with magnetic nanoparticles acting as caps on the porous matrix. In this design, the covalent bond is only broken when the material comes into contact with the natural reducing agents inside the
Figure 32 . Magnetic fi eld-mediated thermoresponsive drug dosing reversiDNA-conjugated mesoporous silica nanoparticles.
Loading
T15 bpTm = 47.3 ºC
AT G A
A AA
A
T
TG GC
CT15 bpTm = 47.3 ºC
AT G A
A AA
A
T
TG GC
C
Cappin
CappingRelease
AA AAT
TT
TT
TC CC
GGAA AAT
TT
TT
TC CC
GG
Magnetic field
cell. With this approach, zero premature drug release may be achieved until the system reaches its goal. Once the magnetic component has been detached from silica nanoparticles, an external magnetic fi eld allows pore uncapping and subsequent release. [ 425 ]
The development of stimuli responsive systems that incor-porate biological mechanisms for a regulated administration of drugs in the body remains a scientifi c challenge. However, the manipulation of materials at the nanoscale allows the coupling of magnetic nanoparticles with nucleic acids. Self-assembly processes of biological molecules can be used for the design of smart materials that respond to external stimuli, which are capable of delivering controlled doses of drugs. [ 426 , 427 ]
Moreover, magnetic nanoparticles are capable of covalently joining nucleic acid strands that break the link with their complementary strand when subjected to an alternating mag-netic fi eld in terms of their dehybridization temperature. This temperature-sensitive system can be used as a gating mecha-nism at the outlet of a porous matrix. [ 428 ] Several parameters such as DNA strands length, nitrogen bases composition or even the amount of oligonucleotides anchored to the nanopar-ticle, may be responsible for variations in the temperature of response. [ 429 ]
For this purpose, thermoresponsive gates based on mag-netic nanoparticles have been suggested. Magnetic mesoporous silica matrices are functionalized with a single strand DNA and loaded with a drug model (fl uorescein). After that, the pores are capped with magnetic nanoparticles which are functionalized with the complementary strand. [ 428 ]
As a result, the DNA bonding of the magnetic particles with the silica matrix is progressively unwound by temperature increases, and this mechanism controls fl uorescein release. With this system, delivery is stopped when temperature is ceased and the magnetic nanoparticles block mesopore open-ings again ( Figure 32 ). In contrast, the release pattern of non-
bH & Co. KGaA, Weinh
ble mechanism of
g
conjugated particles has shown to be inde-pendent of increasing temperatures.
The external control of the system has been assessed by applying intermittent thermal sessions. After reaching a tempera-ture set point, this value was maintained for 40 minutes before increasing the tempera-ture to the next stage. [ 428 ] In this way, a stair-case profi le is shown, which is characteristic of stimuli responsive materials ( Figure 33 ). The possibility of controlling the therapeutic agent release by an external stimulus, and increasing surrounding temperature to levels of hyperthermia, opens up interesting pros-pects for thermosensitive nanomaterials in synergistic treatments of chemotherapy and heat.
6. Future Prospects
The huge advance in the design of function-alized materials as bone tissue engineering matrices should lead in the near future to a
5211eim wileyonlinelibrary.com
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Figure 33 . Thermoresponsive drug release profi le of mesoporous silica capped with magnetic nanoparticles through oligonucleotide linkages.
60
80
10060ºC
55ºC
42ºC
releas
ed (%)
37ºC
0 50 100 150 200 250 300 350
0
20
40
30ºC
Fluo
resc
ein
r
t / min
paradigm shift in orthopedics surgery: from structural bioengi-neering to tissue bioengineering. The profound knowledge of the behavior and properties of bone scaffolds may be accom-panied by a careful observation of guest-matrix-cell interac-tions. As a result, in vitro modeling of bone tissue scaffolds will be fostered. Eventually, minimally invasive techniques are expected to favor the clinical success of the next generation of biomaterials.
The development of novel advanced multifunctional mate-rials for a broad range of technological applications represents a renovated hope in many different fi elds. In particular, recent research breakthroughs in the biomedical arena have emerged as the basis for future personalized treatments and diagnostic techniques with a hitherto unsuspected selectivity.
With the amazing advances in the preparation and character-ization techniques of nanotechnology products, the possibility of manufacturing devices capable of establishing an intimate interaction with the biological world has been opened. This fact represents a precise control over the processes of therapeutic substances release, and means an opportunity to improve the specifi city of the therapeutic action, as well as to reconsider some of the promising drugs for certain diseases that were once discarded by their low levels of tolerance.
The targeting ability of these new nanodevices should lead to tailor-made dosing regimens, with a signifi cant reduction of severe side-effects associated to some diseases, such as cancer, and should also eventually result in a more effi cient allocation of health care resources. Moreover, the use of these engineered products may allow the combination of the ther-apeutic potential and nanoscale diagnosis with low tissue invasiveness.
Acknowledgements We thank the following for funding this work: the Spanish CICYT through project MAT2008-00736, and the Comunidad Autónoma de Madrid via the S2009MAT-1472 program grant. E.R.H. is grateful to Andalusian Regional Ministry of Health Postdoctoral Nanomedicine Fellowship for fi nancial support. We also thank J. M. Moreno and Pilar Cabañas for his technical help. The authors would like to express our deepest gratitude to all our co-workers and colleagues that have contributed over the years with their effort and thinking to these studies, and specially to Daniel Arcos for his work and scientifi c discussion in this manuscript.
Received: April 27, 2011Published online: October 18, 2011
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