4.35 Ordered Mesoporous Silica Materials☆

4.35 Ordered Mesoporous Silica Materials☆

4.35 Ordered Mesoporous Silica Materials☆ M Colilla and M Vallet-Regí, Complutense University of Madrid, Madrid, Spain and Networking Research Center ...

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4.35 Ordered Mesoporous Silica Materials☆ M Colilla and M Vallet-Regí, Complutense University of Madrid, Madrid, Spain and Networking Research Center on Bioengineering, Biomaterials and Nanomedicine (CIBER-BBN), Madrid, Spain r 2017 Elsevier Ltd. All rights reserved.

4.35.1 Introduction 4.35.2 Ordered Mesoporous Silica Materials: General Remarks 4.35.3 Biocompatibility of Mesoporous Materials 4.35.4 Mesoporous Materials for Local Drug Delivery Influence of the Textural Properties Pore diameter Surface area Pore volume Influence of the Matrix Structure Influence of the Particle Morphology Influence of the Matrix Organic Functionalization Functionalization with amino groups Functionalization with hydrophobic groups 4.35.5 Bioactivity of Mesoporous Matrices Incorporation of Phosphorous Into Mesoporous Silica Matrices Addition of Small Amounts of Bioactive Glasses Mesoporous Bioactive Glasses 4.35.6 Hierarchical Macroporous Scaffolds for Bone Tissue Engineering 4.35.7 Mesoporous Silica Nanoparticles for Antitumor Therapy Selective Targeting PEGylation of MSNs Passive targeting: Enhanced permeation and retention effect Nanocarriers with enhanced penetrability in solid tumors Cellular vehicles of MSNs Active targeting Antibodies Peptides Proteins Aptamers Saccharides Small molecules Stimuli Responsive MSNs Internal stimuli-responsive MSNs pH as release trigger Redox potential as release trigger Enzymes as release triggers External stimuli-responsive MSNs Light as release trigger Magnetic field as release trigger Ultrasounds as release trigger In Situ Cytotoxic Drug Generation Acknowledgments References

Abbreviations Ab Antibody ACP Amorphous calcium phosphate

646 647 649 650 651 651 652 652 652 653 653 654 655 656 658 658 658 659 663 663 663 664 664 665 666 666 667 667 668 668 668 668 670 670 670 672 672 672 672 674 675 676 676

AL Alendronate AMF Alternating magnetic field AMS Anionic surfactant templated mesoporous silica

☆ Change History: September 2016. M. Colilla and M. Vallet-Regí have made changes throughout the text. Authors’ affiliation, abstract, keywords, glossary, abbreviations, index, figures, tables and sections have been updated. Few of the references have been added and the reference list has been updated.

This is an update of M. Colilla and M. Vallet-Regí, 4.429 – Ordered Mesoporous Silica Materials. In Comprehensive Biomaterials, edited by Paul Ducheyne, Elsevier, Oxford, 2011, pp. 497–514.


Comprehensive Biomaterials II, Volume 4


Ordered Mesoporous Silica Materials

BSA Bovine serum albumin C8 Octyl chains C8TAB Octyltrimethylammonium bromide surfactant C10TAB Decyltrimethylammonium bromide surfactant C12TAB Dodecyltrimethylammonium bromide surfactant C16TAB Hexadecyltrimethylammonium bromide surfactant C18 Octadecyl chains CHA Carbonate hydroxyapatite CLCT Cooperative liquid-crystal template mechanism cmc Critical micelar concentration CTAB Cetyltrimethylammonium bromide surfactant DMSCs Mesenchymal stem cells derived from the decidua of human placenta DMS Dimethylsilyl groups ECM Extracellular matrix ED Electron diffraction EDS Energy dispersive scattering EPR Enhanced Permeation and Retention F127 Pluronic F127 block copolymer surfactant containing poly(ethylene glycol)-block-poly(propylene glycol)-block-poly(ethylene glycol) units (EO100PO65EO100) FA Folic acid FDU FuDan University FR Folate receptor FSM Folded sheet materials FTIR Fourier transmission infrared spectroscopy HA Hyaluronic acid HPLC High-performance liquid chromatography HRP Horseradish peroxidase IAA Indol-3-acetic acid IFP Interstitial fluid pressure KIT Korean Advanced Institute of Science and Technology LCST Lower critical solution temperature LCT Liquid-crystal templating mechanism MAA Methacrylic acid MABG Meta-aminobencylguanidine MBG Mesoporous bioactive glass MCF Mesostructured cellular foam MCF-NH2 Mesostructured cellular foam material functionalized with aminopropyl groups MCM Mobil Composition of Matter MCM-41-DMS MCM-41 mesoporous material functionalized with dimethylsilyl groups MCM-41-MS MCM-41 mesoporous material functionalized with methylsilyl groups MCM-41-NH2 MCM-41 mesoporous material functionalized with aminopropyl groups MCM-41-SH MCM-41 mesoporous material functionalized with mercaptopropyl groups MCM-41-TMS MCM-41 mesoporous material functionalized with trimethylsilyl groups

MDR Multidrug resistant MGHA Hydroxyapatite/mesoporous bioactive glasses nanocomposite MIBG Meta-iodobencylguanidine MMPs Matrix metalloproteinases MMSNs Magnetic mesoporous silica nanoparticles MS Methylsilyl groups MSCs Mesenchymal stem cells MSNs Mesoporous silica nanoparticles MSU Michigan State University NET Norepinephrine transporter NHMA N-hydroxymethyl acrylamide NMR Nuclear magnetic resonance NPs Nanoparticles OCP Octacalcium phosphate P123 Pluronic P123 block copolymer surfactant containing poly(ethylene glycol)-block-poly(propylene glycol)-block-poly(ethylene glycol) units (EO20PO70EO20) PAA Polyacrylic acid PEG Polyethylene glycol PEI Poly(ethylenimine) PNIPAM Poly(N-isopropylacrylamide) PTHrP Parathyroid hormone-related protein PUF Polyurethane foam PVP Poly(vinylpyrrolidone) SBA-15-C3N þ Me3 SBA-15 functionalized with Ntrimethoxysilylpropyl-N,N,N-trimethyl ammonium chloride SBA-15-C3N þ Me2C18 SBA-15 functionalized with octadecyldimethyl (3-trimethoxysilyl-propyl) ammonium chloride SBA-15-NH2 SBA-15 mesoporous material functionalized with aminopropyl groups SBA Santa Barbara Amorphous SBF Simulated body fluid SEM Scanning electron microscopy SMM Silica-based ordered mesoporous material SPIONs Superparamagnetic iron oxide nanocrystals TEM Transmission electron microscopy TEOS Tetraethylorthosilicate Tf Transferrin THPMA 2-tetrahydropyranyl methacrylate TLCT True liquid-crystal template mechanism TMOS-C3N þ Me3Cl N-trimethoxysilylpropyl-N,N,Ntrimethyl ammonium chloride TMOS-C3N þ Me2C18Cl Octadecyldimethyl (3trimethoxysilyl-propyl) ammonium chloride TMOS Tetramethylorthosilicate TMS Trimethylsilyl groups UV–vis Ultraviolet–visible XRD X-ray diffraction

Symbols DP

Pore diameter


VP Pore volume


Ordered Mesoporous Silica Materials

Glossary Bioactive agent Bioactive agents are those substances able to induce an effect on or to stimulate a response in a living tissue. Biomimetism Reproducing artificially essential properties of one or several biological systems. It finalizes products and processes as they are used in Nature. Endocytosis of nanoparticles Endocytosis is the process by which a nanoparticle is brought inside a cell without having to pass through the cell membrane. The formation of an endocytotic membrane vesicle or tubule is mediated by cytosolic coat proteins, which are nucleated and polymerized at specific sites on the plasma membrane to drive subsequently the deformation of the plasma membrane, ultimately to form an endocytotic vesicle. Simultaneously, coat proteins serve as “selection devices” that determine what nanoparticles are to incorporate into a emerging endocytotic vesicle, such as a receptor bound to its ligand. The culmination of these processes results in the cargo being sequestered within a coated membrane vesicle or tubule. Functionalization Functionalization is the addition of functional groups onto the surface of a material by chemical synthesis methods. The functional group added can be subjected to ordinary synthesis methods to attach virtually any kind of organic compound onto the surface. Local controlled delivery Release of an agent at a controlled rate for an extended time in a specific place to localize agent action. Macropores Macropores are those pores with diameter larger than 50 nm according to IUPAC pores classification. Biomaterials scientific community traditionally considers macropores as those pores with diameter above 100 mm. Mesopores Mesopores are those pores with diameter in the range 2–50 nm according to IUPAC pores classification. Mesoporous bioactive glasses Mesoporous bioactive glasses are materials with chemical compositions similar to those of conventional bioactive sol–gel glasses (SiO2–CaO–P2O5) and the textural properties and ordered mesopore arrangement of ordered mesoporous materials. Micropores Micropores are those pores with diameter smaller than 2 nm according to IUPAC pores classification.


Biomaterials scientific community traditionally considers micropores as those pores with diameter below 10 mm. Nanomedicine Nanomedicine is the application of nanotechnology to the prevention and treatment of disease in the human body. This evolving discipline has the potential to dramatically change medical science. Nanotechnology Nanotechnology is the understanding and control of matter at dimensions between approximately 1 and 100 nm, where unique phenomena enable novel applications. Encompassing nanoscale science, engineering, and technology, nanotechnology involves imaging, measuring, modeling, and manipulating matter at this small scale. Nanoscale The nanoscale has been defined as being of the order of 100 nm or less. Ordered mesoporous materials A mesoporous material is a material containing pores with diameters between 2 and 50 nm forming ordered arrangements of pore networks. They have high surface areas, great pore volumes, and narrow pore size distributions. These materials are excellent matrices to host different molecules, which make them suitable for numerous applications in catalysis, separation, biomedical field, etc. Rapid prototyping Rapid prototyping is the name given to a host of related technologies that are used to fabricate physical objects directly from computer-assisted design (CAD) data sources. These methods are unique in that they add and bond materials in layers to form objects. Such systems are also known by the general names solid freeform fabrication and layered manufacturing and offer advantages in many applications compared to classical subtractive fabrication methods such as milling or turning. Stimuli-responsive drug delivery systems Stimuliresponsive drug delivery systems are those devices able to control drug release in response to specific stimuli, either external (variations in temperature, magnetic field, ultrasound intensity, light or electric pulses) or internal (changes in pH, enzyme concentration or redox gradients). Surfactant Short for “SURFace ACTive AgeNT,” it is a molecule that lowers surface tension. Such molecules contain both hydrophobic and hydrophilic components and are thus semisoluble in both organic and aqueous solvents.


Since their discovery in the 1990s, silica-based ordered mesoporous materials (SMMs) have received much attention from the scientific community because of their outstanding textural and structural properties. During the last few years, SMMs have found many applications in catalysis, sensors, separation, etc. [1,2]. Recently, the biomedical research field has shown a growing interest toward these ordered mesoporous materials. It was in 2001 when MCM-41 was proposed for the first time as a controlled delivery system, using ibuprofen as the model drug [3]. Their stable mesoporous structures and well-defined surface properties make SMMs good matrices to host a wide variety of drugs and biologically active species for local controlled delivery applications [4–17]. The structural and textural properties of mesoporous matrices can be tailored depending on the targeted guest molecule. Moreover, there are a number of silanol groups covering the mesoporous silica walls susceptible of undergoing chemical functionalization with organic functions. The different approaches developed up to date to achieve better control over drug loading and release kinetics are tackled in the first part of this chapter.

Ordered Mesoporous Silica Materials


The in vitro bioactive behavior of SMMs has been also demonstrated [4,18]. Recently, the synthesis of mesoporous bioactive glasses (MBGs), with chemical compositions similar to those of conventional bioactive sol–gel glasses and the textural properties and ordered mesopore arrangement of SMMs has been reported [19–21]. MBGs exhibit fascinating features such as the fastest bioactive response reported till date for bioceramic materials or biomimetic behavior similar to the in vivo bone mineralization process [22]. All these aspects and the different chemical strategies carried out to modulate the time of bioactive response are also averviewed. This double scope, controlled delivery capability and in vitro bioactivity, makes ordered mesoporous materials promising candidates to be used as bioceramics for bone tissue-regeneration technologies [4,23–28]. In addition, it is also possible to organically modify the external surface of silica walls to graft certain biomolecules such as peptides, proteins, or growth factors that act as signals to enhance or stimulate new bone formation. Considering all these properties the question is: do these materials have good potential to be used as bioceramics in clinical medical practice? Their biocompatibility, bioactivity, capability to regenerate bone, and ability to act as controlled delivery systems of biologically active molecules have been confirmed. Actually, the preliminary in vitro and in vivo tests performed using SMMs have been favorable. So, the current challenge is to optimize all these properties to the clinical needs and to evaluate the performance of SMMs in comparison with that of bioceramics that are widely employed in clinical practice such as bioactive glasses. The second part of this chapter deals with mesoporous silica nanoparticles (MSNs), which are receiving increasing attention by the biomedical scientific community due to their groundbreaking potential in modern nanomedicine [29,30]. Among the diverse pathologies that can be treated with MSNs cancer has received growing attention [31–34]. They can be loaded with large amounts of therapeutic molecules and once intravenously administrated accumulate in solid tumours via the enhanced permeation and retention (EPR) effect [35,36]. We describe the more sophisticated stimuli-responsive MSNs developed up to date. MSNs can be loaded and the pore entrances closed using appropriate pore blockers, such as nanocaps or nanogates anchored to the MSNs’ surface via sensitive linkers. These nanocarriers are able to host, protect, transport the therapeutic agents at the target tissues and once there, upon exposure to internal (pH, enzymes, reductive species, etc.) or external (light, magntetic fields, ultrasounds, etc.) stimuli triggers pore uncapping and drug release. The recent insights into MSNs for cancer treatment, including sophisticated cellular vehicles of MSNs or in situ generation of cytotoxic drugs are tackled in this chapter.


Ordered Mesoporous Silica Materials: General Remarks

As previously commented, SMMs were reported for the first time in the early 1990s by Japanese researchers [37,38] and Mobil Oil Corporation scientists [39], who reported for the first time the synthesis of KSW-n and M41S families of mesostructured materials, respectively. Ordered mesoporous materials represent a new generation of structurally unique materials, which exhibit order at the mesoscopic scale (2–50 nm) and disorder at the atomic scale. The mesoporous channels are divided by amorphous silica walls and arranged periodically on lattices such as artificial atoms or molecules in ordinary crystals. The synthesis of mesoporous materials is based on the use of surfactants that act as templates, that is, structure directing agents, for the assembly and subsequent condensation of inorganic precursors [40,41]. SMMs are characterized by regular arrays of uniform channels, whose dimensions can be tuned through the chemical nature of the surfactant, additives, and synthesis conditions (pH, temperature, etc.) in the so-called “liquid-crystal templating” (LCT) mechanism [42]. This method is based on the formation of liquid crystals in mixtures of polar solvents and surfactants with a nonpolar tail group. However, different investigations in the formation procedures of SMMs led to suggest two different variations of the LCT mechanism (Fig. 1). As per the first mechanism, known as “true liquid-crystal template” (TLCT), the surfactant concentration is so high that under certain conditions of temperature and pH a lyotropic liquid-crystalline phase is formed in absence of the inorganic precursor. According to the TLCT mechanism, the amount of surfactant dissolved in the aqueous solution is above the critical micelar concentration (cmc), which allows the surfactant molecules to form aggregates called micelles. These micelles are formed because the hydrophobic tails of the surfactant tend to congregate to minimize contact with water while their hydrophilic heads orientate to maximize the interaction with the aqueous medium. The shape and size of these micelles depend on the composition and chemical nature of the surfactant and on the solution conditions, such as the surfactant concentration, pH and temperature, presence of cosurfactants, etc. These micelles aggregate themselves forming supramicellar structures with hexagonal, cubic, or laminar geometry depending on the experimental conditions. Thus, the geometry of the resulting mesoporous framework will depend on the geometry of the supramicellar aggregates. Then the silica source, which is normally tetraethylorthosilicate (TEOS) or tetramethylorthosilicate (TMOS), is added, which condenses around the supramicellar structures. The second proposed mechanism is the cooperative liquid-crystal template (CLCT), which suggests that it is also possible that the lyotropic liquidcrystalline phase is formed even at concentrations of surfactant molecules below the cmc, existing as a cooperative assembly of the surfactant and the already added silica precursor. Independently of the synthetic route followed, the last step in the synthesis of SMMs consists in the surfactant removal by calcination or solvent extraction. The surfactant removal leaves a network of cavities within the silica framework that will determine the physical–chemical properties of the resulting materials. These materials exhibit outstanding properties, such as stable mesoporous structure, high surface area (B1000 m2 g1), large pore volume (B1 cm3 g1), regular and tunable nanopore size (2–50 nm), homogeneous pore morphology (hexagonal or cubic pores), nontoxic and biocompatible behavior, and the


Ordered Mesoporous Silica Materials

Fig. 1 Schematic representation of true liquid-crystal template mechanism and cooperative liquid-crystal template mechanism used to describe the synthesis of silica-based ordered mesoporous materials.

possibility of being chemically modified with different organic groups grafted to the silica walls. Different mesoporous structures can be obtained depending on the synthesis pathway, such as MCM (Mobil Composition of Matter) series [39,42–46], SBA (Santa BArbara) series [47–49], MSU-n (Michigan State University) series [50], KIT-n (Korean Advanced Institute of Science and Technology) series [51], FSM-n (Folded Sheet Materials) series [52], FDU (FuDan University) series [53–55], and AMS-n (Anionic surfactant templated Mesoporous Silica) series [56]. Fig. 2 displays the structural models corresponding to different mesoporous structures exhibiting different symmetries. Among them, three well-known mesoporous materials are MCM-41, SBA-15, and MCM-48, whose structural models are displayed in Fig. 2. The two MCM-type materials are synthesized by using cationic quaternary ammonium surfactants under basic conditions. They have narrow pore size distributions ranging from 2.5 to 4.0 nm, with 2D-hexagonal p6mm structure (MCM-41) [57] and 3D-bicontinuous cubic Ia3d structure (MCM-48) [45]. SBA-15 mesoporous material is synthesized by using neutral nonionic surfactants (block copolymers) as templates. SBA-15 exhibits large pore sizes in the 6–10 nm range with 2Dhexagonal p6mm structure and complementary micropores. These micropores are arranged in a disordered fashion interconnecting adjacent mesopores [58,59]. SBA-15 is currently receiving much attention because of its possible utilization in applications that require materials with easily accessible, uniform, and large-size pores with good stability [60,61]. Besides, different research approaches have allowed the development of new SBA-15-type mesoporous structures with enhanced textural properties to increase the application fields of such matrices. Thus, SBA-15 materials with really large surface areas have been reported [62].

Ordered Mesoporous Silica Materials


Fig. 2 Models of different mesoporous structures, including their pore diameters (DP) and symmetries.

SMMs can be prepared with different morphologies, such as bulk [39,47], fibers [63,64], rods [65,66], films [67,68], monoliths [69,70], spheres [71], etc. Among them, spherical mesoporous silica nanoparticles (MSNs), with sizes in the 50–300 nm range have found many applications in nanomedicine, highlighting the development of new nanosystems for cancer therapy [72–74], at it will be discussed in Section 4.35.7. The synthesis of MSNs for application in nanomedicines must fulfill two major conditions: (i) well-controlled nucleation and growth kinetics to produce uniform sizes within the suitable range; (ii) non-existence of selfaggregation. There are two main synthesis methods used to produce MSN that meet these requirements. The first one is the modified Stöber method [75], which involves the silica condensation in basic medium and in the presence of cationic surfactants as templates. The second strategy is the aerosol-assisted synthesis, which allows using not only cationic but also anionic and nonionic surfactants to obtain MSNs [76]. The surfactant removal usually leads to materials with cylindrical mesopores exhibiting the 2D-hexagonal arrangement typical of MCM-41 [39].


Biocompatibility of Mesoporous Materials

Bioceramics aimed at bone implant applications must be in contact with physiological fluids. In this sense, long-term biocompatibility, which has been demonstrated to depend on the implantation site, the shape, and size of the material and also on its surface chemistry, is a mandatory requirement that must be satisfied. Among bioceramics, SMMs are good candidates to be used as bone implant materials, because they can host a wide variety of biologically active molecules for the treatment of bone diseases and also, synthesized as small size (50–300 nm) MSNs they are suitable to develop potential nanosystems for antitumor therapy. Amorphous silica has been generally considered as nontoxic, and it has been reported as biocompatible and degradable in living tissue [77–79]. In fact, amorphous colloidal and porous silica are used as adjuvant in pharmaceutical technology. To date, biocompatibility and degradability with time of different silica materials have been described [79–82], but few research works have reported the inherent cytotoxicity and biocompatibility of SMMs. The concentration of SMMs has been revealed as an important factor in their in vitro toxicity. SMMs showed low toxicity at low concentrations [83–89], whereas the toxicity increased at high concentrations [83]. It has been also shown that size and concentration of mesoporous silica particles affect the viability of human monocyte-derived dendritic cells as well as the expression of immune regulatory markers by these cells [90]. The surface area of an SMM has been also reported as an important factor that causes cell death. Thus, the in vitro cytotoxicity of MCM-41 mesoporous material, MCM-41 functionalized with aminopropyl groups (MCM-41-NH2) and mercaptopropyl groups (MCM-41SH) and spherical silica nanoparticles (SiO2), toward human neuroblastoma cells [91] has been investigated. Taking into account that the particles studied were not soluble in aqueous media, authors introduced a new quantity to measure the cytotoxicity of these materials, Q50, which is the number of particles required to inhibit normal cell growth by 50%. This study showed that Q50 increased in the order: MCM-41oMCM-41-SHoMCM-41-NH2ESiO2. These results would point to MCM-41 as the most cytotoxic of the studied materials. Albeit the chemical basis of the cytotoxic effect of these materials remains unclear, the exposed surface area of the particle could play an important role in cellular death. Therefore, the most toxic mesoporous silica materials


Ordered Mesoporous Silica Materials

were those with the largest surface areas. However, the chemical nature of the attached functional groups should be also considered. Moreover, surface area may not be the origin for cytotoxicity as spherical silica nanoparticles, which exhibited relatively small surface area, displayed cytotoxicity similar to that of MCM-41-NH2, which had higher surface area. All these findings suggest that cell death could be governed by other factors, such as size and shape. In vivo studies with rats evidenced low toxicity with small amounts of 1–2 mm SBA-15 [92] surgically implanted into the basolateral amygdale of rats. In addition, mesoporous particles (B100 nm) fused with iron oxide and amorphous silica particles exhibited low toxicity when intravenously administrated in mice [93]. Recently, in vivo studies using MCM-41, MCM-48, and mesostructured cellular foam (MCF) mesoporous materials have been reported [94]. Histological studies performed after subcutaneous injection of such SMMs at the sciatic nerve in rats evidenced a good biocompatibility at all-time points. However, intraperitoneal and intravenous injections in mice resulted in death or euthanasia. The above-mentioned results clearly show that toxicity of SMMs strongly depends on the administration route to the living body. Therefore, when evaluating the viability of SMMs as biomaterials for bone tissue regeneration, the first step should consist in performing in vitro tests with osteoblastic cell cultures. Accordingly, in vitro assays with osteoblastic cell culture using several mesoporous materials in the system SiO2–P2O5 with different phosphorus contents have been reported [95]. These preliminary in vitro assays indicated that the lowest cellular damage was induced by the sample with the highest P-content (5 mol%). This fact was ascribed to the greatest decrease in the amount of silica leached from matrices with the highest phosphorous content. Recently, the in vitro behavior of both unmodified and organically modified SBA-15 materials loaded with osteostatin, which is the sequence 107–111 of the parathyroid hormone-related protein (PTHrP), has been evaluated [96]. The results revealed that the incorporation of osteotatin into the SMMs increased cell proliferation and expression of several osteoblast phenotypic markers (alkaline phosphatase, osteocalcin, osteoprotegerin, receptor activator of nuclear factor-KB ligand, and vascular endothelial growth factor) in mouse MC3T3-E1 osteoblast cell line. Moreover, there was an increase in the matrix mineralization, and the secretion of collagen was stimulated after 12 days. Thus, it has been demonstrated that osteostatin-loaded SBA-15 materials promote osteogenesis in cell culture, which is interesting from the point of view of their potential clinical usage in bone regeneration technologies. Once these bioceramics successfully passed the in vitro tests, they were investigated under in vivo conditions [97–99]. Such experiments were performed in healthy and dexamethasone-induced osteoporotic rabbits. Cavitary defects in distal, medial, and lateral condyle were performed and both SBA-15 and SBA-15 loaded with osteostatin were implanted. The histopathological evaluation was carried out after 2 weeks of implantation. The cavitary defects in healthy animals revealed an efficient reparative process, with neoformed bone tissue around the biomaterials after 2 weeks. However, the incorporation of osteostatin into the mesoporous matrices was found to reduce tissue encapsulation, associated with increased cellularity and the appearance of osteoid deposits surrounding the biomaterial. These results open promising expectation for SMMs to be used as bioceramics for bone repair and regeneration technologies. Regarding MSNs, they can be endocytosed by a high number of mammalian cells, depending on surface functionalization, size and morphology of the NPs [100–103]. In vitro toxicology of MSNs reveals that they are well-tolerated by different cell lines at dosages lower than 100 mg mL1 [94], and only transient metabolic alterations are produced [104,105]. In vivo biocompatibility of MSNs indicates that dosages o200 mg kg1 are well-tolerated [106]. In addition, the good hemocompatibility, which is a condition for intravenous administration, has also been verified [107,108].


Mesoporous Materials for Local Drug Delivery

The excellent structural and textural properties of SMMs made them suitable to act as host systems of a wide range of different drugs for the treatment of bone diseases. This is especially important when dealing with anticancer and antiosteoporosis drugs for the local treatment of osseous tumors and osteoporosis, respectively. The local administration of therapeutic agents by bioceramic matrices presents advantages over conventional therapies, such as oral intake, because a better control over drug loading and release can be achieved. Moreover, the local sustained administration permits to keep constant drug levels in plasma and to reduce undesired side effects [109]. Sol–gel processed silica controlled release materials were first proposed in the 1990s [82,110–113]. Since 2001, when MCM-41 was proposed for the first time as controlled delivery system [3], several SMMs have been proposed as local delivery systems of different drugs, such as ibuprofen [114–121], naproxen [122,123], aspirin [124,125], piroxicam [126], gentamicin [127,128], amoxicillin [129], erythromycin [116,130], vancomycin [131,132], rifampin [132], linezolid [132], cephalexin [133–135], ciprofloxacin [136], levofloxacin [137], diflunisal [131], captopril [138], nimodipine [139], itraconazole [140], alendronate [141–144], zoledronate [145–147], ipriflavone [148], methotrexate [149], Chicago Sky Blue 6B [150], curcumin [151] and cisplatin [152]. In addition, SMMs have been also reported as controlled delivery systems of proteins [8,25,153–156], peptides [96,156–158], and amino acids [159,160]. Molecule adsorption and in vitro delivery tests are remarkable steps in the performance of SMMs as controlled delivery systems (Fig. 3). There are different methods of molecule adsorption, but the most commonly used are impregnation methods, in which the mesoporous carrier is soaked into a concentrated molecule solution. Loading conditions, such as temperature, pH, solvent, etc., must be selected attending to the chemical nature of the targeted molecule [161]. On the other hand, in vitro delivery assays are performed by soaking the loaded SMMs into simulated body fluid (SBF) [162] or saline solution at pH 7.4 and 371C to try to mimic physiological conditions. Normally, the amount of molecule released is monitored by chromatographic and spectroscopic

Ordered Mesoporous Silica Materials


Fig. 3 Schematic representation of drug loading and release processes in ordered mesoporous materials.

techniques, such as high-performance liquid chromatography (HPLC) and ultraviolet–visible (UV–vis) spectroscopy, or spectrofluorimetry for fluorescent molecules. The research performed since 2001 on ordered mesoporous silica indicates that the textural properties, mesoporous structure, morphology, and organic functionalization of the silica walls are the main factors that govern adsorption and release of molecules from SMMs. All these parameters can be tailored for the design mesoporous matrices, opening many paths in this research field.

Influence of the Textural Properties Pore diameter

The adsorption of molecules into SMMs depends on the mesopore diameter of the host matrix, that is, it follows size-selectivity criteria. Commonly, pore diameters slightly larger than the molecule dimensions (pore/molecule size ratio 41) are enough to allow the adsorption of molecule inside the pores. On the contrary, if the mesopore diameter is smaller than the molecule size, the molecule would adsorb only on the external surface of the material. This molecule-sieve effect is especially remarkable when the adsorption of large-size molecules, such as proteins, is targeted. Some authors have investigated the adsorption of globular proteins on MCM-41 materials [163]. Such studies demonstrated that proteins with dimensions larger than the mesopore diameter were only adsorbed in the outer surface of MCM-41. In addition, adsorption tests into SBA-15 of proteins displaying different dimensions were carried out [164]. These assays also evidenced that only proteins with sizes smaller than the mesopore diameter showed significant adsorption and consequently, that SBA-15 was acting as a molecular sieve. Among proteins, serum albumins, which are one of the major components in plasma proteins in humans, are of special interest. Albumin is usually composed of a single chain of 582 amino acids with an average length of 10 nm and width of 6 nm [165,166]. Albumins are capable of binding reversibly to a wide variety of ligands. They can also bind different drugs that can be released to sites where the pharmacological action is needed. Therefore, albumins could be employed to store, protect, transport, and deliver different drugs used for the treatment of bone diseases. SBA-15 has been used as host matrix of bovine serum albumin (BSA) [153]. However, the size of BSA is quite close to the SBA-15 pore diameter and the diffusion of protein molecules into the mesoporous channels is difficult. For this reason, an attractive strategy would consist in increasing the pore size of SBA-15 by modifying the synthesis conditions, such as using different hydrothermal treatments [167,168]. Recently, Vallet-Regí et al. reported BSA adsorption and release tests into SBA-15 with different pore diameters [154]. In such research work, pore diameters ranging from 8.2 to 11.5 nm were obtained for SBA-15 materials synthesized by using hydrothermal treatments ranging from 1 to 7 days. BSA loading tests indicated that there is a direct dependence of protein adsorption on pore size, that is, the greater the pore diameter, the higher is the amount of protein loaded (Table 1). Pore diameter has been also demonstrated to play a key role in molecule release kinetics. This fact was observed when testing MCM-41 matrices of different pore diameters as ibuprofen delivery systems. MCM-41 materials with pore sizes ranging from 2.5 to 3.6 nm were obtained by using short-chain or long-chain cationic surfactants, respectively [169]. Ibuprofen release profiles revealed that the drug release rate was modulated by the pore diameter of SMMs. Thus, the ibuprofen release to the delivery medium was faster as the pore diameter increased, confirming the diffusional nature of drug release.


Ordered Mesoporous Silica Materials

Table 1 Pore diameter and amounts of BSA loaded into SBA-15 materials submitted to different hydrothermal treatments ranging from 1 to 7 days Material

Time of hydrothermal treatment (days)

Pore diameter (nm)

Amount of BSA loaded (mg g1)

SBA-15-1d SBA-15-3d SBA-15-5d SBA-15-7d

1 3 5 7

8.2 9.5 10.1 11.4

151 234 242 270

Table 2 Surface area of MCM-41 materials synthesized using different surfactants and amount of ibuprofen loaded Surfactant

Surface area (m2 g1)

Amount of ibuprofen loaded (mg g1)

85%C8TAB-15%C10TAB 70%C8TAB-30%C10TAB C12TAB C16TAB

768 936 1087 1157

110 190 230 340

Note: The surfactants used during the synthesis are indicated for each case. C8TAB ¼ trimethyloctylammonium bromide; C10TAB¼ decyltrimethylammonium bromide, C12TAB dodecyltrimethylammonium bromide; C16TAB hexadecyltrimethylammonium bromide.

Surface area

The adsorption of molecules into mesoporous matrices depends on the adsorptive properties of the silica surface. Therefore, the chemical interactions between the silanol groups covering the silica surface and the functional groups of the guest molecule will determine the amount of molecule loaded (Fig. 3). Surface of SMMs is the textural parameter that indicates the silica surface susceptible of undergoing chemical interactions with the guest molecule and, consequently, it is expected that the higher the surface area, the higher is the amount of molecule loaded. This fact was corroborated when several MCM-41 matrices exhibiting different values of surface area were tested as ibuprofen delivery systems (Table 2). The same behavior was also observed when carrying out loading and release assays of captopril [138] into several 2Dhexagonal structures with different surface areas. This trend was also confirmed when MCM-41 and SBA-15 matrices, with surface areas of 1157 and 719 m2 g1, respectively, were tested as alendronate delivery system [141]. Alendronate is a potent bisphosphonate commonly used for osteoporosis treatments. The maximum loads of alendronate were 139 and 83 mg g1 for MCM-41 and SBA-15, respectively. All these findings demonstrate the direct dependence of the amount of drug loaded on the surface area of the SMM, that is, the highest surface areas promote the greatest amounts of drug loaded.

Pore volume

The adsorption of molecules into SMMs is a molecule-surface interaction, and consequently, the amount of molecule loaded depends on the pore diameter as the size selective limiting factor, and also on the surface area. However, it has been reported performing successive impregnations of MCM-41 material with ibuprofen solutions is possible, which allows to achieve complete pore filling and leads to higher drug loads [170]. This fact was explained by the increase in the intermolecular ibuprofen interactions, as confirmed by solid-state nuclear magnetic resonance (NMR) spectroscopy measurements [171]. In this case, the amount of ibuprofen adsorbed is limited by the pore volume of the SMM, that is, the larger the pore volume the greater is the amount of drug loaded. The pore volume must be also considered when the confinement of large-volume molecules, such as certain proteins, is aimed. With regard to this topic, MCFs have been proposed as host matrices of several enzymes and proteins [172,173]. The synthesis of MCFs is carried out by employing triblock copolymer surfactants and adding a swelling agent, normally trimethylbenzene, into the structure directing template solution [174,175]. For comparison purposes, loading and release tests of BSA were performed on SBA15 and MCF materials [4,8]. Different transmission electron microscopy (TEM) and N2 adsorption studies of MCF matrix revealed the existence of spherical cells of around 28 nm with windows of around 18 nm, with the pore volume of 2.0 cm3 g1 (Fig. 4). In the case of SBA-15, such studies indicated the presence of 2D-hexagonal pore networks with pore sizes of around 8 nm, with 1.1 cm3 g1 as the pore volume (Fig. 4). The BSA loading tests indicated that the amount of BSA loaded was higher in MCF (240 mg g1) than in SBA-15 (150 mg g1). Certainly, the different BSA loads could be ascribed to the difference in the pore volume, that is, the greater the pore volume, the higher is the amount of BSA loaded. Nevertheless, the contribution of other parameters, such as pore diameter, cannot be overruled. As previously mentioned, the BSA dimensions are very tight to the SBA-15 diameter, which could make the protein adsorption difficult. However, the wide pore size of MCF would be facilitating BSA penetration into the carrier cavities.

Influence of the Matrix Structure

The influence of pore connectivity and structure of SMM into drug release kinetics was observed by using different mesostructures as host matrices. Two SMMs with similar pore size, between 2 and 3 nm, and different structures were selected: MCM-41, which

Ordered Mesoporous Silica Materials


Fig. 4 Schematic illustration of host–guest interactions for large-volume molecules, such as bovine serum albumin (BSA) protein, into SBA-15 and mesocelullar silica foam (MCF) matrices.

contains 2D-hexagonal array of pores, and MCM-48, which consists of a 3D-bicontinuous cubic mesoporous structure with Ia3d symmetry (Fig. 2). Then, loading and release experiments were performed using erythromycin [116] and ibuprofen [117]. MCM48 matrix exhibits unique penetrating bicontinuous networks of channels that allow easy molecular accessibility and fast molecular transport. As a consequence of this, a faster drug release was observed from cubic matrices than from hexagonal materials, although the release profile was similar from both matrices, which exhibits first-order kinetics.

Influence of the Particle Morphology

The particle morphology of the mesoporous matrix represents the extension of the interface between the host material and the body fluid, so the effect on drug release kinetics is predictable. For this reason, MCM-41 was synthesized in the form of irregularly shaped powder and monodispersed spheres ranging from 490 to 770 nm in diameter, and ibuprofen loading and release assays were carried out [118]. Such experiments showed that microsized spheres displayed slower release rate than irregularly shaped powders. In fact, those spherical particles with the smallest sizes exhibited the slowest ibuprofen release kinetics. The control of the particle morphology of SMMs is also expected to lead to a more reliable and reproducible controlled delivery system. This control has been achieved by employing an aerosol assisted synthesis to obtain different mesoporous structures in the form of micrometric spheres as delivery system of triclosan, a widely used antimicrobial agent [176]. On the other hand, the use of hollow mesoporous spheres offers the possibility to store higher amounts of guest molecules than conventional SMMs. It has been reported that hollow spheres with mesoporous shells shells would exhibit more advantages in mass diffusion and transportation compared to conventional hollow spheres of solid shells. Within this porous shell, it is of special interest that porous channels penetrate across the shell for drug storage and delivery systems. This penetrating pore channels can be obtained by using poly(vinylpyrrolidone) (PVP) and cetyltrimethylammonium bromide (CTAB) as co-templates during the synthetic procedure [177]. The authors loaded the hollow spheres with ibuprofen by impregnation, leading to a storage capacity three times higher than that of conventional SMMs, as previously published [178].

Influence of the Matrix Organic Functionalization

The organic modification or functionalization of SMMs using different organic groups has been revealed as the cornerstone to modulate the adsorption and release of molecules from these matrices [4,179–181]. The surface of silica walls contains high density of silanol groups susceptible of organic modification by covalently grafting organic silanes ((RO)3SiR0 ) [182]. In general, there are two main functionalization methods that are applied to organically modify SMMs for controlled delivery purposes: (1) Co-condensation method: Also known as one-pot route, it involves the addition of the organic group, normally a silanecontaining reactive group, such as chloride, methoxy, or ethoxy groups, during the mesoporous synthesis and the entire


Ordered Mesoporous Silica Materials

functionalization process is carried out in a unique way. This route leads to functionalized materials in which the organic groups are grafted to the outer and to the inner surface of the silica walls. For this reason, and to avoid loosening of the mesostructural order, the functionalization degree cannot exceed 40 mol%. Two exceptional examples have been given for the vinyl- and thiol-functionalization up to 62 and 50 mol%, respectively, even though the resulting mesoporous silicas were not well-ordered [183,184]. In addition, functionalization up to 60 mol% of SBA-15 with carboxylic acid groups has been also reported [185]. (2) Post-synthesis method: It involves an organic modification process carried out under anhydrous conditions, once the freesurfactant SMM has been already synthesized. Albeit this method leads to heterogeneous distributions of the functional groups, it results in well-defined structures controlled by the silica matrix. Besides, this route guarantees that the organic groups are placed in the outer surface of mesopore walls, which leads to higher functionalization degrees than that obtained by co-condensation method. The functionalization method must be selected considering the targeted guest molecule. For instance, MCM-41 organically modified with aminopropyl groups through co-condensation and post-synthesis routes was used as host carrier for ibuprofen [119]. As expected, MCM-41 functionalized by post-synthesis route experienced higher decrease in the textural properties (pore diameter, surface area, and pore volume) than MCM-41 functionalized by co-condensation method. Moreover, ibuprofen release tests on these matrices revealed that drug release rate from MCM-41 functionalized by post-synthesis was smaller than that from MCM-41 functionalized using co-condensation route. This fact can be explained considering that aminopropyl groups are placed in the outer surface of the silica matrix, giving up higher functionalization degree capable of interacting with ibuprofen molecules and promoting higher drug retention. Proton nuclear magnetic resonance (1H NMR) spectroscopy measurements confirmed the existence of ionic interactions between amino groups from modified matrix and carboxylic groups of ibuprofen [186]. More recently, the organic modification of SBA-15 with aminopropyl groups and the two above-mentioned functionalization methods has been reported [153]. For comparison purposes ibuprofen and BSA were chosen as model molecules and loading and in vitro release assays were carried out. Such studies indicated that functionalization with aminopropyl groups using post-synthesis route led to the smallest ibuprofen delivery rate, in agreement with the results obtained for MCM-41. Oppositely, the smallest BSA delivery rate was obtained from SBA-15 functionalized by co-condensation. Such behavior was explained by the balance of electrostatic interaction and hydrophilic interaction between BSA and the functionalized SBA-15 matrix that would be more favorable in matrices synthesized by co-condensation synthesis method. The chemical nature of the matrix–drug host–guest interaction can be modulated by the appropriate choice of the organic groups depending on the targeted molecule and its application. For this purpose, two main approaches have been widely reported. The first strategy concerns the use of different organic groups able to induce attracting interactions with the functional groups of the guest molecule. For this purpose, the influence of host–guest interaction on drug adsorption and release was investigated by functionalizing MCM-41 through post-synthesis method with several organic functional groups (chloropropyl, phenyl, benzyl, mercaptopropyl, cyanopropyl, and butyl) and using ibuprofen as model drug [187]. Because of the different interactions of carboxylic group of ibuprofen with the tested organic functions, diverse load and release behaviors were observed. MCM-41 functionalized with polar groups loaded higher amounts of ibuprofen than MCM-41 functionalized with nonpolar groups. Furthermore, the smallest drug delivery rates resulted from matrices modified with mercapto or amino groups because such functions induced the strongest attracting electrostatic interactions with the carboxylic acid of ibuprofen. The second strategy involves the use of hydrophobic groups to functionalize the mesoporous silica walls. The main aspects concerning both approaches are tackled in the next sections, which are mainly focused on the functionalizations using amino groups and hydrophobic groups.

Functionalization with amino groups

The organic functionalization of SMMs with amino groups has been widely reported. For instance, MCM-41 and SBA-15 materials functionalized with aminopropyl groups via the postsynthesis route have been proposed as alendronate controlled delivery systems [141]. These research showed that the amount of drug loaded on amino-modified matrices (MCM-41-NH2 and SBA-15NH2) was almost threefold greater than that on unmodified matrices (Table 3). Concerning alendronate release behavior, aminofunctionalized matrices exhibited slower release kinetics than that of unmodified materials (Table 3). Recently, the organic modification of SBA-15 using different aminopropyl functionalization degrees has been reported as a good strategy to modulate alendronate dosage [143]. The functionalization was performed following the postsynthesis route but improved by adding a catalyst during the functionalization process. The results derived from this work pointed to a linear dependence of the amount of drug loaded on the functionalization degree. Besides, the higher the functionalization degree, the smaller the total amount of alendronate released to the delivery medium. For this reason, this gradual functionalization has been demonstrated to allow better control over drug loading and release. These results are of great interest for bone implant technologies, because the high potency of bisphosphonates requires only small local doses. The organic functionalization of SBA-15 using partially silylated poly(propyleneimine) dendrimers has been also reported as a good strategy to modulate drug dosage [115]. For this purpose, authors employed first, second, and third dendrimer generations and two different nominal degrees of surface functionalization, which led to a wide range of materials with a great variety of amine functionalization and surface coverage. The performance of such systems was tested in vitro using ibuprofen as model drug. Such studies revealed that the higher the dendrimer generation, the higher is the host–guest attracting interactions and, consequently,

Ordered Mesoporous Silica Materials

Table 3


Alendronate (AL) release data from MCM-41 and SBA-15 before and after amine functionalization


AL loading (mg g1)

% AL released after 24 h

AL total delivery time (h)

MCM-41 MCM-41-NH2 SBA-15 SBA-15-NH2

139.2 366.1 82.8 223.0

58 28 56 11

72 264a 264 264b


Incomplete delivery (76%). incomplete delivery (69%).


the higher are the drug loads. Moreover, drug delivery patterns exhibit a sustained drug release and the drug dosage is governed by the dendrimer generation. Amino functionalization of SMMs has been also revealed as a good approach to control adsorption and release of proteins such as BSA [153,154]. As previously commented, SBA-15 materials with different pore diameters, which were synthesized using different hydrothermal treatments, were functionalized by postsynthesis route [154]. It is well-known that organic functionalization following the postsynthesis route always leads to a decrease of pore diameter. In the case of BSA, the protein dimensions are very tight to the SBA-15 mesopore size and consequently, the amount of BSA loaded was smaller in amino-modified matrices than in unmodified materials. Nevertheless, the attracting interactions taking place between amine groups of functionalized matrices and amide groups of BSA strongly influenced protein release behavior (Fig. 5). Thus, the initial burst effect observed in BSA released from unmodified matrices (B90%) was reduced almost three times (B30%) after amino functionalization. Moreover, the BSA release from all SBA-15-NH2 mesoporous materials was incomplete, whereas the total amount of loaded BSA was released after 192 h from unmodified matrices. Amino functionalization of large-pore and large-volume MCF materials has also been probed to increase the affinity toward BSA molecule [4,8]. The results revealed that the amount of BSA loaded was higher in amino-modified MCF (MCF-NH2) than in unmodified MCF (Fig. 5). This fact can be easily explained because despite the reduction of pore diameter as a consequence of functionalization, the pore size of MCF-NH2 (B16 nm) is still big enough to permit the BSA penetration. Regarding BSA release behavior, as expected, functionalization exerts a noticeable effect in BSA delivery patterns. Thus, the initial burst effect of unmodified MCF (B60%) was significantly reduced in MCF-NH2 (B10%). All these findings demonstrate that amino functionalization of SMMs is a suitable strategy for the design of controlled delivery systems of many biologically active molecules exhibiting organic groups which are capable of chemically interacting with amino functions.

Functionalization with hydrophobic groups

The organic modification of SMMs with hydrophobic groups is a commonly used way to control the loading and release of certain molecules. For example, the functionalization of SBA-15 with different alkyl chains, octyl (C8) and octadecyl (C18), has been carried out and controlled delivery tests of erythromycin, a hydrophobic antibiotic, have been reported [130]. As previously commented, the functionalization always leads to a decrease in the textural properties of SMMs, which in this case, led to a decrease in the amount of erythromycin loaded in SBA-15-C8 and SBA-15-C18 materials, that is, 130 and 180 mg g1, respectively, compared to unmodified SBA-15 (340 mg g1). However, erythromycin delivery rate was slower from functionalized matrices than from unmodified SBA-15. This fact can be ascribed to the diminution of the wettability degree of the functionalized surfaces, which makes difficult the penetration of the aqueous release medium inside the mesoporous channels and induces slower drug delivery kinetics. Tang et al. reported the functionalization of MCM-41 matrices with methylsilyl (MS) or dimethylsilyl (DMS) groups and tested their capability as ibuprofen delivery systems [120]. As expected, functionalized materials loaded smaller ibuprofen amounts than unmodified MCM-41. However, MS and DMS groups led to slower drug release caused by the difficult diffusion of the delivery medium inside the mesopore channels. Therefore, the complete drug release was achieved after around 10, 20, or 500 h from MCM-41, MCM-41-MS, and MCM-41-DMS, respectively. These authors also reported the modification of ibuprofen-loaded-MCM41 using trimethylsilyl (TMS) groups [121]. Again, the in vitro release tests revealed that the presence of TMS groups remarkably delayed the drug release compared to unmodified MCM-41 material. SBA-15 materials functionalized using hydrophobic groups have been also proposed as controlled delivery systems of Ltryptophan (L-Trp) [159], a hydrophobic amino acid present in the 3D structure of many peptides, proteins, and growth factors of interest in bone tissue-regeneration technologies [188,189]. L-Trp exhibits a hydrophobic aromatic indol ring that makes necessary to increase the hydrophobicity of SBA-15, which contains high density of silanol groups. Actually, unmodified SBA-15 loaded less than 5 mg g1 of L-Trp, probably due to the extremely different chemical nature of the hydrophobic amino acid and hydrophilic SBA-15 (Fig. 6(a)). Thus, to increase the host–guest interactions, functionalization of SBA-15 using quaternary amines with different alkyl lengths, namely, N-trimethoxysilylpropyl-N,N,N-trimethyl ammonium chloride (TMOS-C3N þ Me3Cl), and octadecyldimethyl (3-trimethoxysilyl-propyl) ammonium chloride (TMOS-C3N þ Me2C18Cl), has been reported. The main forces that governed amino acid adsorption in SBA-15 functionalized with short alkyl chains (BC3N þ Me3) were coulombic attracting interactions between deprotonated carboxylic groups of amino acid (–COO) and protonated quaternary amines (–N þ R4)


Ordered Mesoporous Silica Materials

Fig. 5 Schematic representation of mesoporous matrix-bovine serum albumin (BSA) host–guest interactions for (a) pure mesoporous silica and (b) aminopropyl-functionalized mesoporous silica.

covering the mesoporous surface (Fig. 6(b)). In this case, the amount of L-Trp loaded into SBA-15-C3N þ Me3 matrix increased up to 43 mg g1. When using long hydrocarbon chains (BC3N þ Me2C18) two-thirds of the silica surface was functionalized. This high functionalization degree using hydrophobic chains induced strong interactions between mesoporous surface and indol group of 1 L-Trp (Fig. 6(c)), leading to L-Trp loading amounts of 82 mg g for SBA-15-C3N þ Me2C18.


Bioactivity of Mesoporous Matrices

When considering any material for biomedical purposes, the first consideration that should be addressed is the interaction between the material and the living body. Thus, the surface of such material must be carefully investigated. SMMs are characterized by exhibiting high surface areas and high surface density of silanol groups (Si–OH). It is well known that the bioactive response of conventional bioactive glasses depends on the presence of silanol groups. In addition, porosity has been evidenced as key factor that modulates the kinetics of formation of the carbonate hydroxyapatite (CHA) layer [190–195]. In this sense, SMMs are expected to induce bioactive response because they exhibit a high density of silanol groups that could act as nucleation sites for CHA. Moreover, these materials exhibit outstanding textural properties that could bring enhanced kinetics of CHA formation. The first reports on the in vitro bioactivity behavior of mesoporous materials involved three well-known structures, MCM-41, MCM-48, and SBA-15 [181,196,197]. The evaluation of the bone-bonding capability of bioceramics was performed following accepted in vitro protocols at 371C by soaking disk-shaped pieces of materials into SBF [162,198]. The modifications of SMM surfaces were monitored using different characterization techniques, such as Fourier transform infrared spectroscopy (FTIR) and Xray diffraction (XRD). Scanning electron microscopy (SEM) and TEM together with energy dispersive scattering (EDS) and electron diffraction (ED) are also employed. In the case of bioactive materials, all these techniques allow to confirm the formation of a new

Ordered Mesoporous Silica Materials


Fig. 6 Schematic representation of different chemical interactions between L-Trp and (a) unmodified SB-15: L-Trp is linked by weak hydrogen bond interaction, (b) SBA-15 functionalized with short-chain quaternary amines (BC3N þ Me3), and (c) SBA-15 functionalized with long-chain quaternary amines (BC18N þ Me3).

layer, mainly consisting of CHA, on the materials surfaces. However, MCM-41, MCM-48, and SBA-15 mesoporous materials exhibit different textural and structural features, which induce variations in the rates of formation of the CHA layer. Thus, SBA-15 and MCM-48 were able to induce the formation of a nanocrystalline CHA layer onto their surfaces after 30 and 60 days of assay, respectively. However, no evidence of formation of such layer was observed on the surface of MCM-41. These different in vitro bioactive behaviors could be due to the different concentrations of silanol groups present in each material and their diverse textural and structural properties. In this sense, MCM-41 presents a relatively small silanol density (B2 mmol Si–OH m2) compared to SBA-15 and MCM-48 (B13 mmol Si–OH m2), which could explain its nonbioactive response after 60 days of test. However, SBA-15 and MCM-48, which exhibited similar silanol density, presented very different rates of formation of the CHA layer. Such difference could be related to the different textural and structural properties of the matrices. Hence, mesoporous matrices with large and accessible pore networks should favor the ionic diffusion through the mesoporous framework, which should promote the formation of the CHA layer. This fact is in agreement with the experimental results that indicate that SBA-15, with the greatest pore diameters (B9 nm) and 2D-hexagonal arrangement of mesopores interconnected by micropores exhibited the fastest bioactive behavior among the three tested materials. These results clearly indicate that the textural and structural features of SMMs are important factors that are involved in their in vitro behavior. However, these are not enough conditions to achieve a timely bioactive behavior to be useful in vivo because the formation of the CHA layer on SMMs is much slower than that on conventional sol–gel glasses (around 3 days) [199]. Therefore, the chemical composition of materials, mainly the presence of calcium and/or phosphate, has to be also considered. Thus, the next sections will focus on the different strategies developed to date to accelerate the bioactive response of SMMs.


Ordered Mesoporous Silica Materials

Table 4

Structural characteristics, composition (% in mol) and time of bioactive response of different mesoporous materials



Composition (% in mol)


Time in SBF

MCM-41 MCM-48 SBA-15 P-MCM41 P-SBA-15 G-MCM41 MG-S58m MG-S75m MG-S85m

2D-hexagonal (p6mm) 3D-cubic (Ia3d ) 2D-hexagonal (p6mm) 2D-hexagonal (p6mm) 2D-rectangular (c2mm) 2D-hexagonal (p6mm) 2D-hexagonal (p6mm) þ wormlike 2D-hexagonal (p6mm) þ orthorhombic (p2mm) 3D-cubic (Ia3d)

100SiO2 100SiO2 100SiO2 99SiO2–1P2O5 82.6SiO2–17.4P2O5 95.6SiO2–4.0CaO–0.4P2O5 58SiO2–37CaO–5P2O5 75SiO2–20CaO–5P2O5 85SiO2–10CaO–5P2O5

No Yes Yes Yes Yes Yes Yes Yes Yes

60 Days 60 Days 30 Days 15 Days 14 Days 1 Day 1 Day 8h 1h


Induction time needed for the formation of a carbonate hydroxyapatite (CHA) layer onto the surface of materials.

Incorporation of Phosphorous Into Mesoporous Silica Matrices

In vitro bioactivity assays performed with bioactive sol–gel glasses revealed that the presence of phosphorous in the glass network affected the formation of the CHA layer [200]. This fact inspired the idea of incorporating small amounts of phosphorous, of around 1% in weight, into the MCM-41 silica network [201]. The structural and compositional characterization of this new material, denoted as P-MCM-41, using XRD, N2 adsorption porosimetry, SEM-EDS, and TEM-EDS techniques pointed to the preservation of the 2D-hexagonal p6mm mesoporous structure. It was also revealed that the phosphorous content was heterogeneously distributed into the mesoporous silica network and mainly placed in the outer edge of the particles and not in their inner nucleus. The in vitro bioactivity assays demonstrated that P-MCM-41 developed a CHA layer on its surface after 15 days of assay. More recently, SBA-15 containing around 17 wt% phosphorous, exhibiting a 2D-rectangular structure c2mm plane group was synthesized [144]. In vitro positive responsive was obtained after soaking for 14 days in SBF. Compared to the bioactive behavior of pure silica SBA-15 or MCM-48, the incorporation of phosphorus into the silica network accelerates the bioactive response of SMMs (Table 4).

Addition of Small Amounts of Bioactive Glasses

A widely employed strategy to induce bioactivity to different materials consists in preparing biphasic mixtures by incorporating another phase that acts as an accelerator of the CHA nucleation [202–204]. For this purpose, MCM-41 as the main phase (90% in weight) was mixed with a bioactive sol–gel glass in the system SiO2–CaO–P2O5 (10% of weight) [196], which gave rise to the socalled G-MCM-41 material. The in vitro bioactivity tests of this material revealed that the formation of the apatite layer started just after 3 h of assay and were completed in 1 day (Table 4). These results indicate that the addition of small amounts of bioactive glasses is a good strategy to reduce the time of bioactive response of SMMs materials to just 1 day.

Mesoporous Bioactive Glasses

Recently, another strategy has been developed to accelerate the time of bioactive response. This approach consists in using the synthesis methods of SMMs but with the same compositions of glasses. It was in 2004 when the synthesis of highly ordered MBGs in the system SiO2–CaO–P2O5 was reported for the first time [20]. The synthesis of these MBGs was carried out by using the evaporation-induced self-assembly (EISA) method [205] in the presence of a nonionic triblock copolymer as template (Pluronic P123, EO20PO70EO20) and with the employ of TEOS, triethylphosphate (TEP), and calcium nitrate (Ca(NO3)  4H2O) as SiO2, P2O5, and CaO precursors, respectively. The enhanced textural properties of MBGs promoted by the template led to highly bioactive responses compared to conventional bioactive glasses [206–208]. Moreover, it has been demonstrated that the textural and structural properties of MBGs can be modulated by varying the CaO content of the mesoporous glasses [19]. A gradual evolution from 2Dhexagonal to 3D-bicontinuous cubic structure together with an increase in the textural properties was observed when decreasing the CaO content (Fig. 7). The possibility of tailoring both structural and textural features of MBGs is an interesting advance toward the development of biomaterials able to fulfill the essential requirements for specific biomedical applications. The in vitro assays carried out in SBF demonstrated that this new family of mesoporous glasses exhibits enhanced bioactive behaviors, with even faster apatite phase formation than conventional bioactive glasses. Hence, it has been revealed that the kinetics of formation of the CHA layer in MBGs is controlled by their textural and structural properties [209–215]. In this respect, MBGs differ from bioactive glasses, in which the formation kinetics mainly depends on the composition and textural properties [216]. Thus, the MBG material with the smallest CaO content (MG-S85 m) but having 3D-bicontinuous cubic structure and the highest surface area values exhibited faster bioactive response than MBG materials with the greatest CaO content (MG-S58 m). FTIR and TEM-ED studies revealed an accelerated bioactive behavior for MG-S85 m sample, which was able to form a nanocrystalline CHA layer in just 1 h after coming into contact with SBF [22] (Fig. 8). In this case, the 3D pore system not only provided high surface area and porosity, but also promoted the ions interchange by increasing mass transport and diffusion processes, which resulted in the fastest bioactive behavior reported till date in silica-based bioactive materials.

Ordered Mesoporous Silica Materials


Fig. 7 TEM images and their corresponding FT diffractograms of MBGs in the SiO2–CaO–P2O5 system with different CaO contents. TEM images were taken with the electron beam parallel to the pore channels of 2D-hexagonal structures corresponding to mesoporous glasses with compositions (a) 58SiO2–37%CaO–5%P2O5 and (b) 75SiO2–20%CaO–5%P2O5, respectively. (c) TEM image taken in the [1 0 0] directions of 3Dbicontinuous cubic structure with Ia3d symmetry corresponding to a mesoporous glass of 85SiO2–10%CaO–5%P2O5 composition (see Table 4).

Fig. 8 In vitro bioactive study in SBF of MG-S85 sample, with 3D-cubic structure by TEM.

Recently, a new property derived from the textural and structural characteristics of MBGs, namely biomimetic behavior, has been reported. This biomimetic mechanism, where there is a sequential transition from amorphous calcium phosphate (ACP) to octacalcium phosphate (OCP) to calcium deficient carbonatehydroxyapatite (CHA) maturation, similarly to the in vivo biomineralization process, has been observed for the first time [22–24]. Many bioceramics develop a CHA phase in vitro through the direct crystallization of previously precipitated ACP [217], without the prior formation of OCP phase which is formed in natural bone biomineralization process [218–220]. OCP is a metastable phase and will appear only if the pH in the crystallization system is below 7. MG-S58 m sample, with 2D-hexagonal mesoporous arrangement and high CaO content allows an intense Ca2 þ –H3O þ exchange as a consequence of the open channel array, the high textural parameters, and the relatively high CaO content. This ionic exchange led to local acid pH values of 6.5 during the first stages, which favor the OCP formation (Fig. 9). This phenomenon does not take place in the case of bioactive glasses, because the local pH is kept at pH 7.4 or even slightly higher from the beginning of the bioactive process, which would avoid the existence of OCP phase. This biomimetic behavior was monitored by TEM (Fig. 9). After being soaked in the SBF for 1 h, MG-S58 m sample developed a large amount of newly formed ACP layer with a Ca/P ratio of 1.2. This event has been widely observed in many bioactive compounds, and corresponds to the step 4 of the bioactive process as theorized by Hench [221]. To date, the mechanism of in vitro bioactivity involves the direct crystallization of CHA phase from the ACP. However, MBGs have shown a unique behavior allowing the growth of nanocrystalline oval biphasic nuclei composed of OCP with a small fraction of HA after 4 h in SBF, with a Ca/P ratio of 1.3. The transformation from oval OCP nuclei to needle-like nanocrystalline CHA was finally evident after being soaked in SBF for 8 h. All these results bring up the possibility of controlling the reactive properties of MBGs by tailoring their compositional, textural, and structural properties. This new family of ordered mesoporous materials widens the range of possibilities for the use SMMs as new generation bioceramics with outstanding bone-bonding capability.


Hierarchical Macroporous Scaffolds for Bone Tissue Engineering

The different research results reported so far confirm that ordered mesoporous materials are characterized by two valuable features: controlled delivery capability and in vitro bioactive behavior. It would be also feasible to combine both properties to design new


Ordered Mesoporous Silica Materials

Fig. 9 In vitro bioactive behavior of MG-S58 m mesoporous glass with a 2D-hexagonal structure and 58SiO2–37%CaO–5%P2O5 composition. High-resolution TEM study of mesoporous glass surface after different soaking times in SBF together with the schematic depiction of the sequential transition from ACP to OCP to CHA observed by TEM are displayed.

bioceramics with an added value. Nevertheless, when dealing with ceramics aimed at bone tissue engineering applications, it is essential to keep in mind that the main role will be played by cells. Pore dimensions of mesoporous materials fall in the 2–50 nm range, whereas the dimensions of cells are in the 1–100 mm range. That is, mesopores are too small to allow cell uptake, which need dimensions in the order of microns. Bone porosity, which ranges between 1 and 3500 mm, is necessary for several physiological functions [222,223]. Bone tissue engineering requires the design of hierarchically 3D scaffolds with interconnected macroporosity, within the 20–1000 mm range [224–226]. Such macroporosity is essential to allow bone cell penetration, adherence, growth, and proliferation leading to bone tissue in-growth and eventually vascularization on implantation. SMMs, with ordered mesoporous channels and pore dimensions in the 2–50 nm range, are appropriate for applications in which drugs or biologically active molecules can be loaded and subsequently released where needed. On the other hand,

Ordered Mesoporous Silica Materials


Fig. 10 Schematic representation of the two scales of porosity present in an ideal 3D scaffold for bone tissue engineering. Macroporosity is similar to that of natural bone to allow new bone in-growth and mesoporosity to host drugs and biologically active species for controlled delivery functions.

macroporous materials, with pore dimensions in the microns scale, are adequate as scaffolds for bone tissue engineering. So, the challenge is to design and fabricate scaffolds that combine two scales of porosity; macroporosity to allow appropriate bone oxygenation and vascularization and mesoporosity for controlled delivery purposes (Fig. 10). The aim is to fabricate biomaterials that help the human body to improve its regeneration role by recovering not only the structure of the damaged tissue, but also its function. So the current goal in biomaterials is to fabricate 3D scaffolds with hierarchical interconnected porosity that allows cells to proliferate and form new bone tissue following the natural processes of human tissue. Several recent research works have been reported concerning the fabrication of 3D scaffolds using the highly bioactive MBGs as starting materials. Thus, Yun et al. reported the synthesis of ordered giant-, macro-, and mesoporous bioactive 3D glass scaffolds by using a combination of sol–gel, double polymers templating, and rapid prototyping (RP) techniques [227]. The resulting scaffolds presented three lengths of porosity: mesopores (B5 nm) obtained by using a triblock copolymer template (EO100PO65EO100 (F127)), macropores (10–30 mm) produced by using methyl cellulose, and giant-sized pores (30–1000 mm) created by RP methods [228,229]. The in vitro tests in SBF revealed that the bioactive behavior of MBG was preserved because a CHA layer was formed onto the material surface after 24 h of assay. Li et al. prepared hierarchically MBG-based scaffolds using P123 surfactant as mesostructure directing agent and polyurethane foam (PUF) as template for the macroporous structures [230]. The resulting materials presented hierarchical porosity with interconnected macropores of 200–400 mm or 500–700 mm and uniform mesopores of 3.7 nm. In vitro bioactivity tests showed that these MBG-based scaffolds induced the formation of a CHA layer on their surface after 4 h in SBF. Other authors also have reported the employment of P123 and PUF to synthesize hierarchically structured 3D MBGs-based scaffolds with four different chemical compositions and their in vitro bioactivity and cell adherence evaluation [231]. These scaffolds exhibited similar mesostructural and textural features, showing interconnected macroporous networks with pore diameters in the 200–400 mm range and mesopores of 4.9 nm in size. Cell cultures indicated that primary human bone-derived cells were able to attach and spread to different degrees on the different scaffolds. These authors also indicate that the differences observed in supporting cell growth and differentiation observed for different scaffolds could be related to the CHA formation on the surface of scaffold, which has been reported to affect cell activity [232,233]. Additionally, the release of calcium and silicon ions may be contributing to the modulation of cellular attachment [234]. Recently, Cicuéndez et al. reported the synthesis of hierarchical 3D scaffolds based on new hydroxyapatite/MBG nanocomposite bioceramic (MGHA) exhibiting different scales of porosity [235]: (i) highly ordered mesopores with diameters of 10 nm; (ii) macropores with diameters in the 30–80 mm range with interconnections of 1–10 mm; and (iii) large macropores of ca. 500 mm. 3D scaffolds were prepared by the RP technique using an slurry from a sol precursor of MGHA nanocomposite previously synthesized in poder [236], to which methylcellulose was added. Cylindrical 50  133  38 mm3 scaffolds were designed and fabricated layer-by-layer by direct ink deposition over a plate at room temperature. Finally, after drying at 301C, the


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Fig. 11 Scheme of the requirements for scaffolds aimed at bone tissue engineering. The biological point of view involves the presence of osteogenic cells. Osteogenesis process can be improved throughout by the presence of osteoinductive signals, which represents an added value. The mechanical aspect involves that bone remodels itself as a function of those forces acting on it. The two approaches used in bone tissue engineering are here displayed. Left down side: bone tissue regeneration starts in vitro. Right down side: osteogenic agents are grafted to the scaffold and implanted to allow the in vivo regeneration.

pieces were calcined at 7001C for 6 h. With the aim of enhancing their surface properties, 3D scaffolds were functionalized with amine propyl groups, which remarkably improved preosteoblast adhesion, proliferation, differentiation and further cell colonization of these scaffolds. The observed improvement could be related to these amine groups that would favour early adhesion. The design of 3D scaffolds capable to drive cell in-growth is an important challenge in tissue engineering. The aim is to design and fabricate pieces that support and structure the newly formed tissue and such pieces must be made starting from the most appropriate materials for this purpose. Cells must be cultured onto this scaffold to subsequently give rise to the growth of new tissue. Wolf's law states that the bone remodels itself as a function of those forces acting on it, hence preserving its shape and density (Fig. 11). The mechanical loads of stress, compression, flexion, and torsion in bones and the interstitial fluid contained in them originate stresses and deformations at the microscopical level, which stimulate the cells. The use of finite element calculation methods contributes to the study of interactions between materials, mechanical stimulations, and biological response. Such a method allows simulating the conditions that take place in bioreactors in the scaffolds. Moreover, the effect and deformation at each point of the scaffold can be quantified, for instance, at each node, and the relationship between mechanical loads and cell differentiation can be also quantified [237]. One of the most employed approaches in bone tissue engineering involves the initialization of the regeneration process in vitro by soaking the scaffold in appropriate cell culture and in the presence of osteogenic agents (Fig. 11). Then, the scaffold is implanted in the patient [238,239]. However, the current challenge in bone tissue engineering is to chemically graft such osteogenic agents into the 3D scaffold to be directly implanted in the patient, acting as osteoinductive signals for bone regeneration (Fig. 11). So, turning back to ordered mesoporous materials, the presence of silanol groups in the external surface of silica walls would allow the organic modification of the scaffold. The choosing of the appropriate functional groups could permit formation of strong chemical bonds between the material surface and different osteoinductive agents, such as certain peptides, proteins, or growth factors, as schematically displayed in Fig. 12. This approach permits to “decorate” the scaffold by chemically grafting different active agents, such as certain peptides, proteins, or growth factors, that act as potent osteoinductive signals able to promote the appropriate bone cellular functions in the place where needed [240–242]. At is has been widely reported, SMMs allow the design of controlled delivery systems for drug loading and release. Among other applications, these systems could be very useful to locally treat several pathologies such as bone infection and inflammation, osteoporosis, cancer, etc. However, although such molecules could fit into the mesopore channels, the confinement of osteogenic agents into the mesopores would impede them in carrying out their function. The role of these molecules is to act as attracting signals for bone formation cells, but these cells are too large to fit inside the mesopores (Fig. 12). However, it is possible to apply the acquired knowledge from these systems to the grafting of the mentioned peptides, proteins, or growth factors onto the external

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Fig. 12 Schematic representation of a mesoporous silica material whose external surface has been organically modified to graft different osteogenic agents, such as growth factors or peptides. It is also schematically depicted that these osteogenic agents could fit inside the mesopores but that would obstruct them from playing their role. Such osteogenic agents act as signals for bone formation cells, and these cells (of B100 mm) can never fit inside the mesopores (of 2–50 nm in size).

silica surface to induce bone tissue regeneration. Moreover, the possibility of using mesoporous matrices starting materials for the fabrication of 3D macroporous scaffolds provides these biomaterials an added value, their capability of hosting drugs to locally treat bone pathologies.


Mesoporous Silica Nanoparticles for Antitumor Therapy

Selective Targeting

The lack of specificity of drugs can provoke undesired side effects and decrease the efficacy of the therapy. This is a key problem in cancer therapies, where the poor target selectivity of highly cytotoxic drugs causes severe side effects. Moreover the efficacy of the treatment is compromised since unspecific toxicity to healthy cells avoids the effective dose required to kill tumor cells. Therefore, the administration of high drug doses is required, making the risk-to-benefit ratio frequently uncontrollable [243]. MSNs are promising nanocarriers to selectively transport drugs to cancer cells. Selective drug delivery to tumors can be attained by passive or active targeting, or by a combination of both [244–246].

PEGylation of MSNs

Nanocarriers are usually administered by intravenous injection. Thus nanocarrier must be designed with “stealth” properties towards the cells of the immune system to avoid being removed from the bloodstream before reaching the target. Actually, the optimal range size for MSNs in antitumor therapy is 50–300 nm [247]; particles larger than 2 mm can be trapped within the lung capillaries, which could cause alterations in the pulmonary circulation compromising the patient’s life; on the contrary particles smaller than 5–6 nm are quickly cleared by kidneys. Once MSNs come into the bloodstream, they contact with a highly complex fluid, the blood, which contains more than 3700 proteins and many other biomolecules (sugars, cholesterol, amino acids, etc.) that can interact with nanocarriers varying their pharmacokinetic profile [248]. Thus, after a given time into the bloodstream the surface of MSNs is fully covered with diverse proteins and molecules that produce the corona of the nanocarrier [249]. Among blood proteins, there is protein family called opsonins, whose mission is to stick to the surface of foreign bodies and allow the detection by cells of the immune systems, such as macrophages. Consequently, when the target of MSNs is not immune cells it is essential to prevent opsonins adhesion onto their surface. Several strategies have been developed to hinder opsonins adsorption, usually involving the surface functionalization of MSNs with hydrophilic polymers, which block the proteins attachment on the surface by steric hindrance. The most extensively used polymer is polyethylene glycol (PEG) and the functionalization process is called PEGylation [250]. PEG is soluble in water, although its solubility decreases as its molecular weight increases. It exhibits low toxicity, non-immunogenicity and high chemical/


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Fig. 13 PEGylated MSNs.

thermal stability. Moreover, its terminal hydroxyl group (–OH) can be oxidized or functionalized to link targeting molecules (Fig. 13). Linear PEGs with weights in the 1500–5000 Da range are usually employed to prepare “stealth” nanocarriers [251]. Grafting of PEG chains to MSNs can be carried out via covalent bonding using PEG-silane derivatives and following a cocondensation [252] or postsynthesis [253] method. The incorporation of PEG chains on the mesoporous surface prevents opsonization meanwhile improving the stability of MSNs in biological fluids [254].

Passive targeting: Enhanced permeation and retention effect

One of the major advantages of using NPs for cancer therapy is that they can accumulate at tumor tissues due to the so-called enhanced permeation and retention (EPR) effect, which was discovered in 1986 by Matsumura and Maeda [35,36]. This phenomenon involves that macromolecules greater than 40 kDa or NPs intravenously administrated can leak out from tumor blood vessels but not from normal vessels. This is provoked by the characteristic blood vessel architecture present in the solid tumors. Healthy vessels show a well-organized branching hierarchy from large vessels into smaller ones that feed a regularly spaced capillary bed. On the contrary, when a solid tumor is growing, the necessity of oxygen and nutrients increases. As a consequence, the tumor cells secrete (or for the other cells to do it) angiogenic factors that prompt the formation of novel blood vessels. Nonetheless, the formation of these new vessels is very irregular and they present wide interendothelial junctions, high number of fenestrations, and transendothelial pores with diameters as large as several hundreds of nanometers [255]. Therefore, when the injected NPs reach this area, they are able to escape from the blood vessels through these fenestrations (extravasation) being accumulate in the tumor interstitium [256]. In addition the extravasated NPs are retained within the tissue for long times due to the lack of an effective drainage system into the tumor area provoked by the collapse of the lymphatic vessels obeying to the fast growing rate of the diseased cells [257]. The combination of high permeability (enhanced permeability) and lack of drainage (retention) constitutes the basis of the EPR effect (Fig. 14).

Nanocarriers with enhanced penetrability in solid tumors

One of the main restrictions of using NPs to treat solid tumors is the poor penetration of the nanocarriers within the tumor mass. The deficient drainage together with the higher permeability triggers the presence of an interstitial fluid pressure (IFP) within the solid tumors that hinders the penetration of extravasated NPs and can even lead to their intravasation back into the bloodstream. This originates a remarkable decrease in the efficacy of the nanomedicines and may produce systemic toxicity [258]. In addition, the extreme conditions usually present within the core of a solid tumor (eg, low oxygen pressure and acidic pH values) can cause the appearance of multidrug resistant cancer cells that are refractive to cytotoxic drugs. Thus, an effective nanomedicine should be homogeneously distributed through the tumor regions, highlighting the tumor core to destroy the cells living at this place before they acquire multidrug resistance. This is of foremost relevance as tumorigenic cancer stem cells are commonly located in these deeper áreas [259]. To reach these regions, NPs should be able to displace throughout the tumor extracellular matrix (ECM), which is additionally hindered by the fact that IFP abolish convective transport. This matrix forms a dense network is composed by a complex mixture of different proteins such as collagen or glycoproteins, depending on the tumor type, which makes difficult NPs diffusion. Rich-collagen ECM shows higher retention of negatively charged NPs, while ECM rich in negative proteins such as sulfated glycosaminoglycans tends to retain positive NPs. Therefore the external surface of MSNs should be “decorated” with appropriate moieties depending on the ECM present in the tumor tissue to attain higher penetration within the solid tumor. Thus, it has been reported that the administration in the tumor zone of proteolytic enzymes, as collagenase, capable to digest the ECM prior to the NPs treatment improves the penetration of the nanocarriers within the diseased region and consequently their therapeutic efficacy [260]. These enzymes have also been covalently bonded on the nanocarrier surface [261]. Nevertheless, enzymes are highly sensitive and may suffer denaturalization or degradation under different conditions. Very recently, the research

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Fig. 14 Schematic depiction of the EPR effect.

Fig. 15 Operation mechanism of pH-sensitive polymeric collagenase nanocapsules grafted to MSNs. Blue arrows in the TEM micrograph indicate the collagenase nanocapsules attached to MSNs.

group headed by Prof. Vallet-Regí, reported a new strategy to develop pH-sensitive polymeric nanocapsules entrapping and protecting collagenase, which were attached on the MSNs’ surface with the aim of improving particle penetration in 3D tissues [262]. MSNs transport these collagenase nanocapsules attached on their surface; when the pH decreases to mild acidic conditions, which is typical for many solid tumors due to the raised glycolysis of cancer cells [263], the capsules break and release the collagenase, which degrades the collagen matrix and enhances the penetration of the nanocarrier (Fig. 15). The higher penetration capability of this nanodevice was successfully demonstrated using 3D-collagen gel with human osteosarcoma cells embedded. Indeed, the system is biocompatible, opening promising expectations for its future clinical application.

Cellular vehicles of MSNs

As discussed in the previous section, accomplishing a deeper penetration of NPs within the tumor mass is a major challenge. In this context, human mesenchymal stem cells (MSCs) with migratory capability towards tumors have been used as carriers of NPs [264,265]. These cells are able to accumulate in the tumor tissues and once there, they can cross the diseased tissue reaching deeper zones. A specific type of these MSCs derived from the decidua of human placenta, DMSCs, are a very homogeneous population,


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Fig. 16 Schematic depiction of DMSCs as carriers of MSNs towards tumor tissues (left); SEM, TEM and DLS of MSNs (right hand bottom corner); and DMSCs with fluorescein-labeled MSNs internalized into their cytoplasm (right hand top corner).

easy to acquire, and, more significantly, in addition to their migration capability towards tumors, they are able to inhibit the growth of primary tumors and the development of new ones [266]. Recently, these DMSCs have been used as carriers of MSNs towards tumor cells for cancer therapies [267]. First of all, the potential toxicity of MSNs over DMSCs was discarded; then, the effect of MSNs’s surface chemistry over cellular uptake was evaluated, evidencing that positively charged MSNs were better internalized. Just after internalization, MSNs were placed inside lysosomes, but after 3 days they escaped the lysosomes into the cytoplasm. The time MSNs were retained into the cells was 5 days or longer, which is time enough to allow cells migrating to tumors. The cell migration capacity was evaluated both in vitro and in vivo, finding that DMSCs kept their homing capacity towards tumors when transporting MSNs. In addition, DMSCs with MSNs loaded with the cytotoxic drug doxorubicin induced cell death in NMU breast cancer cells when co-cultured in vitro, which reveals these DMSCs as promising platforms for cancer therapy as carriers of MSNs loaded with antitumor agents (Fig. 16) [267].

Active targeting

The efficacy of MSNs can be improved by “decorating” their outermost surface with molecules that can be selectively recognized by the cancer cells. Solid tumors consist in a highly heterogeneous mixture of both malignant and healthy cells. Therefore, it is possible to increase the efficacy of the therapy by functionalizing the nanocarriers with moieties that promote the uptake only by diseased cells and not by normal ones. For antitumor aims, targeting agents are normally compounds that bind to specific cellular receptors overexpressed on the surface of the tumor cells. There are different entities that can be used as targeting agents, including antibodies, proteins, oligopeptides, aptamers, sugars or small molecules (Fig. 17). Active targeting is considered complementary to passive targeting to increase the efficiency of cancer nanotherapies. The targeting capability depends on both the abundance and function of the receptors overexpressed in the target and the ligand density in NPs. Different approaches have been developed to graft targeting agents to the MSNs’ surface, such as carbodiimidemediated COOH–NH2 coupling, maleimide–SH coupling, etc. Herein, we overview some reports concerning targeting agents conjugated to MSNs for specific recognition and selective internalization by cancer cells. Antibodies Antibodies (Ab) are proteins able to recognize specific antigens present on the surface of pathogens, foreign entities, or tumor cells. The principal advantage of using Ab is their high specificity to the target receptors. Ab have been conjugated to MSNs via electrostatic interactions [268] or covalent bonds [269,270]. The main risk associated to the latter approach is that the antibody can lose its capability to bind to the antigen if the functional groups that form the covalent bond are placed near the recognition site of the protein. To improve the dispersibility of the Ab-conjugated MSNs, the covalent bond can be carried out using PEG as a crosslinker [271].

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Fig. 17 Targeting agents conjugated to MSNs.

The Ab-MSNs ratio affects specific cell recognition. For instance, different amounts of the monoclonal Ab anti-Her2/neu were conjugated to through a PEG spacer [272]. This Ab recognizes HER2/neu glycoprotein, which is over-expressed in 25–30% of malignant human breast cancer. These targeted MSNs were able to be internalized into the breast cancer cells by receptor-mediated endocytosis presenting high selectivity in comparison with non-targeted MSNs. Chen et al. reported for the first time Ab-conjugated-MSNs for in vivo tumor targeting [271]. Thus, the surface of PEGylated MSNs was decorated with the monoclonal Ab TRC105. This Ab binds to CD105 receptors overexpressed in the tumor vasculature of many solid tumors. In this case specific targeting was solely vascular CD105 with no cancer cell targeting involved. Results revealed a B2-fold increase of tumor uptake compared to that of passive targeting alone based on the EPR effect. Peptides One of the major limitations of using Ab as targeting agents is their low stability under physiological conditions and their immunogenic nature. The latter can activate an immune response by the host provoking the rapid clearance of the administered MSNs or even severe side effects. A good alternative consists in using only the small peptide sequences, either linear or cyclic, that can recognize specific antigens located on the cell surface. They provide several advantages compared to Ab; peptides are cheaper and easier to produce than Ab and their use avoids the threat of an immune response by the host. Among peptides, the RGD (Arg– Gly–Asp) family, which binds to endothelial anb3 integrins, has been extensively explored as a targeting agent. The avb3 integrin is upregulated in both tumor and angiogenic endothelial cells. Different cyclic and linear RGD peptides containing free thiol as part of cysteine [273–275] or free primary amine as part of lysine [276–278] have been conjugated to MSNPs and their targeting capability towards diverse tumor cells overexpressing avb3 integrins has been demonstrated. The effect of RGD conformation on cellular uptake evidenced that cyclic RGD showed higher internalization into tumor cells than linear sequences [276]. Another interesting approach regards a cell-penetrating peptide derived from the human immunodeficiency virus (HIV) transactivator protein (TAT), which has been conjugated to MSNs to target the nuclei of tumor cells [279,280]. TAT peptide has proved its effectiveness in translocating MSNs into cell nuclei by binding importin a and b receptors and then targeting the nuclear pore complexes of cancer cells and entering the nuclei. TAT-conjugated MSNs have been used for direct intranuclear drug delivery of doxorubicin (DOX) in multidrug resistant (MCF-7/ADR) cancer cells [281]. TAT-MSNs were able to increase the intracellular, as well as intranuclear, drug concentrations much more efficiently than free DOX or drugs delivered from MSNs in the absence of TAT peptide. Proteins Certain proteins have also been conjugated to MSNs for tumor targeting. Cancer cells require certain nutrients, such as iron, to assist their rapid growing rate. Transferrin (Tf) is the protein that transports iron through the bloodstream. Hence, certain types of


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tumor cells overexpress transferrin receptors (TfR), which working together with Tf, aid iron uptake. In fact, TfR can be overexpressed up to 100-fold on cancer cells, making Tf an attractive targeting agent for antitumor therapies using MSNs. Binding Tf to MSNs helps in attaining selective internalization in pancreatic cancer cells (PANC-1) [273], pre-metastatic breast cancer cells (BT549) [273], human cervical adenocarcinoma cancer cells (HeLa) [282], and fibrosarcoma cancer cells (HT1080) [283]. Aptamers Aptamers are macromolecules able to recognize specific cellular receptors or other important biomolecules. Aptamers are based on single-stranded oligonucleotide chains that exhibit several advantages compared to Ab: higher resistance to degradation under physiological conditions; easier synthetic procedures; low immunogenicity; and small size, which permits conjugating more entities on the MSNs surface. Aptamer AS1411, which is the first molecule entering in clinical trials, has been grafted to MSNs [284]. This aptamer interacts with nucleolin, which is overexpressed by given tumor cells such as breast cancer cells (MCF-7). In vitro assays demonstrated that MCF-7 cells internalized higher amounts of aptamer-functionalized MSNs than non-functionalized ones. Thrombin-binding aptamer has been conjugated with lipid-coated MSNs and loaded with the potent antitumor drug docetaxel [285]. Since thrombin is involved in relevant tumor functions, such as proliferation and angiogenesis, there is a synergistic antitumor effect due to the significant increase in the cytotoxicity capability of docetaxel. Saccharides Saccharides, such as lactobionic acid [286], mannose [287], galactose [288], and hyaluronic acid (HA) [289–292], have been used as targeting agents as certain tumors overexpress carbohydrate receptors. Among polysaccharides, HA is receiving increasing scientific attention as targeting agent because many cancer cells overexpress HA surface receptors, such as CD44, CD168, HARE, etc. In a recent work, HA was conjugated on the surface of doxorubicin-loaded MSNs via disulfide bonds; HA acts both as redoxresponsive nanocap and as targeting agent [292]. In vitro assays indicated that these nanosystems exhibited higher cellular uptake efficacy through CD4-mediated endocytosis by HeLa cancer cells (overexpressing CD44) than by LO2 hepatic cells (nonoverexpressing CD44). Small molecules Small molecules with different structures and characteristics are attractive targeting agents that provide several advantages compared to other ligands; they are normally more stable, less expensive and exhibit lower immunogenicity than macromolecules; due to their small size they can be conjugated on the MSNs surface in higher amounts that big molecules such as Ab or proteins. A widely used targeting molecule is folic acid (FA) [293]. FA is a vitamin employed by diverse enzymes during DNA replication and therefore, many cancer cells overexpress folate receptors (FR) on their surface to capture as much as possible this essential compound. The efficacy of FA conjugated to MSNs to reach tumor using in vivo murine models [294]. A small dose (0.5 mg per mouse) of MSNs loaded with a potent antitumor drug (camptothecin) and decorated with FA was sufficient to provoke strong tumor suppression. These NPs were excreted through urine and feces within a week of injection, confirming the biocompatibility of the treatment. Rosenholm et al. grafted FA onto poly(ethyleneimine) (PEI)-coated MSNs showing a significant improvement in cellular uptake by tumor cells overexpressing FR [295]. FA been also grafted to MSNs via disulfide linkers aimed at playing a dual role, as redox-responsive capping agent and targeting agent for efficient tumor targeting [296]. An innovative research consisted of conjugating methotrexate, an antitumor drug structurally similar to FA, to MSNs acting both as targeting and chemotherapeutic agent [297]. The nanosystems induced death by apoptosis in HeLa cancer cells (overexpressing FR), while no apoptosis was detected in healthy cells. Very recently, Villaverde et al. developed an innovative nanosystem for targeting a particular type of cancer, neuroblastoma (NB). NP is the most frequent extracraneal cancer in pediatric oncology and is exhibit a bad prognosis. More than 90% of NB cells overexpress a membrane protein that re-uptake norepinephrine, the norepinephrine transporter (NET) [298]. Certain synthetic molecules, such as meta-iodobencylguanidine (MIBG) interact with NET in a similar fashion to natural epinephrine. They found that the grafting of a specific MIBG derivative, meta-aminobencylguanidine (MABG) via a PEG crosslinker with a molecular of 2000 Da (PEG2000) increases more than 4-fold nanoparticle uptake by NB cells, enhancing the cytotoxic effect of the nanocarrier against NB [299]. To evaluate the efficacy in vivo of the targeted nanosystem the authors used a xenograft murine model in which the NB cells express luciferase to permit an easy visualization of the tumor mass by luminescence measurements. MSNs were labeled with a near infrared fluorophore to allow tissue localization of the nanosystems within the living organism. Targeted MSNs (group IV) exhibited much higher accumulation within the tumor mass at 48 and 72 h after the systemic administration compared with MSNs only functionalized with PEG2000 and the bare NPs (Fig. 18). On the contrary, PEGylated particles (group III) were essentially accumulated in the liver, probably due to the longer circulation time of these NPs in contrast with the bare ones (group II), which were totally removed from the host. In addition targeted MSNs remained in the tumor after 72 h, while the other ones were quickly cleared from the diseased area.

Stimuli Responsive MSNs

Currently, much scientific effort is being devoted to develop novel stimuli-responsive nanosystems able to release antitumor drugs once in the target tissue. MSNs are excellent candidates due to their unique characteristics that permit them to efficiently entrap, protect and transport therapeutic cargoes to tumor tissues [31,32,247,300]. As it is schematically illustrated in Fig. 19, the drug is

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Fig. 18 In vivo evaluation of the targeting efficacy of MSNs decorated with the MABG derivative through PEG2000.

Fig. 19 Schematic depiction of the performance of stimuli-responsive MSNs.

loaded within the mesopores of MSNs and its output is hindered by using organic or inorganic moieties that act as pore blockers or gatekekeepers that avoid premature cargo release. The stimuli-responsive performance can be accomplished grafting these moieties via cleavable bonds or using gatekeepers that undergo any chemical or physical change in response to given stimuli. Upon exposition to a certain stimulus, which can be internal or external, the pore outlets are open and the release of the entrapped payload is triggered at the target site. Albeit smart MSNs that respond to internal stimuli exhibit the advantage of being less invasive than those that respond to external stimuli, since they do not need external apparatus to trigger drug release, the control on the administered dosage is lower.


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In the next sections we overview the new insights in the development of MSNs as stimuli-responsive drug delivery system attending to the nature of the stimulus used as release trigger.

Internal stimuli-responsive MSNs pH as release trigger The pH is one of the most exploited internal stimuli to trigger drug release from MSNs because different pathologies such as cancer or inflammation present pH changes during their evolution [263,301]. Besides, once internalized inside the cell, MSNs can be exposed to different pH values depending on the cell compartment or organelle [302]. Therefore pH-sensitive gatekeepers can be employed as pore blockers to control drug release in response to these pH variations. Organic molecules such as cucurbit[6]uril [303] and cyclodextrins [304] form inclusion complexes with diverse moieties previously attached to the external surface of MSNs, thus acting as pore caps. These complexes can be dissociated at gien pH values producing the pore opening and triggering cargo release. Inorganic NPs composed of gold [305], iron oxide [306] or cerium oxide [307] have been also tethered to MSNs throughout pH-cleavable linkers (boronate, hydrazine, acetals, etc). When the pH decreases to mil acid value, as that found in endosomes (pH 5–6), the linker breaks triggering the drug release. Another strategy consist on the grafting of polymer chains, such as poly (4-vinylpyridine) [308], poly(2-(diethylamino)ethyl methacrylate [309], chitosan [310], starch [311] or poly(styrene sulfonate) [312], which contains functional groups with acid/base properties. Therefore, the polymer shell can experience conformational changes depending on the external pH. When the polymer is in neutral state, the polymer chains tend to interact with themselves adopting a collapsed hydrophobic conformation that inhibits drug release. In the ionized state, the polymer chains acquire hydrophilic character and mainly interact with water molecules adopting the extended conformation and allowing the entrapped cargo to go out of the mesopores (Fig. 20). Very recently Martínez-Carmona et al. developed an innovative nanosystem for selective topotecan delivery to cancer cells by targeted pH-sensitive MSNs [313]. This drug is a potent antitumor agent whose major limitation is its inactivation at the physiological pH of 7.4. To solve this drawback, topotecan was loaded into MSNs and then coated with a multifunctional gelatin shell that prevents drug hydrolysis and premature release; acts as a pH sensitive layer; and offer multiple grafting points for the anchorage of targeting agents, such as FA, for selective internalization in tumor cells (Fig. 21). In vitro assays revealed that cancer cells overexpressing membrane cell surface makers with affinity towards FA, internalized higher number of NPs than normal cells, which do not overexpress such markers. These MSNs are capable of killing tumor cells, but they do not decrease the viability of healthy cells. The opposite behavior has also been described in MSNs functionalized with polyamines, leading to pore opening when the amino groups of the polymer shell are in neutral state [314,315]. Another approach consists in controlling the drug release throughout intermolecular interactions between the drug molecules and the polymer branches. For instance doxorubicin is a potent antitumor drug bearing an amine group that is protonated in physiological conditions. Thus, polyacrylic acid (PAA) grafted on MSNs can retain this drug via electrostatic interactions between positively charged amino group from the drug and negatively charged carboxylic groups from the polymer [316]. In addition, PAA can be conjugated on the MSNs surface via an acid-sensitive bond with the aim of releasing cytotoxic drug confined inside the mesopores [317]. Lastly, the antitumor drug can be directly grafted on the MSNs surface using alike acid-sensitive linkers [318,319]. Redox potential as release trigger The imbalance between reductive species such as glutathione (GSH) outside and inside the cells [320], and also between normal and tumor tissues [321], has been widely exploited as release trigger. The intracellular compartments (cytosol, mitochondria and nucleus) contain 1000 times more reduced GSH than the extracellular media. Since this compound breaks disulfide (S–S) bonds, a pore blocker tethered on the MSNs surface through this bond will be detached when the nanocarrier enters the intracellular region. Hence, inorganic nanocaps such as CdS [322], Fe2O3 [323] or Au [324] NPs, have been grafted on the MSNs surface using GSHsensitive linkers that contains disulfide bonds. Once the nanosystem reaches the intracellular space, the presence of increased concentrations of reductive species triggers the rupture of the disulfide bonds opening the pores (Fig. 22).

Fig. 20 Schematic illustration of the performance of pH-responsive polymer coated-MSNs.

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Fig. 21 Schematic illustration of the tumor-targeted topotecan (TOP) nanosystem consisting of MSNs entrapping TOP, coated with a gelatin shell and decorated with folic acid (FA) as targeting agent. Folate surface markers (FSM) are surface markers with affinity towards FA, which are usually overexpressed in the membrane of certain cancer cells.

Fig. 22 Schematic depiction of the performance of redox-responsive inorganic-caps on MSNs.

Organic molecules grafted via disulfide bonds on the mesopore openings can also act as redox-responsive gatekeepers. Thus, cyclodextrins [325,326] and poly(propylenimine) dendrimer of generations I and II [327] grafted to the MSNs surface through reducible bonds are able to block premature cargo release until they enter the cytosolic space. An interesting strategy consisted in grafting FA molecules via redox-sensitive linkers on the mesopore channels [296]. These molecules interact with themselves via hydrogen bond, hindering drug diffusion until the break of the disulfide bonds takes place. This approach provides the advantage that the same molecule acts both as capping and targeting agent avoiding the use of redundant systems. Redox-sensitive polymer shells cross-linked with disulfide bonds have been also reported [328]. For instance the development of MSNs en-capped with collagen grafted by disulfide bonds have been reported, aimed at achieving cell-specific targeting and redox responsive drug delivery [286]. In another research work, carbonic anhydrase as enzyme model was covalently immobilized within the mesopore channels through a specific crosslinker that contains disulfide bonds [329]. GSH addition triggered release and the enzyme preserves more than 80% of its activity once release. Finally, cytotoxic compounds can be directly conjugated to the MSNs surface using reducible bonds [330]. It is also possible to transport less-toxic species such as inactive oxidized Pt(IV) complexes by MSNs using a similar approach, showing the relevant advantage that they will be activated only in the presence of the reductive environment within the cell [331].


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The main restrictions of all these approaches rely on the fact that the MSNs cellular uptake occurs usually by endocitosis [332] and endosomes show a less-reductive environment, even being oxidative in certain circumstances [333]. To solve this drawback Sauer et al. have reported the use of photosensitizers to induce a rapid light-induced endosomal escape [334]. Functionalization with polycationic polymers such as polyethylenimine (PEI) favors endosomal escape by proton sponge effect and it has been widely employed for DNA or RNA delivery [335]. MSNs have also been decorated with RGD moieties, observing that these NPs are retained during longer times inside the endosomes than bare NPs. This is probably due to the interaction between the RGD and the integrin receptors present in the endosomes [276]. Thus, this phenomenon evidences that it is essential to study the endosomal escape of targeted MSNs loaded with sensitive molecules. Enzymes as release triggers Some diseases states involves significant overexpression of given enzymes; fact that can be used for triggering the drug release from NPs [336]. In this sense, matrix metalloproteinases (MMPs) exhibiting protease activity, are upregulated in certain tumors and they are involved in tumor progression and metástasis [337]. Thus, novel MSNs able to release cytotoxic drugs in response to increased levels of MMPs have been developed [338]. These nanosystems consist of MSNs coated with a polymeric shell that contains protease-sensitive sequences as crosslinkers. Incorporating variations in the oligopeptide motif it is possible to tune the degradability of the shell. Thus, the presence of certain MMPs degrades the polymeric shell through the hydrolysis of these sensitive sequences triggering the pore opening. Therefore, drug release kinetics from MSNs can be modulated depending on the tumor and its protease expression profile. Cyclodextrins [339] and rotaxanes [340] have also been grafted on the mesopore entrances through designed enzyme-sensitive linkers that permit to trigger pore opening only in the presence of specific enzymes. Schlossbauer et al. reported the use of biotin–avidin assemblies as protease-sensitive blocking caps [341]. On the other hand, biopolymers such as polysaccharides [342], polypeptides [343] or HA [290] and other polymeric chains such as polyesters [344], located on MSNs surface, can impede the drug diffusion until certain hydrolytic enzymes able to decompose these macromolecules are present. Recently, Mas et al. reported developed a new strategy to design 3D smart systems consisting of macroporous scaffolds incorporating enzyme-responsive gated MSNs [345]. The nanogates were able to open in the presence of acid phosphatase (APase), an enzyme whose concentration is employed to measure osteoclast activity in bone remodeling processes and as a biochemical marker for the presence of bone metastases. The APase-triggered drug release mechanism was preserved in the 3D scaffolds, but slower release kinetics were found, probably due to the different enzyme accessibility to the gated MSNs embedded in the scaffold compared to free MSNs.

External stimuli-responsive MSNs Light as release trigger Light was one of the first stimuli used to trigger drug release in MSNs. Mal et al. linked UV-light sensitive coumarin derivatives on the mesopore walls to control the drug release process [180]. The same research group improved the system by placing azobenzene molecules inside the pore channels. These molecules undergo “cis–trans” isomerization under UV irradiation and therefore they can act as nanoimpellers propelling the drugs out of the pores [346]. The fast isomerization of azobenzene has also been exploited to design stalks able to reversibly bind cyclodextrins, which can act as light responsive pore blockers [347]. Light is an external stimulus that exhibit several advantages, such as low toxicity, easy application and fine focalization in the target site. Nonetheless, the main drawback relies on its poor tissue penetration. Regardless of this fact, many MSNs have been decorated with lightsensitive moieties on their outer surface to control drug release by the exposure of light with different wavelengths, from UV to near infrared radiation [348,349]. Recently, Vallet-Regí’s research group has reported the use of a responsible biocompatible protein shell capable of acting as a light-breakable gatekeeper that avoids premature cargo release [283]. In addition, the proteins tethered on the MSNs’ surface are able to act as targeting agents improving the cellular uptake by cancer cells. This nanosystem consists of MSNs containing biotin molecules grafted on their surface via a light-sensitive cross-linker that undergoes cleavage upon irradiation by UV light at 366 nm (Fig. 23). Once the MSNs are loaded with different cytotoxic agents (doxorubicin or oxalilplatin), a solution of biotynilated transferrin (Tf) and streptavidin closes the system due to the self-recognition capability of the biotin–streptavidin bridges. In this system, Tf acts both as sensitive cap of the housed drug and also as targeting agent, enhancing the particle uptake by tumor cells up to ten times more than bare NPs. After the nanosystems were internalized by tumor cells, irradiation with UV light for short time periods (less than 5 min), triggers the cleavage of the protein caps and then drug release. It is remarkable that these smart nanosystems achieved great cytotoxicity since 10% of the cancer cell population entered the apoptotic state using doses as low as 0.01 mg mL1. Magnetic field as release trigger The use of a magnetic field to trigger drug release from nanosystems provides remarkable advantages owing to its low toxicity and exceptional penetration in living tissues. To take advantage of this stimulus, a carrier sensitive to the presence of magnetic fields is needed. To this aim, the most widely used approach consist in the encapsulation of superparamagnetic iron oxide nanocrystals (SPIONs) 5–10 nm in diameter within a silica matrix. The exposure of SPIONs to alternating magnetic fields (AMF) produces a noticeable temperature increase in their surroundings principally due to two mechanisms: (i) Brownian relaxation, which is characterized by the rapid rotation of the particles to align with the magnetic field; and (ii) Neel relaxation, provoked by the spontaneous reorientation of the magnetic moment within the particle [350]. The effective encapsulation of SPIONs within MSNs

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Fig. 23 Light-responsive protein shell in MSNs acting acts both as a targeting agent and as capping system.

can be accomplished using diverse strategies, such as aerosol methods [351], or the sol–gel process [352], among others. Thus, the presence of SPIONs inside MSNs permits the use of magnetic fields as trigger stimuli by locating thermosensitive moieties on the pore openings. These temperature-responsive gatekeepers undergo physical–chemical modifications that produce pore opening in response to temperature rises. An interesting approach consists in using DNA as gatekeeper [353]. When two complementary single DNA strands get in contact, they form a stable double helix structure due to the attractive intermolecular interactions between the base pairs adenine–thymine and guanine–cytosine, in a process called hybridization. Importantly, this process is reversible and the two DNA strand separate at given temperatures owing to the breaking of these intermolecular forces, which trigger the dehybridization of the double helix. The temperature at which this process occurs is known as melting temperature. It can be modulated by modifications in the chain length, guanine–cytosine content, or controlling the density of olinucleotides anchored on the NPs surface [354]. Following this strategy, Ruiz-Hernández et al. developed a magnetic-responsive nanosystem based on MSNs containing SPIONs trapped within the silica matrix, giving rise to magnetic MSNs (MMSNs). The external surface of MMSNs were grafted with 15 base pair DNA single strands designed to present a melting temperature of 471C, which is positioned in the upper limit of the hyperthermia range [353]. This nanodevice was capped with iron oxide nanoparticles functionalized with the complementary DNA single strands through the hybridization between the two oligonucleotide chains (Fig. 24). MSNs can be loaded with drugs before sealing the pores, displaying a zero-alternating release behavior at room temperature once the system is closed. The application of a magnetic field of 24 kA m1 and 100 kHz induces a significant temperature increase in the NPs surroundings, reaching the melting temperature. This produces the dehybridization of the double helix and triggers cargo release. An interesting advantage of this system is its reversibility, because when the magnetic field is ceased, the DNA strands are self-assembled again producing the pore closure. Therefore, this nanosystem exhibits an on–off performance able to release the transported cytotoxic agents in a more precise and controllable fashion. Another strategy consists in using temperature-responsive polymer-based gatekeepers. In fact, thermosensitive polymers have been widely employed in nanomedicine both for the synthesis of temperature-responsive organic nanoparticles [355], and for the development of temperature-responsive gatekeepers placed on the surface of inorganic porous materials [356]. Poly(N-isopropylacrylamide) (PNIPAM) is one of the most employed temperature-responsive polymer. When this polymer is dissolved in water, it exhibits a linear state (coil) at temperature below 321C, while it undergoes shrinkage (globular) if the temperature goes above this value. The temperature at which this transition occurs is known as the lower critical solution temperature (LCST) and it can be adjusted modifying the composition of the polymer chain. Therefore, the co-polymerization of NIPAM with hydrophilic monomers such as acrylamide, acrylic acid or N-hydroxymethyl acrylamide (NHMA), among other, produces an increase in their LCST due to the more favored polymer–water interactions provoked by the presence of these hydrophilic monomers [357]. Thus, by means of variations in the polymer compositions it is feasible to obtain PNIPAM co-polymer in which the LCST is set around the hyperthermia range (42–451C) [358]. Baeza et al. synthetized an


Ordered Mesoporous Silica Materials

Fig. 24 DNA-gated nanodevices (top) and thermosensitive-co-polymer coating (blue shell corresponds to PNIPAM and red chains correspond to PEI) (bottom).

engineered block co-polymer based on PNIPAM and PEI, which was then grafted on the surface of MMSNs, providing two attractive features: acting as temperature-responsive gatekeeper capable of controlling the release of drug entrapped within the mesopores due to the presence of the PNIPAM backbone; and retaining proteins or enzymes within the polymer shell by electrostatic and hydrogen bond interactions between the macromolecules and PEI chains (Fig. 24) [359]. The presence of PEI increases the hydrophilic nature of the polymer chain rising the transition temperature up to ca. 401C. This nanodevice evidenced its ability to transport two species of different nature: fluorescein, as a small model molecule, and soybean trypsin inhibitor, similar to growth factors, as a model protein. The polymer shell collapses when the system is exposed to an AMF producing the release of the entrapped payload. Recently, Guisasola et al. have attached to the MMSNs surface a random thermosensitive PNIPAM copolymer doped with 10% NHMA in order to obtain a precise drug release triggered by magnetic fields when the temperature reached 431C [360]. Indeed, this device was capable of releasing the housed drugs upon exposure to AMF even when maintaining the local temperature at physiological conditions (371C) placing the sample into a refrigerated space. This performance demonstrates that the presence of SPIONs within the silica matrix under an AMF produces enough heat in the particle surroundings to promote polymer collapse without needing to achieve a global temperature rise. This fact is of vital relevance for future clinical application of magnetically triggered nanodevices. Ultrasounds as release trigger The use of ultrasounds (US) to trigger drug release permits spatiotemporal control of the drug release at target site with very attractive characteristics, such as non-invasiveness, absence of ionizing radiation, cost-effectiveness and easy control of tissue penetration depth depending on frequency, cycles and exposure time. What makes US very attractive in the field of nanomedicine is that high-frequency US can penetrate deep into the body with focused beams, allowing local therapies and avoiding any damage to healthy tissues [361]. US waves propagate through living bodies producing diverse physical effects that can be employed as trigger for US-responsive drug release. Thus, US inducing thermal and/or mechanical effects have been used to trigger drug release from different nanocarriers, such as liposomes and polymeric micelles. Mechanophores, chemical bonds which can be cleaved upon exposure to US radiation [362], can be used in the design of US-responsive NPs. This is the case of 2-tetrahydropyranyl methacrylate (THPMA), a hydrophobic monomer with an acetal group sensitive to US, than can be simply be converted into hydrophilic methacrylic acid (MAA) [362,363]. Paris et al. took advantage of the hydrophobic to hydrophilic phase transformation under US irradiation to develop responsive MSNs using such moieties as gatekeepers of the mesopores [364]. In addition, those monomers were combined with thermoresponsive monomers, such as (2-(2-methoxyethoxy) ethyl methacrylate, (MEO2MA), to produce dual-responsive co-polymers, ie, copolymers sensitive to US and heat. Anchoring that copolymer to the surface of MSNs permitted to load the drug into the synthesized nanocarriers at low temperature (41C), as the thermosensitive polymer presented an open conformation under these conditions. This was one of the major advantages of this system, since conventional MSNs load

Ordered Mesoporous Silica Materials


Fig. 25 Release profiles of fluorescein from MSNs functionalized with an US sensitive copolymer in PBS solution over time after 10 min of US irradiation (top) and without US (bottom).

their cargo at the same time than the functionalization process, leading to low efficiency upon loading and functionalization. Afterwards, the increase of the temperature to 371C closed the pore openings as the thermosensitive polymer collapsed at such temperature. This fact, allowed the MSNs to transport the drugs without any premature release, which is of enormous relevance when carrying cytotoxic drugs to tumors. The exposure of the nanodevices to US radiation triggered drug release because the USsensitive polymer changed from hydrophobic to hydrophilic. This phenomenon produced a globular-to-coil transition in the conformation of the co-polymer, which triggered the opening of the pore entrances and the release of the entrapped payload (Fig. 25). This innovative nanosystem has been tested with different model molecules, namely fluorescein and the fluorescent ruthenium complex [Ru(bipy)3]2 þ , revealing that nanocarriers obey efficiently to the stimulus independently of the nature of the cargo vehiculized [364]. Besides, these nanodevices were observed to be non-cytotoxic and after internalization into LNCaP prostate cancer cells, they kept their US responsive performance in the cytoplasm of those cells. Thus, after loading those MSNs with doxorubicin and incubating with LNCaP cells, they induced cell death only after exposure to US. This fact evidences that the cytotoxic was only released after US stimulation, avoiding any type of premature release (no cell death) before the application of the external stimulus.

In Situ Cytotoxic Drug Generation

As previously mentioned, the transport of highly cytotoxic drugs, which is the usual aim of nanocarriers for cancer therapy, requires the design of zero-release nanodevices. It is imperative that the cytotoxic agents are retained until the target tissue or cell is reached. This is the case of the stimuli-responsive nanosystems mentioned in the previous sections. Nonetheless, sometimes the fabrication of the smart nanodevices incorporates high complexity that impedes their clinical application because their approval by regulatory agencies demands an extensive evaluation of all their components [365]. An interesting alternative to tackle this concern relies on developing systems capable of carrying the cytotoxic species in an inactivate state and, once accumulates within the tumor area, activate them producing the toxic compounds only in the diseased tissue. This strategy is used by the enzymeprodrug approach that is normally based on the administration of innocuous prodrugs that are activated by definite enzymes overexpressed by the tumor tissue [366]. Regrettably, this strategy is restricted in many occasions by the low concentration of activating enzymes in the target tissue. What’s more, many prodrugs lack for a natural activating enzyme, which makes necessary to co-administer it or to introduce the gene that encodes it in the target cell [367]. Thus, Baeza et al. have developed a nanosystem able to generate highly cytotoxic species within tumor cells transporting a non-toxic prodrug entrapped within the mesopores and the enzyme needed for its activation covalently linked on the MSNs’ surface [368]. These nanosystems consist of MSNs functionalized with amino groups on the pore walls to retain the prodrug, in this case indol-3-acetic acid (IAA), by electrostatic interactions. Horseradish peroxidase (HRP) converts IAA into indole-3-carbinol generating cytotoxic species in the process, mainly hydroxyl and reactive oxygen species (ROS), which are able to destroy cancer cells by membrane and DNA damage [369]. Enzymes are very labile, which present low stability in living tissues due to the presence of proteases, among other aggressive compounds.


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Fig. 26 In situ cytotoxic compound generation by prodrug-loaded MSNs decorated with horseradish peroxidase (HRP) polymeric nanocapsules.

For this reason, in this device HRP was previously coated on a polymeric capsule designed to enable its anchorage to the MSNs’ surface and protecting the enzyme simultaneously (Fig. 26). Once the nanocarrier is internalized within the tumor cells IAA is slowly released and activated by the HRP, generating enough cytotoxic compounds and causing cell destruction.

Acknowledgments The authors deny any conflicts of interest. This study was supported by research grants from the Ministerio de Economía y Competitividad, Spain (project MAT2015-64831-R) and the European Research Council (Advanced Grant VERDI; ERC-2015-AdG Proposal no. 694160).

See also: 1.9 Bioactive Ceramics. 4.1 The Concept of Biocompatibility. 4.2 Biocompatibility and the Relationship to Standards: Meaning and Scope of Biomaterials Testing. 7.17 Bone Tissue Grafting and Tissue Engineering Concepts

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