Silicon oxide based materials for controlled release in orthopedic procedures.

ADR-12800; No of Pages 20 Advanced Drug Delivery Reviews xxx (2015) xxx–xxx

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Silicon oxide based materials for controlled release in orthopedic procedures☆ Haibo Qu, Sanjib Bhattacharyya, Paul Ducheyne ⁎ Center for Bioactive Materials and Tissue Engineering, Department of Bioengineering, University of Pennsylvania, Philadelphia, PA 19104, USA

a r t i c l e

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Article history: Received 27 February 2015 Received in revised form 21 May 2015 Accepted 25 May 2015 Available online xxxx Keywords: Silica Controlled release Bone

a b s t r a c t By virtue of excellent tissue responses in bone tissue, silicon oxide (silica) based materials have been used for bone tissue engineering. Creating nanoscale porosity within silica based materials expands their applications into the realm of controlled release area. This additional benefit of silica based materials widens their application in the orthopedic fields in a major way. This review discusses the various chemical and physical forms of silica based controlled release materials, the release mechanisms, the applications in orthopedic procedures and their overall biocompatibility. © 2015 Elsevier B.V. All rights reserved.

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Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Classes of sol–gel silica materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1. Sol–gel processed silica (xerogel) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1.1. Hydrolysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1.2. Condensation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2. Mesoporous silica nanoparticles (MSN) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3. Bioactive glass . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4. Mesoporous bioactive glass (MBG) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Controlled release mechanisms and mathematical modeling of release . . . . . . . . . . . . . . . . . . . . 3.1. Controlled release mechanisms . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1.1. Overview . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1.2. Diffusion controlled release . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1.3. Degradation controlled release . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1.4. Mixed release kinetics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2. Effect of synthesis parameters and surface functionalization on controlled release . . . . . . . . . . . 3.2.1. Surface functionalization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2.2. Pore blocking . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2.3. Pore parameter . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Physical application forms . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1. Particles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.2. Micron thin films . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.3. Scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Clinical applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.1. Infection treatment and control . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.1.1. Vancomycin-containing silica sol–gels: synthesis and release kinetics . . . . . . . . . . . . 5.1.2. Sol–gel silica thin films for controlled release of antibiotics . . . . . . . . . . . . . . . . . 5.1.3. Sol–gel/vancomycin micron-thin sol–gel coatings for the treatment of S. aureus infection in vivo

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☆ This review is part of the Advanced Drug Delivery Reviews theme issue on “Drug delivery to bony tissue”. ⁎ Corresponding author. E-mail address: [email protected] (P. Ducheyne).

http://dx.doi.org/10.1016/j.addr.2015.05.015 0169-409X/© 2015 Elsevier B.V. All rights reserved.

Please cite this article as: H. Qu, et al., Silicon oxide based materials for controlled release in orthopedic procedures, Adv. Drug Deliv. Rev. (2015), http://dx.doi.org/10.1016/j.addr.2015.05.015

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H. Qu et al. / Advanced Drug Delivery Reviews xxx (2015) xxx–xxx

5.1.4. Silica film for nitric oxide release . . . . . . . . . . . . . . . . . . . . . . . . . . 5.1.5. Sol–gel silica thin film for the treatment of methicillin resistant S. aureus (MRSA) infection 5.1.6. Bioactive glass for infection treatment . . . . . . . . . . . . . . . . . . . . . . . . 5.1.7. Mesoporous silica particles for infection treatment . . . . . . . . . . . . . . . . . . 5.2. Pain treatment and control . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.2.1. Sol–gel silica controlled release of bupivacaine . . . . . . . . . . . . . . . . . . . . 5.2.2. Sol–gel/polymer composite controlled delivery of local anesthetics . . . . . . . . . . . 5.2.3. Sol–gel silica controlled release of Non-steroidal anti-inflammatory drugs . . . . . . . . 5.3. Tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.3.1. Porous, bioactive ceramic scaffolds for bone tissue formation in vivo . . . . . . . . . . 5.3.2. Bioresorbable sol–gel silica-based large biomolecules delivery systems . . . . . . . . . 5.3.3. Mesoporous bioglass (MBG) scaffolds . . . . . . . . . . . . . . . . . . . . . . . . 6. Biocompatibility . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.1. Biocompatibility of silica granules . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.2. Biocompatibility of silica sol–gel nanoparticles . . . . . . . . . . . . . . . . . . . . . . . . 7. Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1. Introduction Silicon oxides in glassy form have been used by human society for thousands of years. In the past several decades, silicon oxide based materials have been studied extensively for biomedical applications. In the early sixties, silicon oxide glasses, (also called silica glass) were mainly synthesized by melt-quenching: oxides are molten at temperatures in excess of 1300 °C and then quenched at room temperature. A major conceptual advance for the biomedical use of silica resulted from the development of sol–gel processing. Since the mid-1980s, when Avnir showed that sol–gel processing can be used to entrap dyes, the sol–gel process was first considered for functionalization of glass with organic molecules [1]. Shortly after, entrapment or immobilization of molecules was pursued. Since the nineties, enzymes have been entrapped in sol–gel processed materials with high retention of activity [2–4]. When entrapped in the sol–gel material, the enzymes were intended to function as biocatalyzers within the sol–gel matrix. The entrapment of living organisms such as cells or bacteria was later explored using the same rationale [5–8]. However, the use of sol–gel for the controlled release of incorporated, but not entrapped molecules, did not emerge until the mid to late nineties [9,10]. Since then, studies on the incorporation of therapeutic molecules and ensuing release from silica materials have increased in a major way, driven by obvious benefits arising from controlled, prolonged release of therapeutic agents. Studies on sol–gel silica showed great promise for the clinical use of these biocompatible and biodegradable controlled release materials. This review summarizes many of these studies. Over the past two decades many different forms of silica materials have been developed. Before describing recent progress and possible future developments of silica controlled release materials, this review will first provide a description of the synthesis procedures of different forms of sol–gel silica materials. This will be followed by the discussion of functionalization and synthesis parameters to achieve controlled release of incorporated therapeutics; various applications; current state of its clinical application in the field of orthopedic procedures; and the overall considerations of biocompatibility.

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In a typical sol–gel process, silica precursors (organosilanes) such as tetramethoxysilane (TMOS, Si(OCH3)4) or tetraethoxysilane (TEOS, Si(OC2H5)4) are subjected to a hydrolysis–condensation reaction to form a colloidal suspension of silica particles, subsequent to which the particles condense into a new gel phase. The hydrolysis reaction, which can be either acid or base catalyzed, replaces alkoxide groups of organosilane with hydroxyl groups. Siloxane bonds (Si–O–Si) are formed during the subsequent condensation reaction. 2.1.1. Hydrolysis The hydrolysis occurs by the nucleophilic attack of the oxygen contained in water on the silicon atom of the organosilane. The hydrolysis is most rapid and complete when catalysts are employed. Under acidic conditions, the alkoxide group (OR) is protonated, which makes the silicon more electrophilic and susceptible to attack by water. All strong acids behave in a similar way, as the hydrolysis is a first order reaction: weak acids require a longer time for the hydrolyzation. In an acid catalyzed reaction, the hydrolysis rate is increased by substituents that reduce steric hindrance around silicon (i.e. hydrolysis of TMOS is much faster than TEOS). Under basic conditions, it is likely that water dissociates to produce nucleophilic hydroxyl anions in a rapid first step. The hydroxyl anion then attacks the silicon atom and liberates the alcohol. Similar to acid catalyzed hydrolysis, the steric factor also affects the hydrolysis rate and the base-catalyzed hydrolysis reaction also depends on catalyst concentration. No matter what condition, the overall hydrolysis reaction can be expressed as: Si(OR)4 + 4H2O → Si(OH)4 + 4ROH

2.1.2 . Condensation Condensation is the polymerization process that creates siloxane bonds. The condensation reaction is as below:

2. Classes of sol–gel silica materials

Si(OH)4 → SiO2 + 2H2O

2.1. Sol–gel processed silica (xerogel)

The condensation rate is highly affected by the pH condition. Maximum condensation rate is achieved around pH 7.0 where both protonated and deprotonated silanols are highest in concentration. The minimum condensation rate is found around the isoelectric point (pH ~ 2.0). The overall sol–gel process is as follows:

The sol–gel process has been extensively studied during the 1980s to 1990s. In this section we briefly introduce the sol–gel process for silica intended for controlled release applications. Readers interested in the details of sol–gel processing can consult “Sol–Gel Science” by Brinker and Scherer, which describes the physical and chemical principles of sol–gel processing extensively [11].

Si(OR)4 + 2H2O → SiO2 + 4ROH



After condensation, alcohol and water evaporate during drying and leave behind a porous structure. The drying is typically performed as a sintering step around 600 °C. The pore sizes can range from angstroms to tens of nanometers. Eliminating sintering would create the possibility of entrapping or immobilizing therapeutic molecules into sol–gel silica, all while their functionality is maintained. However, all processing steps need to be modified in order to obtain a room-temperature processed solid with appropriate release properties. In general, drugs are added to the liquid sol after hydrolysis. During condensation, the therapeutic molecules are homogenously incorporated within the silica xerogel (a xerogel is a sol–gel processed material not or not yet subjected to sintering). Since the mid 1990s, the use of sol–gel silica for the controlled release of incorporated molecules emerged [9,10,12–15]. By virtue of simple processing conditions that eliminate the sintering step, various therapeutic agents can be homogenously incorporated in nanoporous sol–gel processed silica [10,12,14,16–18]. Sol–gel processing is a glass processing methodology that enables to preserve biological functionality of molecules that are incorporated. In principle, this methodology can be used for a variety of compositions, and the network is not necessarily a silicon oxide network.

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leads to the formation of nonporous silica nanoparticles. This method is intended for the preparation of solid silica nanoparticles. To create a mesoporous structure in the silica particles, the Stöber procedure is modified by introducing surfactants into the mixture of silica precursors, ethanol and water. Above a critical micelle concentration, the surfactants self-assemble into an organized structure, and the shape and size of the organized structure depend on factors that include chemical and physical properties of the surfactant, surfactant concentration, pH and temperature [32]. Silica condenses on the surface of the surfactant's micelle through strong electrostatic interactions. The porous structure remains after removing the surfactant. Mesoporous silica nanoparticles (MSNs) with pore sizes in the range from 2 to 20 nm are then obtained. Fig. 1 schematically shows the difference between sol–gel processed silica and the mesoporous silica. The drug is incorporated into porous tunnels when the MSN is immersed in a highly concentrated drug solution. The unique pore features, such as open pores, adjustable pore size and narrow pore size distribution, render MSNs excellent materials for controlled release [33]. Different drug release kinetics can be achieved by either adjusting the structural properties or chemically modifying the mesoporous silica with organic moieties [34–36].

2.2. Mesoporous silica nanoparticles (MSN) 2.3. Bioactive glass In 1992, a family of highly ordered mesoporous silica materials (pore size in the range 2–10 nm) was synthesized by a hydrothermal sol–gel synthesis process. Long chain cationic surfactants were used as the template or pore forming agent [19]. Depending on starting materials and synthesis conditions, several mesoporous silica oxides with an ordered structure in the form of a hexagon (denoted as MCM-41) [20–27], a cube (denoted as MCM-48) [21–27] or a lamella (denoted as MCM-50) [19,28,29] were formed. A more useful form of mesoporous silica nanoparticles for biomedical applications was first synthesized by Grṻn et al. [30]. He demonstrated the suitability of adapting the Stӧber synthesis [31] of silica spheres for the production of spherical mesoporous silica particles under basic conditions. The Stöber process is the most common method for synthesizing colloidal silica-based nanoparticles that are b 100 nm in size [31]. The process comprises hydrolysis of a silica alkoxide precursor (such as tetraethyloxysilane, TEOS) in a mixture of ethanol and aqueous ammonium hydroxide. During the hydrolysis silicic acid is produced which nucleates homogeneously and forms silica particles of submicron size when its concentration is above its solubility in ethanol. The rapid condensation process

Bioactive glass is a silicate glass in which SiO2 forms the glass network. It also contains other oxides including calcium, phosphate and sodium oxides. Traditionally, bioactive glass was made by meltquenching, where oxides are molten at temperatures in excess of 1300 °C. Later on, the sol–gel route was also used to prepare bioactive glass. The sol–gel processing used produces a SiO2 network by the polymerization of silica precursors. As part of the synthesis, a high temperature sintering step was needed [37]. The polymerization results from hydroxylation/condensation reactions at room temperature. A gel is first formed (the detail of sol–gel chemistry is discussed in Section 2.1: Sol–gel processed silica (xerogel)). The silica network gel is then heated to about 600 °C in order to remove the byproducts of the polymerization reaction and to achieve the reaction with other oxide precursors (such as triethyl phosphate and calcium nitrate tetrahydrate) for the solid to become a bioactive glass. Because of the processing which includes sintering, it is not possible to retain the biological functionality of any therapeutic molecules that would be incorporated within the bioactive glass.

Fig. 1. Schematic representation of difference between sol–gel silica (xerogel) and mesoporou silica.


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2.4. Mesoporous bioactive glass (MBG) Mesoporous bioactive glasses (MBGs) were first described in 2004 by synthesizing the highly ordered mesoporous structure with a silicate-based bioactive glass composition [38]. The synthesis comprises bringing the surfactant (such as pluronic-P123 or pluronic-F127, which also are non-ionic triblock copolymers) into the sol–gel bioactive glass precursor solution containing a silica source (tetraethyl orthosilicate), a calcium source (Ca(NO3)2·4H2O) and a phosphate source (triethyl phosphate). A highly ordered mesoporous structure is created with a SiO2–CaO–P2O5 composition. Depending on the choice of surfactant, the pore size of the MBG varies from 5 to 20 nm. The MBG can be prepared into various forms including 3D scaffolds, while still maintaining a mesoporous structure (Fig. 2).

3.1.2. Diffusion controlled release The diffusion-controlled release is the leaching of drug molecules present in the pores of the silica when liquid penetrates the porous matrix, dissolves the drugs and diffuses into the exterior liquid as a result of a concentration gradient. The drug release is governed by Fick's law of diffusion. An important controlled release equation describing the release kinetics of molecules arising from diffusion controlled processes is called the “Higuchi equation”. This equation has helped to define the mathematical treatment of controlled release drug delivery systems, since the era of development of sustained release dosage forms started [41]. This equation was first formulated for diffusion controlled release from a porous matrix, from which a drug is leached by the bathing fluid that penetrates the matrix through pores and capillaries in the following for:

3. Controlled release mechanisms and mathematical modeling of release

Q¼

Controlled release systems have been devised for many reasons including such as enabling excellent control of patient drug exposure over time, assisting drugs in crossing physiological barriers, shielding drug from premature elimination, and shepherding drug to the desired site of action while minimizing drug exposure elsewhere in the body. One of the first principles underlying controlled drug delivery was to provide for a desirable near constant safe drug concentration for prolonged durations. With controlled release systems, drug levels in the blood can follow “constant” profiles between Cmin and Cmax for extended periods of time. Controlled release also offers other advantages over systemic drug administration, such as protection of fragile drugs and increased patient compliance and convenience. Controlled release systems can maintain the levels of therapeutic agents within the effective range, can reduce spread to undesirable sites, can avoid toxic side effects and can enable more optimized drug dosage by virtue of increasing bioavailability and therapeutic index of the drug molecule. In this section we will first review several of the controlled release mechanisms. Next we will discuss the effect of synthesis parameters and surface functionalization on the controlled release of molecules encased inside pores of silica based controlled release materials. 3.1. Controlled release mechanisms 3.1.1. Overview Typically drugs are released from a matrix by any of three routes. Fig. 3 is the schematic representation of these drug release mechanisms: desorption, resorption and diffusion. Desorption release which accounts for the release of molecules absorbed on the surface of the materials, contributes a negligible fraction of total drug release. In fact the majority of drug released from sol–gel silica is controlled by the diffusion rate of drug molecules from the pores within the silica; or, by the degradation of the silica matrix itself, or by both.

rffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi Dεð2A−εCÞCt τ

ð1Þ

with Q: amount of drug released after time t, D: diffusitivity of the drug in the permeating fluid, τ: the tortuosity factor of the capillary system; tortuosity reflects a nonlinear path of transport of a drug through the porous structure; the tortuosity accounts for the complexity of pore geometry, A: the total amount of drug present in the matrix, C: the solubility of the drug in the permeating fluid, ε: the porosity of the matrix. The porosity is the fraction of pores in the system that is open to the surface. The above Eq. (1) can be expressed in a more simple form as: pffiffi Q ¼k t

ð2Þ

where k is known as the Higuchi dissolution constant. Thus, the cumulative release of drug released is proportional to the square root of time. Today the Higuchi equation is one of the most widely used and best-known controlled-release equations. The rate of release can be increased by increasing the drug's solubility (C) in the host matrix and vice versa. It is important to note that assumptions are made in this Higuchi model, including (i) the initial drug concentration in the system is much higher than the matrix solubility; (ii) perfect sink conditions are maintained; (iii) the diffusivity of the drug is constant; and (iv) the swelling of the matrix is negligible. The sink conditions are achieved by ensuring the concentration of the released drug in the release medium never reaches more than 10% of its saturation solubility. It can be seen from the Higuchi equation that the nanostructural properties of sol–gel, such as pore volume, specific surface area and pore diameter have a major influence on release phenomena. The

Fig. 2. 3D printed MBG scaffold exhibits three scales of porosity: ultra-large pores, macropores and mesopores [39]. Reprinted from Acta Biomaterialia, vol. 7, p1265–1273, Preparation of 3-D scaffolds in the SiO2–P2O5 system with tailored hierarchical meso-macroporosity, García A, Izquierdo-Barba I, Colilla M, de Laorden CL, Vallet-Regí M. Copyright (2011), with permission from Elsevier.



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Fig. 3. Schematic representation of the drug release mechanisms [40]. Reprinted from Qu H, Bhattacharyya S, Ducheyne P. Sol–Gel Processed Oxide Controlled Release Materials. In: P. Ducheyne, K. Healy, DW. Grainger, J. Kirkpatrick, Hutmacher D, editors. Comprehensive Biomaterials: Elsevier Science; 2011. p. 475–95. Copyright (2011), with permission from Elsevier.

pore surface chemistry also can affect the drug molecule release through drug/sol–gel interactions. In addition to the commonly used Higuchi model, other models have been proposed to fit the drug release behavior from silica sol–gels [42,43]. These models were considered with the intent of overcoming some of the limitations of the Higuchi model, one being that the Higuchi model only considers the drug being released from a planar/tabular matrix. In reality, sol–gel oxides can be prepared in various shapes, such as coatings, fibers and particles. Among the various shapes, particles or powders are typical physical forms, because of simple preparation processes and straightforward medical use. Analyzing release kinetics from a spherical matrix, Baker and Lonsdale [42,43] developed a model starting from the Higuchi model and formulated the drug release from spherical matrices with the equation: " 2 # 3 Mt 3 Mt − ¼ kt 1− 1− M∞ M∞ 2

ð3Þ

where Mt/M∞ is the fractional drug released at time t, and k: the constant depending on the structural and geometrical characteristics. To study the release kinetics, data obtained from in vitro drug release studies were plotted as [d (Mt/M∞)]/dt with respect to the square root of the inverse of time. In diffusion-controlled release, the pores are large enough to be thought of as “channels” for diffusion (pore diameter greater than the drug molecule). Then diffusion occurs predominantly through these water-filled pores. However, if the pores are smaller than the drug molecule, then the diffusion controlled mechanism is not applicable to describing drug release kinetics. 3.1.3. Degradation controlled release When implanted in the body, silica sol–gels gradually degrade in the body over time. The dissolved sol–gel silica is excreted in soluble form through the urine (see the section on biocompatibility). The sol–gel silica degradation rate depends extensively on the structure of the silica. In fact, under body fluid conditions (pH 7.4, temperature = 37 °C), the solubility of amorphous silica is 120–150 ppm, while crystalline silica is ten times less soluble [44,45]. Not only is the solubility, also the dissolution rate is dependent on the structure. There are two major parameters that can be used to adjust sol–gel silica degradation rate: degree of condensation and specific surface area. In comparison to the Si–O–Si linkage, the silanol (Si–O–H) group is easier to dissolve through hydrolysis. Thus the extent of silanol groups is related to the degradation rate; and the degradation rate is higher when more silanol groups are present. When the degree of condensation is low, there is an abundance of free silanol groups in the silica surface [44,46].

A well hydrolyzed sol of alkoxides gelled to a low degree of condensation can result in a high number of free silanol groups per surface area. For example, in a typical silica sol–gel process, a sol with low water-toTEOS ratios catalyzed at low pH (pH around 2) can be used to spin fibers. Sol–gel fibers are spun immediately before condensation, leaving no time for condensation reactions. The resulting silica fibers, having a very low degree of condensation, usually degrade within less than a day [46]. Other fast form-giving processes, such as freeze drying and spray drying, can also produce silica with a low degree of condensation. On the other hand, a regular slow, spontaneous gelation process of the same sol gives rise to fully condensed silanol groups and the spontaneously formed silica monolith dissolved at a much slower rate [46]. The specific surface area also affects the degradation rate. The higher the specific surface area, the more silanol groups are accessible to be liberated from the surface into the dissolution medium. Therefore the high surface area makes the silica degrades faster [45]. Typically the silica hydrogel contains very high water contents, about N 90% (by weight) of water and b8% of silica [47–49]. The sols for these hydrogels are generally prepared at high water-to-alkoxide ratios (N 50). The silica hydrogel dissolve completely in sink conditions within 2–3 days in vitro. Hopfenberg developed a mathematical model to correlate the drug release from a surface eroding matrix, as long as the surface area remains constant during the degradation process [50]. Hopfenberg assumed that the rate of drug release from the erodible system is proportional to the surface area of the device which is allowed to change with time. All mass transfer processes involved in controlling drug release are assumed to add up to a single zero-order process (characterized by a rate constant, k0) confined to the surface area of the system. This zero-order process can correspond to a single physical or chemical phenomenon, but it can also result from the superposition of several processes, such as dissolution, swelling, and matrix chain cleavage. Hopfenberg derived the following, general equation, which is valid for spheres, cylinders and slabs: n Mt k0 t ¼ 1− 1− M∞ c0 a

ð4Þ

Mt and M∞ are the cumulative amounts of drug released at time t and at infinite time, respectively; c0 denotes the uniform initial drug concentration within the system; a is the radius of a cylinder or sphere or the half-thickness of a slab; n is a ‘shape factor’ representing spherical (n = 3), cylindrical (n = 2) or slab geometry (n = 1). For sol–gel silica controlled release systems the above equation can be approximated (for n = 1) by: Mt ¼ kt: M∞

ð5Þ


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3.1.4. Mixed release kinetics From the above discussion, it is evident that the release kinetics controlled by diffusion can be described by Higuchi model. The Higuchi model can be further simplified as: Mt ¼ kt 0:5 M∞

ð6Þ

where Mt/M∞ is the fractional drug released at time t, k: the constant depending on the structural and geometrical characteristics. Similarly the release kinetics controlled by degradation can be expressed as: Mt ¼ kt M∞

ð7Þ

where Mt/M∞ is the fractional drug released at time t, and k is the constant depending on the structural and geometrical characteristics. Based on Fick's second law of diffusion and assuming that release proceeds under perfect sink conditions and that the drug distribution is homogeneous, Peppas proposed an exponential relationship to describe mixed drug release kinetics arising from both degradation and diffusion [51,52]. Mt ¼ kt n M∞

ð8Þ

where Mt/M∞: the fractional drug released at time t, k: the constant depending on the structural and geometrical characteristics, n: the indicator of the mechanism of drug release. This equation can be used until 60% of the drug has been released. It was shown that exponent values n = 0.5 are indicative of diffusioncontrolled release; n = 1.0 is zero order drug release kinetics, which can be with the result of degradation-controlled release. The n values between 0.5 and 1.0 are indicative of anomalous transport behavior including both diffusion and degradation. Although this equation (Eq. (8)) has its limitations, it is considered useful in comparing diffusion controlled release with matrix degradation release [51,52]. 3.2. Effect of synthesis parameters and surface functionalization on controlled release 3.2.1. Surface functionalization The chemistry of a silica surface is dominated by the presence of abundant silanol groups (Si–OH). Drug loading and release kinetics can be influenced by the interaction between drugs and these silanol groups on the pore walls (inner surface of the pore tunnel). To adjust release kinetics, the silanol groups on pore wall surfaces can be replaced by other functional groups through two main methods: cocondensation (one-pot synthesis) and post-synthesis grafting. The one-pot synthesis method is based on the co-condensation of a tetraalkoxysilane (siloxane) and one or more organoalkoxysilane precursors with Si–C bonds. Siloxane precursors constitute the main network for the porous silica materials, while the organoalkoxysilane precursors contribute to the building of the network and work as functional groups on the surface [53–56]. The post-synthesis grafting method involves the modification of the as-synthesized silica porous silica materials with organosilane compounds. Sol–gel processed silica materials possess silanol (Si–OH) groups that facilitate the attachment of organic groups to the surface. Silylation is the most commonly used reaction for the surface modification in the post-synthesis grafting method [57]. The direct synthesis has an advantage over the grafting method, because it produces mesoporous materials with a high loading of functional groups [53,54].

Two major interactions, namely drug–wall interaction and water–wall interaction, are altered after pore wall functionalization. Organically modified silanes with amide moiety can adjust the drug release kinetics by taking advantage of electrostatic interactions between protonated aminopropyl groups and carboxylate groups present in the drugs. It has been shown that the release rate of ibuprofen was considerably reduced when the aminopropyl–triethoxysilane was used to functionalize the silica materials [58]. In addition to the benefit of controlling the release rate, strong drug–wall electrostatic interaction can increase the drug payload within the porous structure [59]. A second strategy for effective control of drug release is functionalization of the pore wall with hydrophobic species such as –CH3 or –CH3CH2. In this case, the functionalization is not intended to affect the drug–wall interactions, but to affect the water penetration into the pores. By impeding water penetration due to the hydrophobic nature of the CH3 or CH3CH2 groups present in the silica network, drug transport out of the matrix is also regulated. The drug release rate from hydrophobicly functionalized sol–gel silica is slower compared to that from unfunctionalized ones [60–62]. 3.2.2. Pore blocking Another strategy to modifying the drug release kinetics involves covering the outside surface of the silica materials to restrict water from entering the porous structure. Various materials have been deposited on the surface of silica in order to explore this mechanism. By way of example, the surface of mesoporous silica particles was coated with magnesium oxide (MgO) using a hydrothermal method. Such surface coating is useful for the extending the release of drug substances [63]. Calcium phosphate thin films have also been deposited on MSN particles in order to achieve rapid drug release in acidic intercellular compartments such as endosomes and lysosomes [64]. Hydroxyapatite thin films have also been deposited on MBG to achieve a pH-sensitive drug release ability [65]. Sol–gel silica itself can also be used to coat porous silica particles for additional control of release rates. Drug release rates decreased with the increase in the number of coatings [66]. The gradual degradation of the silica films smoothed the initial burst release and extended the release duration [67]. A similar approach to tailor drug release was used for depositing polymeric films on the MSN [68,69]. Similar to polymer films reducing burst release from MSN, silk coating on MBG scaffolds decreased the burst release of dexamethasone [70]. The polymer composition can further be optimized to release drug under specific conditions, such as pH and temperature [71–73]. Smaller molecules which can change conformation upon environmental changes have also been used to modify the pore orifice for controlling drug release. The so-called gated/capped system can be selected to respond to external stimuli such as temperature [74], light [75,76], pH [77], enzymes [78], and other chemical agents [79]. When the right stimuli are present, the blocked pore opens and allows drug diffusion through the pores. Thus, the release can also be controlled by external stimuli. 3.2.3. Pore parameter Pore diameter is another factor that can control the transport of the drug molecules. Drug release kinetics decrease as the pore size decreases with constant drug size [33]. If large molecules such as proteins or nucleic acids are the therapeutic agents of interest, the pore size needs to be expanded to accommodate them. In general, MSNs synthesized with cationic cetyltrimethylammonium surfactant have a pore size less than 3 nm. This renders it difficult to load proteins or nucleic acids. The chain length of carbon atoms of surfactants affects the pore size. In one study, the pore size was adjusted within the range of 2.5 to 3.9 nm by varying this chain length from 8 carbon atoms to 16 carbon atoms [33]. Adding molecules to act as swelling agents affect the selfassembly process of surfactant micelles and is a principle that was used to expand the pore diameter further. Swelling agents (such as



alkanes/EtOH mixture) can expand the pore diameter up to 6.0 nm [80]. Another principle to expand the pore size of MSN is by a post-synthesis hydrothermal treatment that exchanges surfactant with a longer alkylchain length [81,82]. For sol–gel silica prepared without surfactant, the pore size can be adjusted by varying the pH of the sol [83,84]. At the isoelectric point, the silica colloids aggregate easily due to lesser electric compulsive forces. As a result, densely packed sol–gel silica (xerogel) with smaller pores can be achieved [83]. Another principle used for sol–gel silica is halting the condensation between colloidal silica particles and this is achieved by rapid water removal. Freeze drying and supercritical drying preempt capillary effects and leaves large pore sizes ranging from 10 to 30 nm. As a result, the drug release rate from critical dried silica with large pores is much faster than from conventionally dried xerogels [85–87]. 4. Physical application forms Sol–gel materials can be prepared in a variety of different configurations, including monoliths, crushed particles, microspheres, nanospheres, fibers, thin films and composites. Silica sol–gel monoliths or crushed particles have been used as implantable or injectable materials [9,15,88]. Spray-drying and emulsion chemistry have been used to prepare silica xerogel microspheres and nanospheres with a narrow particle size range for better control of the release kinetics of embedded molecules [42,89,90]. Hydrolyzed silica sol containing drugs either directly sprayed and dried in air or mixed and separated by non-polar organic solvents to form microsphere or nanospheres. Drug-loaded sol–gel films have been produced by applying a solution of liquid sol with drug onto implant surfaces in order to release therapeutic agents at the site of implantation [91,92]. 4.1. Particles In the case of sol–gel-derived silica, after drying, the silica monoliths can be ground and sieved to arrive at micron sized particles [42]. Similarly, after sintering, an MBG pellet can be crushed and sieved to yield particles in micron sizes. Spherical shaped sol–gel microparticles can be synthesized using spray-drying [18] or emulsification methods [42]. These particles have regular surfaces. Micron sized silica particles can be used as bone substitute materials and expedite bone healing processes by releasing appropriate drugs. In addition, micron sized silica particles can also be subcutaneously implanted as drug depot for controlled release of drugs over prolonged periods [93,94]. Nanosized sol–gel silica particles can be prepared through an emulsion sol–gel process [95]. In such process, surfactants with a mid-range HLB value (hydrophilic–lipophilic balance) (between 10 and 15) are used in order to allow silica precursor to diffuse through the surfactant wall and form nano-sized silica particles. As described above, mesoporous silica nanoparticles are synthesized through a modified Stöber procedure, a procedure which is also suitable to produce nano-sized MBG nanoparticles [96]. Nano-sized particles can be injected into the blood stream and be transported to organs without invasive surgery. To increase circulation time, the surface of the silica nanoparticles has been modified with non-fouling molecules such as polyethylene glycol (PEG) [97]. 4.2. Micron thin films Metallic implants possess the requisite mechanical strength for loadbearing applications. The success of implants extensively depends on the optimal interaction between implant surface and surrounding tissue. In this context, surface engineering is often utilized to facilitate the process of bone healing, to prevent infection and biofilm formation, and to interact optimally with bone. Dip-coating is commonly used for the deposition of micron-thin sol–gel films onto metallic substrates. The substrate is dipped and withdrawn at a controlled speed from the

7

sol using a dipping device. The micron-thin film coatings are typically air-dried. Multi-layer sol–gel films can be produced without interlaminar failure of subsequent layers by repeating the dipping and drying procedures [91]. Typical thicknesses are 1–2 μm (5–10 layer film). Vancomycin containing sol–gel silica film showed a sustained release [91]. As will be discussed at greater length below, microbicidal activity of the vancomycin-containing sol–gel coatings was demonstrated using a rat osteomyelitis model [92]. Molecules that can regulate the biological behavior of osteoblasts/ osteoclasts can also be incorporated in sol–gel films. Using a silica– chitosan coatings fibroblast growth factor 2 (FGF2), known to stimulate osteoblast cell proliferation and migration, has been incorporated into ~1 μm thick coating on a titanium surface. The FGF2 was released over more than 4 weeks [98]. MSNs loaded with β-estradiol were coated using polyelectrolyte multilayers (PEM) of gelatin/chitosan via the layer-by-layer (LbL) technique [99]. The titanium substrates were coated by alternatively dipping into polyethylenimine (PEI) solution, gelatin solution and chitosan solution. The β-estradiol was released from the MSNs embedded in the film. Although the thickness of the coating was not reported, the film morphology suggested the film thickness exceeds the size of MSN, which was 180 nm. The coated film not only promoted the biological function of osteoblasts, but also inhibited the proliferation of osteoclasts, which made the authors to suggest that implants coated this way would be suitable to treat patients with osteoporosis [99]. 4.3. Scaffolds Bone tissue engineering treatments oftentimes rely on bioactive materials that inherently foster bone tissue formation; in tandem, design and processing of a porous, biodegradable three-dimensional structure, exhibiting high porosity, high pore interconnectivity and uniform pore distribution allow bone cell proliferation. As a note of caution to the readers: pore sizes of scaffolds are 3 to 5 order of magnitude different from the nanopores in the controlled release materials. Recent advances in rapid prototyping technologies have made it possible to create mechanically competent scaffolds with controlled architectures [100,101]. Such porous structures frequently mimic the native bone architecture. In addition to optimizing scaffold design, tuneable delivery of growth factors and cytokines embedded in these scaffolds is another goal. The principle is to elicit signaling at local injury sites whereby inflammatory and progenitor cells migrate and trigger the healing process [102,103]. Local, long-term release of growth factors from biodegradable scaffolds is therefore of obvious interest. Polymers, because of their ability to be processed at low temperatures, have been used widely for the preparation of 3D scaffolds. Silica nanoparticles have been incorporated into the polymeric scaffold to provide an additional degree of control of drug release, independent from the polymer material itself [104]. MBG scaffolds have been synthesized with the goal of bringing principles of efficient drug delivery to tissue engineering. The concept is that therapeutic molecules are released from mesopore channels in the scaffold matrix [105]. Currently, two methods are used for the preparation of highly interconnected macroporous MBG scaffolds: negative template method (with polyurethane (PU) sponge) and direct 3D printing technique. Direct 3D printing relies on adding either a mesostructure directing agent (methylcellulose) [39] or a binder (polyvinylalcohol) [106] to the MBG precursor to form a mixture with the appropriate viscosity. The viscous slurry is then deposited layer by layer to create highly ordered macroporous MBG scaffolds. Electrospun nanofiber scaffolds mimic micro and nanoscale features of fibrous extracellular matrix (ECM). In order to create nanofiber scaffolds with drug delivery capacity, electrospun composite nanofibers that incorporate surface functionalized MSN have been prepared [107]. When incorporated into nylon-6 nanofiber scaffolds, the surface functionalized MSN maintained its ability to release model molecules


8


under neutral but not acidic pH conditions [107]. Another concept was to add surface-functionalized, but non-drug loaded MSNs into freezedried nanofiber scaffolds. The goal of this concept was to control the drug release from the polymer matrix. The principle upon which one relies in that the surface amine group of the MSN enhances the interaction between the composite release fiber material and drug, specifically by reducing burst release [108]. Yet another method for synthesizing silica/polymer nanofiber scaffolds is mixing the prehydrolyzed sol with surfactant and polymer solution. By electrospinning the mixture, a silica/polymer nanofiber scaffold can be obtained all while mesoporous structure of MSN within the polymer matrix is retained [109].

5. Clinical applications In a historical context, the biomaterials field advanced as “bioactive” materials replaced “bio-inert” materials, this is as materials were designed and used that elicited desirable tissue responses. Silica based materials represent one such class of biomaterials. Many silica based materials display excellent interaction with bone tissue and they are clinically used in a variety of orthopedic and dental applications. The incorporation of pharmaceutical and biological molecules for controlled release enhances and expands the scope of the clinical use of silica materials in orthopedic, dental and maxillofacial applications. Specifically, silica controlled release materials have been investigated for infection treatment and control, pain treatment and control, scaffolds for in tissue engineering procedures, and cancer and tumor treatment. Table 1 provides an overview of the current preclinical status of various applications using silica controlled release materials. Among these studies, several applications stand out as closer to the clinic. These include infection treatment and scaffolds for tissue engineering.

5.1. Infection treatment and control Bone infections are highly localized, although they are oftentimes caused by systemic pathogens. One of the therapeutic challenges is to deliver pharmaceutical agents to these sites. Due to poor vascularization, the bone tissue has lower local concentration of antibiotics than other highly-vascularized organ, when the antibiotics administered systemically. Thus, a biodegradable local delivery of antibiotics is desired in orthopedic surgeries to prevent infection and sol–gel silica based materials show a great promise in this regard. 5.1.1. Vancomycin-containing silica sol–gels: synthesis and release kinetics Vancomycin-loaded silica sol–gels synthesized by a roomtemperature process were developed in the mid-nineties [12,152]. Single-step acid-catalyzed hydrolysis (pH 1.8–2.0) with excess of water over the stoichiometric H2O/Si ratio was used. The sol was prepared by hydrolysis of tetramethylorthosilane (TMOS) in deionized water; 1 N HCl was used as a catalyst. TMOS, DI, and 1 N HCl were mixed in a glass beaker and stirred with a use of a magnetic stirrer for 30 min to obtain a homogenous sol. Then vancomycin–HCl dissolved in DI (deionized water) was added to the sol. Upon vancomycin loading, the sol was cast into cylindrical polystyrene vials. The vials were sealed and the sol samples were allowed to gel, age and then dry until the sol–gel weight became constant. Fig. 4 shows typical silica xerogel monoliths. Vancomycin loaded silica sol–gels were fabricated with varying water/alkoxysilane molar ratios of 10, 6, 4, and 2 and vancomycin concentrations of 10 to 20 and 30 mg/g [12,13]. Release kinetics were measured by immersing the xerogel in a buffered plasma-like electrolyte solution which was exchanged daily.

Table 1 Publications of preclinical investigation of silica controlled release materials. In vitro feasibility testing (a)

In vitro proof In vivo efficacy In vivo efficacy of concept small animal large animal demonstrated (b) (mice or rat) model model

TOWARD CLINICAL USE Infection treatment and control — thin films on orthopedic devices Controlled release of antibiotics from xerogel film Controlled release of antibiotics from mesoporous silica film Controlled release of nitric oxide from xerogel film MRSAc treatments

[91] [111] [113] [116]

Infection treatment and control — particles Controlled release of antibiotics Controlled release of nitric oxide Controlled release of Ag+ ion

[12,13,42,79,117–122] [117,122,123] [126,127] [126,128] [129] [129]

Infection treatment and control — scaffolds Controlled release of antibiotic from MBG

[130,131]

[131,132]

Pain treatment and control Post-surgical pain

[42,84,133–137]

[133]

[9,15,139,140]

[9,15] [141] [143]

Scaffolds for tissue engineering Controlled release of growth factor (BMP-2, TGF-β,FGF) from xerogel Controlled release of growth factors (BMP, VEGF) from MSN Controlled release of growth factors (BMP, VEGF) from MBG Controlled release of ions from MBG Controlled release of steroids (dexamethasone) from MBG Controlled release of dimethyloxallyl glycine from MBG Controlled release of a plasmid DNA (pDNA) encoding growth factor (PDGF, BMP) from MSN Release of a plasmid DNA (pDNA) encoding growth factor (PDGF, BMP) from bioactive glass a b c d e

[142,143] [105] [146,147] [148] [149,150]

[92] [111] [113] [116]

[146,147] [148] [149,150]

[92] [114]

[110] [112] [115]

[124,125]

[138]

[143] [144]d

[145]e

[149] [151]

In vitro feasibility studies: molecule can be incorporated and the release kinetics can be optimized for the application. In vitro proof of concept studies: biological activity of the molecule is retained. MRSA: methicillin-resistant Staphylococcus aureus. Strontium ion. Lithium ion.



9

The results demonstrated that the release kinetics of vancomycin can be modified by varying the water/alkoxide molar ratio and the vancomycin concentration (Figs. 5 and 6). Variation of the water/ alkoxysilane molar ratio and the antibiotic concentration used in the room temperature sol–gel process is a simple, effective, and minimally time consuming technique for controlled drug delivery. This study, in combination with other studies on silica xerogel physical properties such as surface area and pore volume [2,12] and on the in vivo biocompatibility and resorbability of silica xerogel [3], showed early on the major potential of using sol–gel polymerized ceramics (xerogels) as controlled release materials for the treatment of bone infection. 5.1.2. Sol–gel silica thin films for controlled release of antibiotics Whereas the previous studies were performed with microparticles, it was also shown that vancomycin can be incorporated in micronthin films using silica sol–gel solutions. It was shown that coatings of this sol composed of several layers can be applied to metallic implant surfaces [91]. Sol–gel coatings composed of several layers of 0.25 μm thickness were successfully applied on Ti-alloy strips and anodized Ti-alloy Kirschner wires. Both types of samples showed a timedependent, controlled release (Figs. 7 and 8). The data show that the release rate varied to a large degree with vancomycin concentration (Fig. 7) and the number of applied layers (Fig. 8) or both. Using a multi-layer process long-term release can be achieved. The concentrations released from multi-layer coatings with 10% and 20% vancomycin exceeded the Minimal Inhibitory Concentration (MIC) of vancomycin against Staphylococcus aureus (1.56–3.12 μg/ml). The results also demonstrated that the release of vancomycin from micronthin sol–gel coatings can be tailored to therapeutic needs by varying sol–gel processing parameters, vancomycin load and the number of applied layers. Fig. 9 demonstrates that the sol-gel coating with vancomycin (SGV) significantly reduced the numbers of attached bacteria by nearly four orders of magnitude in comparison to so-gel coating without vancomycin (SG) and uncoated titanium rod (Uncoated control). As such, micron thin silica sol–gel films can be used as a drug delivery depots and the concept is moving toward clinical studies for the prevention or treatment of periprosthetic infections. Various pre-clinical studies support this development, as shown next. 5.1.3. Sol–gel/vancomycin micron-thin sol–gel coatings for the treatment of S. aureus infection in vivo A rat osteomyelitis model was used to assess the microbicidal activity of vancomycin-containing sol–gel coatings [92]. Wistar rats weighing between 300–350 g were used. The bacterial inoculum (150 μl suspension of 103 CFU of bacteria) was slowly injected into the

Fig. 4. Silica xerogel monoliths. The samples are glassy, transparent, and crack-free. The thickness of samples varies with the amount cast [12]. Re-printed with permission from John Wiley & Sons, Inc.

Fig. 5. Percentage of cumulative vancomycin release as a function of elution time and vancomycin load for sol–gels with a H2O/TMOS ratio equal to 6 (n = 3, error bars represent ± 1 standard deviation) [153]. Re-printed with kind permission from Springer Science and Business Media.

medullary space. As the bacteria were injected, the needle was slowly retracted from the femoral canal. Sol–gel coated implants with vancomycin, or without (controls), were press fit into the canal in a retrograde fashion and the stab incision was closed (n = 3 per group and implantation time). No systemic antibiotics were used at any time. The progression of infection was demonstrated radiographically at 7, 14, 21 and 28 days. After 4 weeks of implantation, infection on the control side was evidenced by change in size, the periosteal reaction (arrows), lytic lesions, bone abscesses and extensive bone remodeling (Fig. 10: bottom). In contrast, femora with vancomycin-containing sol–gel rods (Fig. 10 top) displayed minimal radiographic changes. The results of this study demonstrate that the sol–gel micron thin films provides predictable release kinetics that are relevant to ongoing surgical procedures and minimizes the need for systemic antibiotics. Moreover, while preventing bacterial proliferation on the rod surface, the sol–gel inhibits biofilm formation. From a clinical perspective, implants coated with antibiotic impregnated sol–gels provide an ideal technology for delivering a wide range of antibiotics to the bone surface during a critical period of bone repair without delaying or inhibiting osseointegration. 5.1.4. Silica film for nitric oxide release Nitric oxide (NO) has been discovered as an important mediator in the immune response. It has been reported that NO gas released from sol–gel films destroyed colonies of bacteria and a persistent concentration of NO around medical device reduced implant-associated infection

Fig. 6. Percentage of cumulative vancomycin released as a function of elution time and H2O/TMOS ratio for sol–gels containing 20 mg/g of drug (n = 3, error bars represent ± 1 standard deviation) [153]. Re-printed with kind permission from Springer Science and Business Media.


10


Vancomycin, µg

300

5% V

250

10% V

200

20%V

150 100 50 0 0

2

4

6

8

Time, days Fig. 7. Mean cumulative vancomycin release versus vancomycin concentration in the sol– gel film composed of three layers [91]. Reprinted from Biomaterials, vol. 28, p1721–1729, Controlled release of vancomycin from thin sol–gel films on titanium alloy fracture plate material, Radin S, Ducheyne P. Copyright (2007), with permission from Elsevier.

[114]. Another study also showed that sol–gel surfaces capable of NO release decreased bacterial adhesion by 30% to 95% relative to controls [113]. By modifying sol–gel silica through thiol nitrosation, the silica is capable of generating nitric oxide (NO) for up to 2 weeks under physiological conditions [154]. The NO releasing silica film showed reduced foreign body collagen capsule thickness by N50% compared to uncoated substrate after 3 weeks. The chronic inflammatory response was also reduced by N30% after 3 and 6 weeks [155]. 5.1.5. Sol–gel silica thin film for the treatment of methicillin resistant S. aureus (MRSA) infection The incidence of methicillin-resistant S. aureus (MRSA) infection has significantly increased. Generally, the success of this bacterium as a pathogen is attributed to its ability to adhere to surfaces and remain there, under the protection of an extracellular matrix known as biofilm. Treatment and eradication of MRSA infection are difficult since bacteria hide within the biofilm. Thus it is difficult to achieve the optimum, relatively high concentration of therapeutics within the biofilm. As such, there is a critical need for identifying therapeutic strategies that are directed toward preventing the MRSA infection in addition to effective treatment once present. Traditionally, vancomycin is the mainstay of treatment, because it provides in vitro activity against all staphylococci and demonstrates clinical response against MRSA infection [156]. However, in vitro susceptibility of MRSA to vancomycin is gradually decreasing (MIC, 32 μg/ml in the United States) [157,158]. To combat MRSA with low doses (nontoxic) of vancomycin, one of the promising techniques is to use combinational therapeutics (antibiotic/adjuvant) to increase the effectiveness of vancomycin. It has been shown that co-delivered antibiotic and adjuvant from silica sol–gel micron thin films was more effective in killing a methicillin-resistant S. aureus (MRSA) strain in vitro than vancomycin

Fig. 9. Bacterial (S. aureus) growth on sol–gel coated anodized Ti-alloy pin. Note that in comparison to control, the sol–gel coating with vancomycin caused a decrease in bacterial growth by nearly four orders of magnitude [92]. Re-printed with permission from John Wiley & Sons, Inc.

itself [116]. Fig. 11 shows that sol–gel films with 10 or 20 wt.% vancomycin (SGV(10) and SGV(20)) were unable to eliminate MRSA growth; but for sol–gel micron-thin films with 40 wt.% farnesol (SGF(40)) there was no bactericidal effect. However, when 30 wt.% farnesol was added to sol–gel films with 20 wt.% vancomycin (SGVF(20, 30)), it resulted in no measurable bacterial growth on the implant surface. This in vitro study demonstrated that thin sol–gel films can be applied to titanium implants for the controlled release of antibiotic and adjuvant. The data show that a multilayer thin sol–gel film can be deposited containing both hydrophilic vancomycin and hydrophobic farnesol, individually and in combination. The release characteristics and degradation of the thin films are load-dependent and timedependent. The farnesol acts as a potent adjuvant in that it increases the effectiveness of vancomycin. The incorporation of vancomycin and farnesol into thin sol–gel films represents a new treatment paradigm for the topical delivery of antibiotics with adjuvant. Preclinical studies to demonstrate the effect in vivo are currently underway. 5.1.6. Bioactive glass for infection treatment Bioactive glasses with a SiO2–Na2O–CaO–P2O5 composition have some antimicrobial activity in aqueous solutions due to the release of their ionic compounds over time [159]. The release of the dissolution

Cumulative release of Vancomycin from anodized Kirschner wires in PBS

Vancomycin, um

25 20 15 10

3l-10V

5

5l-10V 5l-10+20V

0 0

1

2

3 4 Time. days

5

6

7

Fig. 8. Mean cumulative vancomycin release from sol–gel films composed of three or five layers and applied on anodized Ti-alloy Kirschner wires [92]. Re-printed with permission from John Wiley & Sons, Inc.

Fig. 10. Sol–gel/ vancomycin coating (top) inhibits infection in vivo. After 4 weeks of implantation, infection on the control side (bottom) is evidenced by change in size, periosteal reaction (arrows), lytic lesions and bone abscesses(*), and extensive bone remodeling [92]. Re-printed with permission from John Wiley & Sons, Inc.



11

complications. The effectiveness of bioactive glass S53P4 in a condition of osteomyelitis was presumably based on the antibacterial, osteoconductive, and bone bonding properties of the glass. While the number of patients was small, the clinical trial represents a set of encouraging results for the use of S53P4 as bone substitutes in the treatment of osteomyelitis in humans. A mixture of antibiotics (gentamicin) and sol–gel derived bioactive glass granules has been prepared by isostatical compaction; this produced dense bone void fillers. The bioactive glass/gentamicin mixture was implanted in femurs of New Zealand Rabbits for 12 weeks. The released gentamicin maintained concentration above minimal inhibitory concentration (MIC) in the osseous tissue for more than 12 weeks [124]. Other than the bioactive glass granules, MBG 3D scaffolds can also release gentamicin for the treatment of osteomyelitis [130]. Wu et al. reported that ampicillin loaded MBG scaffolds reduced bacterial growth [132].

Fig. 11. The bactericidal effects against MRSA growth of thin sol–gel films on metal implants containing vancomycin and farnesol alone, or in combination. The implants coated with thin sol–gel films were challenged by a 104 CFU/ml solution of MRSA (mean ± s.e.m., n = 3,*: p b 0.05 compared to SGVF(20,30)) [116]. Reprinted from Biomaterials, vol. 35, p509-517, Sol–gel silica controlled release thin films for the inhibition of methicillin-resistant Staphylococcus aureus, Bhattacharyya S, Agrawal A, Knabe C, Ducheyne P. Copyright (2014), with permission from Elsevier.

products result in a high pH environment [160], which is not welltolerated by bacteria [161]. In addition, it has also been hypothesized that the release of silica has a link to an antibacterial effect of bioactive glass [162]. The biocidal effect of bioactive glass 45S5 derived glassceramic substrates against common Gram positive and Gram negative bacteria, and also against yeast have been reported [163]. Bacterial colonies were typically reduced by one to three order of magnitudes [163]. When doped with Ag, a sol–gel derived silicate bioactive glass in the system SiO2–CaO–P2O5 showed favorable antibacterial activity against Escherichia coli (E. coli), Pseudomonas aeruginosa (P. aeruginosa) and S. aureus [164]. The antibacterial effect was attributed to the leaching of Ag ions from the glass matrix and it was claimed that the incorporation of Ag ions into the sol–gel derived glass provided sustained delivery of the antibacterial agent. Another Ag-doped silicate bioactive glass in the system SiO2–CaO–Na2O, prepared by the sol– gel method, showed high antimicrobial activity against E. coli and Streptococcus mutans (S. mutans) when placed in contact with those bacterial strains [165,166]. Whereas there is a bactericidal effect, it is order of magnitude different from the effect arising from potent antibiotics typically used in bone. Copper-doped bioactive glasses have also been investigated for their antibacterial properties. In one study, phosphate bioactive glasses in the Na2O–CaO–P2O5 system, containing 1–10 mol% CuO, were prepared by conventional melting and casting and evaluated for their ability to function as antibacterial delivery systems [167]. A reduction of 1.5 orders of magnitude was found in the number of viable staphylococci attached to the glass fibers and in the surrounding environment. In another study, phosphate bioactive glass containing 10–15 mol% CuO reduced the number of adherent Streptococcus sanguis (S. sanguis) after 24 h of culture [168]. However, this was followed by an increase in the number of S. sanguis to the level in the control group (glass without CuO), which was attributed to the formation of a biofilm. Bioactive glass S53P4 is antibacterial by itself and it has shown promising results in treating bone infections in humans without any signs of toxic reaction [169]. In a multicenter study [170], 11 patients with verified chronic osteomyelitis in the lower extremity and the spine bioactive glass S53P4 was used as synthetic bone graft. At a mean follow-up period of 24 months (range 10–38 months), the glass was found to be well tolerated, and nine patients healed without

5.1.7. Mesoporous silica particles for infection treatment Different morphologies of MSN have different release kinetics. It was observed that hexagonal internal structures release amoxicillin slower than disk forms [171,172]. Also a faster release rate of amoxicillin was observed in powder compared to that from disks [173]. Fig. 12 illustrates the use of a cadmium sulphide (CdS) cap. Due to this covalently capped CdS, premature release of vancomycin from MSN was so negligible that less than 1.0% of premature release was found. This cap can be removed and drug molecules can be released through cleavage of the disulphide linker from chemical stimulation of dithiothreitol (DTT) and mercaptoehtanol (ME) [79]. A change in release kinetics by surface functionalization was observed in erythromycin incorporated in MSNs. For example, by functionalizing the surface of MSNs with hydrophobic long-chain hydrocarbon moieties, such as octadecyl-functionalized SBA-15, the release rate of erythromycin was impeded [120,174]. In another study, linezolid loaded MSNs were kept inside hollow porous 316L stainless steel pins, which had a porous structure with pore size less than 200 nm (smaller than the size of MSN used in that study, which was 400 nm). The pins containing linezolid-loaded MSNs reduced the bacterial growth by two orders of magnitude when challenged with the S. aureus in vitro [123]. In a sheep tibia infection study [125], antibiotic-loaded implants were placed in the tibia of four sheep which were trans-surgically experimentally infected with a biofilm forming strain of S. aureus. After 7 and 9 days post infection, sheep did not show any evidence of infection as demonstrated by clinical, pathological and microbiological findings [125]. These results demonstrate the capability of such an antibiotic-loaded implant to prevent infection in orthopedic devices in vivo. Further research is needed to assess its possible use in traumatology and orthopedic surgery. Significant advances are achieved for using the MSNs as controlled release materials for antibiotics and other therapeutic molecules. However, more studies are necessary to fully develop this deliver concept so it can be used in both human and veterinary medicine. 5.2. Pain treatment and control Pain of any origin is far more difficult to cure once it has become chronic due to changes (plasticity) in the central nervous system. For example, there is currently no generally accepted and effective treatment for the phantom limb pain syndrome that follows traumatic or surgical amputations but there is a growing consensus that the pain mechanism begins with peripheral events that generate a series of subsequent physiological processes in the central spinal and cortical brain structures [175]. The sensitivity of nerves to local anesthetics is a function of nerve fiber diameter, degree of nerve myelination, physiologic firing rate, and anatomic location. Activated pain fibers typically fire rapidly and as a result pain sensation may be selectively blocked by local anesthetics


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Fig. 12. MSNs showing capping structure useful for controlled release [79]. CdS: cadmium sulphide. Reprinted with permission from Lai C-Y, Trewyn BG, Jeftinija DM, Jeftinija K, Xu S, Jeftinija S, et al. A mesoporous silica nanosphere-based carrier system with chemically removable CdS nanoparticle caps for stimuli-responsive controlled release of neurotransmitters and drug molecules. Journal of the American Chemical Society 2003;125:4451–9. Copyright (2003) American Chemical Society.

5.2.1. Sol–gel silica controlled release of bupivacaine Controlled release of bupivacaine was demonstrated with acidcatalyzed xerogel granules containing bupivacaine (BP) [180]. These were synthesized by mixing tetraethoxysilane (TEOS), [Si(OC2H5)4], and de-ionized water with 1 N HCl as a catalyst until a one-phase solution (sol) was formed. Bupivacaine incorporation in the sol was achieved by adding methanol solutions of the drug (70 mg of bupivacaine per ml of methanol) to achieve desired drug loading levels. The sol was dried into xerogel disks that were then crushed into granules and sieved to obtain microparticles within the range of 20– 105 μm. Release studies of bupivacaine from sol–gel particles were conducted in phosphate buffered saline (PBS, pH 7.4). The xerogel microparticles were immersed in PBS, incubated at 37 °C while shaken at 100 rpm, and solutions were exchanged daily. Released bupivacaine concentrations were measured spectrophotometrically at 265 nm. Results demonstrated that the bupivacaine can be released for extended period of time up to 10 days in controlled fashion [180]. 5.2.2. Sol–gel/polymer composite controlled delivery of local anesthetics Composites of copolymers with bupivacaine-loaded sol–gel silica particles were developed by solution blending [180]. This study demonstrates that the composite active wound dressings can provide fully tunable drug delivery kinetics approaching zero-order by carefully integrating the drug binding and drug release properties of polymeric and ceramic biomaterial components that are each biodegradable and non-inflammatory. Active composite wound dressings were studied by dispersing drug-loaded xerogel microparticles into a continuous polymeric phase consisting of tyrosine–poly(ethylene glycol) (PEG)derived poly(ether carbonate) copolymers. Sustained, controlled release of bupivacaine and mepivacaine, two water-soluble local anesthetics from these composites was demonstrated in vitro [180]. Approximately 200 mg copolymer was dissolved in tetrahydrofuran (THF) and xerogel granules containing BP were added, followed by vigorous mixing for 1 min. The suspensions were then poured into

small Petri dishes, dried under nitrogen flow and then in a vacuum oven at 40 °C overnight. The resulting composite films were then compression molded to the desired thickness. When the xerogels and copolymers were combined in composites the BP release kinetics changed dramatically (Fig. 13). The two stage release process with initial burst release from either the xerogel or polymer was replaced by a single stage release over 7 days that is essentially zero-order, as seen by the nearly linear dependence on time. The zeroorder kinetics were only attained with very specific combinations of tyrosine-derived monomer and PEG stoichiometries (DTO-20%PEG50%). 5.2.3. Sol–gel silica controlled release of Non-steroidal anti-inflammatory drugs Non-steroidal anti-inflammatory drugs (NSAID) are also used frequently for pain relief. NSAIDs relieve pain by blocking enzymes and proteins synthesis during inflammation. However, the risk for bleeding is contra-indication associated with their long term usage. A sustained release bypassing the GI tract is therefore very promising. Sustained release of NSAIDs, such as ibuprofen and diclofenac, has been achieved by incorporation in sol–gels [135,136]. Sousa et al. devised a method to tailor the release of NSAID from silica nanotubes 120

Cumulative release, %

that block the sodium channels involved in conduction of neuronal action potentials. Exposure to higher doses of the agent in a region surrounding the nerve may or may not lead to motor blockade due to the requirement of drug penetration to the core of the nerve [176,177]. The combination of the local anesthetic bupivacaine and narcotic agents (e.g., Fentanyl) often yields a variable motor blockade and a more predictable blockade of the sensory pathways postadministration. The use of low dose L-bupivacaine produces less motor blockade than its stereoisomers (D-bupivacaine) [178,179].

100 80 60 40 Copolymer Xerogel Composite

20 0 0

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40

60

80

100

120

140

160

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Time, hrs Fig. 13. Comparison of bupivacaine release from xerogel, copolymer and their composite. The BP release rate from the composite is substantially less than that from either the copolymer or xerogel from which it is made. Triangles: polymer (3.5 wt.% BP in total composition); Circles: xerogel, (7 wt.% BP in total composition); Squares: xerogel– polymer composite (3.5 wt.% BP in total composition) [180]. Reprinted from Biomaterials, vol. 31, p6336–6343, Polymer–xerogel composites for controlled release wound dressings, Costache MC, Qu H, Ducheyne P, Devore DI. Copyright (2010), with permission from Elsevier.



by functionalizing the silica pore wall with aminosilane molecules. Such modification would only allow the NSAID release at chronic inflammatory conditions, which are often acidic [181]. Similar attempts to functionalize pore walls with amino groups have also been tried by others [182]. In an in vivo study, sol–gel silica containing dexmedetomidine, an alpha-2-agonist, was subcutaneously administrated to dogs. Dexmedetomidine is widely used in veterinary anesthesia. This in vivo study showed that the mean maximum safe concentration of dexmedetomidine in serum extended over 5 h and measurable dexmedetomidine concentrations were still found 48 h after implantation [61]. This study proved that anesthetic drugs (dexmedetomidine) incorporated in silica xerogel can be released over extended period of time in vivo [61]. Typically, patients undergoing elective orthopedic surgery are discharged to home on post-operative day 3. The main reason behind an extended hospital course is inadequate pain control. Effective treatment regimen for postoperative pain is also required for dental surgery. To minimize the time spent in the hospital, decrease postoperative narcotic use and increase patients' quality of life, sol–gel silica based controlled release long acting pain medicine can provide better management of postoperative pain relief. 5.3. Tissue engineering Even though osseous tissue has the unique capacity to heal and remodel without scarring, there are several conditions, both congenital and acquired, where bone grafting and bone replacement is needed [183,184]. The holy grail in bone tissue engineering is the pursuit of scaffolds that act as a temporary support for cells to facilitate the regeneration of bone. With requirements for scaffold materials are being very challenging. Scaffolds provide cell anchorage sites, mechanical stability and structural guidance, and in vivo, provide the interface to respond to physiologic and biologic changes. In addition, they serve to facilitate remodeling the extracellular matrix in order to achieve integration with the surrounding native tissue. Focusing on bone grafting first, non-resorbable materials do not remodel and integrate with bone tissues, oftentimes repeat surgeries are needed to remove the materials. While the use of autograft material is a preferred bone grafting technique, there are limitations such as donor site morbidity, limited donor bone supply, anatomical and structural problems, and elevated levels of resorption during healing. Allografts have the disadvantage of potentially eliciting an immunological response due to genetic differences. They also suffered from the risk of potentially inducing transmissible diseases. Discussing the use of growth factors next, the methods currently being studied to incorporate growth factors into synthetic scaffolds include absorbing growth factor to the scaffold [185], blending growth factor containing polymer microspheres into the scaffold [186,187], or directly mixing growth factor containing protein powder into the scaffold during processing [188,189]. Adsorbing growth factor onto the scaffold has the drawback of low loading efficiencies and rapid release. Loading growth factor into polymer microspheres can be associated with a loss in bioactivity due to harsh solvents such as methylene chloride [190]. Silicon oxide based porous materials have gained promise in this regard. Sol–gel silica particles can be incorporated with growth factors without involving any harsh chemicals and it is demonstrated that both high loading and controlled sustained release of growth factors can be achieved by incorporating these particles into porous biocompatible polymer scaffolds. 5.3.1. Porous, bioactive ceramic scaffolds for bone tissue formation in vivo Bioactive glass templates combine several requirements of the ideal template for in vitro synthesis of bone tissue. Especially when made in porous, resorbable form, and conditioned to develop a bone-like surface prior to being seeded with pluripotential cells capable of expressing the

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osteoblastic phenotype, these templates lead to expeditious and abundant in vitro synthesis of extracellular matrix with most important characteristics of bone tissue [191]. In addition, when growth factors with established beneficial effect on bone tissue formation would be incorporated in the templates and be released in a time and dose controlled fashion, it is possible to further optimize in vitro bone tissue engineering concepts [15]. Silica based glasses by themselves, without the delivery of growth factors are actually very useful as scaffold materials. In fact, bioactive glass templates are resorbable and are gradually replaced by bone tissue. This follows from an implantation experiment in which the typical femoral bone defects encountered in revision hip arthroplasties were simulated [192]. In this study, a unicortical window defect was created bilaterally in the femoral diaphysis of adult, male syngeneic rats. The model used was not a critical size defect model, but addressed the question of the effect of adding cells. Porous, surface modified bioactive glass ceramics (pSMC) scaffolds were prepared from 45S5 bioactive glass granules. Two tissue-engineered constructs were synthesized, namely the scaffold with bone marrow stromal cells seeded at the time of surgery (‘primary’), or the same scaffolds with culture expanded cells producing extracellular bone-like matrix (‘hybrid’). Defects were treated randomly with pSMC, primary, hybrid, or left untreated (‘sham’) to compare healing rates at 2, 4, and 12 weeks. The results demonstrated that at 2 weeks, the long bones treated with the hybrid and primary constructs had 40% more bone in the defect than bones with the pSMC. By 12 weeks, the surface modified bioactive glass scaffold was fully replaced by in situ formed calcium-phosphate. The ratio of transformed scaffold to defect size decreased significantly for all groups over time. The ultimate torque to failure and stiffness were affected by treatment and by time. By 2 weeks, defects treated with hybrid had the highest stiffness and this property was not significantly different from the one of intact bone (Fig. 14). Porous, surface modified bioactive ceramic integrated well with bone tissue in the healing of a skeletal defect site and resorbed in concert with bone formation, which allows for improvement of the long bone's structural integrity over time. The osteogenic activity of both tissue-engineered constructs, osteoprogenitor cells seeded onto scaffolds at the time of surgery, or cells expanded to form in vitro synthesized bone scaffold, was similar. Bone formation and the return of normal torsional properties were enhanced for the tissue-engineered constructs as compared to the scaffold alone [192]. 5.3.2. Bioresorbable sol–gel silica-based large biomolecules delivery systems Silica-based ceramic sol–gel materials with tunable porosities capable of controlled delivery of a broad range of hydrophilic and

Fig. 14. Histomorphometrical analysis of the percentage of bone in available pore/open space in the defect at 2, 4, and 12 weeks postimplantation. a) Hybrid is significantly greater than all groups at 2 weeks. b) Primary is significantly greater than PSMC and sham at 2 weeks. c) Significant increase from 2 to 4 weeks. d) Significant increase from 4 to 12 weeks. *The percentage of bone in available pore/open space in the defect (p b 0.05). Error bars are standard deviations [192]. Re-printed with permission from John Wiley & Sons, Inc.


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hydrophobic therapeutic agents such as growth factors, anti-oxidants and antibiotics has been developed [153]. The sol–gels can be prepared in various physical forms, including pellets, thin films or powders, in controlled sizes of 1 μm and larger. Bhattacharyya et al. [69] have studied room temperature processed mesoporous silica nanoparticles (MSN) for controlled release of various biological agents such as drugs, proteins and growth factors. Mesoporous silica nanoparticles included trypsin inhibitor (TI), a model molecule for TGF-β and BMP-2 as it has about the same molecular size [69]. Surface functionalization of MSN with polyethylene glycol helps to achieve quasi zero order release of incorporated TI for a duration up to 4 weeks (Fig. 15). Santos et al. demonstrated controlled release of bone morphogenetic protein (BMP-2) from room temperature processed sol–gel glass. They also documented pronounced effect of the calcium-phosphate surface layer that forms on the surface of the sol–gel on the function of the BMP-2 molecule [15]. In this study, primary cultures of rat stromal marrow cells (rMSCs) were seeded onto the surface of sol–gel glass disks. In addition, cells were cultured on the sol–gel in media supplemented with 10 ng BMP-2. Control groups included cells cultured on tissue culture plastic without and with 10 ng BMP added to the media. The results of this study demonstrated that sol–gel glass affects differentiation of cells with osteogenic potential. The study also suggests that the sol–gel substrate and BMP-2 act synergistically on rat stromal marrow cell differentiation in vitro [15]. The data further show that the normalized phosphatase activity was enhanced in contact with the sol–gel; it was most increased when the BMP was incorporated or adsorbed onto the sol–gel (Fig. 16). However, it is evident that cells on the sol–gel without BMP were more differentiated than cells grown on plastic with BMP, thereby demonstrating the additive effect of a bioactive substrate and BMP on osteoblast differentiation. The results of this study demonstrated that sol–gel glass on which a calcium phosphate layer develops affects differentiation of cells with osteogenic potential. The study also suggests that the sol–gel substrate and bone morphogenetic protein (BMP-2) act synergistically on rat stromal marrow cell differentiation in vitro [15]. 5.3.3. Mesoporous bioglass (MBG) scaffolds Various ions, including monovalent (Ag and Li), divalent (Sr, Zn, Mg, Cu and Co), trivalent (Ce, Ga, B and Fe) and tetravalent (Zr) ions, have been incorporated into MBG scaffolds by adding ionic salts into the

Fig. 16. Average normalized AP activity of rat marrow stromal cells (rMSC) for control © and experimental groups. Control groups: rMSCs cultured on tissue culture plastic without © and with 10 ng BMP (C-BT) added to the media. Experimental groups: rMSCs cultured on sol–gel glass disks: without BMP (SG); with 25 μg BMP incorporated in the glass (SG-BI); with BMP adsorbed onto the sol–gel surface (SG-BP); with 10 ng BMP added to the medium (SG-BT) [15]. Re-printed with permission from John Wiley & Sons, Inc.

MBG precursors during the synthesis step [105]. Ions released from MBG with the exception of silver (Ag) enhanced the proliferation and osteogenic differentiation of human bone marrow stromal cells (hBMSCs) [105], as well as in vivo bone formation [193]. Small molecules such as dexamethasone and dimethyloxallyl glycine can also be incorporated into silica materials. Dimethyloxallyl glycine can upregulate osteogenic and angiogenic differentiation of mesenchymal stem cells. MBG scaffolds have been used for effectively delivering these drugs to enhance bone regeneration application [146, 147]. The release of dexamethasone from MBG scaffolds enhanced alkaline phosphatase (ALP) activity and bone-related gene expression of osteoblasts in vitro [146,147]. A recent report showed that dimethyloxallyl glycine released from MBGs enhanced hypoxiainducible factor 1-alpha (HIF-1α) stabilization, vascular endothelial growth factor (VEGF) secretion, and the expression of bone-related gene by hBMSCs [148]. Growth factors, such as endothelial growth factor (VEGF) [142] and recombinant human bone morphogenetic protein-2 (rhBMP-2) [143], have been also loaded into MBG for controlled release and enhancing bone regeneration. An in vivo study with rhBMP-2 loaded MBG scaffolds showed a gradual MBG resorption and full replacement by new bone at 12 weeks [143]. 6. Biocompatibility 6.1. Biocompatibility of silica granules

Fig. 15. Cumulative TI release from PEG coated pore-expanded MSNs as a function of immersion time up to 4 weeks. (Error bars represent standard deviation (N = 3)). After 4 weeks of immersion the total TI release was 65% of the original protein content [69]. Reprinted from Acta Biomateriala, vol. 8, p3429–3435, Polymer coated mesoporous silica controlled release nanoparticles for macromolecules, Bhattacharyya S, Wang H, Ducheyne P. Copyright (2012), with permission from Elsevier.

The biocompatibility of bioactive glass has been well known for several decades [194,195]. The tissue response to various sol–gel processed silicon oxides has also been studied extensively. When implanted in bone (iliac crest of New Zealand white rabbit), silica xerogel generated a minimal inflammatory response. Extensive trabecular bone ingrowth around the silica was found at 2 weeks of implantation. A gradual resorption of xerogel granule over time was also found, with the size of xerogel granules gradually decreasing with implantation time (Fig. 17) [94]. When xerogels were implanted in non-osseous tissue, the silica xerogel did not cause tissue irritation and wound healing associated acute inflammation subsided within a 7 day span, a typical duration for any wound-healing. Two weeks after implantation, a fibrous capsule was present around the xerogel [16,83]. In these studies, a graduation resorption of xerogel was found and tissue necrosis was absent. To understand the bioresorption of resorbable silica glass, granules were implanted in two different sites: tibiae [196] and paraspinal muscle [197] of rabbits. After simplantation in the tibiae, the silicon concentration was measured in urine and blood for up to 7 months



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Fig. 17. Rabbit bone tissue response to (a, b) control, (c, d) silica xerogel and (e, f) xerogel composition 90% SiO2–5% CaO–5% P2O5 after implantation into the iliac crest defects: for 2 (a, c, e) and 4 (b, d, f). Magnification: 12× (a, e, f) and 10× (b–d). Progressing healing of the defect via trabecular bone growth was observed for both the control and the implant groups. Extensive trabecular bone growth occurred at the opening and the borders of the defects at 2 weeks of implantation. After 4 weeks, the amount and the density of the trabaculae increased for all the experimental groups [94]. Reprinted from Biomaterials, vol. 26, p1043–1052, In vivo tissue response to resorbable silica xerogels as controlled-release materials, Radin S, El-Bassyouni G, Vresilovic EJ, Schepers E, Ducheyne P. Copyright (2005), with permission from Elsevier.

after implantation. The urinary silicon was elevated for 24 weeks and the calculated average excretion rate was approximately 1.8 mg/day. Surrounding bone and muscle tissue as well as kidney, liver, lung, lymph nodes, and spleen were resected for chemical and histopathological analyses. There was no measurable increased concentration of silicon in any of these organs. The urine and blood samples for up to 6 months were again collected for analyzing the silicon release when resorbable silica granules were implanted in muscle. The elevated silicon concentration was recorded for 19 weeks with average excretion rate was 2.4 mg/day. No elevated concentrations of silicon were found at the implant sites at sacrifice. Statistically, all silicon had been excreted through the urine in both studies [197]. Fig. 18 shows the results of the tibiae implantation. 6.2. Biocompatibility of silica sol–gel nanoparticles Given the size of MSN (nanometer range) and the difference in sol– gel processing methodology, the biological response to these materials has recently been the subject of studies. Several reviews suggest that

Fig. 18. Glass granules implanted in rabbit tibiae. The implant group excreted an elevated amount of silicon over approximately 24 weeks. The total amount of silicon excreted was statistically the same as the total silicon content of the implanted granules. Error bars represent standard error of the mean [197]. Re-printed with permission from John Wiley & Sons, Inc.


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silica nanoparticles (sol–gels, xerogels, etc.) are biocompatible and degrade over time in the body [34,90,198,199]. Recently Fisichella et al. studied the cytotoxicity of MSN in HELA and astrocytes cells; they found that these materials are non-toxic up to a dose level of 100 μg/ml. More interestingly, they found that the most widely used MTT assay for evaluation of cytotoxicity produce falls positive results [200,201]. They suggested performing at least three different assay techniques for evaluation of cytotoxicity of MSN. The biocompatibility of MSN with and without surface functionalization has also been tested by different methods. The viability and proliferation of various mammalian cells indicated that this property was not affected by the internalization of MSN at concentrations below 100 μg/ml per 105 cells, even after 7 cell cycles. Cell membrane integrity is also preserved after uptake of MSN as determined by selective DNA staining followed by flow cytometry [202]. Microscopic analysis shows normal cell morphology upon uptake of MSN. Mitochondrial activity remains at normal levels as determined by colorimetric assay with 3-[4,5-dimethylthiazol-2-yl]-2,5-diphenyltetrazolium bromide (MTT) [203,204]. Growth rates for cells exposed to MSN are similar to the ones of cells grown in the absence of MSN [79,202]. Al-Salam et al. studied mice lung specimens incubated at 37 °C in Krebs–Henseleit buffer with and without 200 μg/ml of calcined MSN for 5–14 h [205]. They found normal alveolar morphology in all the studied specimens. There was no significant difference in the number of apoptotic cells between the treated and untreated samples. The particles were abundantly present in pneumocytes, macrophages, endothelial cells, fibroblasts, and interstitium. They were seen in different areas of the cytoplasm, suggesting intracellular movements [205]. They found that the presence of MSNs did not appear to disturb cellular configuration or micro-organelles. Due to their rigidity and surface charges, some were firmly attached to (indenting) the nuclear membrane. The rate of respiration (cellular mitochondrial O2 consumption, in μMO2/min/mg) in specimens exposed to 200 μg/ml particles for up to 12 h was the same as in untreated specimens. Their findings confirm “reasonable” bioavailability and biocompatibility of calcined mesoporous silicas with mouse lung tissue over at least 5–14 h of exposure time. Recently Malvindi et al. [206] performed a systematic in vitro study to assess the biological impact of SiO2 nanoparticles (SiO2NPs), by investigating 3 different sizes (25, 60 and 115 nm) and 2 surface charges (positive and negative) of the nanoparticles in 5 cell lines (3 in adherence and 2 in suspension). They analyzed the cellular uptake and distribution of the NPs along with their possible effects on cell viability, membrane integrity and generation of reactive oxygen species (ROS). Experimental results showed that all the investigated SiO2NPs did not induce detectable cytotoxic effects (up to 2.5 nM concentration) in all cell lines, and that cellular uptake was mediated by an endocytic process strongly dependent on the particle size and independent of its original surface charge, due to protein corona effects. They demonstrated that monodisperse and stable SiO2NPs are not toxic, revealing their promising potential in various biomedical applications. He et al. studied the in vivo biodistribution and urinary excretion of spherical mesoporous silica nanoparticles (MSNs) by tail-vein injection in ICR mice, and the effects of the particle size and PEGylation were investigated [97]. The results indicate that both MSNs and PEGylated MSNs of different particle sizes (80–360 nm) distribute mainly in the liver and spleen, a minority of them in the lungs, and a few in the kidney and heart. The PEGylated MSNs of smaller particle size escape more easily from capture by liver, spleen, and lung tissues, possess longer blood-circulation lifetime, and are more slowly biodegraded and correspondingly have a lower excreted amount of degradation products in the urine. Neither MSNs nor PEGylated MSNs cause tissue toxicity after 1 month in vivo [97]. Lieu et al. studied the systematic single and repeated dose toxicity, biodistribution and clearance of MSNs in vivo after intravenous injection in mice [207]. For single dose toxicity, lethal dose 50 (LD50) of

110 nm MSNs was higher than 1000 mg/kg. Further repeated dose toxicity studies indicated no death was observed when mice were exposed to MSNs at 20, 40 and 80 mg/kg by continuous intravenous administration for 14 days. Their results suggest no or very low toxicity of MSNs, based on the typical experiments with intravenous injection using a single or multiple doses. Inductively coupled plasma optical emission spectrometry (ICPeOES) and transmission electron microscopy (TEM) results show that the MSNs mainly accumulate in mononuclear phagocytic cells in liver and spleen [207]. Recently Fu et al. performed a systematic investigation of the absorption, distribution, excretion and toxicity of silica nanoparticles (SNs) with the average size of 110 nm after four different exposure routes including intravenous, hypodermic, intramuscular injection and oral administration to mice. The results demonstrated that SNs could be hardly absorbed in a short time after hypodermic and intramuscular administration. Pathological examinations illustrated that SNs possessed good tissue biocompatibility after oral and intravenous injection. These findings revealed that oral and intravenous administration were safe routes for possible biomedical application [208]. These findings are valuable for the future development of nanotechnology-based drug delivery systems.

7. Conclusion The excellent bio-degradable and biocompatible behavior of silica based materials has been well documented. In addition, silica materials can be prepared in a wide range of forms, including particles (nanoand micro-sized), thin films and 3D scaffolds. The incorporation of therapeutic molecules into silica based materials for controlled release enhances and expands the scope of the clinical use of silica materials. Controlled release silica materials have been extensively studied for the past two decades. Release kinetics have been tailored through various ways, including by functionalizing the pore walls, modifying the surface, and adjusting the pore parameter. Undoubtedly, the combination of two benefits, namely controlled drug release and excellent biocompatibility, provides a strong stimulus to propel silica based materials along toward clinical application. Pre-clinical studies with silica controlled release materials are well underway. In particular, pre-clinical studies for the treatment and prevention of infection, management of surgical pain, tissue engineering treatments, and cancer treatment stand out, as these fields increasingly appreciate the controlled, sustained, or targeted delivery of therapeutic agents. Among these fields, controlled release silica materials for the treatment of orthopedic infections and for tissue engineering treatments are close to the clinic. Moving forward, clinical studies will be undertaken to ensure the successful application of controlled release silica materials for the benefit of patients.

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