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Nanomaterials: the next step in injectable bone cements
Injectable bone cements (IBCs) are biocompatible materials that can be used as bone defect fillers in maxillofacial surgeries and in orthopedic fracture treatment in order to augment weakened bone due to osteoporosis. Current clinically available IBCs, such as polymethylmethacrylate and calcium phosphate cement, have certain advantages; however, they possess several drawbacks that prevent them from gaining universal acceptance. New gel-based injectable materials have also been developed, but these are too mechanically weak for load-bearing applications. Recent research has focused on improving various injectable materials using nanomaterials in order to render them suitable for bone tissue regeneration. This article outlines the requirements of IBCs, the advantages and limitations of currently available IBCs and the state-of-the-art developments that have demonstrated the effects of nanomaterials within injectable systems.
Young Jung No1, Seyed-iman Roohani-Esfahani1 & Hala Zreiqat*,1 Biomaterials & Tissue Engineering Research Unit, School of AMME, The University of Sydney, Sydney 2006, Australia *Author for correspondence: hzreiqat@ usyd.edu.au 1
Keywords: bioactive materials • bone tissue engineering • injectable bone cement • injectable materials • nanocomposites
Injectable bone cement (IBC) is a class of materials that are initially in either liquid or paste form, which can then either be injected through a channel or molded into shape. After a certain period of time, the IBC solidifies in order to take the shape of the site of implantation and aids the regeneration of new bone tissue. IBCs are commonly used to augment and stabilize weakened or injured bone due to diseases such as osteoporosis or trauma in order to allow screw fixation during the treatment of fractures (Figure 1) [1] . Osteoporosis results in approximately 9 million fractures annually worldwide [2] , with fractures commonly occurring in the vertebrae, proximal femur and the distal radius [3] . The number of osteoporotic fractures requiring surgical treatment is projected to increase mainly due to the general aging of the population [4,5] . Ideally, IBCs should weave themselves into the surrounding cancellous bone and take the shape of the defect upon injection. IBCs should not interfere with the healing process, should provide strong mechanical support
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and should facilitate an ideal biological environment in which new bone can grow into the site of the implant. Almost all of the IBC materials currently in clinical use or those under the subject of research can be classified into three main categories: polymethylmethacrylate (PMMA); calcium phosphate-based bone cement (CPC); and calcium sulfate-based bone cement (CSC). IBCs consist of two main components: a separate powder component as the base material and a liquid part that also acts as a reaction medium. Upon mixing, the liquid and powder chemically react with each other (polymerization for PMMA and crystallization for CPCs and CSCs) in a time-dependent manner until they form a solid structure in vivo. Novel injectable gelbased systems with varying mechanisms for in vivo solidification have also been developed [6,7] . Although these materials were initially designed for localized drug delivery, they have expanded into other applications, such as bone and cartilage repair. Recent
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Fractured femoral head
Figure 1. Schematic of femoral head screw fixation. (A) In good, healthy bone stock, the fixation of a fracture is maintained with metallic screws. (B) In osteoporotic bone, there is a risk of implant loosening and displacement of the fracture. (C) The surrounding osteoporotic bone is strengthened using injectable bone cement (green), reducing the risk of fracture displacement. For color figure, please see online at http://www.futuremedicine.com/ doi/full/10.2217/NNM.14.109
research has focused on incorporating nanoparticles and nanofibers into various injectable materials. This article will discuss the requisite properties of IBCs from a materials perspective and elaborate on the strengths and limitations of the clinically available IBCs. Furthermore, this article will look at the implementation of nanotechnology in order to address the current IBC limitations. Requirements of IBCs An ideal IBC for use in bone defects, and particularly in load-bearing orthopedic applications, should possess the following properties [8,9] : • Easy injectability and consistent homogeneity throughout the injection; • Appropriate setting time after mixing; • Low risk of necrosis; • Adequate tensile, compressive and shear strengths,
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depending on the site of injection (i.e., degree of load bearing and degree of motion, among other factors); • Similar stiffness to the surrounding bone; • High radiopacity, so that the IBC can be readily distinguished from the surrounding soft and hard tissues under conventional x-ray imaging; • Bioactivity; • A resorption rate similar to that of tissue formation; • Sufficient amount of microporosity (100 μm diameter) for nutrient/ waste transfer, angiogenesis and osseointegration during the bone healing process. Each clinically available IBC material is graded relative to each other according to these properties, as listed in Table 1. Firstly, IBCs should be injectable (i.e., be able to flow and pass through a narrow channel and
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μ: Microporous; C: Stronger than cancellous bone in compressive strength; CPC: Calcium phosphate-based bone cement; H: High; L: Low; M: Moderate; N: None; PMMA: Polymethylmethacrylate; T: Stronger than cancellous bone in tensile strength; Y: Yes.
μ H H L Y L Y Calcium sulfate-based (BonePlast®, Y MIIG®)
L
μ M H C Y L Y Y Brushite-based (CPC) (ChronOs®)
L
μ L H C Y L Y Apatite-based (CPC) (Norian® SRS, Y BoneSource®, HydroSet, Calcibon®, α-BSM®, Biopex®)
L
μ N H T, C Y H Y Bisphenol A glycidyl methacrylate- Y based (Cortoss®)
M
μ N H T, C N H Y Y PMMA-based (Simplex® P, PALACOS® LV40, CMW®, CranioplasticTM )
M
Stiffness Resorbability Mechanical strength Bioactivity Radiopacity Risk of necrosis Appropriate setting time Injectable Bone cement
Table 1. Relative grading of clinically available injectable bone cements according to their physical, chemical and mechanical properties.
Porosity
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then fill the intended gap upon ejection). If there are initially different phases in the IBC prior to mixing (e.g., powder and liquid components in PMMA and CPC), these phases should not separate, but rather be homogeneously mixed throughout the whole injection procedure. Once the IBC takes the shape of the defect and interlocks with the pores of the surrounding cancellous bone, it should begin to set. It is imperative that the setting time of IBCs falls within a narrow range. Typically, setting times for clinically available IBCs vary by between 5 and 15 min [8,10–11] . Faster setting times within this range are ideal for reducing operation times and quick implant stabilization, but require experience from the surgeon in order to prevent the cement from setting prior to adequate defect filling. Slower setting times are ideal for surgeons who require a certain degree of flexibility, although longer setting times above approximately 15 min are undesirable, as the material will not be able to withstand surrounding stresses and retain its shape after injection. The currently available IBCs listed in Table 1 possess several drawbacks that prevent them from gaining universal approval for use in load-bearing orthopedic applications (Figure 2) . Ideally, IBCs should be both bioactive and bioresorbable for eventual host tissue regeneration at the site of the implant. However, there appears to be a trade-off between bioresorbability and mechanical stability among the currently available IBCs. The order of bioresorbability rate from highest to lowest, and subsequently the mechanical strength from weakest to strongest, is as follows: CSC > brushitebased CPC > apatite-based CPC >> PMMA (PMMA is not resorbable and remains as a permanent fixture). It has been suggested that an ideal IBC should have a “resorption rate that is neither too high nor too low” [8] . There is good agreement in literature that an IBC with a high degradation rate, such as a CSC, is not suitable because it does not provide short-term in vivo mechanical stability due to having a faster degradation rate than that of bone ingrowth [12,13] . This could be addressed by incorporating bioactive particles with slower degradation profiles, such as hydroxyapatite, certain bioactive glasses and bioactive ceramic powders. Urban et al. demonstrated that, with the incorporation of a certain amount of calcium phosphate into CSCs, they managed to significantly delay the degradation profile and increase compressive strength, as well as improve bone formation in in vivo canine models [14] . Lin et al. also showed that, by adding calcium silicate particles into CSCs, they were able to delay in vitro degradation and show more new bone formation compared with pure CSCs in in vivo rat models [15] . Studies have shown that the growth of newly formed bone is limited by the amount of space left after the degradation of the cement
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Fibrous tissue Bone cement
Cancellous bone
Inflammation due to cement fragments
Crack
Stress/load application
Stiff/rigid
Heat
x-ray x-ray
x-ray x-ray
x-ray
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Radiolucent cement
New bone formation
Cement dissolution/ resorption
Figure 2. Drawbacks of currently available injectable bone cements. (A) High stiffness of injectable bone cement results in adjacent cancellous bone fracture. (B) Low radiopacity makes injectable bone cement difficult to detect with x-ray radiography and fluoroscopy. (C) Lack of bioactivity (i.e., polymethylmethacrylate [PMMA] and most gels) leads to fibrous tissue formation surrounding the implant and, ultimately, to implant loosening. (D) Heat generated from the polymerization of PMMA can lead to necrosis of surrounding bone tissue. (E) Low tensile/ compressive strength, resulting in brittle fractures (PMMA/calcium phosphate-based bone cement/calcium sulfate-based bone cement) or significant plastic deformation (gel-based). Fragments can generate inflammation. (F) Rapid degradation leads to mechanical instability at the injection site (calcium sulfate-based bone cement and some gels). The rate of new bone growth may not be adequate to stabilize the defect. [16–18] .
Nevertheless, a restriction of new bone growth due to a lack of space alone does not appear to be detrimental to the host, as long as the bone cement remains intact and forms a bioactive bond with the host bone
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[16] .
IBCs must remain mechanically robust postinjection; it should not plastically deform or fragment under multidirectional loading. Fragments of PMMA of 1–12 μm in diameter were shown to be engulfed by macro-
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phages; fragments larger than 12 μm had macrophages and fibroblasts adhered to their surfaces, but were not engulfed [19] . Phagocytosis of PMMA fragments leads to the increased production of TNF, which also acts as a bone resorption mediator, as well as inhibition of DNA synthesis of the macrophage, whereas macrophages exposed to PMMA particles larger than 1 mm showed no difference in the production of TNF and DNA synthesis compared with the macrophages not exposed to PMMA particles [19] . In addition, it has been shown that 8 weeks after CPC was injected into the interbody regions of sheep lumbar spine, the presence of CPC fragments as a result of shear fracture led to severe inflammatory responses, with extensive fibrous tissue formation and CPC particles engulfed by macrophages, whereas specimens with good intervertebral body fusion showed no inflammation [20] . It is therefore likely that early fragmentation and the release of microparticles, particularly from brittle materials such as PMMA and CPC, will result in undesired inflammatory reactions. In essence, IBCs should be designed so that they are able to withstand compressive and shear forces so as not to fragment, and/or the osseointegration of the new bone into the IBC should be accelerated in order to prevent early fragmentation [20] . Clinically available IBCs Polymethylmethacrylate
PMMA-based bone cements generally consist of: a powder component comprising PMMA; a radiopacifying agent; and a liquid component comprising the liquid monomer methacrylate. Upon mixing, a doughy paste is formed that can be injected into the site of implantation. After injection, the paste will be fully polymerized and take the shape of the defect. PMMA cements are biocompatible and can be made highly radiopaque by incorporating inorganic compounds, such as zirconium oxide (ZrO2) and barium sulfate (BaSO4), in order to track the flow of bone cement via x-ray fluoroscopy [21] . Commercially available PMMA bone cements often distinguish themselves by the type of incorporated radiopacifying agent, such as BaSO4 and ZrO2, as well as antibiotics, such as gentamicin [8] . Due to their favorable mechanical properties and robustness compared with other available IBCs, PMMA-based bone cements remain the most common injectable materials used in the clinical setting for filling bone defects, particularly for arthroplasty, vertebroplasty and screw augmentation [22,23] . The tensile and compressive strengths of PMMA could reach up to 49 and 114 MPa, respectively [24] , whereas the tensile and compressive strengths of human cancellous bone range from 1 to 5 MPa and 4 to 12 MPa, respectively [25] . The mechanical strength of IBCs should be
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higher than that of the surrounding cancellous bone in order to safely withstand complex, multidirectional loading without plastic deformation or brittle failure at the site of the bone defect. A detailed comparison of the mechanical properties with other materials used for IBCs is presented in Figure 3 [24–29] . Despite its use for more than 50 years in clinical applications, PMMA bone cements have major drawbacks; they do not form bioactive bonds with the surrounding trabecular bone, nor are they resorbed by the body. PMMA-based cements form regions of permanent mechanical fixation as they intertwine through the trabecular pores upon injection [30] , with fibrous tissue eventually encapsulating the surface of the PMMA bone cement. In sheep vertebrae models, PMMA bone cement fillers failed to fuse properly with the surrounding trabecular bone, leading to the formation of a fibrous membrane surrounding the bone cement, which in turn resulted in micromovements and subsequent inflammatory reactions, culminating in a reduction in the vertebral strength and stiffness over a time course of 24 weeks [18] . PMMA surfaces can be rendered bioactive by incorporating a high concentration (~70%) of bioglass powders, as evident by the absence of a fibrous tissue layer at the bone–PMMA interface [31,32] . However, the high particle loading compromises the efficient injection of PMMA. By contrast, lower particle concentrations reduce the exposure to the bioactive components of the bioglass, and hence compromise the overall bioactivity of the PMMA [33] . PMMA surfaces incorporated with bioglass particles were found to be more effective in forming a bioactive bond with the surrounding bone when compared with those incorporated with hydroxyapatite [31] . More recent studies have demonstrated the bioactivity (both in vitro and in vivo) of PMMA incorporated with γ-methacryloxypropyltrimethoxy-silane and calcium acetate [34,35] . This material showed bioactive behavior in both dog and rabbit models, and also displayed a significant reduction in the released heat during polymerization [34,35] compared with that for the unmodified PMMA. Another major drawback of PMMA bone cements is their high stiffness compared with cancellous bone (Figure 3) , resulting in stress-shielding effects and, ultimately, weakening of the surrounding host bone, such as fractures occurring in adjacent vertebrae after undergoing vertebroplasty [36] . Efforts have been directed towards the development of low-modulus PMMA cements; Boger et al. developed an injectable macroporous PMMA (porosity of 56% and average pore size of 260 μm) by incorporating a 40% v/v fraction of 2% w/w sodium hyaluronate solution into the bone cement, reducing the modulus from approxi-
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PMMA
Cortical bone
Unreinforced CPC
CNT-CPC
Strength (MPa)
10
Cancellous bone Polymer fiber-CPC Gel-based Calcium sulfate
1.0
CPC group (E = 1-3 GPa)
Compressive strength Tensile strength 0.1 0.01
0.1
1.0
10
Elastic modulus (GPa) Figure 3. Elastic modulus and strength of injectable systems and bone tissue. Solid rectangles indicate compressive strength, dotted rectangles indicate tensile strength. The three CPC types (unreinforced, polymer– fiber composites and CNT composites) possess elastic moduli in the same range (1– 3 GPa) and have been separated for visual clarity. CNT: Carbon nanotube; CPC: Calcium phosphate-based bone cement; E: Elastic modulus; PMMA: Polymethylmethacrylate.
mately 1.8 GPa to 0.29 GPa, bringing it within the average stiffness of cancellous bone [37] . However, the compressive yield strength significantly dropped from 91 to 7.4 MPa, and while this is within the range of cancellous bone, it is inadequate for load-bearing orthopedic applications. Clinical incidence of bone cement implantation syndrome (BCIS), which is characterized by a rapid decrease in systemic blood pressure and, in rare cases, cardiac arrests, is another reported side effect of using PMMA bone cements [38] . Initially, liquid monomers were thought to play a key role in BCIS. However, it has been speculated that BCIS could be due to the effects of other factors, such as an increase in the intramedullary pressure, leading to embolization [38] . Another concern relates to altering the immune response and subsequent wound healing process due to the presence of unreacted monomer components,
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which are known to be toxic [39] . Due to the exothermic reaction of the polymerization process, there is a risk of thermal necrosis in the surrounding bone tissue, where the temperature of the polymerization was reported to reach 74°C in some clinically available PMMA IBCs [40] , which is significantly higher than 50°C, at which level bone tissue necrosis is reported to occur if exposure occurs for more than 30 s [41] . However, some researchers have argued that the temperature at the PMMA–bone interface does not exceed 50–54°C [41,42] and that the human body can act as a heat sink and therefore render the thermal effect on the surrounding cells clinically insignificant and insufficient to cause extensive thermal injury [41–43] . Nevertheless, the heat produced during the polymerization of PMMA is undesirable and should be minimized in order to reduce the risk of thermal injury.
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Nanomaterials: the next step in injectable bone cements
Calcium phosphate & calcium sulfate cements
Similarly to PMMA, CPCs and CSCs consist of a powder and a liquid (water for CPC and CSC or sodium phosphate solution for CPCs). Upon mixing, a series of crystallization reactions occur that result in the solidification of the cement. The differences between the CPC products are due to the different initial reactants used (which include α-tricalcium phosphate, β-tricalcium phosphate, anhydrous dicalcium phosphate and monocalcium phosphate monohydrate), producing either an apatite-based or brushite-based cement [44] . CPCs are bioactive and can possess various degrees of bioresorbability that, if sufficient, lead to new bone formation, as the CPC degrades within the defect. The setting reactions of CPCs are isothermal or slightly exothermal and hence lead to a minimal risk of thermal necrosis [43,44] . The hydroxyapatite layer that forms upon the setting of CPCs in vivo [8,10–11,26] is thought to be a contributing factor to the bioactivity and osteoconductivity of the CPCs by allowing a direct bonding of CPCs to the surrounding bone tissue without eliciting an immune reaction. By contrast, CSCs are not bioactive, and limited inflammatory responses have been reported [13] . Nevertheless, it has been claimed that the rapid release of calcium and sulfate ions faciliates the growth of new bone into the defect region [13] . The main disadvantage of CPCs and CSCs compared with PMMA-based bone cements is that they have inadequate tensile strength (Figure 3) . CPCs are weak against tensile, bending, shear and torsional stresses, and its tensile strength values are rarely reported. The three-point bending strength of apatitebased CPCs have been reported to be between 10 and 15 MPa [27] , and the tensile and shear strengths have been reported to be between 3 and 10 MPa [26] . In comparison, the flexural and tensile strengths in bulk, sintered hydroxyapatite range from 7.7 to 113 MPa and 18 to 130 MPa, respectively, and are heavily dependent on the presence of pores [25] . The clinical use of CPCs is applied in regions with primarily compressive loading, including the treatment of defects at the tibial plateau and the calcaneus, or where multidirectional loading is negligible, such as the distal radius [1,26] . Due to the weaknesses in tensile and shear strengths, the clinical applications of CPCs have remained largely limited to screw augmentation of the femoral head and vertebral compression fracture treatment via percutaneous injection [18,20,45–48] . Efforts have been directed toward increasing the tensile strength of CPCs through the fabrication of CPC–fiber composites [27,49–51] . The incorporation of fibers such as Vicryl® (Ethicon, NJ, USA) and carbon nanotubes (CNTs) into CPCs enhanced their flexural strength and the work of fracture is directly depen-
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dent on the fiber content, length and strength [27,49] . Buchanan et al. demonstrated that the incorporation of polypropylene fibers into CPC resulted in an increase in the work of fracture and setting time, as well as a decrease in the compressive strength and modulus [50] . However, the major drawbacks of the currently available fibers include low fiber–matrix interface bonding and a lack of bioresorbability and bioactivity. Ideally, the incorporated fibers themselves should be able to support or enhance the growth of new bone. The challenge is to find suitable fibers that are bioresorbable and bioactive while providing a sufficient shear strength and fracture toughness [27] . In order for fibers to augment the mechanical strength of ceramic matrices, there must be sufficient interfacial bonding between the fibers and the ceramic matrix [52,53] . CPCs have a similar chemical composition to human bone and hence a similar radiopacity. Therefore, it is difficult to obtain a good contrast between the bone and CPC with x-ray imaging, especially where there is excessive soft tissue [54] . Unlike PMMA, incorporating such inorganic nonresorbable radiopaque materials into resorbable CPCs may not be ideal, as the radiopaque compounds, such as BaSO4 and ZrO2, have been shown to remain permanently fixed within the remodeled bone [54,55] . Neither of the previous studies, however, reported adverse histological reactions arising from the in vivo incorporation of these nonresorbable inorganic particles (up to 24 months for BaSO4, rabbit model [54] ; up to 12 weeks for ZrO2, rat model [55]), suggesting that it is safe – at least in the relatively short term – to incorporate nonresorbable compounds into CPCs, provided that these compounds do not induce cytotoxicity or immunogenesis. However, long-term in vivo studies are required in order to establish a definitive conclusion regarding the safety of these radiopaque CPC compounds. Calcium silicate cements
Calcium silicate cement (CSiC) is another group of injectable materials that have been used primarily in endodontic treatment. In particular, mineral trioxide aggregate (MTA) is a CSiC that has been approved by the US FDA since the late 1990s for dental pulp repair. MTA consists of tricalcium silicate, dicalcium silicate, tricalcium aluminate (TCA), tetracalcium aluminoferrite and calcium sulfate, with bismuth oxide as the radiopacifying agent, and when all of the constituents are mixed with water, it forms a paste that can be molded to fit the site of implantation [56,57] . Although sintered and preformed calcium silicate ceramic scaffolds have shown promising results for implantation [58,59] , there are major obstacles that need to be overcome if MTAs are to be modified for
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Review No, Roohani-Esfahani & Zreiqat orthopedic applications. Although the final compressive strength for MTA is reported to be approximately 70 MPa, which is relatively high compared with other injectable systems, the setting time is long (∼4 h) and the setting reaction is highly alkaline, reaching up to pH values of 12.5 [56,57] towards the beginning of the reaction and decreasing to pH 8.5 by day 28 [57] . Traditional MTAs have been shown to form surface apatite when immersed in simulated body fluid and other phosphatecontaining solutions [60] , and have been demonstrated to support the in vitro differentiation of human orofacial bone mesenchymal stem cells [61] . MTA in vivo rodent model studies have shown some cases of direct new bone apposition, but most cases have also shown fibrous tissue formation between the MTA and new bone formation, despite the ‘bioactivity’ deduced from the apparent formation of in vitro surface apatite [60] . The setting time of MTA was shown to be reduced from >2 h to approximately 40 min by increasing the concentration of TCA to 15 wt%. However, 15 percentage by weight (wt%) of the TCA hindered the development of its compressive strength, with its maximum strength reaching up to approximately 40 MPa, compared with approximately 55 MPa for MTA with 0 wt% TCA [62] . A novel CSiC system that is different to the MTA system but still has calcium silicate as its major component has been developed, which is more directed towards potential orthopedic applications. Ding et al. have developed a CSiC consisting of SiO2 and CaO powders mixed with ammonium phosphate solution (3.7 M, pH = 7.4) [63] or water [64,65] . Their CSiC samples had clinically acceptable setting times of between 3.3 and 8.5 min [63] , and the pH values of the CSiC samples in culture medium were later shown to reach pH 7.9 after 24 h [64] . In addition, higher proliferation and osteogenic differentiation were observed in human osteosarcoma cell cultures on the CSiC samples compared with cultures with no cement [63] . Adding 5 wt% gelatin and/or chitosan oligosaccharides to a similar CSiC system demonstrated improved proliferation and osteogenic differentiation of human mesenchymal stem cells, but slightly reduced its diametral tensile strength [65] . However, this CSiC system is too weak for orthopedic applications, with diametral tensile strength not exceeding 3.0 MPa [63] , and strategies to improve its mechanical strength should be explored. Gel-based injectable systems based on in situ polymer cross-linking In situ polymeric cross-linking gel-based injectable systems such as polyethylene glycol (PEG)–polycaprolactone (PCL)–PEG, poly(lactic-co-glycolic acid)–PEG, polyvinyl alcohol–acrylamide and alginate are relatively novel materials that are designed primarily for the pur-
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pose of controlled drug-delivery applications [66,67] . These materials are injected directly at the site of action, which allows for increased efficiency and bioavailability of the drug and hence lower doses in order to reduce the risk and intensity of potential side effects. However, they also offer a promising potential for use as IBCs. In contrast to traditional injectable cements, these materials have a unique in situ cross-linking ability that can be initiated by either temperature, pH, light or cationic cross-linkers [6,7] . They remain as a low-viscosity solution for easy injection, but after injection and under physiological conditions, cross-linking starts and the material solidifies. Although these injectable materials are not yet suitable for load-bearing orthopedic applications due to their low inherent compressive strengths and stiffnesses, which do not exceed 6 and 10 MPa in terms of compressive strength and modulus, respectively, and so are significantly below those of PMMA and CPCs (Figure 3), some have been identified for potential use in bone tissue repair as a carrier for the delivery of growth factors, drugs and cells directly to the defect site in order to enhance the growth of new bone (Table 2). Most of the gels indicated in Table 2 have been shown to support the growth and osteogenic differentiation of mesenchymal stem cells [29,68–70] and also in vivo new bone tissue formation in rabbit models [71] . Furthermore, the gels possess interconnected microporous structures that are initially filled with water or other liquid molecules at the time of injection, which can then subsequently act as in situ porous channels for efficient nutrient and waste transfer during the bone regeneration process [72] . The role of nanomaterials in IBCs Recently, nanomaterials have been incorporated into existing IBCs and other novel injectable materials in order to improve their properties for use in bone augmentation and bone tissue regeneration (Supplementary Table 1). Figure 4 outlines the ways in which nanomaterials in IBCs can aid the regeneration of bone. Nanoparticles can: improve mechanical properties, such as tensile, compressive and shear strengths and fracture toughness [75,76] ; release ions or drugs into the surrounding environment in order to enhance osteoblast recruitment, adhesion, proliferation and differentiation [77,78] ; enhance surface energy and protein adsorption and influence integrin binding; and act as radiopacifiers for enhanced visual detection under x-ray imaging. Nanoparticle leaching for improved biological properties Chemical composition
Bioactive ions such as calcium (Ca 2+), silicon (Si2+), magnesium (Mg2+) and strontium (Sr2+), or certain proteins and drugs, can improve the osteoblast recruitment,
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Table 2. Gels systems that have potential for use as injectable materials for bone regeneration. Material
Properties suitable for injectable materials in bone regeneration
Ref.
PEG–PCL–PEG triblock copolymer plus acellular bone matrix granule composite
Temperature-induced gelling at 37°C Biodegradable hydrogel Low cytotoxicity against rat MSCs Inflammation reduced after 4 weeks of implantation in BALB/c mice
[70]
PLGA/PEG + hyaluronic acid hydrogel
Temperature-induced scaffold assembly PLGA particles sintered at 37°C PLGA–PEG particles injected within a hyaluronic acid hydrogel Total porosity approximately 50% Compressive strength
Nanomaterials: the next step in injectable bone cements
Injectable bone cements (IBCs) are biocompatible materials that can be used as bone defect fillers in maxillofacial surgeries and in orthopedic fracture treatment in order to augment weakened bone due to osteoporosis. Current clinically available IBCs, such as polymethylmethacrylate and calcium phosphate cement, have certain advantages; however, they possess several drawbacks that prevent them from gaining universal acceptance. New gel-based injectable materials have also been developed, but these are too mechanically weak for load-bearing applications. Recent research has focused on improving various injectable materials using nanomaterials in order to render them suitable for bone tissue regeneration. This article outlines the requirements of IBCs, the advantages and limitations of currently available IBCs and the state-of-the-art developments that have demonstrated the effects of nanomaterials within injectable systems.
Young Jung No1, Seyed-iman Roohani-Esfahani1 & Hala Zreiqat*,1 Biomaterials & Tissue Engineering Research Unit, School of AMME, The University of Sydney, Sydney 2006, Australia *Author for correspondence: hzreiqat@ usyd.edu.au 1
Keywords: bioactive materials • bone tissue engineering • injectable bone cement • injectable materials • nanocomposites
Injectable bone cement (IBC) is a class of materials that are initially in either liquid or paste form, which can then either be injected through a channel or molded into shape. After a certain period of time, the IBC solidifies in order to take the shape of the site of implantation and aids the regeneration of new bone tissue. IBCs are commonly used to augment and stabilize weakened or injured bone due to diseases such as osteoporosis or trauma in order to allow screw fixation during the treatment of fractures (Figure 1) [1] . Osteoporosis results in approximately 9 million fractures annually worldwide [2] , with fractures commonly occurring in the vertebrae, proximal femur and the distal radius [3] . The number of osteoporotic fractures requiring surgical treatment is projected to increase mainly due to the general aging of the population [4,5] . Ideally, IBCs should weave themselves into the surrounding cancellous bone and take the shape of the defect upon injection. IBCs should not interfere with the healing process, should provide strong mechanical support
10.2217/NNM.14.109 © 2014 Future Medicine Ltd
and should facilitate an ideal biological environment in which new bone can grow into the site of the implant. Almost all of the IBC materials currently in clinical use or those under the subject of research can be classified into three main categories: polymethylmethacrylate (PMMA); calcium phosphate-based bone cement (CPC); and calcium sulfate-based bone cement (CSC). IBCs consist of two main components: a separate powder component as the base material and a liquid part that also acts as a reaction medium. Upon mixing, the liquid and powder chemically react with each other (polymerization for PMMA and crystallization for CPCs and CSCs) in a time-dependent manner until they form a solid structure in vivo. Novel injectable gelbased systems with varying mechanisms for in vivo solidification have also been developed [6,7] . Although these materials were initially designed for localized drug delivery, they have expanded into other applications, such as bone and cartilage repair. Recent
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Fractured femoral head
Figure 1. Schematic of femoral head screw fixation. (A) In good, healthy bone stock, the fixation of a fracture is maintained with metallic screws. (B) In osteoporotic bone, there is a risk of implant loosening and displacement of the fracture. (C) The surrounding osteoporotic bone is strengthened using injectable bone cement (green), reducing the risk of fracture displacement. For color figure, please see online at http://www.futuremedicine.com/ doi/full/10.2217/NNM.14.109
research has focused on incorporating nanoparticles and nanofibers into various injectable materials. This article will discuss the requisite properties of IBCs from a materials perspective and elaborate on the strengths and limitations of the clinically available IBCs. Furthermore, this article will look at the implementation of nanotechnology in order to address the current IBC limitations. Requirements of IBCs An ideal IBC for use in bone defects, and particularly in load-bearing orthopedic applications, should possess the following properties [8,9] : • Easy injectability and consistent homogeneity throughout the injection; • Appropriate setting time after mixing; • Low risk of necrosis; • Adequate tensile, compressive and shear strengths,
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depending on the site of injection (i.e., degree of load bearing and degree of motion, among other factors); • Similar stiffness to the surrounding bone; • High radiopacity, so that the IBC can be readily distinguished from the surrounding soft and hard tissues under conventional x-ray imaging; • Bioactivity; • A resorption rate similar to that of tissue formation; • Sufficient amount of microporosity (100 μm diameter) for nutrient/ waste transfer, angiogenesis and osseointegration during the bone healing process. Each clinically available IBC material is graded relative to each other according to these properties, as listed in Table 1. Firstly, IBCs should be injectable (i.e., be able to flow and pass through a narrow channel and
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μ: Microporous; C: Stronger than cancellous bone in compressive strength; CPC: Calcium phosphate-based bone cement; H: High; L: Low; M: Moderate; N: None; PMMA: Polymethylmethacrylate; T: Stronger than cancellous bone in tensile strength; Y: Yes.
μ H H L Y L Y Calcium sulfate-based (BonePlast®, Y MIIG®)
L
μ M H C Y L Y Y Brushite-based (CPC) (ChronOs®)
L
μ L H C Y L Y Apatite-based (CPC) (Norian® SRS, Y BoneSource®, HydroSet, Calcibon®, α-BSM®, Biopex®)
L
μ N H T, C Y H Y Bisphenol A glycidyl methacrylate- Y based (Cortoss®)
M
μ N H T, C N H Y Y PMMA-based (Simplex® P, PALACOS® LV40, CMW®, CranioplasticTM )
M
Stiffness Resorbability Mechanical strength Bioactivity Radiopacity Risk of necrosis Appropriate setting time Injectable Bone cement
Table 1. Relative grading of clinically available injectable bone cements according to their physical, chemical and mechanical properties.
Porosity
Nanomaterials: the next step in injectable bone cements
Review
then fill the intended gap upon ejection). If there are initially different phases in the IBC prior to mixing (e.g., powder and liquid components in PMMA and CPC), these phases should not separate, but rather be homogeneously mixed throughout the whole injection procedure. Once the IBC takes the shape of the defect and interlocks with the pores of the surrounding cancellous bone, it should begin to set. It is imperative that the setting time of IBCs falls within a narrow range. Typically, setting times for clinically available IBCs vary by between 5 and 15 min [8,10–11] . Faster setting times within this range are ideal for reducing operation times and quick implant stabilization, but require experience from the surgeon in order to prevent the cement from setting prior to adequate defect filling. Slower setting times are ideal for surgeons who require a certain degree of flexibility, although longer setting times above approximately 15 min are undesirable, as the material will not be able to withstand surrounding stresses and retain its shape after injection. The currently available IBCs listed in Table 1 possess several drawbacks that prevent them from gaining universal approval for use in load-bearing orthopedic applications (Figure 2) . Ideally, IBCs should be both bioactive and bioresorbable for eventual host tissue regeneration at the site of the implant. However, there appears to be a trade-off between bioresorbability and mechanical stability among the currently available IBCs. The order of bioresorbability rate from highest to lowest, and subsequently the mechanical strength from weakest to strongest, is as follows: CSC > brushitebased CPC > apatite-based CPC >> PMMA (PMMA is not resorbable and remains as a permanent fixture). It has been suggested that an ideal IBC should have a “resorption rate that is neither too high nor too low” [8] . There is good agreement in literature that an IBC with a high degradation rate, such as a CSC, is not suitable because it does not provide short-term in vivo mechanical stability due to having a faster degradation rate than that of bone ingrowth [12,13] . This could be addressed by incorporating bioactive particles with slower degradation profiles, such as hydroxyapatite, certain bioactive glasses and bioactive ceramic powders. Urban et al. demonstrated that, with the incorporation of a certain amount of calcium phosphate into CSCs, they managed to significantly delay the degradation profile and increase compressive strength, as well as improve bone formation in in vivo canine models [14] . Lin et al. also showed that, by adding calcium silicate particles into CSCs, they were able to delay in vitro degradation and show more new bone formation compared with pure CSCs in in vivo rat models [15] . Studies have shown that the growth of newly formed bone is limited by the amount of space left after the degradation of the cement
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Fibrous tissue Bone cement
Cancellous bone
Inflammation due to cement fragments
Crack
Stress/load application
Stiff/rigid
Heat
x-ray x-ray
x-ray x-ray
x-ray
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Radiolucent cement
New bone formation
Cement dissolution/ resorption
Figure 2. Drawbacks of currently available injectable bone cements. (A) High stiffness of injectable bone cement results in adjacent cancellous bone fracture. (B) Low radiopacity makes injectable bone cement difficult to detect with x-ray radiography and fluoroscopy. (C) Lack of bioactivity (i.e., polymethylmethacrylate [PMMA] and most gels) leads to fibrous tissue formation surrounding the implant and, ultimately, to implant loosening. (D) Heat generated from the polymerization of PMMA can lead to necrosis of surrounding bone tissue. (E) Low tensile/ compressive strength, resulting in brittle fractures (PMMA/calcium phosphate-based bone cement/calcium sulfate-based bone cement) or significant plastic deformation (gel-based). Fragments can generate inflammation. (F) Rapid degradation leads to mechanical instability at the injection site (calcium sulfate-based bone cement and some gels). The rate of new bone growth may not be adequate to stabilize the defect. [16–18] .
Nevertheless, a restriction of new bone growth due to a lack of space alone does not appear to be detrimental to the host, as long as the bone cement remains intact and forms a bioactive bond with the host bone
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[16] .
IBCs must remain mechanically robust postinjection; it should not plastically deform or fragment under multidirectional loading. Fragments of PMMA of 1–12 μm in diameter were shown to be engulfed by macro-
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Nanomaterials: the next step in injectable bone cements
phages; fragments larger than 12 μm had macrophages and fibroblasts adhered to their surfaces, but were not engulfed [19] . Phagocytosis of PMMA fragments leads to the increased production of TNF, which also acts as a bone resorption mediator, as well as inhibition of DNA synthesis of the macrophage, whereas macrophages exposed to PMMA particles larger than 1 mm showed no difference in the production of TNF and DNA synthesis compared with the macrophages not exposed to PMMA particles [19] . In addition, it has been shown that 8 weeks after CPC was injected into the interbody regions of sheep lumbar spine, the presence of CPC fragments as a result of shear fracture led to severe inflammatory responses, with extensive fibrous tissue formation and CPC particles engulfed by macrophages, whereas specimens with good intervertebral body fusion showed no inflammation [20] . It is therefore likely that early fragmentation and the release of microparticles, particularly from brittle materials such as PMMA and CPC, will result in undesired inflammatory reactions. In essence, IBCs should be designed so that they are able to withstand compressive and shear forces so as not to fragment, and/or the osseointegration of the new bone into the IBC should be accelerated in order to prevent early fragmentation [20] . Clinically available IBCs Polymethylmethacrylate
PMMA-based bone cements generally consist of: a powder component comprising PMMA; a radiopacifying agent; and a liquid component comprising the liquid monomer methacrylate. Upon mixing, a doughy paste is formed that can be injected into the site of implantation. After injection, the paste will be fully polymerized and take the shape of the defect. PMMA cements are biocompatible and can be made highly radiopaque by incorporating inorganic compounds, such as zirconium oxide (ZrO2) and barium sulfate (BaSO4), in order to track the flow of bone cement via x-ray fluoroscopy [21] . Commercially available PMMA bone cements often distinguish themselves by the type of incorporated radiopacifying agent, such as BaSO4 and ZrO2, as well as antibiotics, such as gentamicin [8] . Due to their favorable mechanical properties and robustness compared with other available IBCs, PMMA-based bone cements remain the most common injectable materials used in the clinical setting for filling bone defects, particularly for arthroplasty, vertebroplasty and screw augmentation [22,23] . The tensile and compressive strengths of PMMA could reach up to 49 and 114 MPa, respectively [24] , whereas the tensile and compressive strengths of human cancellous bone range from 1 to 5 MPa and 4 to 12 MPa, respectively [25] . The mechanical strength of IBCs should be
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higher than that of the surrounding cancellous bone in order to safely withstand complex, multidirectional loading without plastic deformation or brittle failure at the site of the bone defect. A detailed comparison of the mechanical properties with other materials used for IBCs is presented in Figure 3 [24–29] . Despite its use for more than 50 years in clinical applications, PMMA bone cements have major drawbacks; they do not form bioactive bonds with the surrounding trabecular bone, nor are they resorbed by the body. PMMA-based cements form regions of permanent mechanical fixation as they intertwine through the trabecular pores upon injection [30] , with fibrous tissue eventually encapsulating the surface of the PMMA bone cement. In sheep vertebrae models, PMMA bone cement fillers failed to fuse properly with the surrounding trabecular bone, leading to the formation of a fibrous membrane surrounding the bone cement, which in turn resulted in micromovements and subsequent inflammatory reactions, culminating in a reduction in the vertebral strength and stiffness over a time course of 24 weeks [18] . PMMA surfaces can be rendered bioactive by incorporating a high concentration (~70%) of bioglass powders, as evident by the absence of a fibrous tissue layer at the bone–PMMA interface [31,32] . However, the high particle loading compromises the efficient injection of PMMA. By contrast, lower particle concentrations reduce the exposure to the bioactive components of the bioglass, and hence compromise the overall bioactivity of the PMMA [33] . PMMA surfaces incorporated with bioglass particles were found to be more effective in forming a bioactive bond with the surrounding bone when compared with those incorporated with hydroxyapatite [31] . More recent studies have demonstrated the bioactivity (both in vitro and in vivo) of PMMA incorporated with γ-methacryloxypropyltrimethoxy-silane and calcium acetate [34,35] . This material showed bioactive behavior in both dog and rabbit models, and also displayed a significant reduction in the released heat during polymerization [34,35] compared with that for the unmodified PMMA. Another major drawback of PMMA bone cements is their high stiffness compared with cancellous bone (Figure 3) , resulting in stress-shielding effects and, ultimately, weakening of the surrounding host bone, such as fractures occurring in adjacent vertebrae after undergoing vertebroplasty [36] . Efforts have been directed towards the development of low-modulus PMMA cements; Boger et al. developed an injectable macroporous PMMA (porosity of 56% and average pore size of 260 μm) by incorporating a 40% v/v fraction of 2% w/w sodium hyaluronate solution into the bone cement, reducing the modulus from approxi-
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100
PMMA
Cortical bone
Unreinforced CPC
CNT-CPC
Strength (MPa)
10
Cancellous bone Polymer fiber-CPC Gel-based Calcium sulfate
1.0
CPC group (E = 1-3 GPa)
Compressive strength Tensile strength 0.1 0.01
0.1
1.0
10
Elastic modulus (GPa) Figure 3. Elastic modulus and strength of injectable systems and bone tissue. Solid rectangles indicate compressive strength, dotted rectangles indicate tensile strength. The three CPC types (unreinforced, polymer– fiber composites and CNT composites) possess elastic moduli in the same range (1– 3 GPa) and have been separated for visual clarity. CNT: Carbon nanotube; CPC: Calcium phosphate-based bone cement; E: Elastic modulus; PMMA: Polymethylmethacrylate.
mately 1.8 GPa to 0.29 GPa, bringing it within the average stiffness of cancellous bone [37] . However, the compressive yield strength significantly dropped from 91 to 7.4 MPa, and while this is within the range of cancellous bone, it is inadequate for load-bearing orthopedic applications. Clinical incidence of bone cement implantation syndrome (BCIS), which is characterized by a rapid decrease in systemic blood pressure and, in rare cases, cardiac arrests, is another reported side effect of using PMMA bone cements [38] . Initially, liquid monomers were thought to play a key role in BCIS. However, it has been speculated that BCIS could be due to the effects of other factors, such as an increase in the intramedullary pressure, leading to embolization [38] . Another concern relates to altering the immune response and subsequent wound healing process due to the presence of unreacted monomer components,
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which are known to be toxic [39] . Due to the exothermic reaction of the polymerization process, there is a risk of thermal necrosis in the surrounding bone tissue, where the temperature of the polymerization was reported to reach 74°C in some clinically available PMMA IBCs [40] , which is significantly higher than 50°C, at which level bone tissue necrosis is reported to occur if exposure occurs for more than 30 s [41] . However, some researchers have argued that the temperature at the PMMA–bone interface does not exceed 50–54°C [41,42] and that the human body can act as a heat sink and therefore render the thermal effect on the surrounding cells clinically insignificant and insufficient to cause extensive thermal injury [41–43] . Nevertheless, the heat produced during the polymerization of PMMA is undesirable and should be minimized in order to reduce the risk of thermal injury.
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Nanomaterials: the next step in injectable bone cements
Calcium phosphate & calcium sulfate cements
Similarly to PMMA, CPCs and CSCs consist of a powder and a liquid (water for CPC and CSC or sodium phosphate solution for CPCs). Upon mixing, a series of crystallization reactions occur that result in the solidification of the cement. The differences between the CPC products are due to the different initial reactants used (which include α-tricalcium phosphate, β-tricalcium phosphate, anhydrous dicalcium phosphate and monocalcium phosphate monohydrate), producing either an apatite-based or brushite-based cement [44] . CPCs are bioactive and can possess various degrees of bioresorbability that, if sufficient, lead to new bone formation, as the CPC degrades within the defect. The setting reactions of CPCs are isothermal or slightly exothermal and hence lead to a minimal risk of thermal necrosis [43,44] . The hydroxyapatite layer that forms upon the setting of CPCs in vivo [8,10–11,26] is thought to be a contributing factor to the bioactivity and osteoconductivity of the CPCs by allowing a direct bonding of CPCs to the surrounding bone tissue without eliciting an immune reaction. By contrast, CSCs are not bioactive, and limited inflammatory responses have been reported [13] . Nevertheless, it has been claimed that the rapid release of calcium and sulfate ions faciliates the growth of new bone into the defect region [13] . The main disadvantage of CPCs and CSCs compared with PMMA-based bone cements is that they have inadequate tensile strength (Figure 3) . CPCs are weak against tensile, bending, shear and torsional stresses, and its tensile strength values are rarely reported. The three-point bending strength of apatitebased CPCs have been reported to be between 10 and 15 MPa [27] , and the tensile and shear strengths have been reported to be between 3 and 10 MPa [26] . In comparison, the flexural and tensile strengths in bulk, sintered hydroxyapatite range from 7.7 to 113 MPa and 18 to 130 MPa, respectively, and are heavily dependent on the presence of pores [25] . The clinical use of CPCs is applied in regions with primarily compressive loading, including the treatment of defects at the tibial plateau and the calcaneus, or where multidirectional loading is negligible, such as the distal radius [1,26] . Due to the weaknesses in tensile and shear strengths, the clinical applications of CPCs have remained largely limited to screw augmentation of the femoral head and vertebral compression fracture treatment via percutaneous injection [18,20,45–48] . Efforts have been directed toward increasing the tensile strength of CPCs through the fabrication of CPC–fiber composites [27,49–51] . The incorporation of fibers such as Vicryl® (Ethicon, NJ, USA) and carbon nanotubes (CNTs) into CPCs enhanced their flexural strength and the work of fracture is directly depen-
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dent on the fiber content, length and strength [27,49] . Buchanan et al. demonstrated that the incorporation of polypropylene fibers into CPC resulted in an increase in the work of fracture and setting time, as well as a decrease in the compressive strength and modulus [50] . However, the major drawbacks of the currently available fibers include low fiber–matrix interface bonding and a lack of bioresorbability and bioactivity. Ideally, the incorporated fibers themselves should be able to support or enhance the growth of new bone. The challenge is to find suitable fibers that are bioresorbable and bioactive while providing a sufficient shear strength and fracture toughness [27] . In order for fibers to augment the mechanical strength of ceramic matrices, there must be sufficient interfacial bonding between the fibers and the ceramic matrix [52,53] . CPCs have a similar chemical composition to human bone and hence a similar radiopacity. Therefore, it is difficult to obtain a good contrast between the bone and CPC with x-ray imaging, especially where there is excessive soft tissue [54] . Unlike PMMA, incorporating such inorganic nonresorbable radiopaque materials into resorbable CPCs may not be ideal, as the radiopaque compounds, such as BaSO4 and ZrO2, have been shown to remain permanently fixed within the remodeled bone [54,55] . Neither of the previous studies, however, reported adverse histological reactions arising from the in vivo incorporation of these nonresorbable inorganic particles (up to 24 months for BaSO4, rabbit model [54] ; up to 12 weeks for ZrO2, rat model [55]), suggesting that it is safe – at least in the relatively short term – to incorporate nonresorbable compounds into CPCs, provided that these compounds do not induce cytotoxicity or immunogenesis. However, long-term in vivo studies are required in order to establish a definitive conclusion regarding the safety of these radiopaque CPC compounds. Calcium silicate cements
Calcium silicate cement (CSiC) is another group of injectable materials that have been used primarily in endodontic treatment. In particular, mineral trioxide aggregate (MTA) is a CSiC that has been approved by the US FDA since the late 1990s for dental pulp repair. MTA consists of tricalcium silicate, dicalcium silicate, tricalcium aluminate (TCA), tetracalcium aluminoferrite and calcium sulfate, with bismuth oxide as the radiopacifying agent, and when all of the constituents are mixed with water, it forms a paste that can be molded to fit the site of implantation [56,57] . Although sintered and preformed calcium silicate ceramic scaffolds have shown promising results for implantation [58,59] , there are major obstacles that need to be overcome if MTAs are to be modified for
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Review No, Roohani-Esfahani & Zreiqat orthopedic applications. Although the final compressive strength for MTA is reported to be approximately 70 MPa, which is relatively high compared with other injectable systems, the setting time is long (∼4 h) and the setting reaction is highly alkaline, reaching up to pH values of 12.5 [56,57] towards the beginning of the reaction and decreasing to pH 8.5 by day 28 [57] . Traditional MTAs have been shown to form surface apatite when immersed in simulated body fluid and other phosphatecontaining solutions [60] , and have been demonstrated to support the in vitro differentiation of human orofacial bone mesenchymal stem cells [61] . MTA in vivo rodent model studies have shown some cases of direct new bone apposition, but most cases have also shown fibrous tissue formation between the MTA and new bone formation, despite the ‘bioactivity’ deduced from the apparent formation of in vitro surface apatite [60] . The setting time of MTA was shown to be reduced from >2 h to approximately 40 min by increasing the concentration of TCA to 15 wt%. However, 15 percentage by weight (wt%) of the TCA hindered the development of its compressive strength, with its maximum strength reaching up to approximately 40 MPa, compared with approximately 55 MPa for MTA with 0 wt% TCA [62] . A novel CSiC system that is different to the MTA system but still has calcium silicate as its major component has been developed, which is more directed towards potential orthopedic applications. Ding et al. have developed a CSiC consisting of SiO2 and CaO powders mixed with ammonium phosphate solution (3.7 M, pH = 7.4) [63] or water [64,65] . Their CSiC samples had clinically acceptable setting times of between 3.3 and 8.5 min [63] , and the pH values of the CSiC samples in culture medium were later shown to reach pH 7.9 after 24 h [64] . In addition, higher proliferation and osteogenic differentiation were observed in human osteosarcoma cell cultures on the CSiC samples compared with cultures with no cement [63] . Adding 5 wt% gelatin and/or chitosan oligosaccharides to a similar CSiC system demonstrated improved proliferation and osteogenic differentiation of human mesenchymal stem cells, but slightly reduced its diametral tensile strength [65] . However, this CSiC system is too weak for orthopedic applications, with diametral tensile strength not exceeding 3.0 MPa [63] , and strategies to improve its mechanical strength should be explored. Gel-based injectable systems based on in situ polymer cross-linking In situ polymeric cross-linking gel-based injectable systems such as polyethylene glycol (PEG)–polycaprolactone (PCL)–PEG, poly(lactic-co-glycolic acid)–PEG, polyvinyl alcohol–acrylamide and alginate are relatively novel materials that are designed primarily for the pur-
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pose of controlled drug-delivery applications [66,67] . These materials are injected directly at the site of action, which allows for increased efficiency and bioavailability of the drug and hence lower doses in order to reduce the risk and intensity of potential side effects. However, they also offer a promising potential for use as IBCs. In contrast to traditional injectable cements, these materials have a unique in situ cross-linking ability that can be initiated by either temperature, pH, light or cationic cross-linkers [6,7] . They remain as a low-viscosity solution for easy injection, but after injection and under physiological conditions, cross-linking starts and the material solidifies. Although these injectable materials are not yet suitable for load-bearing orthopedic applications due to their low inherent compressive strengths and stiffnesses, which do not exceed 6 and 10 MPa in terms of compressive strength and modulus, respectively, and so are significantly below those of PMMA and CPCs (Figure 3), some have been identified for potential use in bone tissue repair as a carrier for the delivery of growth factors, drugs and cells directly to the defect site in order to enhance the growth of new bone (Table 2). Most of the gels indicated in Table 2 have been shown to support the growth and osteogenic differentiation of mesenchymal stem cells [29,68–70] and also in vivo new bone tissue formation in rabbit models [71] . Furthermore, the gels possess interconnected microporous structures that are initially filled with water or other liquid molecules at the time of injection, which can then subsequently act as in situ porous channels for efficient nutrient and waste transfer during the bone regeneration process [72] . The role of nanomaterials in IBCs Recently, nanomaterials have been incorporated into existing IBCs and other novel injectable materials in order to improve their properties for use in bone augmentation and bone tissue regeneration (Supplementary Table 1). Figure 4 outlines the ways in which nanomaterials in IBCs can aid the regeneration of bone. Nanoparticles can: improve mechanical properties, such as tensile, compressive and shear strengths and fracture toughness [75,76] ; release ions or drugs into the surrounding environment in order to enhance osteoblast recruitment, adhesion, proliferation and differentiation [77,78] ; enhance surface energy and protein adsorption and influence integrin binding; and act as radiopacifiers for enhanced visual detection under x-ray imaging. Nanoparticle leaching for improved biological properties Chemical composition
Bioactive ions such as calcium (Ca 2+), silicon (Si2+), magnesium (Mg2+) and strontium (Sr2+), or certain proteins and drugs, can improve the osteoblast recruitment,
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Nanomaterials: the next step in injectable bone cements
Review
Table 2. Gels systems that have potential for use as injectable materials for bone regeneration. Material
Properties suitable for injectable materials in bone regeneration
Ref.
PEG–PCL–PEG triblock copolymer plus acellular bone matrix granule composite
Temperature-induced gelling at 37°C Biodegradable hydrogel Low cytotoxicity against rat MSCs Inflammation reduced after 4 weeks of implantation in BALB/c mice
[70]
PLGA/PEG + hyaluronic acid hydrogel
Temperature-induced scaffold assembly PLGA particles sintered at 37°C PLGA–PEG particles injected within a hyaluronic acid hydrogel Total porosity approximately 50% Compressive strength
