**1. Introduction**

Polymethylmethacrylate (PMMA) is widely commercially used bone cement [1, 2]. The most common method of application is the dry mixing of drugs along with the bone cement and administering it into the body [3]. While long considering the 'gold' standard in local antibiotic therapy, it has many disadvantages. A radiolucent fibrous tissue is often observed at the bone/cement interface due to the release of toxic methylmethacrylate (MMA) monomers which damages surrounding tissue [3, 4]. These cements also exhibit a high exothermic setting temperature. Temperatures ranging from 70–120°C have been reported during setting of the PMMA bone cement during implantation [2, 5]. In an experiment conducted by Stańczyk and Rietbergen on bovine cancellous bone, the temperature exposure of 70°C by a fraction (10%) of the bone at the bone/cement interface was recorded [6]. The addition of antibiotics results in reduced mechanical properties in the cement [7]. Furthermore, the release of the antibiotic is short lived and results in less than maximal antibiotic release [7, 8]. Finally, PMMA cements lack elasticity and have dense structure which does not allow bone growth inside the cement [9].

Halloysite nanotubes (HNTs) are commercially inexpensive and two-layered aluminosilicate nanotubes are found naturally as raw mineral deposit. When dehydrated, 15–20 clay layers form a hollow tubule which is capable of carrying drugs [31, 32]. Inside of the lumen is positively charged and the external surface is negatively charged, permitting additional functional modifications of these surfaces. This charge on both outer and inner surfaces affects the efficiency of loading drugs and chemical agents [33, 34]. Substances with overall negative charge can be easily loaded into the lumen when compared to positively charged particles. The size of HNTs varies from 500 to 1000 nm with an inner diameter of 15–100 nm depending on the deposit [35]. Due to physical properties such as nanosized lumens, high L/D (length to diameter) ratio, low hydroxyl group density, low cost, and abundant natural deposits, HNTs have been intensely studied as a controlled or sustained release agent [31, 33, 34]. Loading HNTs with pharmaceuticals showed low initial release concentrations, preventing an initial outburst and uniform drug delivery (particularly with drugs, such as antibiotics, hormones, and growth factors) [35–37]. The drugs dexamethasone, furosemide, and nifedipine gave an extended 6–10 h release profile [29]. Under optimal conditions, a maximum loading of 12% volume (very close to theoretical capacity) was obtained [29]. When HNTs are added to polymeric materials, they improve material and mechanical performance and stability [34]. HNTs have been used to increase surface area, impart high surface reactivity, and improve mechanical strength with a relatively low cost [32, 34]. To achieve an increase in the toughness, mechanical strength, and thermal stability, HNTs have incorporated into a variety of polymers and examples include: poly(methyl-methacrylate) [35], poly(butylene succinate) [38], polyamide [12, 39], styrene-butadiene [40], epoxy [41], and chitosan [42]. In two previous studies using electrospinning, HNTs were used to enhance the polymer material properties including biological, chemical, mechanical, and thermal properties (see Kamble et al., (2012)

Calcium Phosphate/Clay Nanotube Bone Cement with Enhanced Mechanical Properties and…

http://dx.doi.org/10.5772/intechopen.74341

125

Presently, the use of CPCs in regenerative medicine and orthopedic surgery is limited only to non-load bearing regions, that is, cranioplasty [43, 44]. There is a critical need for developing osteogenic and osteoconductive cement that can be used in load-bearing sites. The objective of this study is to improve the mechanical and anti-infective properties of CPCs. The common thread between these two objectives is the use of HNTs to provide a means for sustained release of anti-infective agents (gentamicin sulfate and neomycin sulfate) and to improve the material properties (tensile strength and adhesiveness) of the cement. The choice of CPC com-

P, DCPA), β-tri calcium phosphate (Ca<sup>3</sup>

 CaO8 P2 H2

O), and HNTs (H4

), chitosan (low molecular weight), dexamethasone, gentamicin, neo-

O4

) and calcium carbonate (CaCO3

Al2 O9 Si2 2H2 O8 P2 ,

0, MCPM), chitosan oligosac-

P, TTCP) was ordered from CaP

O) were purchased from

) were delivered

for a more detailed review) [34].

ponent materials was also a critical factor.

Calcium phosphate dibasic anhydrous (HCaO4

Biomaterials, E. Troy, WI. Cupric chloride (CuCl2

β-TCP), calcium phosphate monobasic monohydrate (H<sup>4</sup>

H10CaO6

Sigma-Aldrich, St. Louis, MO. Tetra calcium phosphate (Ca4

× H2

O9 ) n

**2. Materials and methods**

**2.1. Materials**

charide lactate ((C12H24N2

mycin, calcium L-lactate (C6

Brown and Chow were the first to propose the use of calcium phosphate cement (CPC) in bone repair [10]. CPC was approved in 1996 by the Food and Drug Administration (FDA) for repairing craniofacial defects [11]. CPCs have many advantages over PMMA cement. Low shrinkage, durable, dense or porous (depending on the site of injury), and formability (ability to fill cavities of complex configurations) are additional positive qualities of CPCs [12–14]. Due to their similarity to apatite minerals in natural bone, calcium phosphate (CaP) bioceramics, such as hydroxyapatite (HA), are osteoconductive and osteoinductive [15–17]. CPCs implants provide an ideal environment for colonization by osteoblasts to form a functional interface [18, 19]. The end product is easily resorbed by osteoclast cells, leading to new natural bone formation at the bone-implant interface [21].

All CPCs are formulated by mixing a solid and a liquid component. The solid component consists of two or more calcium phosphate salts. The solid phase usually consists of a basic and an acidic salt, which reacts together in an aqueous medium and precipitates HA as a final product [19, 20]. The liquids used can be either water, alginates, chitosan, or sodium phosphates [20–22]. To obtain maximum biological use, these components are mixed in predetermined proportions that will lead to the formation of HA. Resorbability of the CPCs completely depends on its end product [22]. The physicochemical reactions that occur during mixing are complex, and cement setting time depends on factors such as solid-liquid composition, liquid-to-powder ratio, and particle size of the powder. Setting conditions also influence the mechanical properties of the cement [23]. However, due to its brittleness, CPCs are restricted to the reconstruction of non-loading bearing bone [24, 25]. Recently, absorbable fibers [23] and chitosan [26] were used to improve the load-bearing capability of CPC [27, 28]. Chitosan and its derivatives are natural biopolymers that are biocompatible, biodegradable, and osteoconductive [29]. CPCs can also be modified through additives (i.e., silicon, strontium, and zinc) and delivered as paste, putty or in an injectable form and set *in situ* to provide intimate adaptation to complex-shaped defects [23, 24]. Polymer materials added as an organic phase to the CPCs have been shown to improve the biological response, physicochemical, and mechanical properties, such as injectability, cohesion, resorption, and toughness [27, 28, 30].

Halloysite nanotubes (HNTs) are commercially inexpensive and two-layered aluminosilicate nanotubes are found naturally as raw mineral deposit. When dehydrated, 15–20 clay layers form a hollow tubule which is capable of carrying drugs [31, 32]. Inside of the lumen is positively charged and the external surface is negatively charged, permitting additional functional modifications of these surfaces. This charge on both outer and inner surfaces affects the efficiency of loading drugs and chemical agents [33, 34]. Substances with overall negative charge can be easily loaded into the lumen when compared to positively charged particles. The size of HNTs varies from 500 to 1000 nm with an inner diameter of 15–100 nm depending on the deposit [35]. Due to physical properties such as nanosized lumens, high L/D (length to diameter) ratio, low hydroxyl group density, low cost, and abundant natural deposits, HNTs have been intensely studied as a controlled or sustained release agent [31, 33, 34]. Loading HNTs with pharmaceuticals showed low initial release concentrations, preventing an initial outburst and uniform drug delivery (particularly with drugs, such as antibiotics, hormones, and growth factors) [35–37]. The drugs dexamethasone, furosemide, and nifedipine gave an extended 6–10 h release profile [29]. Under optimal conditions, a maximum loading of 12% volume (very close to theoretical capacity) was obtained [29]. When HNTs are added to polymeric materials, they improve material and mechanical performance and stability [34]. HNTs have been used to increase surface area, impart high surface reactivity, and improve mechanical strength with a relatively low cost [32, 34]. To achieve an increase in the toughness, mechanical strength, and thermal stability, HNTs have incorporated into a variety of polymers and examples include: poly(methyl-methacrylate) [35], poly(butylene succinate) [38], polyamide [12, 39], styrene-butadiene [40], epoxy [41], and chitosan [42]. In two previous studies using electrospinning, HNTs were used to enhance the polymer material properties including biological, chemical, mechanical, and thermal properties (see Kamble et al., (2012) for a more detailed review) [34].

Presently, the use of CPCs in regenerative medicine and orthopedic surgery is limited only to non-load bearing regions, that is, cranioplasty [43, 44]. There is a critical need for developing osteogenic and osteoconductive cement that can be used in load-bearing sites. The objective of this study is to improve the mechanical and anti-infective properties of CPCs. The common thread between these two objectives is the use of HNTs to provide a means for sustained release of anti-infective agents (gentamicin sulfate and neomycin sulfate) and to improve the material properties (tensile strength and adhesiveness) of the cement. The choice of CPC component materials was also a critical factor.
