5.1 Introduction

Annually, numerous patients worldwide are treated for skeletal complications in the areas of orthopedic surgery, dental implantology, maxillofacial surgery, neurosurgery, etc. The use of autologous bone grafts is the gold standard in treating these bone defects because they are osteoconductive, osteoinductive, and osteogenic. However, they require invasive surgery techniques, and the resource of graft material is limited. Alternatives based on the synthetic biomaterials for autologous grafts become more attractive from the last decades. Within the different materials (i.e., bioactive glass, polymers), calcium phosphate (CaP)-based ceramics or cements are one of the most desirable candidates for bone replacement and other applications [1]. According to the crystal phase, CaP can be classified, e.g., hydroxyapatite (HA), β-tricalcium phosphate (β-TCP), biphasic calcium phosphate (BCP), amorphous calcium phosphate (ACP), carbonated apatite (CA), and calcium-deficient hydroxyapatite (CDHA) [2]. Based on the composition criteria, CaP cements are classified into apatite cements, apatite-forming cements, and dicalcium phosphate dihydrate (brushite) cements [3]. A list of existing calcium phosphate compounds and their abbreviations is summarized in Table 2.1 of Chap. 2.

CaP-based materials have been widely used for the bone tissue engineering and regenerative therapies because of their natural bone-like compositions, nontoxicity, biocompatibility, excellent osteoconductivity, and good processibility [4,5,6,7,8]. A desirable bone substitute may require the following characteristics: (1) good mechanical properties, which allow bone remodeling within the porous structure [9]; (2) appropriate degradation rate, which should ideally match with the osteogenic rate [10]; (3) osteoconductive, thus able to guide bone tissue growth along the surface and into pores [11]; and (4) osteoinductive, inducing the differentiation of pluripotent stem cells into bone-forming osteoprogenitor cells [12, 13]. All these characteristics of the materials are very crucial for bone tissue engineering purposes, and we will focus on the biodegradation properties of biomaterials in this chapter.

After injection/implantation, it requires the enough space that the bone and vascular tissues can grow into. Hence, the material biodegradation is necessary and critical. Typically, “degradation” represents the decomposition of a compound, especially complex substances such as polymers and proteins. The degradation mechanism can be summarized as simple hydrolysis, enzymatic degradation, or their combination [14]. Biodegradation can be considered as an in vivo process. A material breaks down into biologically acceptable molecules that are metabolized and removed from the body via normal metabolic pathways. Some researchers have proposed the mechanism of biodegradation of CaP based via solution-driven extracellular liquid dissolution and cell-mediated resorption processes [15]. The term “resorption” is often used to describe the breakdown of CPC, essentially signifying the biodegradation taking place via cellular mechanisms. The fate of implanted CaP biomaterials in vivo is dependent on various mechanisms and processes. Therefore, their biodegradation behavior can be regulated via physic-chemical properties of materials, the environment conditions around the implanted materials (i.e., pH level, ionic concentration, implanted locations), and the patient’s health conditions. To make the strategies for accelerating biodegradation of materials, there is a great need to get a thorough understanding of the factors influencing the biodegradation.

5.2 Degradation of Calcium Phosphate Cements

As we all know, calcium phosphate cements (CPCs) consist of a powder phase and a liquid phase. The powder phase comprises one or more CaP compounds, while water or a calcium- or phosphate-containing solution is used as the liquid and may contain chitosan, alginate, hyaluronate, gelatin, chondroitin sulfate, succinate, or citric acid to allow the dissolution of the initial CaP compounds [16]. The attractive property of CPC lies on the self-setting reaction at room/body temperature in which a calcium phosphate precipitate is formed. The CPC mixture can self-solidify in the simulated in vivo environment in minutes. The solidification process is not exothermal without excitant odors and is harmless to the surrounding tissues. Another advantage is the easy shaping. When the CPC fills into bone defects, it is easy to meet the requirements of various shapes of wounds, achieving optimal tissue-to-implant contact and fast operation. Generally, CaP cements can be divided into two subgroups: apatite cements (i.e., HA, CA, or CDHA) and brushite cements (i.e., dicalcium phosphate dihydrate (DCPD)). According to the literatures, microporous apatite cements show slow degrading and good mechanical stability, while brushite cements exhibit a lower mechanical strength but a faster biodegradability than the apatite cements [17]. As contacting with body fluids, brushite cements will eventually change into apatite.

Generally, CPC products can be formulated with different CaP compounds. To achieve the desirable replacement for bones from the clinical view, the property requirements of CPC materials are very critical, especially for the degradation behavior. The rate of material degradation should optimally match the rate of new bone formation, allowing a gradual takeover of mechanical strength by newly formed bone tissue. When the CPC materials are injected/ implanted into the specific site, the extracellular liquid and surrounding cells/tissues will influence the resorption processes of CPC and directly mediate the degradation performances. Compared with other common materials, the CaP-based materials show various mechanisms and processes (Fig. 5.1).

Fig. 5.1
figure 1

The degradation fate of CaP biomaterials after implantation. Here, CaP calcium phosphate, DCPD dicalcium phosphate dihydrate, OCP octacalcium phosphate, HA hydroxyapatite (Reproduced from Ref. ([18] by permission of MDPI, Basel, Switzerland)

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According to these proposed mechanisms for CaP biomaterials, their biodegradation may be a combination of the following processes: (1) physical effects (i.e., abrasion, fracture, disintegration), (2) chemical reactions (dissolution, increases of Ca and P locally at the surface), and (3) reductions in pH values due to cellular and phagocytic activities [18]. Typically, these activities will promote the rate of biodegradation due to the dissolution of the CaP biomaterials. First, the biodegradation of CPC is influenced by the cement properties, which include chemical composition, setting reaction, porosity, crystallinity, and particle size of the calcium phosphate compounds [19, 20]. These factors affect each other and are influenced by the chemical composition of the material. The patient conditions also play an important role in material degradation. Depending on the patient’s age, sex, type of the bone, hormone levels, genetic predisposition, and other metabolic health conditions, the speed of degradation can be different [21, 22]. Furthermore, the implantation site influences the biodegradation and bone formation. The implanted materials are exposed to the microenvironment around, in which the tissues, blood/nutrient supply, and other conditions may play a role in degradation of the materials. In this chapter, different factors that influence the biodegradation of calcium phosphate cements are discussed. These factors also have impacts on the mechanical properties of CPC. Typically, the better biodegradation of materials means poor mechanical strength. However, as a fact, bone implants are required to provide adequate short- and long-term mechanical support in the defect site in clinical applications. Therefore, the feasible strategies are proposed not only for promoting the degradation of CPC, but also for maintaining/improving the stress of the bone.

5.3 Major Factors Influencing CPC Biodegradation

5.3.1 Composition and Structure of Bones

Since the degradation of CPC materials as bone alternatives always occurs after implantation, it is necessary to understand the composition and structure of bones and the surrounding tissues. The bone (Fig. 5.2) in the human body can be defined as a composite of hydroxyapatite, type I collagen, water, cells, and lipids [23]. As can be seen, natural bone derives its unique combination of mechanical properties from an architectural design that spans macroscopic to nanoscale dimensions, with precisely and carefully engineered interfaces [24]. The bone is comprised of two distinct forms: dense cortical bone and porous cancellous bone. Cortical bone is the outer layer, which gives the support for the shape and form of the bone. Eighty percent of the skeleton is composed of cortical bone [25]. Cancellous bone has a lower Young’s modulus and is more elastic compared to cortical bone. Its porous structure consists of pore sizes in the range of 200–500 μm, and 30–90% of the porosity, which depends on the load, age, and health situation of the bone [26]. More importantly, the bone is a nanomaterial composed of organic (mainly collagen) and inorganic (mainly nano-hydroxyapatite) components, with a hierarchical structure ranging from nanoscale to macroscale. To design and produce the desirable bone replacement, many research groups have tried to manipulate the mechanical properties (e.g., stiffness, strength, and toughness) of scaffolds through the design of nanostructures (e.g., the inclusion of nanoparticles or nanofiber reinforcements) to mimic bone’s natural nanocomposite architecture [27, 28]. However, the good mechanical properties are not the only objective for realizing the desirable bone replacement. Its biodegradation has to be matchable for the new bone, allowing a gradual takeover of mechanical strength. Therefore, it is necessary to balance the mechanical and biodegradable properties during the designing and fabricating of nanomaterials.

Fig. 5.2
figure 2

Hierarchical organization of vascularized bone, from single collagen molecules to compact and spongy bone (Reproduced from Ref. ([24] with permission. Copyright © 2014, Rights Managed by Nature Publishing Group)

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5.3.2 Effect of Porous Architecture on CPC Biodegradation

Porosity is a parameter that refers to fraction of the void volume filled with liquid phase within a solid matrix. Unlike porosity, pore size directly refers to geometry of pores. Surrounding implanted materials, cell adhesion and motility depend on size of the pores, rather than porosity. Mean pore size has correlation with porosity for many synthetic polymers of simple composition. However, for natural polymers like collagenous gels, there is no direct correlation, since the diameter of collagen fibrils can vary from few nanometers to a few hundred nanometers [29].

In bone regeneration, permeability for fluid flow and molecular diffusion is important, since low permeability may result in a lack of nutrients and ischemia [30]. The permeability in vitro and in vivo depends on the porous structure of implants, including pore size, size distribution, porous morphology, interconnectivity, and surface area-to-volume ratio [31]. Each of these factors not only influences biological response (i.e., cell migration, proliferation, and thus tissue regeneration), but also plays roles in the resorption of implanted materials including CPCs. As a bone substitute material, CPC also contains an intrinsic high nano-/submicron-sized porosity, which may be related with the particle size of the powder phase [19] and liquid/powder ratio [20]. Typically, the porous size of dense materials is classified in three different types by biomaterial scientists: macropores (>50 μm), micropores (0.1–50 μm), and mesopores (2–50 nm). Therefore, we will distinguish these three different types of pores and discuss their effects on the material biodegradation and the bone ingrowth throughout this section.

5.3.2.1 Effect of Macroporosity

Macroporosity plays a critical role in the regeneration of damaged tissues, allowing cell penetration and tissue/blood vessel ingrowth [32]. At the same time, it is the key factor for controlling CPC biodegradation/resorbability, the important property of implanted materials for clinical usage. Ideally, the CPC will degrade in a short period of time matching new bone formation. Suitable pore size and an adequate interconnectivity should be required [33, 34]. However, there exist debates about the proper pore size for bone repairing materials. Some researchers claimed that a pore size larger than 100 μm is sufficient for tissue ingrowth, since macropore size determines the efficiency at which cells seed into the scaffold [33], whereas other studies suggested that the larger pores (up to 500 μm) are preferred [34]. It is generally acknowledged that the materials with larger pores will lead to a faster degradation. However, the structural/functional integrity of materials should be considered for different uses. Generally, the optimum pore size for scaffolds lies in the range between 100 and 400 μm [35]. This will allow cell ingrowth and vascularization and promote metabolite transport. A scaffold with an open and interconnected pore network and a high degree of porosity (>90%) is ideal for the scaffold to interact and integrate with the host tissue [36]. Figure 5.3 describes the typical degradation behavior of CPC/biodegradable polymer in vivo [37].

Fig. 5.3
figure 3

Micro-CT images of reconstructed cross section of beagle’s molar region and the residual materials implanted samples at 3 months (a–d) and 6 months (e–h); (a) and, (e) CPC control; (b) and, (f) CPCC; (c) and, (g) CPCCB; (d), and (h) CPCCBV; Note: CPC control, dense CPC; CPCC, macroporous CPC with collagen; CPCCB, CPCC with recombinant human bone morphogenetic protein-2 (rhBMP-2); CPCCBV, CPCC with VEGF (Reproduced from [37] by permission. Copyright @ 2014 Elsevier Ltd. and Techna Group S.r.l)

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As can be seen, Fig. 5.3 shows the precise micro-CT scan results about the material degradation and osteogenesis performance in vivo at each determined time. Obviously, all the samples exhibited varying degrees of degradation ability in vivo with the time passing. Specifically, the dense CPC group with several microns or less pore size still maintained the original shape and structure after 6 months of implantation. Lack of macroporous structure is the main reason. In comparison, CPCC groups with average pore size of 250 μm and macroporosity of 45% presented a certain amount of new bone formation [37]. The implanted materials had been partly absorbed and replaced by newly formed bone trabeculae on its peripheral portion. CPCCBV group showed the best new bone formation at both peripheral and central areas of the cement. After 6 months, the residual CPC material volume fraction of CPCCBV was about 20%, further indicating the fastest degradation speed among four groups. These behaviors were attributed to the encouraging effect of macroporous structure and macroporosity.

Some other researchers mentioned that CPC products had superiority over sintered HA because CPC could form macroporous HA in situ in the bone site without machining [38]. The pores of diameter ~20 to 50 μm are expected to provide favorable functionality for physiological liquid exchange, while pores of diameter ~100 to 350 μm are suitable for cell colonization and vascularization [39]. Thus, pristine CPC cannot allow vascular permeation and cell ingrowth. Such bone cement is very stable in the body, and the absorption and degradation usually occur only on the surface and the absorption rate is very slow [40]. To speed up the degradation rate of CPC, one of the feasible ways is to create macropores within CPC by using foaming agents or degradable polymer microparticles.

The degree of macropore interconnectivity is also considered to be critically important in a manner similar to pore size. In biodegradable porous ceramics, the degree of interconnectivity was noted to be seemingly more important than the pore size, while in nonbiodegradable materials (i.e., CPC), interconnectivity and pore size were observed to be equally important. Under in vivo conditions, the penetration of cells and chondroid tissue formation inside macropores occurred when the interconnectivity dimensions were greater than ~20 μm, while mineralized bone formation occurred when the interconnectivity size exceeded 50 μm [41]. The interconnectivity of pores ensures availability of higher surface area for fast resorption, enhanced cell adhesion, and proliferation. For bone tissue engineering, the optimal pore size for the resorption and osteoblast activity within the material matrix is still controversial and undermined.

5.3.2.2 Effects of Microscopic Features (Micropores and Grain Size)

While macropores with pore sizes and interconnections are in the range of hundreds of microns, microporous features in the range of nanometers up to several microns also play pivotal roles in tissue engineering. It was in 1984 that Klein et al. [23] underlined the fact that not only macropores but also micropores (typically close to 0.1–10 μm) were essential to provide a fast resorption. In 1993, Langer and Vacanti [42] devoted to the control of the architecture of CaP bone graft substitutes and to the understanding of the relationship between implant architecture and biological response. Since then, a number of researchers have investigated the effects of microscopic features such as micropores or grain size on the resorption behavior of CPCs (listed in Table 5.1) [43].

Table 5.1 List of several studies devoted to the effect of architectural factors on the in vivo performance of resorbable CPC scaffolds (sintered HA is considered to be non-resorbable). Various aspects were considered: micro- and macropore size, micro- and macroporosity, grain size, and interconnection pore size
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For example, Wei et al. [51] implanted magnesium-CaP macroporous scaffolds with 9% and 27% microporosity and found that the presence of micropores increased the degradation rate and accelerated bone formation. A similar result was found by other several groups [23, 47, 49, 52]. Klein et al. [47] reported that changes of microporosity played a more important role in the resorption process than changes in macroporosity. Recently, the presence of micropores was associated with the biological response of cells and osteoinductive potential for bone ingrowth [53,54,55,56]. These pores are usually few microns in size and are involved mainly on the initial adsorption of proteins on the surface of the materials, thus regulating the cell behavior on the implant surfaces as well as cell-material interactions. Some groups have revealed that microporosity that incorporated into CaP biomaterials can enhance the protein adsorption and cell adhesion due to the larger surface area, the tight interlock between material and tissue. Take an example, the cells spread and adhered better on membranes with smaller micropores (0.2 μm diameter) [53]. Once various proteins (soluble growth factor, serum proteins, and extracellular matrix proteins) and cells are adsorbed onto the surfaces, the interfacial properties will be changed, resulting in enhanced degradation in vivo [57, 58].

The influence of grain size on the CPC resorption might be related to the resorption mechanism. It has been proposed that the biodegradation of CaP-based biomaterial takes place via solution-driven extracellular liquid dissolution and cell-mediated resorption processes [21]. The dissolution is heavily influenced by the solubility of the implanted CaP materials, while the cell-mediated CaP resorption is due to the particle formation as a result of disintegration [59]. It has been reported that the grain size of the CaP materials affects the rate and effectiveness of cellular resorption activity [60]. Typically, the resorption rate increased with a decrease in grain size [11, 47, 48, 61], which can be schematically explained by Fig. 5.4 [62]. It was noted that the macrophages and giant cells actively participate in the resorption process for rapidly resorbing cements [63]. When the fragments and particles are <10 μm in diameter, they are easily internalized in macrophages and rapidly intracellular digested. For larger biomaterials (>100 μm in diameter), the bulk digestion is carried out via extracellular degradation by macrophages and giant cells.

Fig. 5.4
figure 4

Macrophage response to biomaterials depending on the size of the implanted materials. Macrophages respond to small fragments and particles (<10 μm in diameter) by internalization and intracellular digestion. If the particle size is larger than 10 μm and smaller than 100 μm, the macrophages fuse together forming giant cells which in turn engulf the particles and digest them. If the particles are larger, the bulk digestion is carried out via extracellular degradation by macrophages and macrophage-fused giant cells (Reproduced from Ref. ([18, 62] by permission of MDPI, Basel, Switzerland)

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Most studies devoted to microscopic features suggest that an increase of microporosity and a decrease of grain size accelerate resorption. However, there still exist the contradictory results and the difficulty to draw conclusions about the influences of microscopic features, since it is impossible to vary one parameter without changing other parameters. Here, we only discussed the effect of porous architectures on the in vivo resorption behavior of CPCs quantitatively. Recently, more attentions have been paid on the hierarchy structure of tissue scaffolds incorporating mesoporous materials ranging within 2–50 nm. With regard to the biomedical applications, mesoporous materials can easily host drug molecules and, therefore, could be useful to improve the drug adsorption ability and enhance bone-forming properties, as well as promoting resorption of CPCs.

5.3.3 Effect of CPC Composition on Biodegradation

Since commercial CaP bone graft substitutes have been launched for 40 years, many different formulations have been commercially available or are still being developed. Within various CaP compounds, HA, β-TCP, α-TCP, and DCPD can be the typical compositions forming CPCs. Among these compounds, HA is the most biocompatible. An important in vivo characteristic of HA-forming CPC is that it does not dissolve spontaneously in a normal physiological fluid environment, yet is resorbable under cell-mediated acidic conditions. β-TCP is the material more suitable for promoting synthetic osteoconduction, but its resorption remains slow in vivo, which can be regulated via cell-mediated processes. It has a lower strength and poorer mechanical properties than other CaPs. Although α-TCP has the same chemical composition as β-TCP, the difference in crystallinity makes it much more soluble than β-TCP. Therefore, early efforts to speed up the resorption of a CaP material were focused on α-TCP. CaP cements made of α-TCP powder have shown the excellent biocompatibility and biodegradability [64, 65]. Also, some studies have reported the positive performance of implanted α-TCP granules [66]. However, α-TCP is considered to be resorbed too fast and has hardly been investigated as raw material for larger granules and shaped blocks. DCPD is the most easily synthesized CaP compound. DCPD is biocompatible and osteoconductive and is reabsorbed faster than the β-TCP/HA-forming CPCs, allowing faster bone formation. However, the conversion process of DCPD into HA easily releases amounts of acid, leading to a severe inflammatory response.

Rapid resorption does not always indicate better bone formation. Under certain clinical situation, rapid resorption was reported to lead to lower quality bone formation [66]. In order to achieve optimum clinical performance, an appropriate CPC resorption rate needs to be modulated by tuning the composition of CPCs for the intended clinical applications. For periodontal bone defect repairs and sinus lift, the rapid replacement of the implant cement is highly desirable. To promote the resorption rate of certain CPCs with slow degradation, one of the feasible ways is to add the compounds with rapid resorption. These compounds could be small molecules (i.e., basic/acidic compounds, specific ions), degradable polymers, or other additives. For instance, calcium sulfate is biocompatible; however, it is less suitable for bone substitution because of its rapid resorption. Interestingly, mixing calcium sulfate with TTCP and DCPA leads to a good biocompatibility, enhanced degradability, and bone ingrowth as compared to CPC consisting of only TTCP and DCPA [22]. Small acidic compound also plays a role in the active resorption of the cement. It has been reported that a cement produced by a mixture of sodium calcium phosphate, tetracalcium phosphate and β-TCP powder, and malic acid or citric acid showed an increased number of osteoclasts on the cement [67], which indirectly indicates the cement mixed with malic acid has better resorption and is more promising for bone substitution than the cement mixed with citric acid.

5.3.3.1 Effects of Ions on CPC Degradation

The incorporation of some specific ions plays an important role in the biochemistry of bone substitutes, improving some properties such as the stability, solubility, porosity, or cytocompatibility of the bone remodeling process [68, 69]. Numerous studies have revealed that doping inorganic ions including calcium (Ca2+), silicate (Si), magnesium (Mg2+), zinc (Zn2+), strontium (Sr2+), manganese (Mn2+), copper (Cu2+), or (Fe3+) may trigger bone cell response, change CPC resorption behavior, and show important impact on bone healing strategies.

Calcium is the most abundant mineral in the human body. About 99% of the calcium in the body is found in bones and teeth [70]. Ca2+ is a mineral that the body needs for numerus functions, especially for bone formation. The mineral phase of the bone also contains Si, an essential trace element for metabolic processes associated with the development of the bone and connective tissues. This mineral is responsible for the depositing of minerals into the bones, especially calcium. It has been reported that Si can speed up the healing of fractures and also reduce scarring at the site of a fracture. Both Ca and Si ions play an important role in the nucleation and growth of apatite and then influence the biological metabolism of osteoblastic cells in the mineralization process and bone-bonding mechanism. Therefore, both ions can make the important contributions for the clinical field of bone regeneration, as well as material resorption.

Si has been substituted in different CPC formulations (i.e., HA, α-TCP, HA-based cements, β-TCP) to enhance in vivo implant degradation favoring their osteotransduction [71]. In the early stages of mineralization, Si content is higher, while decreasing at advanced stages with the increase of the Ca/P ratio toward that of HA [72]. Pietak et al. [73] have demonstrated that aqueous Si solutions enhanced osteoblast proliferation, differentiation, and collagen synthesis in vitro. They also reported that Si levels below 30 ppm could stimulate the development of osteoclasts, whereas higher levels inhibit osteoclast formation. The degradability of the Si-TCP scaffold was 90–80% after 1 year of implantation, and its complete resorption and replacement by new bone tissue had been achieved after 2 years [71]. Alkraishat et al. developed a silica-gel-modified CPC that increased the amount of remaining graft after implantation in rabbit calvaria [74]. To reveal the in vivo resorption of silicon-substituted CPCs, Aparicio et al. [75] prepared the sintered silicon-doped β-TCP powder first. By well controlling the constant Ca/(P + Si) ratio, a powder/liquid ratio, and the powder size, the physicochemical properties and in vivo degradation of the Si-CPC properties were evaluated. The cement porosity was about 40% with a shift of the average pore diameter to the nanometric range, whereas the degree of absorption was lower with increasing Si content. Interestingly, this Si-CPC provides a high specific surface area of up to 29 m2 g−1 and an enhancement of cell proliferation and cell attachment. Moreover, this bone substitute made of Si-CPC with lower Si content could interact efficiently with the surrounding bone and induce the formation of bone tissue in vivo.

Some evidences have suggested the body cannot assimilate calcium without the presence of silica [76]. Therefore, to improve the bioactivity and degradability of CPC, calcium silicate (CS) was added to CPC. Low-crystalline CS prepared by heat treatment at low temperature had excellent bioactivity and degradability. Adding CS to CPC did not affect the phase composition and chemical structure of CPC and had a little effect on the setting time and compressive strength of the composite cement. However, CS could effectively improve the in vitro and in vivo bioactivity and biodegradation of CPC and enhance the cell proliferation on the CPC material [77]. For designing the ideal bone substitutes, it is more challenging to achieve rapid biodegradability with maintaining/enhancing the mechanical strength. Liu group has explored the development of bioactive mesoporous calcium silicate/calcium phosphate cement (MCS/CPC) scaffolds (shown in Fig. 5.5) with high mechanical strength and fast degradation rate by micro-droplet jetting [78].

Fig. 5.5
figure 5

SEM images of (a1–a4, b1–b4) and digital camera photographs (c1–c8) of MCS/CPC scaffolds (Reproduced from [78] by permission. Copyright @ 2015 Springer)

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Figure 5.5 illustrates the MCS/CPC scaffolds with various architectures and pore structures. The inner pore shape was controlled by adjusting the angle of the process pattern. It could be clearly seen that the macropores in the scaffolds were uniform and completely open. In addition, scaffolds with different pore structures in Z direction were prepared precisely (b1–b4). To meet a wide range of individual needs, MCS/CPC scaffolds with different outer architectures were also designed and fabricated successfully. Moreover, addition of MCS to CPC not only exhibited outstanding printability of MCS/CPC paste but also achieved high mechanical strength and fast degradation rate (Fig. 5.6). Therefore, Si/Ca doping is a potential way to speed up the biodegradation and promote the bone repairing of CPCs.

Fig. 5.6
figure 6

The compressive stress–strain curve (left) and weight loss (right) in TrisHCl solution of CPC and MCS/CPC scaffolds (Reproduced from [78] by permission. Copyright @ 2015 Springer)

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The magnesium and zinc are also important foreign ions into the structure of synthetic calcium phosphate phases used as starting powders. The presence of Mg2+ and Zn2+ can alter a series of structural, physicochemical, and biological properties of CPC, such as crystallinity, solubility in the setting liquid, resorption, and bone-bonding capability [79, 80]. Incorporation of ions into α, β-TCP structures has been experimentally proven through quantitative phase analysis and structural refinement of the powders. Several groups have reported that the incorporation of Mg2+ and Zn2+ into β-TCP structure led to a decreasing trend in the lattice parameter values and a contraction of cell volumes [81, 82]. The reason for this contraction in refined parameters is due to the lower ionic radii of Mg (0.72 Å) and Zn (0.745 Å) than Ca ion. Since the degree of lattice disturbance by the substituted ions had implications in terms of thermal stability of the β-TCP phase, Mg2+ has been also reported to have a stabilizing role of noncrystalline CaPs, preventing crystallization into other more stable CaP phases. Therefore, the presence of Mg had a strong effect on cement composition and strength, namely, by increasing the proportion of brushite and decreasing the compress strength. According to these authors [83], Mg could be used to adjust the composition and rate of hydration of the cement. For instance, magnesium phosphate cement (MPC) composed of magnesium oxide and ammonium phosphate is a kind of fast-setting phosphate cement. Moreover, MPC has high initial strength and rapid resorption rate. When combined MPC with CPC, a novel magnesium-doped calcium phosphate complex with improved properties was achieved. For Mg-doped materials, the resorption process is closely related to the biological behavior of surrounding cells. Numerous research have reported that Mg has acted as a surrogate for Ca in transport and mineralization processes [84, 85], but it also exerts a large number of other actions, including enzyme co-factor function and modulation of the action of hormones, growth factors, and cytokines. Mg also has direct effects on the bone formation processes of resorption and mineral aggregation [86]. For example, our group has demonstrated that the C2C12 cells cultured on 5MCPC/rhBMP-2 substrates exhibited dramatically enhanced in vitro osteogenic differentiation, compared with CPC/rhBMP-2. Moreover, it was the Mg on the underlying substrates that became the main contributor to mediate the adsorption and conformation of rhBMP-2 bound on the matrix, thus enhancing osteogenic activity and the resorption of CPC matrix [87].

Besides the ions above, strontium is an alkaline earth metal accounting for approx. 0.02–0.03% of the earth’s crust. Strontium was discovered to follow the metabolic pathways and signaling principles known for calcium, although the response to Sr2+ tends to be weaker. It has been reported that Sr2+ could exert a dual effect on bone remodeling. Sr2+ enhances protein binding capacity and osteoblast activity, increasing new bone formation, although the exact mechanism of how Sr affects bone cells remains unknown; at the same time, Sr2+ inhibits osteoclast activity and thus reduces cellular bone resorption [88, 89]. Recently, many researchers have focused on the systematic investigation of Sr2+ effect on the biological activities of CPC including resorption behavior. As we described before, the solubility of the cement components affects the CPC degradation. Integration of Sr2+ ions into CPCs could increase solubility, and consequently higher degradation was demonstrated for SrHA and SrHA-forming cements. Sr2+ ions were also introduced into β-TCP to produce brushite-forming cements. A zero-order release of Sr during immersion of SrTCP in water indicated that Sr2+ was released by bulk erosion [68]. In addition, several in vivo studies revealed an increased degradation of SrCPC compared to the respective Sr-free control groups and thus confirmed the above findings [90].

To be summarized, ion-substituted bone cements hold great potential for applications in repair of bony and periodontal defects, owing to their relevant excellent properties (i.e., setting time, injectability, good resorption for clinical uses), especially the Sr-containing cements that exhibited an overall better performance. However, there remains a great need for optimization and prospective studies of ionic-substituted cements regarding specific clinical applications.

5.3.3.2 Effect of Degradable Polymers on CPC Degradation

As discussed above, an ideal substrate for the bone regeneration should not only have adaptive mechanical strength, but also can promote bone repairing with good biodegradation. Adding biodegradable polymers into CPC matrix is a feasible strategy to improve the degradability of CPC and alter their mechanical/physical properties. A number of biodegradable polymers have been considered for CPC degradation promotion and mechanical reinforcement (as listed in Table 5.2).

Table 5.2 Summary of major degradable polymers used in combination with CPCs [17]
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Biodegradable polymers are processed in CPC materials with the specific form of wires, fibers, or microspheres. For instance, as abundant reinforcing materials, PLGA fibers or other degradable polymeric fibers were introduced into CPC to enhance the mechanical properties and improve the degradability. The best mechanical properties reported for fiber-reinforced CPC were achieved with PLGA fibers and HA matrix, which had a highest bending strength of 40–45 MPa [91]. In physiologic environment, the PLGA fibers degraded within 3–4 weeks, leading to a loss of reinforcing effect, while the HA matrix took months to years to degrade.

More interesting, the introduction of PLGA microspheres is proposed for mechanical reinforcement and degradation modulation. Adding PLGA microspheres inside CPCs was first done in 2002, and the resulting paste was moldable and showed good biocompatibility [92]. By well controlling the physical/chemical properties of PLGA microparticles via tuning the molecular weight, L/G ratio, and the content of PLGA, PLGA/CPC composites were successfully fabricated and held different degradability and mechanical strength [92,93,94,95,96,97]. Table 5.3 summarizes the degradable and mechanical properties of typical PLGA microspheres/CPC composites.

Table 5.3 The degradable/mechanical properties of typical PLGA/CPC composites
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As can be seen, the presence of PLGA microspheres can enhance the early compressive strength of CPC [92, 94, 97, 98]. And the acidic degradation products of PLGA could accelerate the dissolution of CPC on account of the acidified surroundings [95]. Moreover, the generation of macropores by prior degradation of PLGA microspheres can provide channels for cell distribution and migration together with the ingrowth of blood vessels, promoting the new bone formation [93]. Therefore, tissues can easily grow into the PLGA/CPC composites along with the interconnections among pores formed after microsphere degradation.

Besides the advantages discussed above, these entrapped microspheres can be applied as drug delivery vehicles for osteoinductive growth factors or anti-inflammatory drugs. Combined these bioactive molecules and degradable microparticles with CPCs, multifunctional performances could be achieved for bone repairing including better resorption and new bone formation, which will be discussed later. Obviously, incorporation of PLGA microspheres is a promising way to design the resorption rate of bone graft matching for the rate of new bone formation in bone defect repair.

5.3.3.3 Effect of Other Additives on CPC Degradation

As mentioned before, one of the advantages of CPC is the low-temperature setting, which allows the incorporation of different useful additives/biomolecules: from antibiotics and anti-inflammatory drugs to growth factors. After implantation, these additives released from CPC cannot only treat the different skeletal diseases or speed up the bone repairing but also alter the microenvironment surrounding the implants, influencing the resorption behavior of implants.

In most cases, it is supposed that CPCs do not degrade while additive is released. However, some authors have shown certain degree of degradation of CPCs during drug liberation. For example, Otsuka [99] showed an increase of porosity in a carbonated HA cement during the release of indomethacin. Bai [50] revealed that the amount of newly formed bone and the rate of biodegradation were higher in porous CPC loaded with rhBMP-2, compared to the scaffolds without rhBMP-2. Therefore, it should not be neglected the effect of the additives on the resorption rate of CPC. Typically, the additives such as antibiotics, even growth factors, have the small size, and their release is considered to be excessively fast. Therefore, the additives may tend to improve the degradation rate of CPC. And the influences of additives on the CPC mechanical strength are still controversial: some researchers reported the minor variations of mechanical strength due to the antibiotic loading, while some found the enhanced mechanical properties due to the effect of loaded drugs with substantial interactions with CPCs. To prolong their release behavior and modify the properties of CPCs, one possible way is to entrap the polymeric microspheres loaded with additives/drugs into CPC matrix. Polymeric microspheres not only provide a carrier for controlled release and better bioactivity of drugs/proteins but also reinforce the mechanical strength of CPC as fillers. This has been confirmed by many research groups. Ruhe et al. [100] mixed hrBMP-2-loaded PLGA microspheres with a CPC. The release rate of rhBMP-2 in PLGA/CPC composite was much slower than the rate in the microspheres alone.

As we know, there are many additives available for CPC applications. Various additives/drugs have various effects on CPC properties, and this represents a serious issue for the implementation of this technology. A lot of work has still to be done to establish the general laws to regulate additives and modify the properties of CPC for different therapeutic needs.

5.4 Typical Strategies for Rapid Biodegradation of CaP-Based Cements

CPCs are frequently used as synthetic bone graft materials in view of their excellent osteocompatibility and clinical handling behavior. However, the degradation rate of plain CPC is limited, especially for HA-forming CPCs, thereby limiting complete bone regeneration. In the previous section, we have discussed many factors influencing the biodegradation of the CPC, mainly including porous architectures and intercompositions. Hence, there are many strategies available for achieving the faster biodegradation of CPCs. The most important one is the macroporosity, the key factor for controlling CPC biodegradation/resorbability. For a better bone regeneration, different approaches have been used to induce macroporosity in CPC, such as inclusion of water-soluble additives, foaming agents, a hydrophobic liquid (oil), and biodegradable polymeric microspheres. However, it is considered that macropores severely degraded the CPC strength. After macroporous materials were implanted in vivo, the strength significantly increased once new bone started to grow into the macropores [101]. Therefore, it is in the early stage of implantation when the macroporous implant is in the most need of strength [102].

5.4.1 Water-Soluble Additives for Macroporosity Enhancement

The inclusion of water-soluble molecules in CPC is a common approach to induce macroporosity. Also known as the particle leaching technique, it is often used to create porosity in bone substitute materials such as preset ceramics or polymeric scaffolds. Saccharides such as sucrose, chitosan, as well as the salts sodium chloride and sodium phosphate are water-soluble compounds to generate macroporosity in CPC. Depending on the solubility of these sacrificial materials, these are dissolved before implantation or gradually degraded in vivo. Sucrose is a disaccharide derived from glucose and fructose. Sucrose crystals with the size in the range of 125–250 μm have been included in CPC, and macroporosity was readily generated after dissolution of the crystals during incubation of the samples in aqueous media. However, the mechanical strength of CPC decreased with increasing macroporosity [103]. Another approach is the creation of microspheres from a saccharide. As a linear polysaccharide, chitosan has been processed in the form of fibers and microspheres, embedded in or trapped on the CPC matrix in order to enhance its mechanical properties in an early stage and to generate porosity after the fiber degradation [104]. Several salts have also been used to induce porosity in CPC, such as sodium chloride crystals, NaHCO3, CaSO4-2H2O, and Na2HPO4. Similarly to when saccharides are employed, porosity is generated after the salt is leached out from the CPC in an aqueous environment. The degree of solubility of the particular additives during the setting reaction of the cement is responsible for the content and dimension of the macroporosity [105]. However, the lack of strength of the resulting cement (especially for the additives with quickly dissolving) and compromise requirements between CPC handling and its bioactivity due to a large amount of additives limit the applications of this technique.

5.4.2 Foaming Agents for Macroporosity Enhancement

The use of foaming agents is another way to generate gas bubbles and form macropores. Typical foaming agents involve hydrogen peroxide and carbon dioxide (CO2). CO2 bubbles from sodium bicarbonate have been used during the cement setting to generate porosity [106]. Besides, citric acid was employed as porogens for CPC for good injectable properties and higher porosity [107]. Furthermore, citric acid has combined with sodium bicarbonate as effervescent agents to generate porosity [108]. Since no potentially toxic gas is released after implantation of the cement, this technique holds great potential for macroporosity enhancement and fast biodegradation. However, the low initial strength and poor reproducibility still remain the issues for this technique.

5.4.3 CPC/Synthetic Degradable Polymers

To overcome the issues including the initial strength, degradability, and lack of interconnectivity, synthetic degradable polymers in forms of fibers and microparticles have been added in cements. Typical degradable PCL and PLGA fibers entrapped in CPCs had the function of reinforcing the cement, providing the needed short-term strength and toughness and gradually dissolving afterward, leaving behind macropores suitable for bone ingrowth [109, 110]. The fiber parameters (i.e., length, volume fraction, type of fiber, and the mass percentage) affected the physiochemical properties of cements, especially the resorption behavior. The advantages of long fibers over particulates and short fibers are the fact that, once resorbed, they can form interconnected pores inside the CPC structure facilitating bone tissue regeneration [111]. In vivo results also proved that the introduction of ultrafine degradable fibers within a CPC matrix improved macroporosity efficiently and enhanced CPC degradation and bone ingrowth largely [110].

Of special interest are also the various works that incorporated PLGA microspheres within the cement paste [92,93,94,95,96,97]. The rationale was to create a macroporous structure through degradation of the PLGA microparticles/microspheres. The mechanical strength of the microspheres contributes to the initial strength of CPCs, and the degradable nature of microspheres offers the better resorption performance of CPCs. Since the degradation of PLGA by nonenzymatic hydrolysis of its ester linkages depends on its molecular weight, this allows the easy modulation of macropore formation.

These microspheres can be applied as delivery vehicles for osteoinductive growth factors or other biofunctional additives. Besides the mechanical enhancement, the microspheres could simultaneously release the osteogenic factors for accelerating the new bone formation. The in vivo results from subcutaneous implantation in rats showed that the PLGA microspheres were completely degraded after 12 weeks of implantation, allowing blood vessels to colonize the macropores [100, 112].

Other types of microspheres also showed the similar effects on cement properties. For examples, self-setting CPCs were produced upon the introduction of microspheres composed of gelatin [113] and poly(trimethylene carbonate) (PTMC) [114]. In vitro incubation of all these composites in phosphate-buffered saline or enzyme-containing media resulted in an interconnected, macroporous calcium phosphate matrix after microsphere degradation. Gelatin and PTMC microsphere/CPC composites were found to show gradual degradation of the microspheres from the outer to the inner parts, as the cement delayed enzyme diffusion through the material [113]. On the other hand, PLGA/CPC degraded simultaneously throughout the whole composite as a result of hydrolytic cleavage of the polymer chains. Furthermore, compared with the bulk erosion mechanism of PLGA and gelatin microspheres, surface erosion of the PTMC microspheres [114] resulted in a rapid decrease in compressive strength as they detached from the cement skeleton. This trend is also confirmed by subcutaneously implanting of CPC composite in vivo [94].

Acid-producing microparticles are also attractive to create macroporosity since CPCs degrade by acid dissolution. Félix Lanao et al. [115] incorporated glucono delta-lactone (GDL) into CPC for accelerating CPC degradation. After 2 weeks of implantation, CPC containing 10% of GDL degraded faster and was replaced by more bone tissue than CPCs containing either PLGA or gelatin microspheres. For matching the medical use, the degradation rate of CPCs was turned by formulating mixtures of PLGA microspheres and GDL microparticles [116]. The animal studies revealed that incorporation of 43% PLGA, 30% PLGA-5% GDL, and 30% PLGA-10% GDL in CPC significantly increased bone formation and resulted in higher bone height compared with both 10% GDL- and 20% GDL-containing CPC samples. Besides the biodegradable polymers listed in Table 5.2, there exist many potential additives, which can be used for tuning the CPC degradation, as well as other properties. More contributions should be made for fulfilling the therapeutic needs.

5.4.4 Strategy Combinations

In some studies, researchers combined the previously mentioned methods for accelerating the degradation behavior of CPCs. For example, CO2 foaming has been used to induce porosity in combination with PLGA microparticles embedded in CPC for the creation of secondary porosity at a later time point [117]. Acid-producing GDL was formulated with PLGA microspheres within CPC matrix for adjusting the CPC degradation [116]. Mannitol crystals were used to generate early macroporosity in combination with slower resorbable chitosan fibers, which improved the mechanical properties of the cement at initial time points [118].

Recently, it has already circumstantiated that both micrometer and nanometer scale features of a material have marked influence on cell behaviors in vitro and in vivo. Additionally, materials organized on multiple length scales have better conformity to biological matrices than those with single-scale features. Therefore, some attempts for fabricating synthetic scaffolds with hierarchical macro-/micro- or macro-/nano-architectures have been carried out by using properly mesostructured materials. The purpose of such scaffolds is twofold, as they combine the properties of traditional CPC-derived scaffolds, i.e., mechanical support in the defect zone, bioactivity, favored osteointegration, and bone tissue regeneration, with the unique features supplied by mesoporous materials, such as enhanced bioactivity and controlled drug adsorption/release ability for drug therapy in situ.

Our group developed the bioactive mesoporous calcium silicate/calcium phosphate cement (MCS/CPC) scaffolds by using micro-droplet jetting. The 3D printed MCS/CPC scaffolds with hierarchical architectures (350 μm micropores and 10–20 nm mesopores) showed fast degradation rate, high mechanical strength, and good cytocompatibility [82]. Combined the growth factor rhBMP-2 with the mesoporous bioactive glass (MBG)/CPC composite, the resultant scaffold not only presented a hierarchical pore structure (interconnected pores of around 200 μm and 2–10 μm) and a sufficient compressive strength (up to 1.4 MPa) but also exhibited excellent drug delivery behavior. Moreover, this composite scaffold presented a significant improvement of osteogenic efficiency, especially at the early stage in vivo. Moreover, better resorption was obtained in the rhBMP-2-loaded MBG/CPC scaffold compared to the others [119]. Obviously, it is very challenging to improve the properties of CPC satisfying all the requirements for clinical applications via altering one simple factor. Many research studies have demonstrated that the strategy combination will be feasible and hold great potentials.

5.5 Future Perspectives

The ultimate goal of bone reconstruction is the regeneration of the physiological bone that simultaneously fulfills both morphological and functional restorations. Within the different materials, CPCs are one of the most desirable candidates for bone replacement. However, the poor degradation and low mechanical strength limit their clinical applications. To overcome the limitations, especially for material degradation, different strategies (water-soluble additives, foaming agents, ions introduction, and biodegradable polymeric microspheres) can be applied. To meet the requirements for bone substitutes, strategy combinations are also suggested to obtain the attractive biological performances.

The injectability and moldability of the CPC are promising for less invasive and faster surgery as compared to other bone substitutes. The improvement of resorption behavior of CPCs is still an “old new” topic. Many strategies including the introduction of polymeric wires, fibers, or microparticles have been applied for accelerating the CPC resorption with reinforcing their mechanical properties. Simultaneously, functional biomolecules are loaded/immobilized into the CPC matrix. However, the biocompatibility issues caused by the use of large volumes of degradable polymers and the difficulties to combine polymers with CPC without compromising the physic/chemical properties with biological response still prevent the CPC applications for desirable clinical uses. Recent studies in the area of hierarchical scaffolds with immobilized growth factors have clearly revealed that the architecture of complicated design with well-defined structures has shown the excellent properties (i.e., mechanical, degradable, osteoconductive, osteoinductive features) and outstanding in vivo medical performances. However, in vivo studies and clinical trials have not yet been investigated to their maximum extent. There are still lots of optimization work to do for developing desirable CPC products, which have superior clinical performance and are easily applicable in clinical practice.