Calcium Phosphate Cements: From the 1980s to the 2020s

The groundbreaking discovery

In 1983, Brown and Chow of the American Dental Association Foundation discovered CPC during their studies on CaPs aimed at developing mineralizing pastes to repair early dental carious lesions. Based on the solubility properties of CaPs, they formulated mixtures containing tetracalcium phosphate [TTCP, Ca₄(PO₄)₂O], and dicalcium phosphate anhydrous (DCPA, CaHPO₄), or TTCP and dicalcium phosphate dihydrate (DCPD, CaHPO₄.2H₂O). They observed that some of these pastes, when left in test tubes for a few hours, became a hardened mass. ¹⁹

This led them to realize that the mixtures they had created, primarily consisting of CaPs, formed HA, Ca₅(PO₄)₃OH, which is the major component of tooth mineral.

So, these scientists inadvertently discovered a new type of self-hardening cement that consisted of only CaPs and formed HA as the only product, referred to as apatite cement. ²⁰ This was the first formal CPC and led to the filing of the first patent by Brown and Chow in 1985. ²¹ The simplified chemical equation for apatite cement reaction (TTCP/DCPA) proposed by Brown and Chow was expressed as follows:

C a 4 ( P O 4 ) 2 + CaHP O 4 → H 2 O C a 5 ( P O 4 ) 3 O H

(1)

This reaction is a typical example of a classical acid–base interaction, where a relatively acidic calcium orthophosphate (DCPA) reacts with a relatively basic one (TTCP) in an aqueous suspension to form a poorly crystalline, slightly basic HA paste. ²² An equimolar molar ratio of TTCP and DCPA formed stoichiometric HA (Equation 1), while a molar ratio of 0.5 produced calcium-deficient hydroxyapatite (CDHA) (Equation 2). Mixtures with molar ratios between 0.51 and 2.0 result in cements containing both apatite and residual starting compounds. ²³

3 C a 4 ( P O 4 ) 2 O + 6 C aHP O 4 → H 2 O 2 C a 9 ( H P O 4 ) ( P O 4 ) 5 O H

(2)

Equations 1 and 2 show that there is no consumption of the liquid phase in the setting reaction; the presence of water is simply to allow the formation of a workable cement paste and aid ion transport.

Insights from a first decade of CPC research

Investigation of the setting reaction in the initial CPC formulation

Early studies on the setting mechanism of CPCs based on TTCP and DCPA or DCPD employed thermodynamic analysis, X-ray diffraction (XRD), scanning electron microscopy (SEM), and transmission electron microscopy (TEM). Fukase et al. ²⁴ for example, used XRD to evaluate the cement setting reaction and to identify the products formed over time in a CPC based on TTCP/DCPA, while SEM provided insight into the microstructure of the cement specimens.

The recorded XRD patterns showed that the only reaction product detected was consistent with HA phase obtained at various time intervals, indicating successful conversion of the starting materials. The extent of the cement setting reaction was quantified by rate of formation of HA phase demonstrating a linear relationship during the first four hours. The reaction reached its maximum extent in the 24 h samples, though differences between the 4 h and 24 h samples were minimal, suggesting stabilization after a certain period. ²⁴

SEM analysis confirmed the formation of crystalline structures during the setting reaction, with morphological changes observed over time. One hour after mixing, micrographs showed outlines of the original particles alongside some amorphous materials in the interparticle spaces, which appeared as small petal-like crystals at higher magnification. After 2 h, more interparticle materials appeared, and rod-like crystals replaced the platy ones. At 4 h, the particle outlines were less distinct, and crystal growth indicated increased crystallinity and interparticle adhesion. ²⁴

Doi et al. ²⁵ investigated the setting mechanism of an apatite cement based on TTCP and DCPD by thermodynamic and XRD analyses together with TEM. The thermodynamic analysis of calcium and phosphate in the cement slurry suggested that the solution composition remained for the first 10 min in equilibrium with DCPD. Over time, the solution became undersaturated with respect to DCPD and shifted toward the solubility line of HA, which suggests the direct formation of the thermodynamically stable HA phase. XRD confirmed that HA formed directly from the dissolution of DCPD and TTCP, and TEM images showed that the newly formed needle-like apatite was entangled and intermeshed, creating a hard mass like the structure of plaster of Paris.

These findings provided strong evidence that the setting mechanism of this CPC followed a direct transformation from precursor powders to HA without intermediate phases. Figure 1 schematically represents this process in four stages: (I) mixture of CaP particles derived from the powder precursors (TTCP and DCPD) and liquid phase, along with their illustrative XRD patterns. (II) CaP particles partially dissolve, forming a plastic paste. (III) Dissolved calcium and phosphate ions recombine, leading to the onset of HA precipitation, where initial particle outlines became less distinct due to crystallization and interparticle adhesion. (IV) HA nuclei grow into an interconnected network of needle-like or plate-like crystals, with the XRD pattern confirming HA (red symbol) as the predominant phase.

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FIG. 1.

Processing and microstructure of CPCs.

These findings revealed crucial details about the chemical reactions and mechanisms involved in the setting process, serving as a starting point for exploring new formulations and understanding the factors that influence the physicochemical, mechanical, and biological properties of CPCs based on TTCP and DCPD or DCPA.

Development of new formulations

In the 1980s decade, several researchers developed new CPC formulations that differed from the original of Brown and Chow. Lemaitre et al., for example, developed in 1987 a brushite cement by combining β-tricalcium phosphate (β-TCP) with monocalcium phosphate monohydrate (MCPM) that, when mixed with water to a paste, sets at room temperature within 2 min, according to equation 3. ²⁶

β - C a 3 ( P O 4 ) 2 + C a ( H 2 P O 4 ) 2 . H 2 O + 7 H 2 O → 4 C aHP O 4 . 2 H 2 O

(3)

Although the solubility of DCPD and its resorption rates in vivo and in vitro were greater than those of β-TCP and HA, ²⁷ this cement faced challenges for medical use due to its very short setting time when mixed with water, low diametral tensile strength (∼1 MPa, i.e., 50% of the bending strength of a conventional Portland cement), and decrease of mechanical properties under physiological conditions (37–40°C, 100% r.h.) ²⁸ when compared with Brown and Chow’s original formulation that had a long setting time (approximately 30 min) and a compressive strength of 36 MPa. ²⁴

A third type of CPC formulation was developed by Monma et al. in 1988. ²⁹ They prepared mixtures of α-TCP with DCPD and added water to form pastes which appeared to set within 9–30 min as measured with a Vicat needle. The reaction product in these mixtures was octacalcium phosphate [OCP; Ca₈H₂(PO₄)₆·5H₂O]. Below is the equation representing the formation reaction of this cement.

2 α - C a 3 ( P O 4 ) 2 + 2 C aHP O 4 . 2 H 2 O + H 2 O → C a 8 H 2 ( P O 4 ) 6 . 5 H 2 O

(4)

Unlike the TTCP/DCPA and TTCP/DCPD-based cements, this was a hydraulic cement system, meaning that the product formation was due only to consumption of water. This cement was derived from an α-TCP (α-tricalcium phosphate) cement already developed by Monma and Kanazawa in 1976, ³⁰ by maintaining α-TCP pastes with water at temperatures between 60 and 100°C, where the hydrolysis of α-TCP leads to the formation of CDHA (Equation 5), and according to Driessens et al., ³¹ they could hardly be classified as cements, since the setting temperatures were impractically high.

3 α - C a 3 ( P O 4 ) 2 + H 2 O → C a 9 ( P O 4 ) 5 ( H P O 4 ) O H

(5)

According to Monma et al., ²⁹ the combination of α-TCP and DCPD showed a high hydraulicity, probably due to the complementarity that the hydrolysis of TCP to OCP was accelerated by the coexistence of DCPD and vice versa, and the reaction products were OCP only in the range of 1.20–1.47 molar ratio Ca/P. Hardened bodies soaked in 0.9% NaCl solution at 37° for 1 day had a porosity of 49 ± 3% and a wet compressive strength of 14–19MPa.

In 1993, Driessens et al. ³¹ tested over 100 formulations to evaluate their setting ability when mixed with water. Most formulations included both acidic and basic CaPs, which reacted upon mixing. They found that various reactant combinations were possible within CaP systems. In addition, they discovered that adding CaPs with elements like sodium, potassium, magnesium, zinc, carbonate, or chloride could be beneficial. They also identified that compounds such as biopolymers, organic acids, biodegradable synthetic polymers, growth factors, and osteocalcine could be used as accelerators, retarders, or promoters of bone ingrowth.

Thus, from its discovery until the end of the first decade of research, there was an expansion in the range of formulations and properties of CPCs. Various new formulations containing different CaP sources were developed. In addition, CPCs composed of multiple CaP sources, along with organic and inorganic additives, were also being investigated.^31–34

Investigation of the biological performance of CPCs

Since CaP ceramics were already known to cause no systemic adverse reactions, researchers focused on studying local reactions in the tissues and cells around implanted CPCs, concentrating on the direct impact of this material on the tissue where it was applied. ⁴⁷ In the early 90s, research into the cellular reactions of CPCs began, sparking growing interest in understanding the biological responses and potential applications of these materials.

Initial tests showed that CPC was not mutagenic or toxic. Most of the experiments used TTCP+DCPA or TTCP+DCPD-based CPCs. For instance, when the CPC was used as root canal filler or sealer in a dog study, the CPC was found to be well accepted by the periapical tissues. ⁴⁸ In a monkey study where a large amount of cement was entered into the periapical area, only a mild tissue irritation was observed. ⁴⁹ Craniofacial application of CPC in cat studies demonstrated that the CPC was osteoconductive when in contact with bone and did not evoke an inflammatory response^50,51 In addition, the CPC was found to degrade slowly while being replaced by bone. For this, at that time, Driessens referred to CPCs as osteotransductive materials, meaning that they can be gradually replaced by bone or support bone formation that ultimately replaces the material. ⁵⁰ And according to Driessens, the osteotransductive nature of CPCs was especially interesting because, unlike other bioceramics such as β-TCP and rhenanite (CaNaPO₄), they offered the advantage of being moldable by the surgeon during surgery. This allowed them to conform precisely to the shape of the bone cavity at the time of implantation, enabling immediate osseointegration. A proof for the osteotransductivity of apatite CPCs was found by several authors.^51–55 Friedman et al, ⁵³ for instance, proved the osteotransductivity of an HA-type CPC based on TTCP and DCP (DCP = dicalcium phosphate) or DCPD mixtures in the cat frontal sinus. At 6 months, very little implant resorption or new bone deposition was seen. At 12 months an average of 30% of the reconstructed area was filled with woven bone. By 18 months bone ingrowth reached an average of 63%. At that time approximately 10% of the implant was replaced by fibrovascular material.

Figure 2 is an illustration of the osteotransductive behavior of CPCs, where in (B) it is possible to see the ability of the CPC to contour the osseous defect immediately after being injected, and in (C) to support the growth of new bone while it is gradually replaced without forming a gap between the material and the bone, thereby maintaining mechanical stability throughout the healing process.

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FIG. 2.

Illustration of the osteotransductive behavior of CPCs. (A) A bone defect, (B) CPCs being injected into the defect site, (C) supporting new bone growth while the overall volume of the augmented region remains stable.

To explain the osteotransductive behavior of CPCs, Chow ⁵⁶ hypothesized that CPC that reprecipitates to HA during its setting reaction was not fused to a solid mass, unlike sintered HA. The resulting microstructure (i.e., porosity) allows osteoclasts to gradually resorb the material. In addition, he suggested that the osteocytes appeared to see such a CPC as disorganized bone rather than a synthetic implant. As a result, resorption proceeds gradually, compared with other calcium-based cements, which allows the formation of woven bone during the resorption of the CPC.

Considering the excellent biocompatibility of CPC, three centers were already in 1991 granted a Food and Drug Administration (FDA) Investigational Device Exemption (IDE #910073). This authorization enabled the experimental use of HA cement in clinical studies, specifically to investigate its effectiveness in reconstructing defects in the human craniofacial skeleton. This rapid transition from discovery to in vivo use underscores the significant impact of CPCs in regenerative medicine and its potential to significantly enhance bone repair and bone regeneration.

Summary of the first decade of CPC research

Despite these considerable advances, CPCs were still in their developmental stage during their first decade of research, and several limitations had already begun to emerge—limitations that researchers and clinicians were becoming increasingly aware of. Figure 3 summarizes the advantages and disadvantages of CPCs identified during this initial period.

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FIG. 3.

Fundamental (a) advantages and (b) disadvantages of CPCs verified in the first decade of research.

Optimization and development of new formulations

A significant milestone in optimizing and developing new formulations of CPCs was achieved by Driessens et al. ⁵⁷ just over ten years after the discovery of CPCs. In a highly relevant investigation, they explored over 450 different combinations of reactants for CPCs. The aim of their study was to replicate earlier performed studies and to develop additional formulations for CPCs. Their assessment of the various combinations was based on the following criteria: (a) Formation of the intended reaction product, (b) Final setting time had to be ≤60 min, and (c) The compressive strength after soaking for 1 day in Ringer’s solution at 37°C had to be ≥2 MPa. They found that, of the investigated formulations, only 14 CPC formulations met all the assessment criteria. These cements could be categorized into four classes on the basis of their reaction product: (1) DCPD, (2) calcium and magnesium phosphate, (3) OCP, and (4) CDHA. Ten of the 14 formulations consisted of two components, three of three components, and one of one component. Seven of the formulations contained precipitated HA seeds. The maximum compressive strength as observed was 30 MPa and achieved with a cement in the OCP group. This cement was composed of DCP + α-TCP + PHA.

Following this line of thought, from the mid-1990s to the early 2000s, various formulations emerged. Several researchers investigated self-setting formulations consisting of powder mixed formulations of calcium orthophosphates and other calcium-based substances like calcium carbonates, ⁵⁸ calcium oxide,^59–62 calcium hydroxide,^63,64 calcium aluminate,^65,66 among others. In addition, other chemicals such as Sr-containing compounds, ⁶⁷ Mg-containing compounds,^68,69 as well as cements made of ion-substituted calcium orthophosphates [e.g., Na₃Ca₆(PO₄)₅], ⁷⁰ and cements based on calcium orthophosphates and polymers^71–75 were tested. These different formulations also resulted in different setting times, handling, cohesion and injectability of CPCs.

Enhancement of setting times

Between 1993 and 2003, a significant effort was made in controlling the setting behavior of CPCs. First, Driessens et al., ⁵⁷ decided to discern two different criteria to define the setting behavior of CPCs, i.e., the initial (IST) and the final (FST) setting time, expressed in minutes as measured from the beginning of mixing. The IST reflects the time for which the cement must not be touched, because otherwise the structure that is forming will be destroyed by application of even very small loads. The FST is the time after which the cement can be touched, and closure of the tissue wound is allowed. As a result, a single measurement was considered insufficient to fully describe the setting process.

Two main methods were used to measure the setting times: the Gillmore needles method (ASTM C266-89) and the Vicat needle method (ASTM C191-92). Both methods involved placing a weighted needle on the cement’s surface at different time points to check for indentation, and the cement was considered set when no full indentation was made. Specifically, Gillmore needles were used to measure IST and FST, with a light, thick needle for the initial setting time and a heavy, thin needle for the final setting time. ³¹ This method proved reliable when the applied load was low (10–20% of the cement’s maximum compressive strength) and when the cement’s compressive strength increased steadily during setting. On the basis of the above mentioned, Driessens and Ginebra et al. decided that the IST had about 4 ≤ IST ≤ 8 min and the FST 10 ≤ FST ≤ 15.^76,77

Subsequently, various strategies were used to adjust setting times to meet these clinical requirements. These included: (a) adding calcium or phosphate ions to the liquid phase,^77–79 (b) incorporating seed materials,^64,77,80 (c) using additives to accelerate or inhibit reaction rates,^70,80 (d) modifying the L/P ratio,^59,68,70 (e) altering the particle size of the CaP precursor compounds,^59,68,81 and (f) the effect of temperature.^82,83 Organic acids such as malic and citric acid also proved effective in this regard, with citric acid showing superior performance under specific conditions. ⁷⁰ The concentration of HA seeds had a dual effect, i.e., low concentrations promoted setting time reduction, while higher concentrations counteracted this effect.^64,77 In addition, increase of the room temperature from 22°C to 37°C reduced the setting time, ⁸² which can help in the clinical application of CPC. For example, lowering of the temperature of the cement liquid will allow preparation of the cement paste a long time before use. ⁸³ After application in a bone defect, the cement setting will proceed rapidly due to the human body temperature.

It is crucial to note that a common issue in the studies involving setting time control was the trade-off between rapid setting and mechanical strength. For instance, Fernandez et al. ⁷⁸ showed that use of a P-containing aqueous solution accelerated the setting time but weakened mechanical strength. Similarly, Yang et al. ⁶⁴ noted that excessive HA seeds prolonged the setting time and reduced strength beyond a certain threshold (e.g., 20% content). This delicate balance between rapid setting and mechanical performance has remained a challenge in CPC development.

Enhancement of cohesion

After application into a bone defect, the CPC paste encounters blood or other physiological fluids. Consequently, a good cohesion is an essential factor of the paste. However, it was observed that CPCs can disintegrate before set when they encounter fluids. Therefore, a cohesion time (CT) was defined, which is the time that the CPC does not longer disintegrate when it is immersed in Ringer’s solution. ⁸⁴ The CT must be less than initial setting so that the cement can be applied into the defect and molded. The CT of CPC to allow its clinical application must be ≤2 min. ⁸⁴ For example, CPC composed of α-TCP and precipitated HA and TTCP and DCPA/DCPD, and both mixed with aqueous disodium hydrogen phosphate solution, were found to be stable when they were immersed in water after setting.^84,85 However, if exposed to fluids like water or serum/blood before setting was initiated, these cements began to break down, ⁸⁵ releasing CaP microparticles, as shown in Figure 5. This leaching of CaP microparticles at the implantation site has been associated with a compromised biocompatibility and can evoke an adverse in vivo response. ⁸⁴

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FIG. 5.

Illustration of the reaction of apatite cement paste when exposed to (A) water or serum/blood immediately after component mixing and molding, (B) onset of the CPC structure disintegration process, (C) CPC fully disintegrated.

To enhance cohesion and prevent disintegration in biological fluids before complete setting, sodium alginate and/or chitosan were incorporated into the liquid phase, allowing the paste to remain intact even when immersed in distilled water.^85–88 In addition to sodium alginate and chitosan, the gelling agents hydroxypropyl methylcellulose (HPMC) or carboxymethylcellulose were added to the cement liquid (water) of CPC composed of TTCP/DCPD powder to improve the cohesion properties. ⁸⁹ Although these additives improved the cohesion of the cement paste, the setting time became significantly increased. A similar study was done for α-TCP-based CPCs. ⁹⁰ The researchers demonstrated that other cohesion promoters, like hydroxyethyl starch, sodium dextran sulfate, and polyvinylpyrrolidone, also reduce the CT. In this study, the effect of the used cohesion promoters on other CPC characteristics (e.g., initial/final setting time and compressive strength) was small and could be adjusted by changing the % Na₂HPO₄ in the cement liquid. ⁹⁰

Enhancement of injectability

As highlighted earlier, various CPC compositions were developed in the 1990s, with numerous animal studies confirming their suitability for bone repair in various surgical procedures. However, CPC was applied into the bone defect in most of the animal experiments in a paste-like form with a spatula or as set discs. However, a syringe must be used for delivery of the CPC in narrow defects and sites of limited accessibility. ⁹¹ Lack of viscous flow had been a serious limitation for CPCs in terms of injectability. The cement mix can be made as loose paste by increasing the wetting ratio, but this will be of limited help. The liquid part will get pushed out through the needle, and major portion of the particulate mass remains inside the syringe. Moreover, a high wetting ratio increases the setting time to unacceptable values. ⁹¹

Despite these problems, injection of CPC in paste form was described. The most notable among such reports is that of Constantz et al., ⁹² who successfully used percutaneous administration of a commercial cement, i.e., Norian Skeletal Repair System (Norian SRS®), for the minimally invasive treatment of acute radius fractures. Supplied as a dual-component system—powder and liquid—the product required a dedicated mixing device and injection gun. Although the setup was complex and occasionally unreliable, its high-pressure system allowed the cement to infiltrate fine trabecular structures, such as those found in osteoporotic bone. ⁹²

Advancements in the field of cement injectability occurred during the second decade. In 1995, Andrianjatovo and Lemaître ⁹³ enhanced the injectability of brushite CPCs by incorporating small amounts of polysaccharides, such as chitosan or HPMC, which increased the viscosity of the liquid phase. Nonetheless, their study lacked quantitative data on injectability.

To enable quantitative comparisons across formulations, Khairoun et al. ⁹⁴ in 1998 developed a standardized method for evaluating injectability (I%) using a commercial syringe. The method consisted of mixing the cement in a mortar, then loading 2.0–4.0 g of cement paste into a 20 mL syringe with a 2 mm opening. Exactly two minutes after combining the liquid and powder, the paste was extruded either manually or with a maximum force of 100 N. Injectability was defined as the percentage by weight of the cement paste that could be successfully extruded from the syringe.

The injectability was determined by this method for L/P ratios of 0.35 and 0.40 mL.g⁻¹, 1.5 min after the beginning of mixing powder and liquid. Results showed that (1) increasing the L/P ratio improved injectability significantly, and (2) phase composition of the CPC powder has a major effect on the injectability. Also, it was observed that there is a practical upper limit to injectability of as determined using the described method, which is the maintenance of a minimum amount of 190 ± 10 mg cement paste in the syringe. A side effect of the study was that the injectability measurements can also be used to accurately determine the dough time (also called working time) of CPC, which is the period during which the cement paste can be molded without damaging the structure as formed during cement setting.

Furthermore, the injectability of these α-TCP-based CPCs was measured in the presence of polymeric additives, intended to promote cohesion, but authors concluded that these additives did not significantly affect the injectability of these cements. ⁹⁰ Other approaches were also used. For example, Leroux et al. ⁹⁵ tested the influence of additives such as sodium glycerophosphate (NaGP), lactic acid, glycerol, and chitosan on the injectability of a CPC marketed under the name of Cementek®. Injectability was measured with an apparatus that consisted of a syringe with a catheter (diameter 3 mm) and a T-connector leading to a manometer, which indicated the injection pressure to extrude the cement paste out of the syringe. The in vitro studies indicated that injectability of Cementek® cement was improved by the addition of NaGP, lactic acid, and glycerol. The researchers suggested that this is because Cementek® cement without or with these adjuvants matures in HA, while Cementek® cement with chitosan matures in OCP.

In 2002, Sarda et al. ⁹⁶ investigated the effect of citric acid (up to 5 wt%) addition to aqueous Na₂HPO₄ (2.5 wt%) liquid on the injectability of CPC composed of 100% α-TCP. Injectability was found to increase until the addition of 1.5 wt% citric acid and decreased for higher values. One year later, Komath and Varma ⁹¹ developed a fully injectable CPC by incorporating biocompatible gelling agents, such as glycerin, cellulose derivatives, and alginic acid salts, into a TTCP/DCPD-based system with aqueous Na₂HPO₄ liquid. Alginic acid salts proved most effective to achieve suitable injectability. Although the cement mass transformed into HA, the setting was delayed, but setting times were still clinically acceptable. Also, Ishikawa ⁹⁷ investigated the injectability of CPCs. He studied the effect of powder shape of spherical tetracalcium phosphate, and better injectability of spherical TTCP was observed compared with irregular-shaped TTCP.

All these efforts resulted in the marketing of several injectable CPCs during this period, i.e., Cementek®, Biopex®, and α-BSM®. In summary, various strategies were employed to improve CPC injectability, including adjustments in L/P ratio, incorporation of polymeric additives, use of gelling agents, modifications to particle morphology, and injection methodology. While these efforts led to notable progress, they also revealed the inherent trade-offs between injectability and other key parameters, such as setting time, phase transformation, and mechanical strength.

In vivo and clinical investigation of biocompatibility and biodegradation in commercial CPCs

Between 1993 and 2003 extensive research was conducted to advance CPCs into clinically viable biomaterials. Table 1 lists the FDA-regulated apatite and brushite cements, which became commercially available during this period.

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Table 1.

List of Commercially Available Calcium Phosphate Cements between 1993 and 2003

Cement name	Company	Composition/end product	FDA status	Common use
α-BSMTM (α-Bone Substitute Material; also known as Biobon®)	ETEX corporation	Powder: ACP (50%), DCPD (50%) Solution: Unbuffered aqueous saline solution 98 End product: Hydroxyapatite (poorly crystalline)	510(k) cleared	Bone defect filling, dental, craniofacial
Calcibon® (originally known as Biocement D)	Biomet-Merck Biomaterials	Powder: α-TCP (61%), DCP (26%), CaCO3 (10%), PHA (3%) Solution: H2O, Na2HPO4 90 End product: Hydroxyapatite	510(k) cleared	Bone defect filling
BoneSource®	Stryker-Leibinger Corp.	Powder: TTCP (73%), DCPD (27%) Solution: H2O, mixture of Na2HPO4 and NaH2PO4 56 End product: Hydroxyapatite	510(k) cleared	Craniofacial augmentation and restoration
ChronOS InjectTM	DePuy Synthes	Powder: β-TCP (73%), MCPM (21%), MgHPO4·3H2O (5%), MgSO4 (<1%), Na2H2P2O7 (<1%) Solution: H2O, sodium hyaluronate (0.5%) 99 End product: Brushite	510(k) cleared	Fracture repair
Norian SRS®	Norian Corporation	Powder: α-TCP (85%), CaCO3 (12%), MCPM (3%); Solution: H2O, Na2HPO4 92 End product: Carbonate apatite	510(k) cleared	Distal radius factures, craniofacial
Norian CRS®	Synthes	Powder: α-TCP (85%), CaCO3 (12%), MCPM (3%); Solution: H2O, Na2HPO4 92 End product: Carbonate apatite	510(k) cleared	Distal radius factures, craniofacial

FDA approval requiring evidence of safety and effectiveness; 510(k): clearance based on substantial equivalence to a predicate device.

DCP, dicalcium phosphate; DCPD, dicalcium phosphate dihydrate; FDA, Food and Drug Administration; MCPM; monocalcium phosphate monohydrate; PHA, precipitated hydroxyapatite; TCP, tricalcium phosphate; TTCP, tetracalcium phosphate.

In agreement with the in vivo findings discussed in Section “INVESTIGATION OF THE BIOLOGICAL PERFORMANCE OF CPCs”, animal studies involving these commercially available CPCs confirmed their biocompatibility in both trabecular and cortical bone. These cements did not provoke a significant inflammatory response when in contact with surrounding soft tissues.^98–102 Moreover, the histological studies demonstrated clear osteoconductive properties, effectively supporting new bone formation and integration, and also provided valuable insights into the resorption behavior of CPCs.^{100,101,103–110} It was reported that the commercial apatite CPCs underwent creeping substitution, in which osteoclast-mediated resorption was followed by osteoblast-driven bone formation, highlighting the osteoconductive nature of CPC.^{102,110–112}

However, it was also observed that CPCs can exhibit distinct resorption mechanisms and rates depending on the final setting product of the cement. For instance, ChronOS Inject™ (brushite) and α-BSM® (hydroxyapatite) were implanted for 2, 4, and 6 months into full-thickness cranial defects in sheeps. ¹⁰⁹ After 6 months, the bioresorption rate of ChronOS Inject™ in the bone defects was significantly higher than that of α-BSM®, resulting in less remaining cement matrix within the ChronOS Inject™ filled defects. Histological analysis of ChronOS Inject™ specimens revealed large areas of fibrous tissue containing macrophages with phagocytosed cement particles. In contrast, α-BSM® specimens showed osteoclasts at the interface between newly formed bone and partially degraded residual cement. Newly formed bone was observed within the degraded α-BSM® regions, with osteoclasts located at the interface between the deposited bone and the remaining α-BSM®. This difference in degradation behavior was attributed to the initial drop in pH following defect creation, which promoted the dissolution of ChronOS Inject™ and the release of fine cement particles.

These studies demonstrated that CPCs undergo two presumed decomposition mechanisms: (i) passive resorption through chemical dissolution or hydrolysis in body fluids and (ii) active resorption regulated by living cells such as macrophages or osteoclasts ¹¹³ although the specific cell type involved varied by cement type. Passive CPC degradation depends on the solubility product of the cement setting reaction, while active degradation is mainly dependent on the solubility product of the connecting agents of the powder particles after crystallization. ¹¹³ Brushite CPCs, due to their high solubility in body fluids, were primarily resorbed via the passive mechanism, with macrophages and giant cells breaking the released particles down. In contrast, apatite CPCs (resorbing over months to years) required active resorption, with macrophages and osteoclasts locally reducing the pH to levels where apatite becomes soluble. ¹¹⁰

In view of the above mentioned, it must be emphasized that resorption rates and bone-forming capacity of CPCs varied not only depending on the cement formulation but also on the use animal model and anatomical defect site.^{100,102,106,108,111–116} Also, mechanical loading was demonstrated to influence CPC degradation. Frankenburg et al. ¹⁰⁶ reported that Fracture Grout™—a CPC similar to Norian SRS®—resorbed more rapidly in loaded tibial defects than in unloaded femoral ones in dogs. At 4.5 years, the cement volume had significantly decreased in the tibia, whereas nearly twice as much remained in the femur, except in one case where near-complete resorption occurred at both sites. Cement porosity was another factor affecting resorption. For example, Del Real et al. ¹⁰⁸ demonstrated that a porous Biocement D exhibited 81% resorption within 10 weeks in goat femoral condyle implants, while the nonporous version remained largely intact over extended periods.

Overall, resorption rates and simultaneous bone replacement have been identified as critical properties of CPCs. As emphasized by Jansen et al., ⁹⁹ CPC degradation and concomitant replacement by bone are required for a clinically successful application. Several studies on different CPCs have described that CPCs degrade in the same way that natural bone remodels, i.e., through resorption by osteoclasts followed by new bone formation by osteoblasts. However, the synchronous process of cement breakdown and bone regeneration has been regarded as a significant topic of CPC improvement.

Some of the newly developed CPCs from this decade were subsequently evaluated in clinical trials. For example, Norian® SRS was extensively used in humans for fracture stabilization, particularly in distal radial fractures, tibial plateau fractures, and calcaneal fractures. It was also employed to augment pedicle screw fixation and compression screws. These trials found that Norian® SRS provided strong mechanical support, allowing early mobilization and reducing malunion rates. Kopylov et al. ¹¹⁷ demonstrated that Norian SRS® improved mobility outcomes compared with prolonged immobilization techniques. Schildhauer et al. ¹¹⁴ found that Norian® SRS allowed early weight-bearing in joint depression fractures while maintaining structural integrity. Norian CRS was also used for craniofacial reconstruction.^6,115 The use of the Norian CRS CaP bone cement in the restoration of craniofacial skeletal defects was also described by Losee et al. ¹¹⁶

Besides Norian® CRS and SRS, BoneSource® was used in craniofacial surgery for skull defect reconstructions and augmentations, effectively restoring cranial integrity while serving as a scaffold for new bone formation with minimal complications. ¹⁰⁷ It was also applied to seal cerebrospinal fluid leaks during transtemporal surgeries, where it successfully set and prevented leakage. ¹¹⁸ In addition, BoneSource® was utilized to enhance transpedicular screw constructs and treat metaphyseal bone voids caused by trauma, helping to minimize functional deficits associated with surgical procedures. ¹¹⁹ However, its application in periodontology was less successful. Brown et al. ¹²⁰ and Rupprecht et al. ¹²¹ found that BoneSource® failed in furcation defect treatments, as granulation tissue interposed between alveolar bone and the cement, preventing proper integration.

Summary of the second decade of CPC research

By the late 1990s and early 2000s, CPCs evolved into more versatile materials with enhanced setting mechanisms, improved cohesion, and greater injectability. Furthermore, preclinical and clinical studies validated their safety and efficacy in orthopedic, dental, and maxillofacial applications, contributing to their regulatory approval and commercialization.

Improvement of the mechanical properties

In the early 2000s, Charrière et al. ¹²⁴ evaluated the mechanical behavior of brushite and apatite CPCs using compression, tension, and torsion tests, modeled with cone-wise linear elasticity and the Tsai-Wu failure criterion. The study demonstrated distinct mechanical profiles for a brushite and an apatite cement. Under tension, brushite CPC exhibited a modulus of 6.6 ± 0.4 GPa and a tensile strength of 1.3 ± 0.3 MPa, whereas HA cement presented higher values, with a modulus of 12.3 ± 0.8 GPa and a strength of 3.5 ± 0.9 MPa. In compression, brushite CPC showed a modulus of 7.9 ± 0.3 GPa and a strength of 10.7 ± 2.0 MPa, while HA cement showed a modulus of 13.5 ± 0.8 GPa and 75.0 ± 4.2 MPa. Poisson’s ratios were similar (brushite 0.16 ± 0.03; HA 0.14 ± 0.02) for both CPC types. For shear, brushite displayed a modulus of 2.7 ± 0.4 GPa and a shear strength of 2.9 ± 0.4 MPa, while HA cement showed superior results with a modulus of 4.8 ± 0.3 GPa and a strength of 9.8 ± 2.6 MPa. Overall, these data showed that HA-based CPCs exhibited consistently higher stiffness and strength than brushite-based ones under all loading conditions. The authors found that brushite CPC had a Young’s modulus comparable with that of human cancellous bone, while HA cement fell within the range of human cortical bone. However, both were significantly stiffer than cancellous bone. Based on these findings, they concluded that brushite CPC is more suitable as a bone-filling material for cancellous bone, whereas apatite CPC appears more promising for potential weight-bearing applications.

Considering these findings of Charrière et al., during the third decade of CPC research, a wide range of strategies were employed to enhance the mechanical properties of CPCs. In addition to the approaches adopted in previous decades, new efforts emphasized the development of CPC composites through the incorporation of diverse fillers or reinforcements. These included fibers,^125–137 particles,^138–142 whiskers,^143,144 carbon nanotubes,^145–148 polymers,^149–159 bone chips, ¹⁶⁰ and even biological additives like bovine serum albumin. ¹⁴⁵

Compressive strength was among the most frequently reported mechanical property in these studies, and several reinforcement strategies resulted in substantial improvement across different CPC formulations; however, the values achieved generally remained below those required for reliable load-bearing applications. To the best of our knowledge, only one of these reinforcement strategies has achieved compressive strength values close to the lower limit of cortical bone. Specifically, Gbureck et al. ¹³⁸ demonstrated that adding silica or titanium oxide to TTCP/DCPA-based cements significantly increased their compressive strength, reaching approximately 80–100 MPa. However, this work does not provide a quantitative analysis of porosity, which could be crucial for understanding the material’s mechanical and biological performance. In most studies, the compressive strength of modified cements remained below the threshold required for load-bearing applications.^{126,130,134,139–142,144–147,149,150,153–161}

In addition to compressive strength, some reinforcement strategies also improved flexural strength and fracture toughness,^{126,131,136,137,162–166} with fiber-based reinforcements standing out as particularly effective in enhancing these mechanical properties. Xu and Simon, ¹³¹ for instance, reinforced TTCP/DCPA cement with absorbable polyglactin mesh sheets, significantly enhancing flexural strength and inducing macroporosity upon degradation. CPC with 13 and 6 mesh layers achieved a flexural strength of 24.5 ± 7.8 MPa and 19.7 ± 4.3 MPa, respectively, versus 8.8 ± 1.9 MPa for unreinforced CPC. The work-of-fracture also increased substantially, i.e., from 0.021 ± 0.006 kJ.m⁻² to 3.35 ± 0.80 kJ.m⁻² (13 meshes) and 2.95 ± 0.58 kJ.m⁻² (6 meshes). For comparison, cancellous bone typically exhibits flexural strengths between 10 and 20 MPa, ¹⁶⁷ while cortical bone presents higher values, from 135 to 193 MPa^167,168 along with a fracture toughness ranging from 0.36 to 5.53 kJ.m^−2.169

Xu et al. ¹³⁶ investigated the synergistic effect of incorporating polyglactin mesh with chitosan. The incorporation of the absorbable mesh was intended to create macropores in the CPC after mesh degradation, while the combination of reinforcing agents was hypothesized to have a synergistic effect on the mechanical properties of the CPC. Indeed, they observed that this was true, and the flexural strength was raised to 43.2 MPa and work-of-fracture to 9.77 ± 0.75 kJ.m⁻², while the absorbable mesh created macropores after degradation.

Zhang and Xu ¹⁶² also confirmed the strong synergistic effect of combining chitosan with absorbable suture fibers with a CPC based on TTCP and DCPA. The flexural strength of CPC increased from only 2.7 ± 0.8 MPa (control) to 40.5 ± 5.8 MPa when both reinforcements were incorporated, while the work-of-fracture rose from 0.009 ± 0.003 kJ.m⁻² to 10.85 ± 2.05 kJ.m⁻². These results not only exceeded the reinforcement obtained from either chitosan or fibers alone but also demonstrated that the combined effect was greater than the sum of the individual contributions. According to the authors, this strengthening effect, characterized by an increase in flexural strength from 2.7 to 40.5 MPa and a two-order-of-magnitude improvement in toughness, could extend the applicability of apatite cements to moderate stress-bearing orthopedic applications.

With regard to fiber-reinforced CPCs, Canal and Ginebra ¹²⁹ emphasized in their review published in 2011 that important challenges must be addressed before clinical application of these materials. Specifically, they highlighted the need to (1) improve the homogeneous distribution of fibers within the cement paste, (2) increase fiber–CPC matrix adhesion, (3) enhance injectability, and (4) achieve a better understanding of both in vitro and in vivo behavior. Pérez et al. ¹⁵² described in their review comparable limitations for polymer-reinforced CPCs, where reinforcement is achieved by adding polymers as a second solid phase to the powder phase of CPC or by dissolving the polymer in the liquid phase of CPC. These authors noticed that polymers offer broad potential to enhance rheological behavior, mechanical strength, and even biological functionality, but further investigations are required to optimize the interactions between organic and inorganic phases and to ensure consistent biological outcomes.

On the other hand, following the reasoning of Bermúdez et al., ⁴³ Bohner ¹²² presented an important methodological and conceptual critique of how the mechanical strength data of CPCs have been reported in the literature, noticing that biomechanical loading involves not only compression but also tension and shear. He emphasized that the tensile strength of these cements is approximately one order of magnitude lower than their compressive strength, and, therefore, tensile and shear properties should also be evaluated and reported by the authors. He also pointed out that another important aspect to consider when analyzing the mechanical properties of CPCs is their potential to vary considerably after in vivo implantation. In view of this, it has to be underlined that after three decades of extensive research on CPCs, only a few large-animal studies have evaluated their mechanical performance following implantation into trabecular bone. Most of these investigations, however, focused on the performance of commercial cement formulations, rather than experimental or newly developed ones.^{105,106,112,170–174}

Tuning of degradation rate

Achieving a degradation rate aligned with bone formation remained a critical challenge in the third decade of CPC research. Both slow and rapid degradation impaired bone restoration, and tailoring this behavior to specific clinical needs was difficult, which resulted in extensive research on modulation of CPC biodegradation. Table 2 summarizes the key factors influencing CPC biodegradation and the strategies used to enhance this.

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Table 2.

Factors Affecting Biodegradation and Strategies to Enhance the Biodegradation of CPCs (as Deduced from the Review of He et al. ¹⁶⁹ )

Biodegradation influencing factors	Impact of biodegradation	Enhancement strategy
Macroporosity	Larger pores increase surface area and penetration of body fluids and cells, accelerate cement dissolution.	Addition of porogenic agents within the cement paste, which after the setting degrade faster than the cement itself, giving rise to the macroporosity or foaming the cement paste during its viscous stage, followed by setting to form a solid macroporous construct.
Microporosity and grain size	Higher microporosity and smaller grain size increase surface area, enhance interaction with body fluids and accelerate degradation.	Control grain size distribution and adjust microporosity via synthesis processes (e.g., modifying LPR).
CPC composition	Different compositions (e.g., HA, β-TCP, α-TCP, DCPD) affect degradation rates. Adjusting the material composition can optimize degradation rates.	Incorporate fast-resorbing compounds like DCPD or additives such as calcium sulfate and small acidic compounds to accelerate degradation. 170
Presence of ions	Ions play a crucial role in CPC degradation. They affect CPC dissolution by modifying the local ionic environment, which in turn influences the chemical reactions responsible for material degradation.	Creating ion-substituted or ion-doped calcium phosphates which do not only have a different solubility than the un-doped material, but may provide beneficial bone formation effects due to the release of the doping agents such as Si, Mg²+, Zn²+, Sr²+, Mn²+, CO3. 2–174
Presence of other additives	Additives (e.g., antibiotics, growth factors, anti-inflammatory drugs) modify the microenvironment, impacting CPC degradation positively or negatively.	Integrate bioactive compounds to regulate degradation rate based on therapeutic needs. 175–177

CPC, calcium phosphate cement; DCPD, dicalcium phosphate dihydrate; HA, hydroxyapatite; α-TCP, α-tricalcium phosphate.

As shown in Table 2, various factors influenced the biodegradation of CPCs, and several strategies were developed to enhance their degradation behavior, with porosity being a central factor. CPCs naturally exhibited high microporosity, but this alone was insufficient to ensure effective biodegradation. Since the microporosity of hardened CPCs is influenced by the L/PR of the unset cement—where water acts as a pore-forming agent—adjustment of the L/P ratio became a common method to modify both microporosity and degradation behavior. ¹⁷⁵

In addition, the particle size of the powder phase played a crucial role in microporosity formation. Ginebra et al. demonstrated that increasing the specific surface area of the α-TCP starting powder by a factor of five significantly reduced the size of the precipitated crystals, thereby doubling the specific surface area of the set cement. ¹⁷⁶

However, this microporosity was too small to allow for angiogenesis and rapid tissue colonization ¹⁷⁷ and synchronism between biodegradation and new bone formation. Consequently, the introduction of macroporosity (pores ranging from 50 to 125 µm in diameter) in CPCs was considered essential to facilitate bone ingrowth not only at the external surface but throughout the entire material. ¹⁷⁷ Initial efforts emerged in the late 2nd decade with the incorporation of pore-generating agents (porogens) such as mannitol, sucrose, or frozen sodium phosphate solution particles into the cement paste.^178–181 Upon CPC injection into the defect, these porogens dissolved faster than the cement matrix, thereby creating macropores within the ceramic structure. ¹⁷⁷ In the third decade, this research intensified, with continued use of water-soluble porogens,^{164,182–188} water-insoluble porogens, ¹⁸⁹ and the introduction of degradable pore-forming microparticles.^{185,190–202} Although porogen-based approaches were effective in producing macropores under laboratory conditions, they remained limited by the requirement of large amounts of porogen to ensure pore interconnectivity, which could negatively impact biocompatibility, mechanical strength, and handling properties. ¹⁷⁷ To address this, Lopez-Heredia et al. ²⁰³ recommended the incorporation of degradable polymer microparticles with a minimum size of 40 µm at 30 wt% in CPCs, which enhanced the mechanical response while preserving the overall structural integrity of the material.

An alternative approach to porogen agents involved creating macroporosity prior to setting by foaming the cement paste while it was still in a viscous state. Two commonly used foaming agents were hydrogen peroxide^204,205 or sodium bicarbonate.^{108,193,206–209} When these compounds were mixed into the cement paste, they reacted chemically and released gas (such as oxygen or carbon dioxide), forming bubbles within the paste. Once the paste had set, these bubbles remained as macropores. ¹⁰⁸

Though this approach proved effective in generating macropores, some important limitations were reported. For example, Klijn et al. ¹⁹³ evaluated three distinct strategies for introducing macroporosity into CPCs based on α-TCP and tested their performance in a rat skull bone augmentation model. In one formulation, macropores were created by CO₂ bubble formation during the setting reaction. In the other formulations, porosity was generated through the degradation of incorporated hollow or dense polylactic-coglycolic acid (PLGA) microspheres. Results showed that CPC-foaming promoted significantly greater bone formation and achieved higher bone apposition height compared with both PLGA-containing CPCs. However, morphological analysis revealed that the foaming resulted in uncontrolled porosity, as the pores in this CPC were irregular in size and distribution due to the upward migration of gas bubbles during cement setting.

Although highly porous CPC materials degraded faster and are related to the enhancement of bone regeneration mechanisms, such as angiogenesis and tissue ingrowth, their mechanical strength was significantly reduced by macropore formation. This was confirmed by both finite-element and mathematical models ²¹⁰ and by experimental studies.^{177,203,211–213} Consequently, an identified challenge was to achieve an optimum balance between porosity/biodegradability and sufficient mechanical strength. Furthermore, it has to be emphasized that the used animal model as well as the local bone loading conditions play a role in the CPC degradation rate. ¹⁶⁷

Enhancement of bone regenerative capacity

By the early 2000s, CPCs were recognized for their clinical suitability due to injectability, moldability, and biological properties like biocompatibility and osteoconductivity. However, their ability to induce bone formation in nonbone tissues (ectopic sites) showed mixed results, i.e., some studies reported that they promoted bone growth formation and bone cell differentiation,^214–216 while others demonstrated a limited effect on the osteogenic properties of CPCs.^217–219 This variability led to the concept of “intrinsic osteoinductivity” (the ability of the material itself to induce bone formation), which is driven by material parameters like topography, composition, and macro/microporosity. ²²⁰ Consequently, enhancing the bone regenerative capacity of CPCs became a research focus in the 3^rd decade, with incorporating growth factors emerging as a promising approach to improve their efficacy in bone regeneration.

Some extensively studied growth factors used to stimulate bone regeneration and enhance the osteogenic and angiogenic properties of CPCs include bone morphogenetic proteins (BMPs), transforming growth factor beta (TGF-β), fibroblast growth factors (FGF), ²²¹ and vascular endothelial growth factor. ²²² These biomolecules play an important role in bone repair and regeneration. ²²³ For example, BMPs are well known for their ability to induce bone formation, TGF-β for their regulating role in osteoprogenitor cell growth as well as differentiation, and FGF for playing a key role in skeletal vascularization and bone healing. ²

Table 3 lists studies dealing with the use of growth factors incorporated in CPCs.

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Table 3.

Growth Factors Used in CPCs to Enhance Osteogenicity in Experimental Animal Research

Growth factor	CPC formulation/sample/animal model	Outcomes
bFGF 224	β-TCP-CaSO4 cement Femoral Trepanation Defect Sprague-Dawley Rats	The addition of 0.25, 2.5, and 25 µg bFGF had no influence on the bone formation or resorption process after 3 weeks.
bFGF 225	β-TCP-CaSO4 cement Distal femoral Epiphysis Defect Merino-Mix sheeps	After 3 months the addition of bFGF inhibited the ingrowth of bone inside the defects, while the resorption of the cement was not influenced.
TGF-β1 226	Porous Calcibon® Cranial Defect Wistar Rats	Osteoid-like tissue formation was demonstrated in both experimental groups (TGF-β1-loaded and unloaded implants). TGF-β1 did not significantly increase osteoid-like tissue formation at 4 weeks. A greater number of pores showed inflammatory infiltrate in TGF-β1-loaded implants. No sign of CPC degradation in both experimental groups.
TGF-β1 227	Calcibon® + GM Femoral Defect Rabbits	Similar results for CPC + GM and CPC + GM + TGF-β1 regarding histological aspect, new bone formation and bone/implant contact. TGF-β1 loading demonstrated a significant effect on CPC degradation after 12 weeks of implantation.
BMP-2 175	α-BSM® Bilateral Tibial Osteotomy Canines	Larger callus formation in α-BSM/rhBMP-2-treated osteotomies: 86% greater than controls at 10 weeks.
BMP-2 176	α-BSM® Bilateral Tibial Osteotomy Canines	Higher AMC and increased bone formation in α-BSM/rhBMP-2 treated limbs compared with untreated or CPC-alone limbs at 4- and 8-weeks post-surgery.
BMP-2 228	Porous Calcibon® Cranial Defect New Zealand White Rabbits	Increased bone formation after 10 weeks: pore fill by bone (%): 53.0 ± 5.4% in defects treated with rhBMP-2/CPC versus 43.1 ± 10.4% with CPC alone. Enhanced early callus-like bone formation outside the implants observed in 2-week rhBMP-2-loaded CPCs. There was no evident degradation of the cement in either the rhBMP-2 loaded or unloaded implants.
BMP-2 229	Commercial CPC (CeredexTM) Periodontal Defect Labrador Dogs	Increased cementum regeneration in rhBMP-2/Ceredex™ sites compared with control (CeredexTM), but without functional periodontal ligament formation, and with higher incidence of ankylosis and root resorption.
BMP-2 230	Porous Calcibon® Ectopic Bone Defect New Zealand White Rabbits	Increased pore fill by bone (18% in site treated with rhBMP-2-loaded CPCs versus 0% in unloaded CPC). Degradation of CPCs could not be confirmed after 10 weeks of implantation.
BMP-2 231	α-BSM® Critical-size Mid-diaphyseal Radius Defect New Zealand White Rabbits	Complete cortical bridging and marrow space regeneration in all defects treated with 0.166 mg/mL rhBMP-2/α-BSM at 8 weeks. No complete bridging in defects treated with 0.033 mg/mL rhBMP-2/α-BSM or α-BSM alone. Accelerated α-BSM degradation in the presence of 0.166 mg/mL rhBMP-2 compared with that in the other two groups.
BMP-2 232	Calcibon® or Calcibon® + PLGA microparticles Critical-sized Cranial Defect Wistar Rats	Bone formation was significantly higher in the high-dose group (±10 µg of rhBMP-2 per cement disc) when compared with the plain implant after 12 weeks. Release out of rhBMP-2 from low-dose implants (±2 µg of rhBMP-2 per disc) was not sufficient to enhance bone formation. Implant degradation (≈50%) only seen in high-dose group at 12 weeks.
BMP-2 233	Commercial CPC (Rebone Biomaterial Ltd Co) + GM Trabecular Bone Defect Osteoporotic goats	In vitro tests showed that the rhBMP-2/GM/CPC system released larger amounts of rhBMP-2 compared with CPC without microspheres. Bone mineralization rate was greater in rhBMP-2/GM/CPC than in rhBMP-2/CPC 45 days post-implantation.
BMP-2 and bFGF 234	Commercial CPC (Rebone Biomaterial Ltd Co) Ectopic bone defect Nu/Balb-c nude rats	rhBMP-2 and bFGF at a ratio of 2:1 (100:50 ng/mL) enhanced BMSC proliferation, differentiation, and bone formation more than either growth factor alone.
BMP-2 and bFGF 235	Commercial CPC (Rebone Biomaterial Ltd Co) Bone defect around Branemark Implant Beagle dogs	Bone labeling data showed significantly elevated MAR values with BMP-2/bFGF in the defects, as compared with BMP-2 or bFGF. BMP-2 and bFGF synergistically enhance bone repair.

AMC, area of the mineralized callus; BMSC, bone marrow stromal cells; CPC, calcium phosphate cement; GM, gelatin microspheres or microparticles; MAR, mineral apposition rate, pore fill by bone (%): percentage of pore surface area occupied by bone.

Several studies, as summarized in Table 3, reported a beneficial effect on bone regeneration of incorporating growth factors into CPC. For example, Edwards et al. and Seeherman et al.^236–238 installed α-BSM cement loaded with rhBMP-2 in bone defects created in monkeys, dogs, and rabbits and reported a significant increase in bone formation for the BMP-2 loaded specimens. Furthermore, Seeherman et al. ²³⁸ observed that the addition of rhBMP-2 enhanced the degradation rate of the cement compared with control samples without the growth factor because of the possible ability of rhBMP-2 to affect osteoclast activity. ²²⁴

Ruhe et al. ²²⁵ and Kroese-Deutman et al. ²²⁶ conducted pioneering studies in which they incorporated rhBMP-2 into a macroporous commercial CPC (Calcibon®), with porosity generated by CO₂ foaming. The growth factor was deposited and lyophilized on the scaffold surface one day prior to implantation. In rabbit cranial and subcutaneous models, rhBMP-2-loaded cements promoted greater bone formation within the pores after 10 weeks compared with controls, although no material degradation was observed. In the cranial defect model, ²²⁵ an additional in vitro release assay showed that less than 10% of the rhBMP-2 was released over four weeks. This result contrasted with the significant bone formation observed in vivo, revealing a gap between the measured release and the biological outcome. According to the authors, this discrepancy might have been due to the inability of in vitro tests to accurately replicate in vivo conditions or to the possibility that rhBMP-2 remained biologically active even when bound to the cement.

Bodde et al ²²⁷ incorporated PLGA microparticles loaded with rhBMP-2 into Calcibon® cement and implanted it into rat skulls. Two doses were tested: a high dose (∼10 μg) and a low dose (∼2 μg) of rhBMP-2 per cement disc. The results showed that bone formation was dose-dependent, with significantly greater bone formation in the high-dose group compared with the plain implant after 12 weeks, but rhBMP-2 release from low-dose implants was not sufficient to enhance bone formation. Scintigraphic imaging of the longitudinal in vivo release by using 131-I-labeled rhBMP-2 showed a sustained release of both rhBMP-2 doses with a retention of about 70% of the initial dose for both rhBMP-2 formulations.

Similarly, Link et al. ²²⁸ used a composite of Calcibon® and gelatin microparticles loaded with rhTGF-β1 to enhance the bone response and mechanical strength of rabbit femoral defects. They observed a gradual increase in mechanical strength over time, but the addition of TGF-β1 did not improve bone regeneration. Histological and histomorphometric analyses showed similar results for both formulations (CPC loaded and unloaded with gelatin microspheres) in terms of new bone formation and bone–implant contact. On the other hand, the TGF-β1-loaded specimens showed significantly more degradation than the nonloaded specimens, which was supposed to be due to an enhanced bone remodeling activity around the cement caused by the released TGF-β1.

Also, Li et al. ²²⁹ investigated the effect of rh-BMP2 release from CPCs on bone healing. In their study, a cement composed of TTCP and DCPA was used, and rhBMP-2 was incorporated either by adding it directly to the setting liquid or by embedding it into gelatin microspheres. These formulations were installed into lumbar bone defects created in osteoporotic goats to evaluate the potential of a sustained rhBMP-2 release system for enhancing bone regeneration. In vitro tests indicated that the rhBMP-2/GM/CPC system released larger amounts of rhBMP-2 compared with CPC without microspheres. The in vivo experiments confirmed that after 45 days of implantation, there was a higher rate of bone mineralization, and the mechanical strength measured by push-out tests was superior at both 45 and 140 days postimplantation. However, the authors highlighted certain limitations, notably the small sample size (n = 6 per experimental group), which limited the ability to draw a definitive conclusion regarding the overall efficacy of the treatment in promoting bone repair.

The synergistic effect of combining multiple growth factors in CPC with the same composition was demonstrated in two studies performed by Wang et al. In the first study, ²³⁰ bone marrow stem cells (BMSCs) were extracted from beagle dogs and seeded onto a commercial porous CPC loaded with different ratios of rhBMP-2 and bFGF. The constructs were implanted into ectopic sites in rats. The results showed that a 2:1 ratio of rhBMP-2 to bFGF (100:50 ng/mL) significantly enhanced BMSC proliferation and osteogenic differentiation compared with the application of either growth factor alone. In the second study, ²³¹ the authors investigated the effects of a similar amount (50 ng/mL) of rhBMP-2 and bFGF incorporated into porous CPC on the repair of peri-implant bone defects in beagle dogs. The combined application of both growth factors resulted in more effective bone regeneration, which supported the presence of a synergistic interaction between rhBMP-2 and bFGF.

In contrast with the above mentioned, several studies reported neutral or even negative effects of incorporating growth factors into CPCs for bone repair. For instance, Niedhart et al. ²³² and Maus et al. ²³³ investigated the use of β-TCP–CaSO₄ cement loaded with recombinant human bFGF (rhbFGF) in femoral defects of rats and sheep, respectively. In the rat model, histomorphometric analysis showed that different doses of rhbFGF did not enhance bone ingrowth compared with the control. The authors suggested that this outcome might be due to an inadequate release profile of bFGF (too rapid or too slow) to be effective. In the sheep model, the presence of rhbFGF appeared to inhibit bone formation. It was hypothesized that this could be due to the release of rhbFGF at low concentrations that may have had an inhibitory rather than a stimulatory effect on bone healing.

Huse et al. ²³⁴ observed that the addition of rhTGF-β1 did not enhance the osteogenic potential of a microporous Calcibon® CPC rigidly fixed by a titanium and dense sintered HA “umbrella” screw on the rat skull. Even a detrimental effect was noted, as most pores in the rhTGF-β1-loaded implants were occupied by fibrous tissue with associated inflammatory infiltrate, while the specimens without rhTGF-β1 showed fibrous tissue with decreased cellular activity. The authors emphasized that the lack of positive effect from TGF-β1 does not necessarily mean that the growth factor is ineffective. They highlighted that several factors can influence the osteogenic potential of bone grafts. One important factor might be the degree of contact between the graft and the recipient bone surface. Their histological analysis showed that in their case, there was no direct contact between the CPC and the skull bone, which may have compromised bone regeneration.

Overall, the incorporation of growth factors into CPC seems to demonstrate an enhanced bone healing, but the effect depends on which growth factors are used, their concentration and delivery kinetics, and the specific cement formulation. Therefore, controlled release strategies and optimization of dosage and release are critical for safe and effective outcomes.

Proposed applications of CPCs in drug delivery systems and tissue engineering

CPCs were initially developed as injectable bone fillers, primarily valued for their biocompatibility and osteoconductivity. However, research over the decades progressively transformed the perception of CPCs, repositioning them as multifunctional platforms that served as carriers for cells, biomolecules, ions, and drugs. This evolution expanded their scope from passive defect fillers to active therapeutic systems, capable of enhancing bone regeneration and delivering targeted therapies. Several processing strategies were developed to optimize CPCs for regenerative medicine (Table 4). These included the creation of scaffolds with interconnected macropores, granules with high porosity, microspheres for injectable applications, and “ready-to-use” premixed formulations. These diverse processing strategies were critically reviewed by Ginebra in 2010. ²³⁵ highlighting approaches that optimized CPCs for use in regenerative medicine, and they are summarized in Table 4, along with their descriptions and potential clinical benefits.

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Table 4.

Various Processing Techniques That Enhance the Properties of CPCs for Regenerative Medicine with Their Respective Objectives and Benefits for Clinical Outcomes ²³⁹

Technique	Objective	Clinical benefits
CPC-based scaffolds	Create interconnected macropores using techniques like leaching, foaming, emulsion, freeze drying, templates, and rapid prototyping.	- Enhances cell ingrowth and vascularization, improving bone formation. - Supports colonization by cells, crucial for effective bone regeneration.
CPC-based granules	Break larger CPC blocks into granules (100–400 µm).	- High porosity (35–60%) of CPC-based granules facilitates bone integration. - Improves surgical handling and adaptability to defects, enhancing regenerative outcomes.
CPC-based microspheres	Develop spherical beads for bone grafting, cell microcarriers, or microscaffolds.	- Serve as microscaffolds for cell distribution and tissue regeneration. - Can be injected with gel-like carriers, improving surgical application. - Enable creating hybrid cell-microcarrier scaffolds for optimized cell growth. - Function as drug delivery systems.
“Ready to use” CPCs	Develop premix formulations with liquid or paste phases for better homogenization.	- Reduces mixing errors, ensuring predictable implant performance. - Shortens surgery time and improves implant precision. - Enhances mechanical properties by reducing bubbles and ensuring uniform mixing. - Supports cell survival during the setting phase, crucial for regenerative success.

CPC, calcium phosphate cement.

In this same literature review article, Ginebra also identified critical challenges in the field of regenerative medicine and outlined future trends in CPC research for tissue engineering (see Table 5).

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Table 5.

The Main Critical Challenges Identified by Ginebra ²³⁹ Regarding the Use of CPCs in Regenerative Medicine and Future Trends in Field

Challenge	Description of challenge	Future trends
Cell incorporation	Difficulty in incorporating cells within cement pastes, aiming to use injectable CPCs for osteogenic cell delivery.	Research on optimal cell types and improved methods for incorporating osteoprogenitor cells within CPCs for better cell delivery.
Cell survival	Harsh pH, high ionic strength, and cement hardening stiffness threaten cell viability.	Developing formulations of CPCs with near-physiological pH and adding neutralizing additives in the cement paste.
Nutrient exchange	Limited nutrient and waste exchange between cells and the media hinders cell survival.	Creating interconnected macroporous structures for better fluid flow.
Bioactive molecule integration	The incorporation of bioactive molecules like growth factors into CPCs is complex and can be influenced by the type of molecule and its concentration.	Exploring encapsulation techniques and optimizing molecule release profiles.
Long-term stability	Concerns over CPC stability, degradation, and tissue integration over time.	Investigating long-term degradation behavior and improving CPC formulations for enhanced stability.

CPC, calcium phosphate cement.

As shown in Table 5, challenges remained in adapting CPCs for tissue engineering due to the difficulty of incorporating cells within cement pastes, maintaining viable cells in the matrix, incorporating bioactive molecules such as growth factors, and persisting issues of long-term stability. Thus, the future research in CPCs for regenerative medicine, according to Ginebra, was expected to focus on three areas: (i) control of the cell viability and cell incorporation, (ii) integration of bioactive molecules and autologous or allogenic osteoprogenitor cells into the CPC formulations, and (iii) investigation of the stability and long-term function of CPCs.

Beyond their role in tissue engineering, CPCs have also been investigated as carriers for ions and various therapeutic agents. Their intrinsic porosity, combined with the fact that hardening occurs at room or body temperature ² and the ability to undergo controlled dissolution, makes them suitable for the incorporation and release of (1) antibiotics (i.e., gentamicin, tetracycline), (2) osteoporosis-related agents (i.e., bisphosphonates), and (3) chemotherapeutics (i.e., cisplatin) ²⁴⁰ to induce a local therapeutic effect. In addition, ions influencing bone remodeling (Sr²⁺, Si⁴⁺, Zn²⁺, Mg²⁺) and antimicrobial ions (e.g., Ag⁺) were incorporated to further enhance biological performance.

Regarding investigations of CPCs as carriers for therapeutic agents and ions in 2012, Ginebra et al. ² published an overview of various strategies employed already for using CPCs as drug delivery systems. These approaches are summarized in Table 6, along with the main findings, conclusions, and future perspectives identified by Ginebra for each approach.

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Table 6.

The Main Findings, Conclusions, and Future Perspectives Outlined by Ginebra ² Concerning the Strategies Employed for Using CPCs as Drug Delivery Systems

Approach	Findings/conclusions	Recommendations
Incorporation of LMW drugs	Antibiotics were the most studied, often displaying burst release profiles consistent with the Higuchi model and maintained activity after CPC incorporation.	Increase release studies using unset, drug-loaded cements to better simulate surgical conditions.
Incorporation of HMW biomolecules	Incorporation of growth factors and other HMW biomolecules has been achieved, but concerns persist about potential denaturation during setting and balancing bound vs. released states due to protein affinity toward CPCs.	Conduct detailed studies on protein binding and release mechanisms to design more targeted carriers for protein delivery.
Incorporation of ions	Research focused ions influencing bone remodeling (Sr²+, Si4+, Zn²+, Mg²+) and antimicrobial ions (Ag+). It remained unclear whether the improvements resulted from the ions’ active roles or from changes in physicochemical properties of CPCs.	Quantify ion release kinetics and correlate them with bone healing to clarify the underlying mechanisms of improvement.
General release studies	Both in vitro and in vivo release studies have shown promising results, although evaluation methods vary considerably.	Standardize evaluation methods and conduct more in-depth investigations of release mechanisms to improve comparability and understanding of results.

CPC, calcium phosphate cement; HMW, high molecular weight; LMW, low molecular weight.

According to Ginebra’s overview of CPCs in drug delivery, the proposal to use CPCs in this field showed promising results but still required further refinement of release mechanisms. Collectively, these advances illustrate the potential of CPCs as therapeutic systems that address both regenerative and pharmacological needs. A schematic representation (Fig. 6) can help to visualize the evolution of CPCs from passive fillers to multifunctional therapeutic systems.

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FIG. 6.

Evolution of CPCs from passive fillers to multifunctional therapeutic systems.

Summary of the third decade of CPC research

Major achievements in CPC development between 2003 and 2013 were (1) strong movement toward reinforcement strategies primarily through polymer and fiber incorporation (including polymeric and ceramic fibers), aimed at improving compressive and flexural strength, as well as fracture toughness, without losing the favorable biological properties of CPC on bone healing. This trend is quantitatively illustrated in Figure 7, which compiles reported values for compressive strength, flexural strength, and fracture toughness from the cited studies during this period, benchmarking them against the characteristic property ranges of cortical and cancellous bone, (2) The use of soluble porogens (e.g., mannitol/sucrose, PLGA, and gelatin microspheres) to produce interconnected macroporosity, which improves vascularization and bone ingrowth into the CPCs after leaching/degradation of the microspheres, (3) techniques to load and control release of osteoinductive biomolecules (e.g., rhBMP-2-loaded PLGA microspheres) to enhance in vivo bone regeneration while reducing burst release and side effects of the biomolecules, and (4) the evolvement of CPCs from bone void filler into multifunctional drug delivery platform. Major achievements included antibiotic-loaded CPCs for infection control, bisphosphonate- and strontium-modified CPCs for osteoporosis management, and chemotherapeutic-loaded CPCs for bone tumors.

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FIG. 7.

Strength and toughness of reinforced CPCs relative to cortical and cancellous bone ranges.

Maturation and therapeutic agents

Between 2013 and 2023, the number of publications dealing with CPCs seemed to have reached a platform, as shown in Figure 8.

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FIG. 8.

Annual number of publications by type (article, review article, meeting abstract and conference papers), extracted from the Web of Science Database by searching the keywords “calcium phosphate cement” or “calcium phosphate bone cement” or “hydroxyapatite cement” or “brushite cement” in All Fields. 1st decade: initial phase, 2nd decade: first growth phase, 3rd decade: peak growth phase and 4th decade: sustained high activity.

This suggests that research interest in CPCs has reached a mature and stable phase. Literature analysis revealed that research during the fourth decade still encompassed a broad range of topics. However, major research focus was on osteogenesis and angiogenesis, enhancement of mechanical and handling properties, development of CPCs with antibacterial properties, optimization of degradation rates to match new bone formation, and use in 3D printing.

Regarding osteogenesis and angiogenesis, recent studies have demonstrated that inclusion of inorganic elements in CPCs, such as Sr, Si, Mg, and Zn, can effectively stimulate major cellular signaling pathways involved in bone regeneration.^{134,239,241–243} In addition, the addition of other types of biological materials, such as proteins, polysaccharides, and blood components, has been reported to be effective in addressing the osteogenic potential of CPCs. Two very recent review articles covering this topic have been published by the research group of Sok Kuan Wong.^17,244

In terms of mechanical properties, research continued during the fourth decade to elucidate and further enhance CPC performance.^35,245–247 A comprehensive overview of CPCs, emphasizing their biological behavior, mechanical performance, and degradation characteristics after implantation into trabecular bone, was presented by Schröter et al. in a 2020 publication. ²⁴⁸ According to the presented data, brushite-based CPCs exhibit lower compressive strength, reaching up to 52 MPa, whereas apatite-based CPCs can achieve values as high as 80 MPa, and these properties were generally maintained after in vivo implantation in the trabecular bone of large animals. In addition to mechanical strength, reported tensile strengths were approximately 3.5 MPa for apatite-based and 1.3 MPa for brushite-based cements, while shear strengths averaged around 9.8 MPa and 2.9 MPa, respectively. These values are comparable with those of trabecular bone, but they remain considerably lower than those of cortical bone. Based on these data, Schröter et al. emphasized that the major limitation of CPCs (both brushite and apatite) lies in their intrinsic brittleness, which often results in the formation of cracks, as was observed also in some of the large animal studies.

In light of these mechanical limitations, research in fourth decade continued to explore reinforcement strategies aimed at improving CPC toughness. Besides the incorporation of polymeric fibers into CPCs, ²⁴⁹ the incorporation of graphene and its derivatives as a secondary phase has attracted particular attention, as it might improve the mechanical properties (toughening) and maintain or even stimulate the biological properties of CPCs.^250–258 The presence of this secondary phase in the cement matrix had shown the ability to hinder crack propagation, resulting in improved fracture toughness and mechanical strength of the CPCs. ²⁵⁰ However, these studies are still in an early stage, since many works involving the incorporation of graphene report only compressive strength data, without including, for example, fracture toughness values.

Also, attention was paid to the methods as used for mechanical testing. For example, Paknahad et al. ²⁵⁹ evaluated the failure behavior of an α-TCP cement reinforced with PVA fibers. By integrating experimental testing with numerical modeling, the researchers developed a 3D gradient-enhanced damage model. This model was validated through three-point bending and tensile experiments and accurately predicted the mechanical response of the reinforced CPCs under different loading conditions using a single set of parameters. This approach provided a robust foundation for future experimental and computational studies aimed at developing CPCs suitable for load-bearing applications.

Progress during this period was also made in integrating CPCs with therapeutic agents for localized drug delivery. A significant advancement in this area was marked by two comprehensive reviews: one by Fosca et al. (2022) ²⁶⁰ and another by Pylostomou et al. (2023). ²⁶¹ Fosca et al. provided an in-depth overview of a broad range of studies on CPC–drug systems and classified them based on structural and compositional characteristics to identify specific factors affecting drug release. Pylostomou et al. focused specifically on factors influencing local drug release from CPCs in the context of bone cancer therapy. Together, these reviews offer a base for guiding future research and optimizing CPC-based drug delivery strategies.

The incorporation of antiosteoporosis and chemotherapeutic agents was also described. Van Houdt et al. ²⁶² investigated the development and efficacy of a composite based on CPC combined with PLGA and loaded with antiosteoporosis agent alendronate (ALN). The study demonstrated that CPC formulations showed negligible ALN release, while CPC/PLGA formulations exhibited a 2-week lag phase followed by a biphasic sustained release over 148 days. This controlled release was attributed to the degradation of PLGA, which created a porous network for drug diffusion. Using a rat femoral condyle bone defect model in osteoporotic animals, ALN-loaded CPC/PLGA composites demonstrated stimulatory effects on bone formation both within and outside the defect region.

Farbod et al. ²⁶³ described CPCs incorporating drug-loaded HA nanoparticles (HA NPs) as a method for releasing chemotherapeutic platinum-bisphosphonate complexes, thereby enhancing drug delivery. It demonstrated that these injectable cements not only ensure effective drug release but also exhibit antiproliferative effects on cancer cells. They suggested that the binding affinity of bisphosphonate ligands can be tuned to facilitate efficient loading of high doses of therapeutic complexes onto HA NPs, which may optimize drug release profiles. These findings indicated that injectable CPCs can be rendered chemotherapeutically active, providing a novel approach to treating large bone defects while minimizing damage to surrounding healthy tissues.

In addition, efforts continued to optimize the degradation rates of CPCs (mainly by enhancing their macroporosity) to better align with new bone formation. A comprehensive review published in 2021 by Lodoso et al. ¹ discussed the main strategies to enhance CPC macroporosity, along with their respective advantages, limitations, and recent developments. For example, the use of foaming agents and introduction of water-soluble porogens have major disadvantages that (1) the final CPC construct has a low initial strength and due to the lower density of the gas compared with the CPC, there is a tendency for higher porosity to concentrate in the upper regions of the material. They also noted that direct in vivo injection was not advisable, as the gas could diffuse into surrounding tissues and trigger adverse reactions; (2) high amounts of water-leachable porogen have to be included in CPC to obtain an interconnected porosity, which might reduce the mechanical properties as well as endanger the biocompatibility of the CPC due to the release of dissolution products as well as reduce the mechanical properties.

Lodoso et al. ¹ consider rapid prototyping (RP) and the incorporation of degradable polymeric porogens as particularly promising techniques to create macroporosity in CPC. RP allows the fabrication of CPC scaffolds with highly controlled pore architectures. This represented a significant improvement over earlier methods developed in the 2000s, which had faced challenges with pore interconnectivity and scaffold reproducibility. However, a major limitation emphasized by the authors is that RP-based scaffolds are not injectable and therefore cannot be delivered through minimally invasive procedures. This drawback can be overcome by the incorporation of degradable polymeric porogens, such as PLGA, into CPC. Yet, this strategy also has drawbacks: PLGA requires several days to weeks before hydrolysis begins. Finally, the authors highlighted that although enhancing CPC macroporosity is crucial for bone regeneration, it may inevitably compromise handling properties and/or mechanical strength.

Furthermore, research was performed to add antimicrobial properties to CPC for the prevention and treatment of bone infections. Recent trends and future perspectives for the use of antibacterial CPC strategies have been reported by Liu et al. ²⁶⁴ Antibiotics, antimicrobial peptides, and antimicrobial metals (like, silver, copper, zinc) can be incorporated into CPC for the treatment of bone infections using the same approach as used for the incorporation of growth factors. Such antimicrobial CPCs can be applied as surface coating on or as slurry around an orthopedic prothesis to have a local antimicrobial effect.

Additional research to 3D printing of CPCs for bone reconstruction was done.^265–269 This approach allows the preoperative fabrication of implants that fit precisely to the patient’s anatomy, leading to the optimization of esthetical results and a reduction in risks and surgery time. Bertol et al., ²⁶⁶ for instance, evaluated the feasibility of manufacturing customized implants for craniofacial reconstruction via 3D printing processes using α-TCP as the powder phase and an aqueous solution of Na₂HPO₄ as the binder phase. The implants were previously generated in a computer-aided design based on the patient’s tomographic data. The fabrication of the implants was carried out in a commercial 3D powder printing system. The fit of the 3D-printed implants was measured by three-dimensional laser scanning and by checking the right adjustment to the patient’s anatomical biomodel. The printed parts were shown to present a good degree of fitting and accuracy.

Patent filings during the fourth decade reflected a trend comparable with those found for scientific publications. This alignment can clearly be seen through reading the work of Vezenkova et al., ²⁷⁰ who did a patent search ranging from 2017 to 2022 and compiled a table listing the reported benefits and the intended applications for the filed patents. The patent search and analysis by Khairoun et al., ²⁷¹ indicates that the field of CPCs remains under active development and notices opportunities for further research. However, the number of patent applications for CPCs has declined since 2013, which corroborates with the leveling out of the number of publications.

Clinical trials for CPC products were still ongoing during this fourth decade, with notable progress. An example is HydroSet® as developed by Stryker Corporation. This material was intended to be used in the repair of neurosurgical burr holes, contiguous craniotomy cuts, and other cranial defects, as well as the augmentation or restoration of the bony contour in the craniofacial skeleton. This material was granted 510(k) clearance by the FDA in 2016, while CE approval was received in 2014. ²⁷² Another development was the approval in 2018 of a self-setting CPC containing rhBMP-2, developed by Shanghai Rebone Biomaterial Co., Ltd., PR China. ²⁷³

Summary of the fourth decade of CPC research

Fourth decade was characterized by maturation and stabilization of CPC research. Specific attention was put onto improving mechanical properties of CPC as well as the incorporation of therapeutic agents into CPC.