A review of calcium phosphate cements and acrylic bone cements as injectable materials for bone repair and implant fixation

Abstract

Treatment of bone defects caused by trauma or disease is a major burden on human healthcare systems. Although autologous bone grafts are considered as the gold standard, they are limited in availability and are associated with post-operative complications. Minimally invasive alternatives using injectable bone cements are currently used in certain clinical procedures, such as vertebroplasty and balloon kyphoplasty. Nevertheless, given the high incidence of fractures and pathologies that result in bone voids, there is an unmet need for injectable materials with desired properties for minimally invasive procedures. This paper provides an overview of the most common injectable bone cement materials for clinical use. The emphasis has been placed on calcium phosphate cements and acrylic bone cements, while enabling the readers to compare the opportunities and challenges for these two classes of bone cements. This paper also briefly reviews antibiotic-loaded bone cements used in bone repair and implant fixation, including their efficacy and cost for healthcare systems. A summary of the current challenges and recommendations for future directions has been brought in the concluding section of this paper.

Keywords

Calcium phosphate cement acrylic bone cement fracture repair bone void filler implant fixation

Introduction

Osteoporosis is a condition characterized by low bone mass and can lead to increased susceptibility to fractures in the elderly.¹ Worldwide, fractures resulting from osteoporosis affect approximately one woman in three and one man in five over the age of 50 years.² In addition, approximately 2 million cases of trauma or disease-related bone fractures occur every year in the United States alone, with an estimated annual direct cost of $10 billion.³ The repair rate of a bone defect is dependent on the wound size. When the defect size is greater than the healing capacity of bone, the fibrous connective tissue becomes dominant in the bone defect.^4,5 Bone grafting is considered to be the gold standard for the treatment of traumatic bone defects.⁶ However, bone grafting has the drawback of donor site morbidity and limited availability,⁷ and may be associated with post-operative complications.^8,9

As an alternative to autologous bone grafts, a wide variety of synthetic inorganic/organic or biological materials have been explored as bone graft substitutes for the treatment of bone defects.¹⁰ For thousands of years, humans have made use of seashells, nuts, and so forth as potential bone substitute materials.¹¹ With the introduction of tissue-engineering approach and its successful application in clinical trials,^12,13 the concept of bone graft substitutes has evolved significantly. The scaffold-based tissue engineering aims to restore and maintain normal function in injured and diseased bone.^14–16 A fundamental requirement for tissue-engineered bone grafts is the ability to integrate with the host bone, while providing the capacity for load-bearing and remodeling.¹⁷

Under certain circumstances, such as age-related bone disintegration and osteoporosis, bone reconstruction involving solid scaffolds may not be a feasible approach.^18–20 Therefore, minimally invasive alternatives are sought in certain orthopedic and maxillofacial procedures.^18,21,22 Osteoporotic fracture of vertebral bodies is another example where minimally invasive procedures are sought. Vertebral fractures deform spine and cause chronic pain while limiting patients’ freedom of movement.^23,24 Vertebroplasty (VP) makes use of an injectable bone cement (IBC) to restore the mechanical strength and stiffness of the vertebral body,²³ although its effectiveness as a treatment modality for vertebral compression fractures is yet to be established.²⁵

Injectable biomaterials with appropriate functional properties and self-setting characteristics at physiological temperatures are essential for the success of these interventions. IBCs have been used for bone reconstruction and in a wide range of bone augmentation procedures, such as orthopedic and maxillofacial surgeries.^26–28 Bone cements are obtained by mixing a biomaterial in powder form with a liquid, which can set once implanted as a paste within the body.²⁹ In terms of their chemical composition, IBCs may be regrouped as calcium phosphate cements (CPCs), calcium sulfate cements (CSCs), acrylic bone cements (ABCs), and filamentary composite materials.²⁶ As an alternative, these IBCs can be classified on the basis of their orthopedic applications. While ABCs are used for high and medium load-bearing applications, CPCs and CSCs are generally used for medium and low load-bearing applications.²⁶

Unlike the first-generation biomaterials that were mostly inert, the second-generation bone substitute materials intended for orthopedic applications are expected to elicit a response from the body to favor osseointegration.^30,31 This is true whether the biomaterial is designed for mechanical fixation (e.g., metallic screws) or for the repair or replacement of diseased or injured bone (e.g., bone graft extenders and substitutes).³² Calcium phosphates have been substantially used in the past few decades as second-generation biomaterials.^31,33,34 The chemical composition of CPCs is similar to the natural bone, which means that the release of $C a^{2 +}$ and $P O_{4}^{3 -}$ ions may promote osteoconduction and osteogenesis.^35,36 In addition, calcium phosphates offer fast or slow degradation rates depending on their Ca/P ratio³¹ and the type of filler used.^37–39 Scaffolds made of calcium phosphates have been used in dentistry since the early 1970s and in orthopedics since the 1980s.⁴⁰ Hence, calcium phosphates are gaining a special interest as bone graft substitutes in periodontal and osteochondral regeneration and as potential carriers in controlled release systems.^41–43

As for ABCs, poly(methyl methacrylate) (PMMA) has evolved from ophthalmological and dental applications^44,45 to orthopedics,^46,47 for the fixation of prosthetic implants as well as for remodeling osteoporotic and vertebral fracture repair.^48,49 In addition, PMMA has been used for total joint replacement procedures, including hip, knee, ankle, elbow, and shoulder.^26,50–53 It has been shown that the injection of PMMA can enhance the stability of hip fracture fixation.⁵⁴ Also, it has been shown that an osteoporotic proximal femur can be mechanically stabilized through the injection of PMMA into it.⁵⁵ In general, PMMA bone cement has a relatively high mechanical strength and offers a suitable curing time and ease of application.⁵⁶ The ISO standard 5833 for bone cements requires a compressive strength of ⩾ 70 MPa.⁵⁷

There is a growing body of literature on the development of new bone graft substitutes, making use of synthetic inorganic/organic compounds as well as composites of demineralized bone with various self-setting pastes. Several review papers have provided overviews of the current state-of-the-art in bone cement development and their clinical applications.^48,58–66 This paper briefly covers the main classification of bone graft substitutes for bone reconstruction. The main body of the paper describes the most common injectable bone substitute materials for orthopedic use, placing the emphasis on CPCs and ABCs. Some alternative formulations for these cements have been listed at the end of each section, while discussing the prospect of the new formulations and future directions in the field, including the advantages and drawbacks of antibiotic-loaded bone cements (ALBCs).^28,67–69

Bone graft substitutes

Bone is an organic–inorganic tissue with a hierarchical architecture that makes it a micro-/nano-composite.^11,70 The mineralized bone matrix consists of an organic phase mainly composed of collagen (~35% dry weight), a mineral phase of carbonated apatite (~65% dry weight), as well as other non-collagenous proteins that form a stimulating microenvironment for cellular functions.¹¹ Collagen, a triple helix with a diameter of ~1.5 nm,⁷¹ has a Young’s modulus of 1–2 GPa and an ultimate tensile strength of 50–1000 MPa, whereas the mineral hydroxyapatite (HA) possesses a Young’s modulus of ~130 GPa and an ultimate tensile strength of ~100 MPa.¹¹ The collagen phase is responsible for the rigidity, viscoelasticity, and toughness of bone, while the mineral phase provides structural reinforcement and stiffness.¹¹ The mechanical properties of bone depend on its composition and structural organization, such as mineralization, collagen fiber orientation, porosity, as well as on trabecular vs. cortical bone architecture.^11,72

Autologous bone grafts are considered as the gold standard in orthopedic surgery, but they are limited in availability and associated with post-operative complications.^73–75 A more future-oriented alternative to bone grafting is to make use of synthetic bone graft substitutes.⁷⁶ Despite decades of research attempts to design bio-inspired materials, mimicking the properties of natural bone tissue is not a trivial undertaking.^42,72,77 Nevertheless, according to animal model studies, it has been reported that materials with compressive properties on the low end of human cancellous bone can encourage new tissue growth.^78–80

Synthetic bone graft substitutes are often in the form of granules, typically with diameters ranging between 0.1 and 5 mm, or porous blocks (or sponges) (Figure 1(a) and 1(b)).⁷⁶ Hydraulic cements represent a class of bone graft substitutes that harden after in situ implantation or injection (Figure 1(c)).⁷⁶ This includes cements such as calcium sulfate hemihydrate (plaster of Paris) and CPC,⁷⁶ as well as magnesium phosphate cement.^81,82 Besides granules, porous blocks, and hardening pastes, non-hardening pastes (putty) have also been proposed as bone graft substitutes. These pastes typically consist of a mixture of granules and a highly viscous hydrogel (Figure 1(d)).⁷⁶

Figure 1.

The four main types of bone graft substitutes: (a) granule; (b) block or pre-form; (c) hydraulic cement; and (d) putty. The presence of pores in the size range of 0.1–1.0 mm (as seen in (a) and (b)) is essential for a rapid bone ingrowth. Reproduced with permission from Bohner.⁷⁶ Copyright © 2010 Elsevier Inc.

Table 1(a) provides a summary of some commercially available bone graft substitutes and their mechanisms of action.⁸³ Commercial bone graft substitutes can be classified as (i) allograft based, (ii) ceramic based, (iii) factor based, and (iv) polymer based. Allografts such as demineralized bone matrix (DBM) are osteoinductive⁸⁴ and provide a biological stimulus (proteins and growth factors),⁴¹ but may pose the risk of disease transmission and immunogenicity.^85,86 Ceramics such as calcium phosphates, calcium sulfates, and bioactive glasses have been extensively used as bone graft substitutes.^16,40 Growth factors such as bone morphogenetic proteins (BMPs) have the ability to recruit and signal to mesenchymal progenitor cells to differentiate toward a bone-forming lineage.^40,83 Finally, non-degradable/degradable polymers as well as a variety of their organic/inorganic composite systems make up the market for some commercially available strips and injectable pastes (Table 1(a)). Compositions of some commercially available IBCs are summarized in Table 1(b). This review paper focuses on the two main classes of injectable bone graft substitutes: (i) ceramic based (CPCs) and (ii) polymer based (ABCs).

Table 1.

(a) Summary of selected commercially available bone graft substitutes. Reproduced with permission from Ricciardi and Bostrom.⁸³ Copyright © 2013 Elsevier Inc. (b) Compositions of some commonly used commercially available injectable bone cements. Modified from Lewis.²⁶ Copyright © 2011 Wiley Periodicals, Inc.

(a) Selected commercially available bone graft substitutes.

Class	Commercial Product	Composition	Claimed Mechanism of Action	Formulations
Allograft based	DBX®(Synthes)	DBM with sodium hyaluronate carrier	Osteoinduction, osteoconduction	Putty, paste, mix, injectable, strips
	Grafton®(Medtronic)	DBM fibers w/ w/o cancellous chips	Osteoinduction, osteoconduction, osteogenesis	Putty, strips, crunch, gel, paste
	Orthoblast®(Citagenix)	DBM and cancellous chips in reverse phase medium	Osteinduction, osteoconduction	Putty, paste
Ceramic based	Norian SRS®(Synthes)	Calcium phosphate	Osteoconduction, bioresorbable	Injectable paste
	Vitoss®(Stryker)	Beta-tricalcium phosphate and bioactive glass	Osteoconduction, osteoinductive w bone marrow aspirate, bioresorbable	Putty, strips, injectable, morsels, shapes
	ProOsteon®(Biomet)	Corraline hydroxyapatite and calcium carbonate	Osteconduction, bioresorbable	Injectable granules or block
	BonePlast®(Biomet)	Calcium sulfate	Osteoconductive, bioresorbable	Paste
	Osteoset®(Wright)	Calcium sulfate	Osteoconductive bioresorbable	Pellets
Factor based	Infuse®(Medtronic)	rhBMP-2 on bovine collagen carrier	Osteoinductive	Bovine collagen carrier
Factor based	OP-1®(Stryker)	rhBMP-7 with type I collagen carrier	Osteoinductive	Putty (with carboxymethyl-cellulose addition), paste
Polymer based	Healos®(DePuy)	Crosslinked collagen and HA	Osteoconductive, osteogenic with bone marrow	Strips
Polymer based	Cortoss®(Stryker)	Non-resorbable polymer resin with ceramic particles	Osteoconductive	Injectable paste

(b) Selected commercially available injectable bone cements.

Cement Type / Brand Name	Composition / Constituents^a,b	Manufacturer / Supplier
Acrylic bone cements
CMW™1	Powder (40.00 g): 35.54 g PMMA, 3.64 g BaSO₄, 0.82 g BPO Liquid (18.37 g): 18.22 g MMA, 0.15 g DMPT, 25 ppm HQ	DePuy CMW Blackpool, UK
Palacos® R	Powder (40.00 g): 33.55 g poly(methyl acrylate, MMA), 6.13 g ZrO₂, 0.32 g BPO, 1.00 mg chlorophyll Liquid (18.78 g): 18.40 g MMA, 0.38 g DMPT, 0.40 mg chlorophyll	Heraeus Kulzer GmbH, Hanau, Germany
Surgical Simplex® P	Powder (40.00 g): 29.40 g poly(MMA, styrene), 6.00 g MMA, 4.00 g BaSO₄, 0.60 g BPO Liquid (18.79 g): 18.31 g MMA, 0.48 g DMPT, 80 ppm HQ	Stryker, Michigan, USA
Calcium phosphate cements
Biopex®	α-TCP, TTCP, DCPD	Mitsubishi Materials Corp., Tokyo, Japan
Cerament®	Powder: 40% hydroxyapatite (HA), 60% calcium sulfate (CaS) Liquid: iohexol and saline	BoneSupport, Sweden
chronOS Inject®	Powder: 42 wt.% β-TCP, 21 wt.% MCPM, 31 wt.% β-TCP granules, 5 wt.% Mg hydrogen phosphate trihydrate, < 1 wt.% sodium hydrogen pyrophosphate Mg(SO₄)₂ Liquid: 0.5% solution of sodium hyalurone	Oberdorf, Switzerland
Graftys® QuickSet Graftys® HBS	Powder: calcium phosphate salts and HydroxyPropylMethylCellulose (HMPC) Liquid: phosphate-based (Na₂HPO₄) aqueous solution	Graftys, France
Norian® Skeletal Repair System	α-TCP, CaCO₃, MCPM	Synthes, Inc.,West Chester, PA, USA

MMA: methylmethacrylate; BPO: benzoyl peroxide; DMPT: N, N-dimethyl-p-toluidie; HQ: hydroquinone; TCP: tricalcium phosphate; TTCP: tetracalcium phosphate; DCPD: dicalcium phosphate dihydrate; MCPM: monocalcium phosphate monohydrate.

Compositional details for the acrylic bone cements and chronOs Inject^® cement were taken from products’ brochures.

Calcium phosphate cements

Composition and cement formation

Some available calcium phosphates, with their composition, standard abbreviations and solubility data are listed in Table 2.⁸⁷ The clinical potential of calcium phosphate materials further increased when a self-setting CPC was developed in the early 1980s.^88,89 CPCs are generally formed by combining one or more calcium phosphate powders with a liquid phase, which is usually water or an aqueous solution,^90–92 although water-immiscible liquids have also been used to improve handling and cement properties.⁹³ Upon mixing, the produced paste is able to set and harden within the body. Unlike ABCs, which undergo polymerization during hardening, CPCs set as a result of dissolution and precipitation, while the entanglement of the precipitated crystals is responsible for hardening.⁹⁰ A recent study by Mellier et al.⁹⁴ made use of ovine whole blood as a liquid phase for CPCs. The formation of a 3D clot-like network and its interaction with the precipitated apatite crystals resulted in a microstructure that was more sensitive to biological degradation and promoted new bone formation.⁹⁴

Table 2.

Molar ratio (Ca/P)	Compound	Formula	Solubility at 25°C		pH stability (range)
Molar ratio (Ca/P)	Compound	Formula	−logK_s	g l⁻¹	pH stability (range)
0.5	Monocalcium phosphate monohydrate (MCPM)	Ca(H₂PO₄)₂·H₂O	1.14	∼18	0.0–2.0
0.5	Monocalcium phosphate anhydrous (MCPA or MCP)	Ca(H₂PO₄)₂	1.14	∼17	^a
1	Dicalcium phosphate dihydrate (DCPD), mineral brushite	CaHPO₄·2H₂O	6.59	∼0.088	2.0–6.0
1	Dicalcium phosphate anhydrous (DCPA or DCP), mineral monetite	CaHPO₄	6.9	∼0.048	^a
1.33	Octacalcium phosphate (OCP)	Ca₈(HPO₄)₂(PO₄)₄·5H₂O	96.6	∼0.0081	5.5–7.0
1.5	α-Tricalcium phosphate (α-TCP)	α-Ca₃(PO₄)₂	25.5	∼0.0025	^b
1.5	β-Tricalcium phosphate (β-TCP)	β-Ca₃(PO₄)₂	28.9	∼0.0005	^b
1.2–2.2	Amorphous calcium phosphates (ACP)	Ca_xH_y(PO₄)_z·nH₂O, n = 3–4.5, 15–20% H₂O	^c	^c	∼5–12^d
1.5–1.67	Calcium-deficient hydroxyapatite (CDHA or Ca-def HA)^e	Ca₁₀−_x(HPO₄)x(PO₄)₆−x(OH)₂−x (0 < x < 1)	∼85	∼0.0094	6.5–9.5
1.67	Hydroxyapatite (HA, HAp or OHAp)	Ca₁₀(PO₄)₆(OH)₂	116.8	∼0.0003	9.5–12
1.67	Fluorapatite (FA or FAp)	Ca₁₀(PO₄)₆F₂	120	∼0.0002	7–12
1.67	Oxyapatite (OA, OAp or OXA)^f	Ca₁₀(PO₄)₆O	∼69	∼0.087	^b
2	Tetracalcium phosphate (TTCP or TetCP), mineral hilgenstockite	Ca₄(PO₄)₂O	38–44	∼0.0007	^b

Stable at temperatures above 100°C.

These compounds cannot be precipitated from aqueous solutions.

Cannot be measured precisely. However, the following values were found: 25.7 ± 0.1 (pH 7.40), 29.9 ± 0.1 (pH 6.00), 32.7 ± 0.1 (pH 5.28). The comparative extent of dissolution in acidic buffer is: ACP >> α-TCP >> β-TCP > CDHA >> HA > FA.

Always metastable.

Occasionally referred to as “precipitated HA” (PHA).

The existence of OA remains questionable.

As Figure 2 shows, CPCs can be classified by the number of components in the solid phase (single or multiple), type of setting reaction (hydrolysis or acid–base reaction), setting mechanism and microstructure evolution during setting, and the type of end product.⁹⁰ Although a large number of formulations has been used to produce CPCs,^87,95–100 the CPCs developed to date have only two different end products: precipitated HA or brushite (dicalcium phosphate dihydrate, DCPD).⁹⁰ This is anticipated, since HA is the most stable calcium phosphate at pH > 4.2, whereas brushite is the most stable one at pH < 4.2.⁹⁰

Figure 2.

Classification of calcium phosphate cements, with examples of the most common formulations. From top to bottom the cements are classified by the type of end product (apatite or brushite), number of components in the solid phase (single or multiple), type of setting reaction (hydrolysis or acid–base reaction), setting mechanism and microstructure evolution during setting. Scanning electron micrographs of set apatite and brushite cements obtained by the hydrolysis of α tricalcium phosphate (α-TCP) and by reaction of β-TCP with MCPM (monocalcium phosphate monohydrate) respectively, are also shown. Reproduced with permission from Ginebra et al.⁹⁰ Copyright © 2012 Elsevier Inc.

Changing the powder-to-liquid ratios can lead to CPCs with a variety of self-setting times.^101,102 According to ISO/DIS 18531 for CPCs, the setting time is the “time required from the start of powdered agent and liquid agent blending until hardening of the cement.”¹⁰³ Reducing the powder-to-liquid ratio of CPCs increases the injectability¹⁰⁴ while increasing the setting time and affecting the mechanical properties of CPCs after setting.¹⁰⁵ Although CPCs have demonstrated self-setting both in simulated body fluid (SBF) and in vivo in rats, the properties of CPCs set in SBF are expected to differ from those of CPCs set in vivo.¹⁰²

From a biological perspective, CPCs are attractive because they have proven to be biocompatible, osteoconductive, and bioresorbable,^106–110 while they offer an intrinsic microporous structure¹⁰⁰ for the transport of nutrients and metabolic waste products.¹¹¹ Given the excellent biocompatibility and osteoconductivity of CPCs, these cements are great candidates for various clinical applications. For example, CPCs are considered as the most suitable injectable biomaterials to accommodate narrow and irregular bone defects.¹⁸ In addition, CPCs can be injected to form a bioactive scaffold in situ under physiological conditions.¹⁰¹ The lack of macropores is considered a major limitation for the widespread use of CPCs.¹¹² Introducing macropores in CPCs has shown to be beneficial, as macroporous structures can facilitate bone ingrowth and aid in fast resorption of CPCs.¹¹¹ For example, the porosity of CPCs can be enhanced via the use of calcium carbonate,^100,113 polymers,^114,115 and foaming agents.^116,117 In addition, the porosity and microstructure of CPCs can be altered by adjusting the process parameters, such as the liquid-to-powder ratio and the particle size of the powder phase.⁹⁰ Ease of handling is also of paramount importance for the commercial success of injectable bone graft substitute materials.⁷⁶ Some examples of poor handling include the use of non-reproducible or complicated mixing procedures,⁷⁶ poor injectability of a cement paste,^38,91 or the release and migration of fine particles caused by the lack of paste cohesion.^91,118

Cement injectability

Separation of the solid and liquid phases during cement delivery is the primary cause of poor injectability of CPCs.⁸⁹ Some studies have developed theoretical models to gain an in-depth understanding of the solid–liquid phase separation.^119–121 Figure 3 depicts the schematics of the phase separation during paste extrusion and their location in the extruder.⁸⁹ Phase separation mechanisms identified in different studies include: (i) filtration in the barrel, where the drainage of the liquid phase caused by the exerted pressure leads to consolidation of the solid phase;¹²² (ii) suction, caused by dilation of the powder network while flowing into the die;¹¹⁹ and (iii) filtration in the needle (or die land), exacerbated with the formation of solid mats.^89,123 (Figure 3).

Figure 3.

Phase separation mechanisms observed during extrusion of pastes and their location in the extruder. Schematic diagram is not to scale. Reproduced with permission from O’Neill et al.⁸⁹ Copyright © 2017 Elsevier Inc.

Two key aspects should be considered when testing the injectability of CPCs:³⁸ (1) defining a suitable experimental setup for injectability measurements, and (2) investigating adequate compositions and parameters that could affect injectability. The injectability of CPCs is commonly assessed using the syringe ejection method.^124–127 For example, Wang et al.¹²⁷ used a needle with an inner diameter of 1.6 mm for injectability measurements. The as-prepared paste was poured into the syringe 2 min after mixing the cement powder and liquid. Then, 1 kg vertical compressive load was applied on top of the plunger and the expelled paste was collected for 2 min. The following equation was used to calculate the injectability:

I n j e c t a b i l i t y % = \frac{\begin{array}{l} P a s t e v o l u m e e x p e l l e d \\ f r o m t h e s y r i n g e \end{array}}{\begin{array}{l} T o t a l p a s t e v o l u m e b e f o r e \\ i n j e c t i n g \end{array}}

(1)

Nevertheless, this method should be used with caution and may lead to inconsistent results, particularly when certain additives are present in the formulation. For example, adding polyvinyl alcohol (PVA) into a cement paste has been shown to highly enhance the injectability of CPCs.¹²⁸ However, this increase of cement injectability in the presence of PVA has been attributed to the increase of setting time only.³⁸ Hence, both the injectability and setting time should be reported, while making sure that the injectability of a cement is measured after a specific time interval (e.g., 50% of the setting time).

In a review paper, Habraken et al.⁹¹ highlighted major research achievements related to calcium phosphate materials in the past 15 years: as biologically active agents, carriers for gene or ion delivery, and bone graft substitutes. In another review paper, Zhang et al.¹²⁹ discussed the role of key processing parameters (e.g., particle size, cement composition, and additives) on the setting properties of CPCs and their handling and mechanical performance.

Some alternative formulations

O’Neill et al.⁸⁹ elaborated on the properties of CPCs that are essential for their clinical success. The paper discusses how the presence of powder or liquid additives can have a positive or detrimental effect on the delivery process of CPCs. For example, the size and proportion of solid additives relative to the bulk calcium phosphate powder can affect its packing ability.⁸⁹ It has been shown that the addition of fine fillers (~1 µm in diameter) alters the packing ability, reduces the water demand, and increases injectability.^89,97 The addition of large glass beads or microspheres of poly(lactic-co-glycolic) acid (PLGA) either decreases or increases injectability depending on the size and wt% of the beads.^95,96 Other additives, such as gelatinized starch, can increase the compressive strength, compressive modulus, and strain energy density of CPCs.¹³⁰

Some studies have aimed at improving the cohesion of CPCs, arguing that these cements can get disintegrated in the presence of water and blood.^18,131 For example, the setting properties of a cement are affected when blood or biological fluids are present at the injection site.¹³¹ Gelling agents have been used in some studies to improve the handling properties and cohesion of CPCs,^131,132 including glycerin, gelatin, cellulose derivatives, alginic acid salts, and chitosan.^131,133 Alternatively, water-immiscible carrier liquids have been proposed to improve the cohesion of injectable CPCs.⁹³ Table 3 shows a selection of studies investigating a few alternative CPC formulations and summarizes the key findings of these studies.

Table 3.

Selection of studies investigating alternative CPC formulations.

Authors	Primary Powder	Additive(s)	Additive Amount (wt%)	Findings
Heinemann et al.⁹³	α-TCP	Castor oil ethoxylate 35 and hexadecyl-phosphate	14.7 and 4.9	The oil-based liquid generated a water-immiscible cement that was injectable and showed improved cohesion, a higher compressive strength, and prolonged shelf life compared with the conventional cement. The product of the setting reaction was identified to be nanocrystalline HA. A resorbability similar to the conventional cement was anticipated because of the identical mineral structure of the set products.
van Houdt et al.¹³⁶	α-TCP	PLGA	30 and 50	PLGA microspheres or milled PLGA particles led to similar CPC porosity and degradation. Increasing the amount of PLGA porogen accelerated the in vitro degradation of CPC/PLGA. Using a distal femoral condyle in vivo model in rats, the CPC/PLGA formulations prepared with milled PLGA showed favorable bone responses, particularly at 50 wt% PLGA.
Smith et al.¹³⁵	α-TCP	Glucose microparticles (GMPs)	10, 20, 30, 40	Glucose microparticles (GMPs) with two sizes ranges (100–150 µm and 150–300 µm) were used as porogens to introduce macroporosity within CPCs. Setting times increased for GMP/CPC formulations compared with control CPC. The local pH was maintained at a neutral pH range over 8-week of in vitro degradation study, but the inclusion of GMPs had a detrimental effect on the compressive strength.
Tödtmann et al.¹³⁸	α-TCP	Mineralized collagen type I	2.5	Cement formulations with mineralized collagen type I, either in combination with osteocalcin or phosphoserine, resulted in an increase of new bone formation in cyst-like jaw defects of minipigs. An increase in the resorption rate was reported for the cement formulations containing osteocalcin and phosphoserine.
Schumacher et al.^98,157	α-TCP	Strontium (Sr)	0.72–2.21	Strontium (II) $(S r^{2 +})$ was introduced into the cement either by adding SrCO₃ to an α-TCP-based cement precursor mixture (A-type) or by substitution of CaCO₃ by SrCO₃ during precursor composition (S-type).The S-type cement (SrCPC) prolonged the setting time and increased the compressive strength by up to 90%, while enabling the release of $S r^{2 +}$ .⁹⁸ SrCPCs appeared to significantly enhance cell proliferation and osteogenic differentiation of hMSCs.¹⁵⁷
Zhu et al.¹⁵⁸	β-TCP/ HA	Strontium (Sr)	3–12 (molar ratio)	The presence of strontium and its concentration played a significant role in the phase compositions, setting time, compressive strength, in vitro degradation rate and cytotoxicity of the biphasic cement. Pure biphasic cements exhibited a higher stiffness compared with Sr-modified cements. Also, a higher amount of Sr-β-TCP in biphasic cement was associated with a higher degradation rate.
Zheng et al.³⁵	Calcium phosphate silicate (CCPSC)	Carbon fiber (1 µm diameter)	2	Carbon fiber-reinforced CCPSC was reported to possess moderate biodegradability and favorable osteoconductivity. Carbon fibers significantly improved the mechanical properties of the CCPSC cement. A small amount of osteoid found at the heterotopic site 1 month after implantation was attributed to potential short-term osteoinductivity of the cement formulation.

The resorbability of calcium phosphates can be controlled through the regulation of Ca/P ratio (see Table 2). Compounds with Ca/P ratio of less than 1 are not suitable for biological implantation because of the higher speed of hydrolysis with decreasing Ca/P ratio.³¹ While β-tricalcium phosphate (β-TCP) with Ca/P ratio of 1.5 has been classified as a resorbable calcium phosphate material, the resorbability of HA depends on several factors. For example, sintered HA with a stoichiometric Ca/P molar ratio of 1.67 may not show resorbability. Nevertheless, HA becomes resorbable in the presence of certain impurities and structural defects, or when its grain size is reduced to nano-scale.¹³⁴ To enhance resorbability of CPCs, pore-forming additives, such as water-soluble polymers, glucose,¹³⁵ biodegradable polymers (e.g., PLGA^136,137), collagen,¹³⁸ and biphasic calcium phosphate (BCP) granules,¹³⁹ have been proposed. However, these additives may alter the physiochemical characteristics of CPCs, including the setting time, viscosity, dispersibility, and compressive strength.^139,140

Self-setting gelatin-based HA foams have been proposed for the treatment of bone defects.¹¹² The team reported a suitable injectability and cohesion for the formulation, an interconnected porous structure, and good in vivo biocompatibility in rabbits. Additional functionalization with soybean extract slightly reduced the degradation rate and positively influenced bone formation in a critical-size bone defect in rabbits.¹¹² Similarly, an injectable cement made of calcium-deficient HA and foamed gelatin has been shown to promote bone ingrowth.¹⁴¹ The bubble volume caused by foaming appeared to increase the ductility of the cement. Hence, the brittle characteristic of the ceramic was reduced in the presence of the organic phase.¹⁴¹ Incorporation of polylactic acid (PLA) and phosphate glass fibers into CPCs has also been proposed to achieve a less brittle CPC matrix.¹⁴²

A combination of nano-hydroxyapatite (nHA) as the solid phase and a modified sodium phosphate solution as a mixing liquid has been proposed to improve the cohesion of CPCs in aqueous media as well as in human blood.¹⁸ HA has a broad range of clinical applications, has been used in bone cements for the repair of craniofacial and dental defects,^106,143 and has proven to be osteoconductive as a 3D scaffold and bone graft substitute material.^144–151 In particular, nHA has a higher surface area and a higher reactivity¹⁵² favoring cell adhesion and proliferation of mesenchymal stems cells (MSCs), alkaline phosphatase activity, calcium deposition, and osteogenic gene expression.^153–155 Hence, nHA has a greater potential for stimulating new bone growth compared with conventional HA, making it superior for bone reconstruction.^43,150,156

Stimulation of bone formation at the defect site in osteoporotic patients has motivated the development of some new cement formulations. In particular, ionic additives are often considered as viable alternatives to growth factors, not only because they are considerably less expensive, but also there is a lower risk that their delivery would result in adverse effects.¹⁵⁹ For example, strontium (II) $(S r^{2 +})$ can promote bone formation and inhibit bone resorption, and is frequently used in the treatment of osteoporosis.^98,160 A local release of strontium ions into the bone defect site has proven to be a more effective approach than orally administered strontium.^98,161 In vitro trials of strontium-modified cements (SrCPC) have revealed an enhanced cell proliferation and osteogenic differentiation of human bone-marrow-derived MSCs (hMSCs).¹⁵⁷ In a study by Thormann et al.,¹⁶¹ SrCPC-treated rats showed a statistically higher bone volume fraction (BV/TV) compared with the defect areas filled with CPC. Also, there was more bone formation at the bone–cement interface for the SrCPC compared with CPC.

The efficacy of the local bisphosphonate (BP) delivery via calcium-deficient apatite (CDA) granules and CPCs has also been investigated.^110,162 Systemic administration of BPs is the main pharmaceutical treatment option for osteoporosis, as BPs effectively increase bone density and prevent bone loss, and thereby reduce the risk of vertebral and non-vertebral fractures.^110,163 The efficacy of a CPC loaded with BP was evaluated by Verron et al.¹¹⁰ for the local BP delivery in a preclinical large animal model (ewes). To quantify bone formation, the team used three regions of interest (ROI) in microcomputed tomography (µCT). Implantation of CPC ± BP significantly improved bone volume fraction (BV/TV) in each ROI of implanted vertebral bodies compared with non-implanted ones. The BV/TV appeared twice as large on the largest ROI (1.2 mm) in the presence of CPC-BP compared with non-implanted vertebrae. These effects on BV/TV were slightly superior with CPC-BP (e.g. 85% at 1.2 mm) compared with CPC alone (79.5% at 1.2 mm).

Acrylic bone cements

Composition and cement formation

For several decades, ABCs have played an important role in orthopedic surgery.⁴⁸ The commercial ABCs are marketed as a two-component system (see Table 1(b)).²⁶ The powder phase primarily consists of PMMA (82–89 wt%), an inorganic radiopacifying agent such as barium sulfate or zirconium dioxide (10–15 wt%), as well as benzoyl peroxide (BPO; 0.5–2.6 wt%) that acts as a catalyst for the polymerization reaction. The liquid phase is largely methyl methacrylate (MMA) monomer (98 wt%), with 2 wt% N, N-Dimethyl-p-toluidine (DmpT) to accelerate polymerization.¹⁶⁴

The polymerization reaction of an ABC is schematically shown in Figure 4(a–c).¹⁶⁴ During this reaction, the DmpT causes BPO to decompose and produce a benzoyl radical and a benzoyl anion (Figure 4(a)). Then, the benzoyl radicals initiate the polymerization of MMA (Figure 4(b)). The active centers formed at this step combine with additional molecules to form a polymer chain (Figure 4(c)).¹⁶⁴ The paste formed during the reaction is a viscous fluid, which allows the polymerizing cement to be injected by the surgeon into the site of interest (for example, in VP and balloon kyphoplasty (BKP), into a fractured vertebral body),^48,165 or into the prepared bone canal prior to implanting an implant (hip replacement).¹⁶⁴

Figure 4.

Schematic diagram showing (a) the decomposition of BPO leaving a benzoyl radical and a benzoyl anion; (b) how these benzoyl radicals initiate polymerization of MMA; and (c) formation of a polymer chain. Reproduced with permission from Dunne and Ormsby.¹⁶⁴ Copyright © 2011 Dunne and Ormsby. (d) A typical curing curve for PMMA bone cement where T_max is the maximum temperature reached, T_set is the setting temperature and T_amb is the ambient temperature. Reproduced with permission from Dunne and Ormsby.¹⁶⁴ Copyright © 2011 Dunne and Ormsby. (e) Viscosity during setting of three commercial PMMA cements at 37°C. Reproduced with permission from Nicholas et al.¹⁷² Copyright © 2007 Springer.

A typical temperature-versus-mixing time curve during the curing of a PMMA bone cement is shown in Figure 4(d),¹⁶⁴ revealing a highly exothermic polymerization reaction. The peak temperature (T_max) is the maximum temperature attained during polymerization (curing) process.¹⁶⁶ This is considered a drawback, as the high temperature during curing may reach up to 110°C.^56,164 High temperatures can lead to thermal necrosis of the bone cells and extensive bone damage,^164,167 since collagen denatures with prolonged exposure to temperatures in excess of 56°C.^168,169 In addition, PMMA bone cement is not absorbable, with no functions of bone conduction or induction.¹⁷⁰

The dough time shown in Figure 4(d) represents the time elapsed between the initial mixing and the time the paste reaches a homogeneous dough-like state. As specified in the British Standard BS 7253 (ISO 5833),⁵⁷ at this point, the cement dough no longer sticks to powderless surgical gloves (typically 2–3 min after initial mixing).¹⁶⁴ The working time is the period between the end of the dough time until the cement can no longer be manipulated.¹⁶⁴ Finally, the setting time of the cement is usually defined as the time when the temperature rise is halfway between the maximum temperature (T_max) and the ambient temperature (T_amb), as described in ISO 5833.⁵⁷

Cement characteristics

Commercially available brands of ABC have similarities and differences on the basis of their chemical composition, bead size of the prepolymerized PMMA in the powder, powder particle size distribution, molecular weight of the powder, and molecular weight of the cured cement (see Table 1(b)).²⁶ The viscosity rise as a function of mixing time differs among the various brands.^26,171–173 Figure 4(e) shows how the viscosities of three commercial PMMA bone cements at 37°C increase with time and reach a maximum of ~75 kPa.s in all three cases, at which the cements turn into an elastic solid.¹⁷² The viscosity can influence the injectability of the cement, as well as its leakage, extent of retention within the vertebral body, and final mechanical properties.^174–176 Several research and review papers have reported on the kinetics of cure reactions for thermosetting resins.^177–181 Studying the effect of experimental factors on the reaction kinetics may guide the development of ABCs, and therefore, the design of experiments can be an effective approach to bone cement design, by enabling the identification of the dominant factors that influence its properties.¹⁸²

There is a limited number of literature reviews on viscoelastic properties of IBCs.^26,183–185 The rheological properties of a cement may play a significant role in the formation of pores during cement mixing and delivery. Such pores could contribute to crack formation and implant loosening over time.¹⁶⁴ In addition, properties such as creep, stress relaxation, and damping can influence the long-term performance of a cemented implant. A recent review paper by Lewis²⁶ has identified the key properties of ABCs for six commercial formulations in orthopedic use. The paper investigated how the viscoelastic properties of the cements were influenced by several relevant parameters, such the monomer-to-polymer ratio, polymerization pressure, cement mixing method, duration of cement mixing, length of aging time, test medium composition, test frequency, and temperature.²⁶ It was reported that the particle size distribution of the powder, as well as the temperature and frequency of the tests, was among the most influential variables affecting the damping behavior of the ABCs.²⁶

Some alternative formulations

Despite the widespread use of commercially available plain PMMA bone cement, mainly due to its outstanding compressive strength and functional performance, it has several major drawbacks. PMMA is not remodeled in the body,^1,6 and the exothermic reaction of PMMA cement fixation can impair fracture healing. Moreover, its high elastic modulus can result in stress shielding and implant loosening.¹⁸⁶ Hence, the use of PMMA cement to stabilize and/or reinforce fractured vertebral body (for example, in VP and BKP) can lead to extensive bone stiffening and potential fractures at the adjacent vertebral bodies.¹⁸⁷ In addition, the mechanical failure of the cement can lead to premature failure of an implant.¹⁸⁸ In a recent review of the literature, Lewis¹⁸⁹ has summarized the properties of some nanofiller-loaded PMMA bone cement composites.

Many studies have aimed at developing alternative formulations that could eliminate these drawbacks. On the basis of the shortcomings addressed in these studies, the new formulations can be grouped into 16 categories as described by Lewis.¹⁹⁰ This review paper focuses on the formulations that aim to reduce the peak temperature, improve the mechanical properties, and impart bioactivity to PMMA cements. For example, adding chitosan or starch-stabilized polyethylene glycol to PMMA bone cement has been shown to reduce the maximum curing temperature of the cement.^56,166 Although chitosan did not appear to affect the tensile properties, it increased the compressive strength of the chitosan/PMMA cement.¹⁶⁶ Chitosan is expected to degrade over time in vivo as the new bone tissue forms, which could enable the generation of a porous network over time for bone ingrowth.

Some studies have investigated the copolymerization of MMA with other materials, such as the hydrophilic acrylic acid (AA)¹⁹¹ as well as 2-hydroxyethyl methacrylate (HEMA) and diethyl 2-(methacryloyloxy)ethyl phosphate (DMP).¹⁹² Copolymers of MMA and AA, modified using 4-iodo phenyl isocyanate and 3,4,5-triiodo phenyl isocyanate as the radiopacifying agents, generated radiopaque composites as bone cement materials.¹⁹¹ Silk sericin has also shown to enhance hydrophilicity of MMA-based copolymers while imparting antibacterial ability.¹⁹³ In another study, the presence of DMP/PHEMA segments improved the thermal behavior of the MMA-based copolymer via reducing the glass transition temperature.¹⁹² Moreover, the incorporation of a commercial block copolymer, Nanostrength^® (NS), into the liquid phase of the PMMA bone cement has been shown to improve the fracture toughness of the cement.¹⁹⁴ Table 4 lists some other alternative PMMA cement formulations and briefly summarizes key findings of the cited studies.

Table 4.

Selection of studies investigating alternative PMMA bone cement formulations to improve the mechanical properties or to impart magnetic or bioactive characteristics.

Author	Additive	Quantity of the additive (wt%)	Findings
Kawashita et al.¹⁹⁵	Magnetite nanoparticles (Fe₃O₄)	40, 50, 60	Increasing the Fe₃O₄ content increased the setting time of the cement and decreased the peak temperature during setting. The cements had a compressive strength of ~90 MPa, adequate for clinical applications. Unlike the standard PMMA cement, the surface temperature of the samples containing 40% and 50% Fe₃O₄ increased rapidly in a magnetic field of 300 Oe (over 70°C within tens of seconds).
Li et al.¹⁹⁶	Magnetite nanoparticles (Fe₃O₄)	30 and 50	The addition of 30% Fe₃O₄ did not significantly alter the compressive strength of the cement or the proliferation of rat fibroblast cells on the cement. The mean diameter of Fe₃O₄ (11–35 nm) appeared to affect the heating capability in AC magnetic fields, depending on the magnetic field (40–300 Oe) and frequency (100 kHz and 600 kHz).
Miola et al.¹⁹⁷	Silver-containing bioactive glass	30	The silver release was greater on the first week, when the rate of infection development was higher, and enabled significant antibacterial effect toward S. aureus strain. The glass acted as a nucleating agent for the precipitation of hydroxyapatite on the surface of the cement by a biomimetic mechanism. Mouse fibroblast L-929 cell line was used to verify the non-cytotoxic effect of the composites.
Mousa et al.¹⁹⁸	Apatite-wollastonite glass ceramic (AW-GC)	50 and 70	Adding 70 wt% dry-silanated AW-GC to PMMA cement maintained or increased the bending strength, fracture toughness, and osteoconductivity, but reduced the elastic strain. Small-diameter PMMA beads improved the handling properties, the mechanical properties and the bioactivity of the PMMA-based cement.
Verné et al.⁶²	Ferrimagnetic bioactive glass ceramic (SC-45)	10–20	The composites supported the growth of hydroxyapatite on their surface after 28 days of immersion in a simulated body fluid (SBF). Iron leaching test revealed a negligible release of iron, while cytotoxicity assays with human osteoblasts indicated good cell viability and biocompatibility.
Serbetci et al.¹⁹⁹	Hydroxyapatite (HA)	7.7 and 14.3	Very few foreign body giant cells were reported for HA-containing cement, which was attributed to the improved in vivo biocompatibility (in rabbits) in the presence of HA. Addition of HA into the PMMA cement also increased the viscosity, improved the workability, decreased the polymerization temperature, and increased the modulus. The changes in the compressive strength depended on the PMMA brand.
Li et al.¹⁸⁶	Nano-HA coated recombinant human collagen	15	The new formulation reduced the elastic modulus from 1.91±0.08 GPa to 1.21±0.12 GPa. The in vivo results in rabbits suggested a significantly higher bone growth for the modified PMMA group when compared with their standard PMMA counterparts.
Arens et al.¹⁸⁷	Bone marrow from sheep and hydroxyapatite (HA)	15 (HA)	The maximum temperature during curing decreased from ~61°C to 38°C and the Young’s modulus decreased from 1830 to 740 MPa by adding 7.5 ml bone marrow to 23 ml cement. The samples had a porosity of up to 51%, with pores ranging between 30 and 250 μm.
Hernández et al.²⁰⁰	Strontium hydroxyapatite (SrHA)	10 and 20	Incorporation of SrHA negatively influenced the handling and mixing properties, emphasizing the need for surface treatment of SrHA particles. The surface-treated particles improved the compressive strength of the cements, although this effect was significant for 20 wt% SrHA only.
Slane et al.²⁰¹	Silver nanoparticles (AgNPs)	0.25–1.0	The cement formulations modified with AgNPs significantly reduced biofilm formation on the surface of the cement, demonstrating a potential for preventing bacterial adhesion, but showed no antimicrobial activity against planktonic bacteria. Mechanical properties were not substantially changed when compared with the standard cement.

Various magnetic materials, such as magnetite (Fe₃O₄), have been shown to generate heat in an alternating magnetic field.^125,195 Hence, these materials have been widely studied for the minimally invasive treatment of cancer through hyperthermia of metastatic bone tumors.^202,203 Adding magnetite has shown to increase the setting time of PMMA cement and decrease the maximum temperature during setting.¹⁹⁵ The addition of 30–50 wt% Fe₃O₄ did not seem to significantly impact the compressive strength of PMMA-based bone cements.^195,196

Reduction of bacterial adhesion on the cement surface can be achieved by favoring the direct bonding of the cement to the bone.¹⁹⁷ One strategy for improving the bone-bonding ability is through the addition of bioactive fillers such as bioactive glasses and glass–ceramics^{62,197,198,204,205} as well as HA.^{199,200,204,206} Bioactive glass acts as a nucleating agent for the precipitation of HA on the surface of PMMA cement.^62,197 It has been well established that HA is an osteoconductive material,^31,147 and thereby can improve the in vivo biocompatibility of PMMA bone cement.^49,199 Biological materials, such as collagen¹⁸⁶ and bone marrow,¹⁸⁷ have also been investigated in combination with bioactive fillers.

Arens et al.¹⁸⁷ studied the effect of adding freshly harvested bone marrow from sheep to a PMMA cement formulation containing 15 wt% HA. Besides a reduction in modulus in the presence of bone marrow, a higher initial cement viscosity immediately after mixing with bone marrow was deemed beneficial, as a result of a reduced risk of leakage upon injection.^{23,187,207,208} In vertebral body augmentation procedures, such as VP and BKP, cement leakage can result in life-threatening cardiac injury²⁰⁹ as well as pulmonary and cerebral embolism.^210–212

Li et al.¹⁸⁶ added 15 wt% recombinant human collagen, coated with HA nanoparticles, to a commercial PMMA cement (C-PMMA) during the early dough stage. Both C-PMMA and the modified cement (MC-PMMA) were implanted in rabbits to examine bone ingrowth and bone affinity index after 4, 12, and 24 weeks. The team reported a significantly higher bone growth for the MC-PMMA group, when compared with their C-PMMA counterparts (p<0.05).¹⁸⁶ Load transfer at the cement–bone interface is mainly influenced by direct contact at this zone. Improving the cement–bone bonding can increase the cement–bone contact area and, hence, lead to a more uniform load transfer in VP and BKP.¹⁸⁶

Drug-loaded bone cements

Silver-containing bioactive agents have been shown to reduce the risk of infection^197,201 while enhancing the bone-bonding ability of the cement.¹⁹⁷ Some new generations of bone cements offer antibiotic-loaded PMMA and CPC formulations.^{26,67,213–217} PMMA bone cement is a good carrier for sustained release of antibiotics in the site of infection.²¹⁴ This is particularly important since chronic infection of joint prostheses requires surgical removal of the implant so as to eradicate the infection.⁴⁸ Several recent papers have reviewed the efficacy and safety of ALBC and highlighted the current state-of-the art in drug-eluting cements.^{28,59–61,218,219} Nevertheless, a recent systematic review comparing deep prosthetic joint infections between total knee arthroplasty patients, treated with either ALBC or plain bone cement, has argued that ALBC would not substantially reduce prosthetic joint infections, and thereby could be considered as an unnecessary cost to the healthcare system.⁶⁸ The cost efficiency profile of ALBC in routine primary hip and knee arthroplasty seems to be different in the United States compared with Europe.²²⁰ A significant hospital overhead cost has been reported in the United States with the use of ALBC, compared with plain bone cement, as ALBC could cost as high as $350–$400 per batch compared with $60–70 per batch of plain bone cement.⁶⁹

It should be noted that CPCs have also been used with other biologically active products (e.g., analgesics and contrast agents). For example, Dupleichs et al.²²¹ have investigated the efficacy of analgesic CPCs loaded with either bupivacaine or ropivacaine. The bupivacaine-loaded cement demonstrated an earlier return to full functional recovery than the ropivacaine-loaded cement, while CPCs retained their mechanical and biological properties. In another study, Le Ferrec et al.²²² loaded CPCs with Xenetix® radiopaque agent and reported that incorporating up to 70 mg/mL of Xenetix® into CPCs did not affect the injectability, setting time, and cohesion of the composite. Upon injection in a bone defect in rats, the Xenetix®-loaded CPC appeared to be biocompatible.²²² A recent review paper by Parent et al.²²³ has elaborated on the design of CPCs for drug delivery applications, with an emphasis on the parameters affecting the loading and release of therapeutic substances. In addition, a paper by Ginebra et al.⁹⁰ provides an overview of drug release from CPCs for low molecular weight drugs (e.g., antibiotics, non-steroidal anti-inflammatories, anti-cancer drugs, and anti-osteoporotics) and for high molecular weight molecules (e.g., growth factors and other proteins). The paper also provides a summary of some CPC formulations intended for ion release (e.g., calcium, phosphate, strontium, silicate, zinc, and magnesium).

Conclusions

An aging population and sports-related injuries have led to a dramatic increase in bone-related diseases and bone fractures.⁷⁶ When biomaterials and medical devices are designed to replace a degenerated or diseased joints, these materials may highly benefit from multi-functional characteristics that can meet the biomechanical and biological requirements of the bone.¹⁶⁴ ABCs composed of PMMA have adequate compressive strength, but suffer from high elastic moduli,¹⁸⁶ high polymerization temperatures,^56,164 and lack of remodeling in the body.^1,6 In addition, cardiac complications and embolism caused by PMMA cement injection and subsequent cement leakage have been reported in several studies.^209–212 Hence, alteration in cement formulation is a rapidly evolving field.²⁰⁹ Bone cement formulations combining PMMA with bone marrow¹⁸⁷ or nano-HA coated human collagen¹⁸⁶ have been proposed to reduce the modulus of PMMA cement while improving its biocompatibility. In addition, porous PMMA cements loaded with bioactive glass have been used to improve osseointegration and provide a better mechanical stability and biological integration.⁴⁸

CPCs are osteoconductive,^106–110 release no heat,¹⁷⁰ and can be shaped arbitrarily because of their self-setting characteristic. However, they are associated with the problems of low strength, high brittleness, and low cohesion in aqueous environment.¹³⁰ These shortcomings hinder their applications as loading-bearing bone substitutes in clinical settings and minimally invasive orthopedic surgeries. Hence, there is an unmet need for tougher and biologically active bone cement materials. Reinforcement of injectable CPCs using different materials such as gelatinized starches,¹³⁰ mixtures of CPCs/PMMA cement,⁹⁹ and composites of nano-HA/PMMA cement⁴⁹ have been considered. The addition of carbon nanotubes has been shown to improve the mechanical properties of CPCs, particularly their fracture strength and toughness.^39,164 In another study, an injectable bone substitute material made of calcium-deficient HA and foamed gelatin has been proposed to promote bone ingrowth.¹⁴¹ Recently, a copper-doped CPC formulation has demonstrated that copper could be a dually effective ion: toxic for bacteria while being beneficial for the healthy cells. This is based on the increased viability of human glial E297 cells, murine osteoblastic K7M2 cells, and human primary lung fibroblasts in the presence of the cement.¹⁵⁹

Both ABC and CPC classes are relatively inadequate in terms of key biological requirements and functional mechanical properties. Taking advantage of the benefits offered by the other class and making use of nanostructured materials can potentially lead to superior bone cement formulations for orthopedics applications. Manufacturing, safety, and regulatory issues should be kept in mind when it comes to nanomaterials.¹⁶⁴ Making use of osteoinductive materials, such as DBM, has also proven to offer a biological stimulus to promote osteogenic differentiation of MSCs and other osteoprogenitor cells.^41,84

Most CPCs have apatite precipitation, which reduces the resorbability under physiological conditions because of the low degradability of HA. Enhancing the bioresorbability of bone cements can pave the way to the development of new bone substitute materials and drug delivery systems.³⁹ The current market for resorbable bone graft substitutes is small compared with the medical device market or the drug market. However, it is a fast-growing field, with materials ranging from metallic alloys to polymers and from injectable cements to complex porous structures.⁷⁶ Rigorous preclinical and clinical trials are necessary to confirm the cost-effectiveness of new injectable bone graft substitutes and assert their benefit-to-risk ratios.⁸³ As for ALBCs, the use of ALBC in routine primary hip and knee arthroplasty appears to remain controversial because of the high cost and limited efficacy reported in some recent studies.^68,220

Footnotes

Declaration of Conflicting Interests

The author(s) declared no potential conflicts of interest with respect to the research, authorship, and/or publication of this article.

Funding

The author(s) received no financial support for the research, authorship, and/or publication of this article.

ORCID iD

Azizeh-Mitra Yousefi

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