1. Introduction
Bone injuries caused by accidents or bone-related diseases (i.e., bone tumor, infection, and osteoporosis) are common
[1]. In many cases, surgical procedures are required to restore the function of the impaired bone. Minimally invasive surgeries (MISs) involve the use of smaller incisions during surgical procedures, which can simplify the operating process, relieve pain, and reduce the overall incidence of complications. The recognition and spread of MISs have brought new requirements for bone graft substitutes. Ideally, bone graft substitutes for MISs should be biocompatible, resorbable, and osteogenic. Sufficient mechanical and anti-infective properties are highly desirable for materials used for load-bearing and vulnerable-to-infection sites. In addition, such materials should be easy to handle, so the repairing materials can be precisely delivered to the target sites and can fill irregularly shaped bone defects. Thus, new techniques and materials that can promote surgery with a minimal intervention approach are in high demand
[2].
At present, autologous bone grafts remain the gold standard in the bone reconstruction field. However, their applications are limited by a short supply and the possible occurrence of a second injury during transplanting. Allografts are available in relatively large quantities but carry the risk of transmission disease and immune rejection reactions, and their healing effects are not as good as those of autologous bone grafts
[3]. More importantly, while bone-substitute materials are commonly shaped into various forms to be less invasive and more regenerative, it is not easy to process autografts or allografts into different sizes or shapes that are suitable for MIS
[4]. Therefore, further development of synthetic materials that are easily handled and have the capacity to promote bone regeneration is still in high demand for various applications (
Fig. 1). In the past few years, many studies have focused on minimally invasive biomaterials, and reviews have been published on minimally invasive biomaterials for different clinical applications (i.e., bone defects and infections) and components of biomaterials (i.e., natural polymers and synthetic polymers)
[5],
[6],
[7]. With the deepening of research on minimally invasive biomaterials, increasing attention is being paid to their bone-promoting mechanisms.
A variety of synthetic materials have been proposed as candidates for bone reconstruction, primarily including ceramics, polymers, and metals. Although metals such as titanium (Ti) and magnesium have wide applications in orthopedics, the focus of this article is on injectable ceramics and polymers that can be implanted in a minimally invasive manner. According to clinical needs for a bone-reconstruction approach, ceramics and polymers can take on the forms of porous blocks, pastes, and particles. To fulfill the requirements of MISs, such materials are commonly processed into pastes or particles, to be injected to the injured site through needles
[8]. Injectable bone-defect repair materials that cause little trauma and have good plasticity can carry growth factors (GFs), seed cells, and so forth, and can also fill irregular bone defects. These materials are considered to be a suitable choice for MIS for orthopedics and have attracted widespread attention. A typical example of an injectable polymer for MIS is poly(methyl methacrylate) (PMMA). Prior to reaction, PMMA is aqueous-like with a high viscosity; it then hardens onsite after being injected to the target tissues. Because of its self-setting and excellent mechanical properties, great success has been achieved using PMMA in vertebral augmentation, bone screw fixing, and so forth
[9],
[10]. When applied in vertebral augmentation, PMMA can restore the height of the vertebral body, enhance its mechanical strength, and relieve pain. Despite its popularity, PMMA has a bioinert and non-degradable nature that limits its further applications. These disadvantages result in the formation of fibrous capsules between the bone tissues and PMMA, which greatly reduce the bond strength. In addition, PMMA’s high stiffness and long-term contact with bone tissues cause a stress shielding effect
[11]. Therefore, there is increasing interest in the development of a new generation of injectable biomaterials with enhanced bioactivity and biodegradability.
Bioceramics such as bioactive glass (BG), calcium phosphates, and calcium sulfates have been demonstrated to be able to bond directly to bone tissues
[12]. The most well-known bioceramics are calcium phosphate cements (CPCs). Being injected as pastes to fill bone defects, CPCs have demonstrated excellent biocompatibility and bone-regenerative effects
[13]. Given their self-hardening capability and tunable injectability, CPCs have been proposed as alternatives to PMMA for bone fracture healing and implant-fixation applications. Injectable hydrogels are another material with great potential for minimally invasive intervention. Possessing a three-dimensional (3D), highly hydrated polymeric network, a hydrogel can flow like liquid during injection and then crosslink to solidify once it has been delivered to the target site. Seed cells are also an important factor in tissue engineering. When combined with cells, hydrogels and microspheres can serve as scaffolds for bone tissue engineering
[14]. Clinical needs continuously drive the development of bone-substitute materials, from bioinert to bioactive materials, and now toward materials with multiple functionalities and bioresponsive properties. These engineered materials are not only injectable and bioactive but also angiogenic and capable of treating bone-related disease
[15]. Moreover, some biomaterials can respond to environmental changes in order to actively participate in the bone repair and regeneration process
[16]. Different injectable materials can achieve bone regeneration in different ways, mainly through matrix biomineralization, angiogenesis, and immunomodulation. The evolution of materials has greatly expanded the reservoir of tools for MIS, giving clinicians more choices from which to select the systems that can best meet specific clinical needs.
In this review, we focus on implantable biomaterials used for MIS. The most commonly used bone-substitute materials for MIS, such as bioceramics and polymeric materials, are introduced. Their applications in various orthopedic fields, recent progress in multifunctional and bioresponsive materials, and different ways of promoting osteogenesis are also summarized. Finally, perspectives on future biomaterials for bone reconstruction are discussed.
2. Categories of minimally invasive implantable biomaterials for bone regeneration
2.1. Bioceramics
Bioceramics are ceramics that are applied in biomedical applications. The first generation of bioceramics includes bioinert materials such as zirconia and alumina, which are used as artificial femoral heads
[17]. Although these materials are biocompatible and strong enough to support the body, they elicit a foreign body reaction, resulting in the formation of acellular collagen capsules that isolate them from surrounding tissues
[18],
[19]. The invention of BG by Hench launched a new era of bioactive materials. BG, which contains 46.1% SiO
2, 26.9% CaO, 24.4% Na
2O, and 2.6% P
2O
5, is the first biomaterial that can bond to bone without forming fibrous tissue
[20]. Later, glass–ceramic apatite–wollastonite (A–W), which exhibits high bending strength and compressive strength, was invented by Kokubo
[21]. Since then, an increasing number of bioactive ceramics, including calcium-phosphate-based, calcium-sulfate-based, and calcium-silicate-based bioceramics, have been used in biomedical applications
[22]. In the orthopedic field, BG, hydroxyapatite (Ca
10(PO
4)
6(OH)
2; HA), and β‐tricalcium phosphate (β-Ca
3(PO
4)
2; β-TCP) have become important alternatives to autografts
[23],
[24]. HA and biphasic calcium phosphates (BCPs) have been used to fabricate porous bioceramic granules with core–shell structures that can enhance the osteogenic ability of bone-marrow mesenchymal stem cells (MSCs)
[25]. All these materials are said to be osteoconductive, which means that the materials can attract the growth of osteoblasts and osteoclasts on the surface and promote new bone formation. Some materials have exhibited osteoinductivity, stimulating osteoprogenitor cells to differentiate into osteoblasts
[26].
Bioceramics can be processed into the forms of particles, scaffolds, or cements to fulfill various clinical needs. In MIS, bioceramics are usually injected to the injured site in the form of pastes. Current efforts are being devoted to the development of advanced delivery systems to facilitate the surgery process, and to the synthesis of bioceramics with enhanced osteogenic properties. At present, research in this field is mainly attempting to solve the problem that bioceramics with high mechanical strength have poor degradation performance, while bone cement with a good degradation performance cannot reach the appropriate mechanical strength. Therefore, finding degradable bone cement with good mechanical properties has become the focus of investigation.
2.1.1. Injectability of bioceramics
The friction force between the inorganic particles hinders the flow of bioceramics in the syringe; therefore, a liquid phase is indispensable for the injection system. Aqueous solutions or non-aqueous solutions such as glycerol and poly(ethylene glycol) (PEG) can act as the liquid phases for injection
[27]. After being mixed with a liquid to form a paste, bioceramics can be delivered to the fracture site via one- or two-paste syringes. In a one-paste system, the bioceramics are combined with a liquid phase and injected directly into the bone defects. In a two-paste system, the reactive components are separated into two parts, and the reaction is triggered after the two pastes are mixed in the syringe; they then harden onsite after delivery
[28]. To achieve the desired injectability and curing effect, the liquid-to-powder (L/P) ratio, particle size and shape, and selection of the liquid phase are essential factors to be considered. For example, a higher L/P ratio usually results in higher injectability
[29]. Moreover, it has been shown that particles with round shapes are more injectable than those with irregular shapes
[30]. To further facilitate the handling of the bioceramics, premixed (or ready-to-use) bioceramics have been developed
[31]. These bioceramics are usually mixed with non-aqueous solvents before injection; after delivery to the body, the non-aqueous solvents exchange with the body fluid, which leads to the hardening of the bioceramics
[32]. Premixed bioceramics can greatly facilitate handling in the clinic and decrease surgical risks in a minimally invasive manner (
Fig. 2)
[33].
2.1.2. Compositions of bioceramics
According to their chemical composition, bioceramics can be classified into different groups: calcium phosphates, calcium silicates, zirconia, alumina, and phosphate-based or silicate-based glasses. Of these materials, calcium phosphate-based bioceramics are most widely used for bone reconstructions. These bioceramics include materials that contain calcium ions (Ca
2+) and orthophosphates (PO
43–), metaphosphates (PO
3–), or pyrophosphates (P
2O
74–). Among these phases, calcium orthophosphates—some of which are listed in
Table 1 [34],
[35]—are studied most extensively. The popularity of calcium phosphate-based bioceramics for bone regeneration is mainly due to their chemical similarity to human bone mineral. HA and β-TCP are among the most commonly used calcium phosphates for bone reconstructions
[36]. HA is the least soluble calcium orthophosphate, while β-TCP is more soluble
[37]. The combination of these two compounds results in so-called BCP ceramics, which have gained increasing popularity for their osteogenic properties
[38]. Recently, much attention has been paid to other calcium-phosphate-based bioceramics, including α-tricalcium phosphate (α-Ca
3(PO
4)
2; α-TCP), dicalcium phosphate dihydrate (CaHPO
4·2H
2O; DCPD, also known as brushite), and amorphous calcium phosphates (ACPs)
[39],
[40],
[41]. Calcium-silicate-based materials are another important group of bioceramics. The most common phases are calcium silicate (CaSiO
3), dicalcium silicate (Ca
2SiO
4; C
2S), and tricalcium silicate (Ca
3SiO
5; C
3S). Porous β-calcium silicate ceramics exhibited excellent bone-formation capability in a rabbit calvarial defect model
[42]. Further study showed that β-calcium silicate significantly enhanced the vascularization and osteogenic differentiation in a co-cultured system containing different cells
[43]. The most well-known calcium-silicate-based bioceramic is mineral trioxide aggregate (MTA), which has been applied as a root canal filling material in dentistry
[44]. The main components of MTA are tricalcium silicates and dicalcium silicates. It exhibits excellent capabilities for sealing, biocompatibility, and apatite formation, which are beneficial in dentistry
[45]. Some commercial MTA products are dispersed in capsules with a tip, enabling mixing and direct injection into the cavity
[46].
2.1.3. Types of bioceramics
Bioceramics in non-biodegradable bone cements. PMMA was first applied in a clinical setting in 1958, as a bone cement for a total hip replacement
[47]. Since then, it has been considerably developed, becoming the current gold-standard material for vertebral body augmentation. Many attempts have been made to further improve the performance of PMMA, aiming to promote its bioactivity, reduce its elastic modulus, decrease its polymerization temperature, and increase its porosity to allow bone growth
[48],
[49],
[50]. For example, PMMA cements modified with castor oil and linoleic acid have been demonstrated to have decreased yield strength and a lower Young’s modulus
[51]. Such modified cements, which have a mechanical strength closer to that of osteoporotic cancellous bone, can greatly reduce the incidence of adjacent vertebral fractures after augmentation. Other studies have focused on improving the bioactivity of PMMA cements, using bioactive materials such as
γ-methacryloxypropyltrimethoxysilane (MPS) and calcium acetate
[52], HA and bone morphogenetic protein-2 (BMP-2)
[53], silicate-based bioceramics
[54], or BG
[55]. These bioactive materials promote the osteogenicity of PMMA cements, resulting in better integration between the cement and host bone. Despite PMMA’s clinical success and the significant improvement achieved with its use in the past decades, the nondegradable nature of PMMA has hardly been changed; this is because PMMA’s natural properties are intrinsic, making it very difficult to alter via modification. As ideal bone-filling materials should be gradually replaced by bone tissues, more efforts are required to explore resorbable cements for use in the future reinforcement of osteoporotic vertebrae.
Biodegradable bone cements. (1) Calcium phosphate-based bone cements. CPCs are promising materials for bone substitution, although their relatively low strength and fragility limit their use for vertebral augmentation. The mechanical strength of CPCs is influenced by factors such as the porosity, filler content, particle size of the precursor, and component of the liquid phase
[29],
[56]. Decreasing the L/P ratio reduces the porosity of set cements, resulting in cements with higher mechanical strength. However, a minimum amount of water is required in order to retain the injectability of the cement
[57]. To modulate the mechanical behavior of CPCs, fillers have been incorporated, including mesoporous BG
[58], anhydrous dicalcium phosphate and titanium dioxide
[59], and nanosilica
[60]. Moreover, the addition of fiber-like and whisker-like filler contents has been demonstrated to be efficient in reinforcing CPCs
[61],
[62]. For example, the incorporation of poly(vinyl alcohol) (PVA) fibers can increase the flexural strength and toughness of CPCs
[63]. The addition of 5 wt% calcium silicate fibers increased the compressive strength of CPCs from 14.5 to 50.4 MPa
[64]. Another study showed that chitosan fibers improved the toughness of CPCs by several hundred-fold without affecting their elastic modulus and while maintaining their bending strength
[65]. It has been reported that reducing the particle size of the precursor can improve the mechanical properties and injectability of CPCs
[66]. However, fine particles have a larger surface area and faster rate of hydrolysis than coarse ones, resulting in a shorter setting time
[67].
The liquid component is also an essential factor in the mechanical behavior of CPCs. CPCs consist of a solid phase and a liquid phase, where the liquid phase is usually water. When water is used as the liquid phase, however, the cement will solidify quickly and cannot be evenly mixed. If retarders such as citric acid or its salt, sodium citrate, are added, the hydration process can be delayed, allowing the operation time to be increased; then the two phases can be fully mixed
[68]. CPCs with citric acid or sodium citrate as liquid components are easier to handle and possess a decreased L/P ratio, which can significantly increase the compressive strength of the cement
[69],
[70]. Recently, many studies have focused on incorporating polymeric substances into the liquid phase in order to reinforce the inorganic network of CPCs. Polyacrylic acid
[71], poly(γ-glutamic acid) (γ-PGA) and its strontium salt
[72], and silanized-hydroxypropyl methylcellulose (MC) have been incorporated to enhance the mechanical properties of CPCs. A dual-setting cement with an optimal compressive strength of over 98.3 MPa was prepared by the
in situ crosslinking polymerization of glycidyl methacrylate-derivatized dextran in CPCs
[73]. The flexibility of a brushite cement was greatly increased by combing PEG-based hydrogels with a CPC network
[74]. In these studies, polymeric matrices were integrated to create a double-network structure, which greatly enhanced the mechanical performance of the CPCs. Recently, a brushite cement with a high compressive strength of 74.4 MPa was developed
[75]; moreover, the compressive strength of apatite cement was reported to reach around 100 MPa without precompression
[74]. Despite these improvements, current CPCs are generally fragile and have poor strength; thus, further investigations are required to develop CPCs with sufficient mechanical strength. In combination with their biodegradability and osteoinductivity, novel CPCs with enhanced mechanical properties would be promising materials for application in vertebral body augmentation.
(2) Calcium sulfate-based bone cements. Calcium sulfate is one of the oldest bone-substitution materials, with its first use as a bone void filler dating back to 1892 in a report by Bahn
[76]. It is considered to be a promising material for orthopedic application due to its good bioactivity, injectability, and osteoconductivity
[77]. The setting of calcium sulfate cement is based on the reaction of calcium sulfate hemihydrate (CaSO
4·0.5H
2O) with water, forming calcium sulfate dehydrate (CaSO
4·2H
2O) as the end product. The setting reaction is moderate, with low exothermic heat; therefore, it is friendly to the surrounding tissue when used as a bone void filler. However, due to its low mechanical strength and high resorption rate, calcium sulfate alone is inappropriate for vertebral body augmentation
[78]. Calcium sulfate cement can be reinforced by particles such as mesoporous bioglass
[79], bismuth ferrite
[80], and nano-HA with collagen
[81]. Another strategy is to combine calcium sulfate cement with other types of bone cements to overcome its weakness. For example, tricalcium silicate was introduced into calcium sulfate cement to improve its mechanical properties and degradability
[82]. Furthermore, CPC
[83] and magnesium phosphate cement (MPC)
[84] were respectively combined with calcium sulfate cement to develop composite cements with enhanced physico-chemical properties. Despite these efforts, the mechanical strength of calcium sulfate cement is inadequate for load-bearing applications, so further development is required to broaden its application to vertebral body augmentation.
(3) Other biodegradable bone cements. Novel inorganic and composite bone cements have been developed to meet the requirements of vertebral augmentation. The major components of calcium silicate cement, which include tricalcium silicate ((CaO)
3·SiO
2, C
3S) and dicalcium silicate ((CaO)
2·SiO
2, C
2S), have superior bioactivity, sealing ability, and marginal adaptation. After mixing the powder with water, a viscous colloidal gel is formed that eventually solidifies into a hard structure. Injectability is essential for the clinical application of calcium silicate cements. Although an increase in the L/P ratio improves the injectability, this can affect the other properties of the cement. Additives such as gelatin can convert the cement into a more injectable paste
[85]. The presence of gelatin was found to appreciably improve the anti-washout and brittle properties of the cement without adversely affecting its mechanical strength, and the fracture characteristics of the cement were changed from brittleness to toughness fracture
[86]. By combining tricalcium silicate with sodium alginate, a new composite cement with enhanced washout resistance, formability, injectability, and higher compressive strength (54 MPa) was obtained
[87]. A novel fast-setting calcium silicate cement was prepared with a solid phase (silicate powder) and a liquid phase (ammonium phosphate solution), and the setting time was reduced to within 9 min
[88].
MPCs are essentially acid-base cements in which magnesium oxide and ammonium phosphates are respectively used as the basic component and preferred acidic component
[89]. MPCs can react at room temperature and have the characteristics of fast setting and high early compressive strength. Additives are important to improve the injectability of MPCs; in particular, lactic acid, glycerol, chitosan, and citric acid can increase the viscosity of the liquid phase and improve the wettability of the cement particles
[90]. MPC with chitosan showed enhanced water resistance, and the compressive strength of the composite cement reached 42 MPa after 28 days’ incubation
[91]. The incorporation of metakaolin (MK) into MPC led to a longer setting time (to a maximum of 52 min) and less-exothermic reactions
[92]. The use of citric acid as a setting retarder in MPC prolonged the setting time appropriately (from 11 to 17 min) and increased the compressive strength (to a maximum of 76 MPa)
[93]. By adjusting the granularity of the phosphate salt and using sodium borate as a retardant, the exothermic and setting kinetics of the MPC were improved
[89]. A tricalcium silicate/MPC (C
3S/MPC) composite bone cement showed the highest compressive strength (87 MPa), compared with C
3S (25 MPa) and MPC (64 MPa)
[94].
Recently, injectable inorganic/organic composites have been explored in order to overcome the disadvantages of traditional cements
[95]. For example, composite cements were prepared to mimic the reversible interactions and exceptional strength and toughness of natural tissue. In these cements, calcium phosphate nanoparticles and bisphosphonate-functionalized hyaluronic acid were used to form an injectable, robust, and cohesive nanocomposite hydrogel
[96]. When calcium phosphate nanoparticles were integrated with gelatin nanospheres to prepare injectable colloidal organic–inorganic composite gels, the gels exhibited strongly enhanced gel elasticity, shear thinning, and self-healing behavior
[97]. The utilization of laponites with silated hydroxylpropyl methyl cellulose to prepare composite hydrogels led to the formation of a hybrid interpenetrating network
[98]. By combining BG with PVA and PEG, an injectable composite cement with high compressive strength (57 MPa after mineralization for 14 days) was prepared (
Fig. 3)
[99]. A new injectable borate BG cement prepared with borate BG particles and a chitosan solution exhibited high injectability and compressive strength (31 ± 2) MPa
[100]. In addition, a set of elastomeric nanocomposites were made from PEG and HA nanoparticles. 15% HA nanoparticles added to the PEG hydrogel resulted in a significant increase in toughness compared with the pure polymer hydrogel. This result can be attributed to the ability of hydroxyapatite nanoparticles (nHA) to reinforce the covalently crosslinked PEG network
[101]. These studies show that combining inorganic particles with organic networks is a promising strategy for preparing injectable and biodegradable bone cements.
2.2. Hydrogels
A hydrogel is a hydrophilic polymer network system with a 3D structure. To date, injectable hydrogels are widely applied in bone tissue engineering because of their unique advantages, by which anomalous defects can be easily targeted with noninvasive therapy. Hydrogels can enhance the physico-mechanical properties of the injured bone, delivering a variety of biomolecules for therapeutic application
[102],
[103]. An ideal hydrogel for bone repair would be easily produced, biocompatible, and injectable, with the sustain release of an appropriate active GF. Injectable hydrogel systems are beneficial for bone defect repair and can be solidified
in situ at the defect site.
Hydrogels can be fabricated using various polymers. Natural polymers possess good biocompatibility and biodegradability and can interact with cells, providing a suitable microenvironment for cell adhesion, migration, and proliferation. Compared with natural systems, the advantages of synthetic hydrogels lie in their controllable physical and chemical properties. The poor mechanical properties of hydrogels limit their use in load-bearing parts of bone. Thus, there is an urgent need to improve the mechanical properties of hydrogels while maintaining their biocompatibility, injectability, and other advantages. However, many studies have shown that the mechanical properties and biocompatibility of most hydrogels are independent of each other, and the use of various chemical reagents might affect the biocompatibility of hydrogels. Current studies mainly focus on regulating the mechanical properties of hydrogels through the Hoffmeister effect of small molecular inorganic salts. However, a high salt content in hydrogels may reduce their biocompatibility and limit their applications in biomedical fields. Therefore, researchers have also proposed improving the mechanical properties of hydrogels by changing the crosslinking mode or the degree of crosslinking, constructing double-network hydrogels, and adding nanoparticles, while maintaining good biocompatibility
[104].
2.2.1. Hydrogels from natural polymers
The components of natural polymers are similar to the natural extracellular matrix (ECM), with the capacity to promote proliferation and adhesion
[105]. Natural polymers incorporated with osteogenic factors, bioactive molecules, or cells show excellent bone repair ability
[106],
[107],
[108],
[109],
[110],
[111]. Common natural polymers include hyaluronic acid, chitosan, alginate, fibrin, heparin (Hep), and gelatin. As a glycosaminoglycan (GAG), hyaluronic acid can be found in the ECM and is broadly distributed in various parts of the human body. It consists of alternate disaccharide linked units and is the only non-sulfated GAG. Hyaluronic-acid-based hydrogels are biocompatible, non-immunogenic, and non-inflammatory, due to their natural origin. Thermosensitive hydrogels can be obtained through the combination of hyaluronic acid and thermosensitive polymers, such as poloxamers
[112], pluronic F127
[113], and poly(
N-isopropylacrylamide) (PNIPAAm)
[114]. Moreover,
in situ chemically crosslinked hyaluronic acid hydrogels can be obtained through the reaction of a Schiff base
[115],
[116], Michael addition
[117], and enzyme catalysis
[118],
[119]. Chitosan, which has excellent biocompatibility and immunostimulatory activity, is obtained via the deacetylation of chitin, which is widely present in nature
[120]. When applied as an injectable hydrogel, chitosan can undergo thermal
[121],
[122] and pH-triggered
[123] gelation and can be enzymatically degraded
in vivo by lysozyme
[121] and chitosanase enzymes
[124]. Alginate has also been widely used in biomedical applications
[125] due to its mild gelation process and degradability
[126]. It can form an injectable gel via the chelation of cations such as Ca
2+, Ba
2+, and Sr
2+. However, because of the diffusion of cations in the hydrogel under physiological conditions, the mechanical properties and degradation performance of the ion-crosslinked alginate hydrogel are insufficient
[127], limiting its application in biomedical fields. In addition to ion-crosslinking, injectable alginate hydrogels can be obtained via enzyme crosslinking
[128],
[129] and Schiff base crosslinking
[126].
Fibrin is a protein formed by the action of thrombin on fibrinogen and is a key element in the final stages of clotting, which stops bleeding. Injectable fibrin gels can be formed by gelling soluble fibrinogen in various ways. The mechanical properties can be tailored by the concentrations of the fibrinogen, thrombin, or enzyme used
[130]. Hep is a GAG with a negative sulfated charge that is mainly distributed in the liver. Hep-based hydrogels can be obtained via physical
[131] and chemical crosslinking (i.e., via enzyme
[132], Michael addition
[133],
[134], or Schiff base
[135]). Gelatin is the product of the partial hydrolysis of collagen
[136], and proteinase K can rapidly hydrolyze gelatin hydrogel
[137]. The stability of the gel can be enhanced through enzyme crosslinking
[138] and Schiff base crosslinking
[139]. A ternary gel that was stable after gamma irradiation was prepared with gelatin–water–glycerol (GWG) and showed potential for use as a broad-spectrum injectable carrier. The researchers attributed these features to the addition of glycerol, which increased the chemical potential of the gelatin via hydrogen bonding. When loaded with demineralized bone matrix, GWG scaffolds exhibited excellent osteogenesis performance
[140].
One disadvantage of hydrogels made of natural polymers is their poor mechanical strength, which limits their applications in load-bearing sites. Numerous methods have been used to strengthen hydrogels and increase their bioactivity in order to expand the scope of their clinical applications. For example, double network (DN) hydrogels, composed of two interpenetrating or semi-interpenetrating polymer networks, exhibit superior mechanical performance in comparison with single-network hydrogels
[141]. One example is a chitosan-based DN hydrogel, which was fabricated by combining a short-chain chitosan and a covalent crosslinking polyacrylamide (PAM) network by means of hydrogen bonding. The prepared DN hydrogel showed enhanced mechanical performance
[142]. The mechanical properties of hydrogels based on alginate
[143] and hyaluronic acid
[144] can be enhanced via DNs as well. Moreover, the incorporation of nanoparticles into hydrogels has been demonstrated to be an efficient way to enhance the mechanical performance of natural polymer-based hydrogels. Such nanoparticles include ferroferric oxide nanoparticles
[145], nanosilica
[146], gold nanorods
[147], graphene oxide
[148], and more. The use of click chemistry
[149],
[150] and supramolecular chemistry
[151] can also improve the mechanical properties of natural polymers.
2.2.2. Hydrogels from synthetic polymers
The most commonly used synthetic polymers include peptides, PEG or poly(ethylene oxide) (PEO), poly(galacturonic acid) (PGA), and their copolymers, such as PEG–polylactide (PLA) and PEG–
b-polycaprolactone (PEG–PCL). The properties of synthetic polymers can be adjusted for specific applications. In comparison with natural polymers, synthetic polymers have reliable raw material sources and long shelf lives
[152],
[153]. However, the biocompatibility of synthetic polymers is not as good as that of natural polymers, and the degradation of some synthetic polymers can produce toxic byproducts
[154]. Hydrogels with synthetic polymers can be used as an excellent carrier for drugs, bioactive factors, and biomolecules. For example, a PEG hydrogel comprising four-arm PEG macromers functionalized with terminal maleimide groups (PEG-4 MAL) and loaded with lysostaphin was able to kill bacteria efficiently and promote bone fracture healing
[155]. Injectable polymeric hydrogels produced by crosslinking oxidized pullulan and 8-arm PEG hydrazine and loaded with dexamethasone inhibited inflammation and induced osteoblastic differentiation (
Fig. 4)
[156]. Through the selection of appropriate segments, thermo-gelling copolymers can be further fabricated
[157]. For example, PEG–polylactide-glycolide (PLGA) copolymers exhibit excellent thermo-gelling behavior, forming a hydrogel at body temperature while existing as a liquid at room temperature. As an
in situ gel-forming carrier, PEG–PCL hydrogels were used to realize the sustained release of dexamethasone and promote bone regeneration
[158]. Moreover, by combining a thermosensitive PEG–PCL–PEG (PECE) copolymer with collagen and nHA, osteoactive scaffolds have been obtained. This composition, which has good biocompatibility, can be used to guide bone regeneration
[159]. Strong bone adhesives are becoming increasingly popular
[160]. An injectable hydrophobic laminous adhesive with liquid hydrophobic photo-crosslinkable poly(lactide-
co-propylene glycol-
co-lactide) dimethacrylates (PmLnDMA) as the cartilage phase and PmLnDMA-encapsulated methacrylated HA nanoparticles (PmLnDMA/MH) as the mineralized subchondral bone phase was used to treat osteochondral defect, osteoarthritis, and osteoporosis
[161].
Despite the advantages mentioned above, the further application of hydrogels is limited by their poor biological activity, toxic byproducts, and other shortcomings. Therefore, biological and chemical entities have been conjugated to synthetic polymers to improve the comprehensive properties of hydrogels
[162]. To improve their osteogenic and osteoinductive properties, hydrogels can be modified with arginine-glycine-aspartic acid (RGD)
[163]. RGD-modified polymers efficiently promoted cell adhesion, and RGD-functionalized hydrogels led to increased MSC viability
[164]. In another study, hydrogels loaded with RGD were shown to enhance alkaline phosphatase (ALP) activity
[165]. Chemical modification of the biomaterial structure can also promote mineralization. For example, adding phosphate groups to hydrogels can not only maintain biodegradability but also promote the mineralization of encapsulated MSCs
[166],
[167].
2.3. Cell-loaded injectable biomaterials
Seed cells play an important role in tissue repair. Recently, studies have suggested that cell transplantation holds great promise for treating bone defects. MSCs may play a key role because of their unique proliferation and differentiation capacities
[168]. Stem cells from different sources, such as bone-marrow stromal cells (BMSCs) and adipose-derived stem cells (ASCs), are widely used to repair bone injuries. Tissue-derived MSCs are not only capable of multidirectional differentiation but also have the following advantages: ① stem cells are primitive and have stronger proliferation and differentiation ability; ② the immunogenicity of cells is low; ③ stem cells are easily available; ④ the conditions of amplification culture are stable and uniform, which is convenient for large-scale amplification and quality control; ⑤ they are suitable seed cells for tissue engineering and can be used many times, with little cell loss after freezing; and ⑥ they permit convenient collection, easy storage and transportation, and fewer ethical disputes
[169],
[170]. Previous studies have shown that various biomaterials containing stem cells are capable of inducing osteogenesis and bone defect healing
[14],
[171].
2.3.1. Cell-loaded microspheres and microgels
In recent decades, the use of microspheres and microgels in bone tissue engineering has attracted widespread attention, as they can be used to carry GFs, exosomes, and cells. Micro-invasive surgery using microspheres or microgels to treat irregular bone defects has the advantages of a small scar, short operation time, few complications, and improved patient comfort and satisfaction. Because of their good biocompatibility, polymers can be widely used in the preparation of microspheres. Injectable microspheres loaded with ASCs were shown to promote the osteogenic differentiation of ASCs and repair mouse femoral non-union
[172]. Injectable apatite-coated atorvastatin (AT)-loaded PLGA microparticles also supported the osteogenic differentiation of ASCs
[173]. Porous shape-memory cryogel microspheres (CMS) prepared from methacrylated gelatin (GelMA) promoted the proliferation and adhesion of human BMSCs (hBMSCs) and human umbilical vein endothelial cells (HUVECs), thereby realizing the development of vascularized bone-like tissue
[174]. Encapsulating cells in microgels—which can mimic a 3D microenvironment, supporting the viability and function of cells and protecting cells from environmental stress—has been widely used in tissue regeneration and cell therapy. In one study, the osteogenesis of a cell-loaded microgel was obviously enhanced, and the mineralization of the microgel was accelerated. Moreover, the microgel-loaded MSCs caused significant enhancement of bone formation in a rat tibial ablation model, compared with a cell-mixed microgel and an acellular microgel
[175]. PVA microgels loaded with human MSCs (hMSCs) and BMP-2 also enhanced the osteogenic differentiation of hMSCs
[176].
2.3.2. Cell-loaded hydrogels
Cell-based therapies are providing new potential therapeutic methods for bone regeneration. Hydrogels with porous structures to enhance the natural ECM microenvironment can be used as cell carriers. MSCs were encapsulated in an injectable cell carrier (Pluronic F-127) loaded with recombinant human BMP 4 (rhBMP4), which improved cell osteogenic differentiation
[177]. An injectable bioactive hydrogel composed of alginate, gelatin, and nanocrystalline HA and loaded with osteoblasts was also prepared and was shown to promote the osteogenic differentiation of cells
[178]. Combining an injectable hydroxypropyl-β-cyclodextrin (HPβCD) crosslinked gelatin-based hydrogel with BMSC was shown to promote bone regeneration
[179]. To treat irregular bone deformities, a photo-crosslinked composite bioactive scaffold consisting of GelMA, BMSCs, and BMP-2 was fabricated; it promoted the osteogenic differentiation of BMSC and demonstrated remarkable bone-regeneration ability
[180]. BMP-2 and vascular endothelial GF (VEGF)-loaded microcarriers seeded with MSCs were incorporated into an injectable alginate-RGD hydrogel laden with endothelial cells (ECs). The composite hydrogel promoted vascularized osteogenesis
[181]. A biohybrid hydrogel based on crosslinked decellularized bone ECM and fatty-acid-modified chitosan was shown to deliver human amnion-derived stem cells and promote bone repair (
Fig. 5)
[182]. Unique cell-infiltratable and injectable gelatin hydrogels that encapsulate MSCs and icaritin efficiently prevented a decrease in bone mineral density and promoted
in situ bone regeneration by creating a microenvironment that favored the osteogenic differentiation of MSCs
[183].
The controlled delivery of therapeutic agents is critical to bone regeneration. A bioactive nanocomposite hydrogel based on hyaluronic acid and self-assembled pamidronate-magnesium nanoparticles realized the localized elution and on-demand simultaneous release of bioactive ions and small molecule drugs. Magnesium ions released from the hydrogels promoted osteogenic differentiation of the encapsulated hMSCs and activation of ALP. The activated ALP subsequently catalyzed the dephosphorylation of dexamethasone (Dex) phosphate and expedited the release of Dex from hydrogels to further promote hMSC osteogenesis
[184].
3. Injectable bioresponsive materials
Bioresponsive materials, which are sensitive to external stimuli, are appealing materials for bone reconstruction, as they can respond to variations in temperature, pH, or stress. In the forms of particles, hydrogels and composites, bioresponsive materials have shown great potential for cancer therapy, disease diagnostics, and as substituted materials for bone repair (
Table 2 [185],
[186],
[187],
[188],
[189],
[190],
[191],
[192],
[193],
[194],
[195],
[196]).
3.1. Thermal-sensitive biomaterials
Thermal-sensitive biomaterials are an important class of bioresponsive materials whose physical or chemical properties change under an external temperature change
[197],
[198]. One such material is thermal-sensitive hydrogel, which has been widely applied in bone tissue engineering for minimally invasive therapy. Thermal-sensitive hydrogels often possess a critical phase-inversion temperature and can achieve a sol–gel state transition based on temperature changes. In response to temperature variations, their swelling behavior, network structure, and/or mechanical properties change significantly, giving them certain advantages as injectable scaffolds
[199]. Currently, PNIPAAm is one of the most commonly used thermal-sensitive materials for the preparation of hydrogels due to its phase change between ambient and body temperature and its ability to copolymerize with different types of monomers
[200]. At room temperature, PNIPAAm is a free-flowing solution; when the temperature is raised, it solidifies to form an elastic hydrogel
[201]. In addition, crosslinked PNIPAAm can even be injected through small-gauge needles due to its highly swellable nature
[136]. Recently, thermal-sensitive PNIPAAm was grafted onto gelatin via atom-transfer radical polymerization. The thermal-sensitive composite hydrogel was found to be an excellent delivery vehicle of cells for bone repair
[185]. One disadvantage of synthetic thermal-sensitive hydrogels is that their biocompatibility is not as good as that of natural materials, and they lack the groups that are compatible with cells. Therefore, attempts have been made to form thermal-sensitive hydrogels through the modification of natural hydrogels. For example, agarose is a natural cooling hydrogel that transforms into a gel state when cooled to below 37 °C within a certain concentration range. HA/agarose gel composites were investigated and were found to be easy to handle and to closely connect with the surrounding tissues. Implanted into the medial femoral condyle of rabbits, excellent bone regeneration was observed
[202]. A chitosan-based thermal-sensitive hydrogel was prepared with chitosan and β-glycerol phosphate; it served as a biocompatible substrate for culturing rat bone-marrow stem cells
in vitro, and the injected cells survived in the gel scaffold for 28 days
[203]. A novel thermosensitive hydrogel based on blended MC with Persian gum (PG) was found to enhance the mechanical and biological characteristics of the hydrogel
[204]. A PLGA–PEG–PLGA/HA thermosensitive hydrogel loaded with vancomycin showed sustained release of antibiotics and promoted bone tissue repair (
Fig. 6)
[205].
3.2. pH-sensitive biomaterials
pH-sensitive materials undergo physical or chemical changes in response to variations in alkalinity
[206],
[207]. These pH-responsive properties can be attributed to the protonation of ionizable groups or the degradation of acid-cleavable bonds
[208]. pH-responsive injectable materials mainly include organic/inorganic nanoparticles and hydrogels, and one application of these materials is for cancer treatment. Many efficient pH-responsive anti-cancer drug-delivery systems have emerged for bone tissue engineering, due to the subtle differences in pH between healthy tissues and the extracellular environment of solid tumors. One pH-responsive delivery system composed of type I collagen and calcium carbonate loaded with cerium dioxide (CeO
2) nanoparticles and the anticancer drug doxorubicin (DOX) was used to treat osteosarcoma
[209]. In another study, researchers developed a multifunctional nanodevice using DOX-loaded mesoporous silica nanoparticles as nanoplatforms; the nanodevice exhibited selectivity to human osteosarcoma cells and pH-responsive anti-tumor drug-delivery ability (
Fig. 7)
[210].
Bone is a common site of tumor metastasis, and the local microenvironment is very conducive to tumor growth
[211]. Bortezomib has been demonstrated to be an efficient anticancer drug. Currently, a variety of pH-responsive materials have been developed to deliver bortezomib for cancer treatment
[212],
[213]. In one study, bortezomib was loaded onto a tripeptide RGD-targeted dendrimer via a boronate catechol linkage with pH-responsiveness in order to treat metastatic bone tumors
[212]. In another study, bortezomib-loaded micelles with bone-targeting capability were used to treat breast cancer bone metastases. The micelles, which were alendronate-decorated and bortezomib-catechol-conjugated, significantly inhibited cancer growth
in vivo [213].
In addition to their application in treating bone tumors, pH-responsive materials can be used to treat other bone diseases. For example, an injectable CPC embedded with pH-responsive microspheres was prepared and was able to control the release of drugs to treat osteomyelitis
[214]. In another study, multifunctional receptor-targeting and pH-responsive nanocarriers composed of PLGA–PEG-folic acid, poly(cyclohexane-1,4-diylacetone dimethylene ketal) (PCADK), and lipids were used to deliver methotrexate in order to treat rheumatoid arthritis
[215].
In addition, various pH-responsive hydrogels have been developed as potential bone replacements for bone repair and autogenous bone transplantation. The protonation/deprotonation of chitosan hydrogels via primary amine groups (NH
2) allows the hydrogel to respond to changes in external pH, granting them considerable potential for bone remodeling
[216]. Recently, a pH- and thermo-sensitive hydrogel based on chitosan and HA loaded with Hep was prepared
[217]. The results showed that the bioactive chitosan/HA/Hep hydrogels were able to induce angiogenesis. Anionic hydrogels such as carboxymethyl chitosan (CMCh) swell at a higher pH due to the ionization of the acid groups. A composite nanohydrogel composed of CMCh and ACP was developed
[218]. The hydrogel was osteoinductive and markedly enhanced the efficiency of BMP9-induced bone regeneration.
In addition to chitosan hydrogels, there are many other types of injectable pH-responsive hydrogels. For example, an injectable pH-responsive microgel based on methacrylate was developed to accelerate tissue repair to enhance mechanical properties
[219]. In another study, a novel injectable pH/thermo-sensitive biodegradable block copolymer hydrogel was synthesized
[220]. Mineralized tissue formation and high levels of ALP activity were observed, indicating great potential as an injectable material for bone tissue engineering.
3.3. Stress-sensitive biomaterials
Another important class of smart materials is stress-sensitive material, which can sense and respond to environmental stress by physically or chemically changing when mechanical force is applied
[221]. Shear-thinning hydrogels are typical stress-sensitive materials. These materials can be manually injected into tissues, flow under modest pressure, and set rapidly at the target site; therefore, they are widely used in bone repair. Self-assembly is the main pathway for shear-thinning hydrogel crosslinking, and the mechanism of the self-assembly process is specific to the shear-thinning system. Here, the “self-assembly process” refers to the balance between the competitiveness that facilitates the assembly and the forces that counteract the assembly. These interactions are usually relatively weak individually, but together they can lead to the formation of a stable network structure
[222]. Due to the dynamic nature of these weak physical associations, the formed network can be dissociated by the applied shear force. The shear-thinning process is highly nonlinear. After removing the shear forces, these networks reassemble into a hydrogel
[223].
Recently, a hydrogel with a Herschel–Bulkley fluid nature was developed by incorporating CaSO
4 and fibroblast GF-18 (FGF-18) into a chitin-PLGA composite hydrogel. The results showed that the composite hydrogels had excellent osteogenic differentiation ability
in vitro. Furthermore,
in vivo experiments revealed early and almost complete bone healing, indicating the great potential of this hydrogel for craniofacial bone defect regeneration
[224]. In another study, an nHA-incorporating chitin-PCL-based injectable composite microgel with the same Herschel–Bulkley fluidity was developed. The study confirmed that the chitin-PCL-nanohydroxyapatite (nHAp) microgel elicited an early osteogenic differentiation compared with the control gel
[225]. Hydrogels with shear-thinning rheological properties are capable of self-assembly and viscoelastic recovery after mechanical disruption. A study showed that medium-molecular-weight GAGs can be used to modify the rheological properties of hydrogels. GAG-HAp colloidal mixtures exhibited good osteogenic effect in calvarial defects
[226].
Shear-thinning and self-healing hydrogels also permit the homogeneous encapsulation of payloads; these hydrogels can be injected through a needle without clogging and then return to their original state, making them ideal for controlling the release of small molecules such as proteins, drugs, and genes
[227]. For example, a cohesive colloidal gel was prepared using oppositely charged PLGA nanoparticles. As the applied shear force was increased, the colloidal gels loaded with dexamethasone exhibited shear-thinning behavior.
In vitro drug-release tests showed that the encapsulated dexamethasone was fully released within two months, and
in vivo observations demonstrated massive bone formation
[196]. Moreover, to obtain a supramolecular hydrogel with a rapid self-integrating capability and excellent biocompatibility in order to promote the regeneration of tissue complexes, a novel injectable (shear-thinning) and self-integrating hydrogel was developed by grafting a significant number of multiple-hydrogen-bond units onto a biocompatible hydrophilic polymer. When the hydrogel was implanted into mice, the regeneration of cartilage-bone tissue was observed (
Fig. 8)
[228]. For healing non-union defects, an injectable silicate-based shear-thinning hydrogel incorporating PCL nanoparticles was introduced, which delivered cells and vasculogenic GFs. The injected hydrogel was capable of filling any irregularly shaped defects in bone, making it a good bone-repair material
[229].
Aside from hydrogels, other materials such as bone cements also have shear-thinning fluid properties. The rheological properties and injectability of an original CaCO
3 self-setting paste were studied via steady shear measurements, which revealed that the paste exhibited shear-thinning behavior over an extended period of time after paste preparation. Thanks to this shear-thinning behavior, the CaCO
3 self-setting paste was easily injectable and could be well applied to the treatment of bone defects
[230]. In another study, calcium silicate slurry was successfully prepared by means of 3D printing technology; the slurry also exhibited shear-thinning properties and had a compressive strength equivalent to that of human cancellous bone
[231]. A series of HA pastes with injectability and biocompatibility were shown to have potential for use as injectable bone graft substitutes
[232]. In another study, an injectable biomaterial based on isosorbide-alicyclic diol was prepared, consisting of two novel dimethacrylic monomers and bioactive nano-sized HA in the form of a two-paste system. The biomaterial exhibited non-Newtonian shear-thinning behavior and could be used as an alternative bone cement, overcoming some of the drawbacks typical of conventional acrylic bone cements
[233]. To repair complex bone defects, a composite scaffold of poly(trimethylene carbonate) with bone-forming nano-HA and noricaritin (derived from bone-growth-stimulating icariin) was prepared. The composite exhibited a sharp crystallization response and shear thinning and achieved the sustained release of noricaritin, indicating possible application in the reconstruction of orbital floor defects
[234].
4. Various ways to achieve osteogenesis with minimally invasive biomaterials
Natural bone consists of compact cortical bone and trabecular cancellous bone, which are respectively composed of densely packed cylindrical osteons and a porous network of trabeculae. Osteons and trabeculae consist of lamellae with different collagen fiber patterns. The collagen fibers are composed of bundles of mineralized collagen fibrils, with HA nanocrystals deposited in the gaps between collagen molecules
[235]. Bone defect repair is a complex process involving many factors. Improving biochemical functions (i.e., biomineralization, angiogenesis, and immunomodulation) is the main way to promote osteogenesis. For bone regeneration, biomineralization of the matrix is needed to deposit HAP and collagen. In addition, angiogenesis is necessary because blood vessels can transport cells, nutrients, and oxygen. Immune cells also play an indispensable role in bone regeneration; in particular, the appropriate transformation of macrophage phenotype can promote osteogenesis. Therefore, we classify matrix biomineralization, angiogenesis, and immunomodulation as ways to promote bone regeneration. Matrix biomineralization is the main process of bone repair and should be supported by appropriate angiogenesis and immunomodulation (
Fig. 9).
4.1. Targeting angiogenesis
Angiogenesis is essential for bone formation
[236],
[237],
[238],
[239]. Biomaterials can promote angiogenesis by delivering pro-angiogenic molecules
[240] or supporting cell proliferation
[241],
[242]. Injectable GFs are expected to gain popularity as a medical approach in orthopedics
[243]. Two key GFs in the process of angiogenesis are VEGF and basic FGF-2
[244]. The administration of VEGF is correlated with higher densities of blood vessels and tissue sparing
[242]. FGF-2 can also contribute to angiogenesis and bone repair
[243]. Angiopoietin-1 (ANG-1) is a vital protein in vascular development and angiogenesis
[245], and the synergistic action between VEGF and ANG-1 has been explored in angiogenesis
[246].
Over the past decades, drug-delivery systems have attracted widespread interest and investigation
[243]. An amorphous non-fibrous hydrogel composed of hyaluronic acid was found to be capable of acting as a VEGF delivery system
[242]. Study have shown that the local injection of gelatin hydrogel loaded with FGF-2 can realize the sustained delivery of FGF-2, protect the biological activity of GFs, and promote the induction of biological tissue regeneration
[243]. The combined use of different GFs has also been found to greatly enhance the effect of angiogenesis
[247].
4.2. Targeting immunoregulation
Bone regeneration is an extremely complex process that requires the cooperation of cells in different systems and tissues. Immune cells play an important role in bone repair. After bone injury, the inflammatory cells that first enter the bone microenvironment are immune cells, such as neutrophils, macrophages, and T cells. In the early inflammatory microenvironment, immune cells release various cytokines and chemokines in order to attract all kinds of cells into the bone injury area to participate in the inflammatory reaction; at the same time, they recruit BMSCs in bone tissue and regulate their proliferation, differentiation, or apoptosis. However, excessive inflammatory reaction will damage engineered cells and promote macrophages and fibroblasts to form fibrous capsules on the surface of implanted biomaterials. This will separate the biomaterials from the tissues, leading to the failure of engineering materials implantation. In recent years, the development of bone immunology has revealed the important roles played by immune cells in regulating bone regeneration, maintaining bone homeostasis, and regulating bone remodeling. The immune system and skeletal system share many regulatory molecules. After implantation, biomaterials inevitably lead to a host reaction, which causes the destruction of normal tissues around the implanted biomaterials and eventually leads to failure of the integration of the implants with the surrounding tissues. Moreover, the degree of inflammation affects the performance of implanted biomaterials, especially the osteogenic performance
[248]. Further study on increasing the immunomodulatory function of bone biomaterials would be beneficial in order to regulate the immune response in the process of bone repair.
Through in-depth investigation of the early inflammatory response, researchers have found that the polarization state of macrophages has a key impact on the process of inflammatory response. Macrophages are activated into the M1 type through the classical activation pathway; this predominantly has a pro-inflammatory effect and promotes the activation of early inflammatory reaction. Macrophages are also activated into the M2 type via alternative ways; this mainly results in anti-inflammatory activity, promotes the elimination of inflammatory reaction, and is conducive to osteogenesis and angiogenesis. An injectable periosteal ECM (PEM) hydrogel that dynamically integrates multiple biological functions was found to induce macrophage M2 polarization and further promote mature bone formation
[249]. Interleukin-4 (IL-4) is widely used in the construction of bone tissue engineering for macrophage regulation. IL-4-loaded calcium-enriched gellan gum hydrogel was also found to be able to regulate macrophage polarization
[250]. Composite hydrogels loaded with IL-4 and BMP-2 induced macrophages to differentiate into the M2 type and enhanced bone formation
[251]. It was also reported that nano-fish-bone-loaded hydrogels enhanced the mechanical properties of a hybrid hydrogel and modulated the immune microenvironment
[252]. Metal ions such as Zn
2+, Ca
2+, Li
+, and Mg
2+ are considered to be therapeutic ions with the ability to regulate macrophage polarization. A GelMA hydrogel loaded with lithium (Li)-modified bioglass regulated macrophages in a high-glucose (HG) microenvironment (
Fig. 10)
[253].
4.3. Targeting mineralization
Biomineralization refers to the combination of a biomineralization mechanism and material preparation to mimic the organism environment; it is a process of introducing inorganic particles such as calcium phosphate compounds and calcium carbonate into biomaterials by certain means. Through this process, the structure and composition of the materials can be closer to those of natural bones, so that composite materials with unique micro-structural characteristics and excellent biological properties can be prepared. Studies have demonstrated that the preparation of a biomineralization layer on the surface of a material can improve the material’s biocompatibility and cell affinity; in addition, the formation of a bone-like apatite layer on a material can promote direct osteogenesis—that is, the original biological characteristics of a material can be significantly improved by the formation of a biomineralization layer (e.g., an a HA layer) on the material’s surface
[254].
As ideal minimally invasive implantable biomaterials, mineralized biomaterials are widely used for bone reconstruction, especially mineralized hydrogels. Hydrogel mineralization involves the addition of calcium and phosphorus compounds, calcium carbonate, and other inorganic particles into the hydrogel matrix by different methods. Through this process, the structure and composition of the gel more closely resemble those of natural bone. A bioinspired mineralized hydrogel consisting of nHA, polyacrylic acid, and sodium carbonate was able to maintain good mechanical properties while exhibiting good osteoconductivity
[255],
[256]. A common mineralization strategy is to add prefabricated inorganic particles to hydrogel precursors; the particles then become trapped in the hydrogel network during gel formation. This preserves the injectable properties of the hydrogel while promoting the mechanical strength, growth, and osteogenic differentiation of bone-forming cells. A mineralized hydrogel loaded with a high concentration of Ca
2+ ions showed the rapid formation and high crystallization of apatite after soaking in simulated body fluid
[257].
Mineralized HA nanofibers were also incorporated into a GelMA hydrogel to form a composite hydrogel, improving the bone-repair effect
[258]. A mineralized hydrogel was obtained by adding 2,2,6,6-tetramethylpiperidine-1-oxyl (TEMPO)-oxidized cellulose nanofibrils into polymer solution. The crosslinking reaction of the polymer chains initiated by a calcium chloride reaction was optimized, and the
in situ mineralization of calcium phosphate was mutually promoted
[259]. It was also reported that an
in situ mineralized hydrogel scaffold loaded with rhBMP-2 realized bone tissue regeneration. Due to the opposite charge structure, a zwitterionic component templated the well-integrated dense mineralization of calcium phosphate throughout the scaffold
[260].
5. Biomedical applications of minimally invasive implantable biomaterials in orthopedics
In general, large bone defects require interventional therapy to recover. However, as the gold-standard bone graft material, autogenous bone is limited by supply. Therefore, many studies have focused on tissue engineering strategies to promote bone regeneration
[261],
[262],
[263]. Scaffolds are considered to be a key part of tissue engineering, providing a 3D environment for cell adhesion and proliferation. The application of implantable biomaterials is the main means of treatment for new bone formation and bone regeneration
[264]. The use of injectable biomaterials to load living cells and bioactive molecules is an ideal strategy, and its minimally invasive delivery can reduce inflammation and avoid the bone loss caused by invasive surgery
[265],
[266]. Genes, cells, and GFs can be effectively delivered to the target tissues through injectable biomaterials, making such materials desirable for application in bone tissue engineering (
Fig. 11)
[267]. Minimally invasive biomaterials have wide clinical application prospects, as they can be used to repair bone defect after trauma, bone defect caused by osteomyelitis, osteoporotic compression fracture, and bone destruction caused by bone tumor. Minimally invasive biomaterials can also assist in fixing metal internal fixtures such as screws and orthopedic Ti-alloy plates
[5].
5.1. Minimally invasive implantable biomaterials for the healing of bone fractures and defects
The repair of bone defects or fractures caused by accidents, bone tumors, congenital bone abnormalities, and trauma has always been a challenge in clinical bone repair. At present, the main repair materials are autogenous bone, allogeneic bone, and artificial bone graft biomaterials
[268],
[269]. The specific choice of bone-repair material for transplantation is mainly based on the following four aspects: the mechanical stability of the fixation, the blood supply at the fracture site, the range of the bone defect, and the size of the bone defect. Synthetic bone grafts are expected to mitigate the huge demand pressure caused by the lack of suitable autograft and allograft materials
[270],
[271],
[272].
The transplantation of bone-substitute materials to bone defects by means of minimally invasive injection is a clinically desired method of bone grafting with minimal trauma to patients
[273],
[274],
[275]. In clinical applications, some common fractures, such as calcaneal fractures, distal radius fractures, vertebral compression fractures, and tibial plateau fractures, often require only an injection of bone substitutes for treatment. For fresh fractures of the distal radius, the percutaneous injection of bone substitutes can fully fill the fracture defect, prevent the displacement of broken fractures, and promote fracture healing
[276]. Such injections can also be used to treat bone defects after bone tumor surgery, as they can effectively fill bone defects and induce the production of new bone
[273]. By carrying GFs and drugs, injectable bone substitutes may also be used to treat infectious bone lesions and repair large sections of long bone defects
[277],
[278].
As mentioned earlier, the application of injectable biomaterials has become an inextricable part of treatment for new bone formation and bone regeneration
[279],
[280]. Most clinically used injectable bone grafts belong to three major groups: acrylic bone cements, CPCs, and calcium sulfate cements
[274]. GFs and antibiotics have been increasingly developed as part of the composition of bone graft substitutes in order to promote bone formation and reduce infections in clinical practice
[281]. BMP have been shown to have osteoinductive properties and play an important role in regulating the expression of the target genes involved in bone physiology
[282],
[283]. Several clinical cases have used BMP-2 for maxillary reconstruction in cleft-lip and palate patients
[284],
[285]. Hissnauer et al.
[286] reported that a locking plate loaded with rhBMP-2 was successfully applied to treat non-union of the femur. To suppress infection, antibiotics have also been used for the prevention and treatment of bone and joint infections, such as open fractures and osteomyelitis
[287]. Local antibiotic use is an interesting approach due to the advantage of high local concentrations and low systemic concentrations
[288]. This has led to a good outcome not only in active infections
[289] but also when antibiotics are used prophylactically
[290]. Antibiotics doped into PMMA have been widely used in clinical treatment
[291].
5.2. Minimally invasive implantable biomaterials for vertebral body augmentation
Osteoporosis is prevalent among the old, featuring bone loss, degeneration of bone microstructure, and reduction in bone strength
[292],
[293]. Osteoporotic vertebral compression fracture (OVCF) is one of the complications of osteoporosis. OVCF can cause severe pain, affecting patients’ general health and quality of life. However, vertebral augmentation can quickly and efficiently relieve this pain and restore the strength and stiffness of the vertebral body
[294],
[295]. During surgery, bone cement is delivered to the injured site through a needle. Ideally, the injected materials used for OVCF should be degradable and have sufficient mechanical strength to sustain the loading. At present, the most widely used bone cement for vertebral augmentation is based on PMMA
[296] and consists of a solid phase and a liquid phase. The two phases are mixed to form a paste and injected into the body during the operation. The solid phase contains PMMA powder, initiator, and radiopacifier, while the liquid phase comprises methyl methacrylate (MMA) monomer, stabilizer, and accelerator. In current OVCF treatment, biomimetic mineralized collagen-modified PMMA bone cement has been found to achieve good vertebral height restoration. It can also significantly reduce the incidence of postoperative adjacent vertebral fracture
[297]. Although PMMA can provide adequate support and quickly restore the integrity and function of the vertebral body, it is nondegradable and bioinert. Moreover, its high elastic modulus can cause a stress-shielding effect, which results in bone loss
[298]. Therefore, degradable materials such as calcium phosphate-based and calcium sulfate-based bone cements have been under development for vertebral body augmentation. CPCs are also widely used in the clinic and have been reported to demonstrate good biocompatibility and osseointegration in clinical use
[299]. While CPCs exhibit acceptable compressive strength, their brittleness remains a major limitation for clinical applications. Poor mechanical properties and relatively unpredictable degradation behavior are still important problems that require further study
[300].
5.3. Minimally invasive biomaterials for bone implant fixation
Metal screws, steel plates, intramedullary needles, steel wires, or bone plates are frequently applied to fix bone fractures. In the clinic, debonding at the bone-material interface is often responsible for the failure of implants. For example, pedicle screws are more commonly used than other implants for internal fixation
[301]. However, the fixation strength of pedicle screws can be insufficient, or mechanical overload of the spine can occur after repair. In particular, osteoporosis can lead to loosened or ineffective screws, which results in non-union of bone and the need for surgical revision after implantation
[302]. To strengthen the implants and prevent them from falling off, adhesive materials are often required to increase the bonding between the prosthesis and bone tissues. At present, clinically used cementitious materials for bone implant fixation include PMMA
[303], calcium sulfate cements
[304], and CPCs
[305], which can be implanted into the body in a minimally invasive way, stabilize the bone implants, and promote bone healing.
The most commonly used adhesive material for bone implant fixation is PMMA, which can provide sufficient mechanical support. Although great success with PMMA has been achieved, it is difficult for PMMA to be absorbed; therefore, its use can result in a barrier at the broken end of the fracture that affects fracture healing
[306]. A number of bioactive materials that permit direct bonding with the bone tissues have been developed to replace PMMA as cementitious substances. One of these alternative materials is calcium sulfate cement. When used as the cementitious material, calcium sulfate cement can significantly strengthen the interface of the screw and bone and enhance the long-term stability of pedicle screws. The weak acidic environment around calcium sulfate cement can prevent the growth of fibrous tissue inwards while promoting the aggregation of osteoblasts and the formation of osteoid around the cement
[307]. Another study has shown that injectable calcium sulfate cement can improve the pull-out strength of pedicle screw fixation, indicating that it might be good alternative to PMMA
[308]. Calcium phosphate bone cement can be used as a binder to fill cracks and cavities in order to increase screw stability. Studies have shown that calcium phosphate bone cement fixed
in vivo by injection significantly increased the strength of screw extraction before and after implantation
[309]. An HA cement based on the reaction of tetracalcium phosphate and dicalcium phosphate was applied for screw augmentation and significantly increased the initial screw pull-out force
[310]. A novel CPC modified with phosphorylated amino acid monomers of phosphoserine was developed as a bone adhesive
[311]. Experiments showed that the addition of phosphoserine improved the adhesive force of the bone cement with a variety of surfaces, including biomaterials (metals, polymers, etc.), calcification, and soft tissues.
More recently, bone adhesives based on organic–inorganic interactions have been attracting increasing attention and are expected to revolutionize the clinical treatment of bone repair. A new type of bone binder based on tetracalcium phosphate and phospholipid was found to solidify in a few minutes in a watery environment and provided high bone–bone binder strength. The new material was measured to be 10 times more adhesive than bioabsorbable calcium phosphate bone cement and 7.5 times more adhesive than nonabsorbable PMMA bone cement
[312]. In another study, an inorganic–organic hybrid hydrogel was prepared via the spontaneous copolymerization of tannic acid with silk fibroin and HA. The strong affinity between tannic acid and silk fibroin, as well as the binding of the sacrificed coordination bond, significantly improved the toughness and adhesive strength of the hydrogel by increasing the energy dissipation at the nanometer level, which contributed to adequate and stable fixation of fractures in moist biological environments
[313].
5.4. Injectable biomaterials for bone tumor therapy
In general, patients with osteosarcoma are treated with the intravenous injection of anticancer drugs after tumor resection. However, due to the residual tumor cells at the tumor resection site, cancer recurrence often occurs. Moreover, the large bone defect left after the osteosarcoma operation exceeds the self-healing ability of bone tissue, bringing long-term pain to patients and even leading to surgical failure
[314]. Successful bone regeneration and control of residual cancer cells pose challenges to tissue engineering
[315].
Based on the synergistic improvement that can result from the interaction of two or more treatments, clinical research on tumors has gradually shifted from single treatments to combined therapy in recent years in order to improve the therapeutic effect
[316],
[317]. Near-infrared (NIR) laser photothermal therapy (PTT) has the advantages of deep penetration, accurate irradiation, and good ablation effect
[318]. Injectable AgBiS
2 nanoparticles have been designed for computed tomography (CT) imaging and the phototherapy of tumors. These AgBiS
2 nanoparticles can effectively convert light into heat under NIR laser irradiation and significantly increase the production of intracellular reactive oxygen species (ROS)
[319]. The combination of PTT/photodynamic therapy (PDT) successfully inhibited the growth of osteosarcoma
in vivo [320]. Bovine serum albumin-iridium oxide (BSA-IrO
2) nanoparticles were prepared, with high drug loading of DOX, high photothermal conversion ability, and good photostability, making it possible to achieve effective synergistic chemotherapy-PTT for osteosarcoma (
Fig. 12)
[321]. Using injectable magnetic nanoparticles (MNPs), magnetothermotherapy can convert electromagnetic energy into thermal energy in order to achieve targeted thermotherapy, heat the tumor from inside to outside, and avoid damage to the surrounding normal tissues
[322]. In addition to directly leading to the death of osteosarcoma cells, magnetothermotherapy may induce the remaining cancer cells to differentiate into more mature cell types, thereby inhibiting their self-renewal ability
[323].
The use of chemotherapy drugs improves the survival of patients suffering from osteosarcoma, but the tumors continue to relapse and undergo metastasis. In addition, these drugs sometimes lead to multidrug resistance (MDR)
[324],
[325]. When a patient is receiving systemic chemotherapy, the rapid removal of anticancer drugs in the bloodstream through non-specific protein binding will result in a poor therapeutic effect and adverse side effects
[326]. Therefore, successful treatment of osteosarcoma requires new and advanced chemotherapy methods
[327]. The appropriate design of a drug-carrier system is necessary in order to improve the effect of combined chemotherapy
[328],
[329]. An injectable hydrogel based on a local drug-delivery system was able to deliver anticancer drugs to cancer tissues through local injection, without the need for treatment in the bloodstream, and reduced MDR to a minimum via combination chemotherapy
[330]. In addition, compared with systemic chemotherapy, local administration can overcome the challenges of side effects and a short drug half-life introduced by systemic administration
[331],
[332]. A DOX hydrochloride (DOX·HCl)/cisplatin (CP)-loaded hydrogel/(2-hydroxypropyl)-beta cyclodextrin (GDHCP) drug-delivery system showed continuous DOX·HCl and CP release over a period of 7 days. Compared with DOX·HCl, CP, and glycol chitosan hydrogel, GDHCP exhibited a better anticancer effect
[333]. In addition, an injectable hydrogel-based drug-delivery system improved the maximum tolerated dose (MTD) of chemotherapeutic drugs to some extent and reduced systemic toxicity
[334].
The combination of vascular blockers and cytotoxic drugs can improve the antitumor effect and induce the whole-area apoptosis of a tumor
[335]. Microspheres containing combretastatin A-4 (CA-4) and docetaxel (DTX) were embedded into injectable hydrogels to form hydrogel microspheres (Gel-MPs) constructed for the sequential release of drugs in order to cooperate in the treatment of osteosarcoma. More specifically, CA-4 was preferentially released from the degraded hydrogels, disrupting the vascular structure of the tumors, reducing the exchange of nutrients between the tumors and surrounding tissues, and creating gaps for DTX to penetrate tissues, thereby inhibiting the proliferation of tumor cells
[332].
Injectable particles are another promising drug-loading media for the treatment of bone tumors. For example, pH-responsive polymers were synthesized as injectable particles for the targeted delivery of bortezomib to metastatic bone tumors. The particles loaded with drugs remained stable in normal tissue and rapidly released bortezomib in the acidic microenvironment, which effectively inhibited the progression of metastatic bone tumor and significantly inhibited tumor-related osteolysis
[212]. Similarly, using alendronate as a bone-targeting ligand and embedding bortezomib-catechol conjugate, injectable pH-responsive micelles were developed that significantly inhibited the growth of metastatic bone tumors
in vivo and reduced bone destruction at metastatic sites
[213]. To improve the specificity of chemotherapeutic drugs, dual bone/tumor-targeted injectable nanoparticles were developed as a paclitaxel carrier and effectively delivered the drug. These nanoparticles were able to accumulate in bone metastases
in vivo and inhibited 4T1 tumor growth and lung metastasis
[336].
5.5. Injectable biomaterials to treat bone defect-related infections
According to clinical research, if infection occurs at the implantation site, then the therapeutic effect will not be good, especially in large and medium-sized orthopedic operations. When infections occur, the patient mortality rate rises, resulting in an extended hospital stay and a significantly increased hospitalization cost
[337]. Due to the irrational use of antibiotics, antibiotic-resistant bacteria are gradually increasing
[338]. Therefore, it is urgent to develop materials that can effectively resist infection and promote bone regeneration
[339].
5.5.1. Injectable biomaterials loaded with antibiotics
Hydrogels are widely used for tissue regeneration because of their biocompatibility
[340],
[341]. From naturally derived materials to synthetically derived materials, hydrogels are promising for bone tissue engineering
[342],
[343]. Many injectable hydrogels that are crosslinked
in situ and can deliver antibacterial agents/antibiotics are widely used in bone regeneration
[344],
[345]. Recently, dual-function injectable hydrogels have been widely adopted, as these gels can kill bacteria and release antibiotics into the surroundings
[346]. The antibiotics are selected depending on their bacteria-killing efficacy against Gram-positive bacteria and Gram-negative bacteria. The most efficient drugs, such as vancomycin, teicoplanin, quinupristin/dalfopristin, oxazolidinones, daptomycin, telavancin, and ceftaroline, are known to be highly active against Gram-positive bacteria
[347]. Cephalosporins, fluoroquinolones, aminoglycosides, imipenem, and broad-spectrum penicillins—with or without β-lactamase inhibitors—are known to be effective against various Gram-negative bacteria
[348]. For example, a functional injectable hydrogel loaded with vancomycin was able to deliver antibiotics locally. The antibiotics were primarily encapsulated by means of a reversible imine bond formed between vancomycin and dextran aldehyde in the hydrogel; this realized the sustained release of vancomycin in a pH-dependent manner and achieved good results in
in vivo experiments
[346].
In recent years, the research hotspot of novel domestic and international antibacterial medicine has started to include the biological engineering of enzymes with unique bactericidal activity, such as staphylococcus lysozyme and lysozyme
[349]. These enzymes can cut off pentaglycine peptide-bond bridge construction in the aureus cell wall peptidoglycan, thereby reaching, rapidly dissolving, and killing the targeted bacteria
[350]. As enzymes usually have high protein activity but poor stability, they are incorporated into carriers to enhance stability and improve the release rate of the protein. In one study, lysostaphin-loaded CPC controlled the release of the enzyme and exhibited good antibacterial activity against methicillin-resistant
Staphylococcus aureus, demonstrating its potential for the treatment of bone defects and infections.
5.5.2. Injectable biomaterials loaded with bacteria-killing particles
The nanoparticle-mediated drug-delivery approach has recently been widely investigated as a means of delivering drugs to treat and respond to biofilm-related infections
[351]. Many nanoparticles have good antibacterial properties, and nanomaterials such as nano-ZnO have attracted a great deal of attention in various fields. Nanoparticles have a large surface-to-volume ratio, and some (e.g., Ag, TiO
2, ZnO, and chitosan) exhibit good antimicrobial activity
[352],
[353],
[354],
[355]. ZnO exhibits bacteria-killing efficacy against both Gram-positive bacteria such as
Staphylococcus aureus and Gram-negative bacteria such as
Escherichia coli [356]. For example, ZnO particles were synthesized via a sol–gel method and added to alginate-based injectable hydrogels to treat bone-defect-related infections during the process of cartilage regeneration
[357]. Recent research suggests that the antibacterial mechanism of ZnO particles may involve Zn
2+ release and oxidative stress
[358]. It is believed that microorganisms carry a negative charge, whereas metal oxide carries a positive charge. This creates an “electrostatic” interaction between bacterial infection and metal oxide. Once contact is made, the bacteria are oxidized and die instantly. The antibacterial activity of ZnO mainly results from the adhesion of bacteria onto the ZnO surface. The ions released from the ZnO react with the thiol groups (–SH) of the protein presenting on the cell surface of the bacteria. After that, the proteins are inactivated by the ZnO, which decreases the membrane permeability and eventually causes cellular death
[359].
Silver-based nanoparticles are another important antibacterial substance; they can be used as antimicrobial filler in injectable biomaterials (
Fig. 13)
[360]. Polymer-silver nanocomposites are widely used because of their easy modification and biodegradable properties
[353].
6. Conclusions and perspectives
Great success has already been achieved with implantable biomaterials for MIS in the field of orthopedics. These materials not only facilitate operations and reduce complications but have also revolutionized the surgical process to some extent. In the future, the continuous development of implantable biomaterials—ranging from bioceramics to polymers—will provide an increasing number of options and possibilities for MISs.
However, the fabrication of perfect bone-substitute materials that can match with natural bone tissues remains a major challenge. Therefore, additional expectations remain to be met by next-generation bone substitutes. Based on our understanding, future trends in implantable biomaterials for MIS might lie in the following aspects:
(1) Many biodegradable materials, such as calcium phosphates and calcium sulfates, have been proposed as alternatives for PMMA. However, thus far, PMMA remains as the dominant synthetic bone-substitute material in clinical applications. This is especially evident for biomaterials applied in load-bearing areas such as vertebral bone augmentation. Ideally, in these applications, the degradation rate of the bone graft material should match the rate of new bone formation, and the mechanical strength of the material should be sufficient to support the bone structures during the process. Currently, control of the mechanical strength and biodegradability of implantable materials remains a major challenge. In addition, most studies use static compressive and tensile strength measurements to evaluate the mechanical performance of bioceramics. Such measurements might be arbitrary, since bone-substitute materials are in a more dynamic environment after implantation. Therefore, other measurements such as fracture toughness and fatigue performance tests need to be applied in order to better predict the mechanical behavior of biomaterials in vivo.
(2) As the field of tissue engineering continues to expand, scaffolds that can promote bone regeneration are greatly desired. Scaffold engineering requires a deep understanding of cell-materials interaction and rational design of the architecture to best mimic the environment of the ECM. Recently, 3D-printing technology has been providing numerous possibilities for personalized and minimally invasive treatments. This technique permits the architectural design of scaffolds at different levels, which might pave the way for the clinical application of tissue engineering. In addition, to achieve optimal curing effects, scaffolds should protect cells from destructive shear stress during the injection process. Therefore, further investigations need to be performed to design novel scaffolds that can increase the survival rate of cells.
(3) Biomaterials achieve osteogenesis in various ways, including angiogenesis, immunomodulation, and mineralization promotion. An understanding of the detailed mechanisms involved is not only important for the improvement of novel fracture-healing therapies but also critical for the development of methods to fabricate biomimetic bone substitutes for MISs. This is a rather complex issue that requires close collaborations among chemists, biologists, and materials scientists to reveal mechanisms.
(4) In addition to the biocompatibility and bioactivity of implantable biomaterials, increasing attention has been paid to other properties, such as bone implant fixation, anti-infection, and anti-cancer properties. Thus, there is a demand for materials with multifunctional and bioresponsive properties. Thus far, only small quantities of such materials are available, and most are based on hydrogel and microsphere systems. Further development of minimally invasive implantable biomaterials is required in order to meet various clinical needs.
Acknowledgments
This work was supported by the National Natural Science Foundation of China (81925027, 82002275, and 32271421) and the Priority Academic Program Development of Jiangsu Higher Education Institutions.
Compliance with ethics guidelines
Feng Han, Zhao Liu, Qiang Wei, Luguang Ding, Li Yu, Jiayuan Wang, Huan Wang, Weidong Zhang, Yingkang Yu, Yantao Zhao, Song Chen, and Bin Li declare that they have no conflict of interest or financial conflicts to disclose.
PDF
(4294KB)
