1 Clinical need for degradable bioceramics

The demand for new bone substitutes with the physical, mechanical and biological properties matching that of bone is of great importance. Starting from the early 1960s, a wide variety of biomaterials such as metals1, polymers2, ceramics3, 4 and composites5 have been used in the biomedical field. Among all, ceramics are increasingly used to treat bone defects, small fractures of tibia, dental, maxillofacial reconstruction and spinal discs6. Bioceramics, particularly calcium phosphate-based ceramics, namely hydroxyapatite (Ca10(PO4)6(OH)2) (HA), dicalcium phosphate dihydrate (CaHPO4·2H2O) (DCPD) and tricalcium phosphate (Ca3(PO4)2) (TCP), have been widely used as bone replacement substituents because of their similarity in chemical composition to the natural bone, good biocompatibility and osseointegration. Limited solubility of HA and TCP also raises a question on the true nature of the degradable bioceramics7,8,–9. Keeping the degradability in mind, newer bioceramics have been developed for lower life span in the human body with similar biocompatibility and bone regeneration ability.

The chemical composition of bioceramics is an important parameter that governs their degradability, biocompatibility and bioactivity when implanted in vivo. Further, it has been reported that the dissoluted or degraded ions, from the implanted material, stimulate the surrounding environment for new bone formation10. Recently research efforts have focussed on biodegradable implant materials that degrades completely in the body leaving no toxic products and, hence, no secondary surgery is needed for implant removal11. For this purpose, the implant materials are made with an interconnected porous structure which allows the diffusion of essential nutrients, and cell proliferation which allows the bone tissue to grow forming a strong bond with the implant material12,13,–14. Recently, magnesium-based silicate ceramics, forsterite (Mg2SiO4) (MgS), a young field of research, have been investigated because of the beneficial role of Mg and Si ions for bone regeneration15, 16. The enhanced degradation rate, good biocompatibility and solubility of forsterite gained significant interest in the field of bone tissue engineering10, 17. Research findings of Naghiu et al. proved that the dissoluted products of forsterite improved the osteoblast proliferation of U20S-type cells without cytotoxic effect18. Similar studies of Ni et al. also showed the well-covered proliferated cells on the surface of the forsterite ceramics, confirming that the degraded ions from forsterite bioceramics increased the proliferation rate than control samples over 7 days19.

This review focuses on the growth and achievements of forsterite bioceramics for bone regeneration applications till date. We first discussed the importance of Mg and Si ions in bone regeneration metabolism. Further, we moved to the synthesis methods using various techniques followed by complete in vitro biological behaviour and in vivo biocompatibility of the forsterite ceramics. Finally, we conclude that the forsterite bioceramics are suitable compared with the calcium phosphate-based ceramics in the future directions and the clinical implementations of it.

2 Significance of Mg and Si ions in bone metabolism

In addition to calcium (Ca), magnesium (Mg) is also an essential element in the human body as it plays an important key role in bone metabolism, DNA stabilization, and skeletal development20, 21. Research findings have showed that magnesium increases the proliferation and stimulation of osteoblastic cell growth22,23,–24. The released Mg ions from porous magnesium scaffold showed good cytocompatibility and increased the ALP activity and expression of osteogenic differentiation. Further, the in vivo studies in rabbit model reported mature bone regeneration at the bone–implant site25. The deficiency of Mg leads to loss in bone mass, decreased bone growth, increased skeletal fragility and a risk factor for osteoporosis26,27,28,–29.

Similarly, silicon (Si), an important element in the metabolic processes, helps in the development and calcification of bone tissue30. Studies have showed that the trace amount of Si helps in bone repair, regulates the production of collagen type 1, and plays an essential role as a cross-linking agent in the connective tissue31. It also improves bioactivity, promotes osteoblast differentiation and mineralization of bone32. Substitution of 0.8% wt of Si in hydroxyapatite enhanced the metabolic activity, stimulated the expression of type I collagen and increased the proliferation of human osteosarcoma cells (HOBs)33. Similar studies of Balamurugan et al. showed that the incorporation of 5 mol % of Si into hydroxyapatite enhanced the dissolution rate of the material and increased the proliferation rate of human osteoblast cells compared to undoped HA34. The bioactivity of undoped and Si-doped (0.8 wt% and 1.5 wt%) hydroxyapatite was analysed by immersing all the samples in simulated body fluid (SBF). The (0.8 wt%) Si-doped HA showed higher bioactivity and solubility, with increased proliferation of osteoblast cells35. Patel et al. reported the in vivo studies of pure HA and 0.8 wt% Si-doped HA in the femoral condyle of rabbits for a period of 23 days. The percentage of new bone growth was higher for Si-doped HA (37.5 ± 6%) compared to pure HA (22 ± 6%) confirming the significant role of Si in new bone formation36. In another study, pure HA and Si-substituted HA (0.8 and 1.5 wt%) were implanted in an ovine defect model for 6 and 12 weeks. The new bone regeneration was significantly higher for Si-doped HA without any inflammation making it a suitable material for bone graft substitutes37. Thus considering the essential role of Mg and Si ions, the bioceramics based on the above material will play an essential role for bone tissue engineering.

3 Synthesis of Magnesium Silicate Bioceramics

Various synthesis routes, such as mechanical activation38,39,–40, sol–gel18, 19, 41,42,–43, and hydrothermal44 methods have been reported (Table 1) for synthesis of magnesium silicate bioceramics for bone regeneration applications. The common difficulty faced during the synthesis of forsterite ceramics was the formation of the intermediate enstatite (MgSiO3) and periclase (MgO) as impurities45. To decrease these by-products, researchers have used solid-state method with extended ball milling (9 h). The synthesized powder when compacted and sintered at 1200 °C for 2 h showed phase pure forsterite10, 17, 46. In another study, nanocrystalline forsterite powders were prepared via solid-state reaction, using talc (Mg3Si4(OH)2) and magnesium oxide (MgO) as the starting precursors. Here, the authors have studied the mechanical properties and microstructure of the prepared ceramics. In the first method, the samples were prepared by mechanical activation method via sonication, and then it was ball milled with varying time and heat treated at different temperatures. In the second method, the same procedures were followed, but the samples were not heat treated before sintering. The authors have showed that formation of secondary phases such as MgO and MgSiO3 can be decreased with increase in ball milling time, sintering temperature and higher amplitude of ultrasonication47. Anovitz et al. combined both sol–gel and surfactant method to obtain fine particles of forsterite without any contamination by burning the powder at 800 °C overnight48.

Table 1: Review of forsterite synthesis methods, raw materials and sintering temperature.
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In a similar study, Tavangarian and Emadi showed that pure forsterite bioceramics can be prepared via the mechanical activation method using ammonium chloride as a catalyst, annealed at 1000 °C for 1 h. The results showed that the synthesized nanostructured forsterite was bioactive and had the ability of apatite formation49. Choudhary et al. used sol–gel combustion method to prepare forsterite powder. They used urea and glycine as two different fuels and studied the degradation properties and antibacterial activity of the synthesized and compacted forsterite bioceramics structures45. Saberi et al. synthesized nanocrystalline forsterite powder via cost-effective low temperature polymer matrix method50. In a related studies, Tavangarian et al. prepared nanocrystalline forsterite powder by varying the milling time from 5 h up to 60 h and annealed the powder at two different temperatures (1000 °C and 1200 °C) for 1 h51 [Fig. 1 (reproduced with permission)]. The forsterite phase formation was noticed with increase in milling time up to 5 h and further increase in temperature showed no significant effects on phase formation. The particle size and crystallite size were found to be less than 1 µm and 60 nm, respectively38, 40, 51.

Figure 1:
figure 1

X-ray diffraction patterns of forsterite powders with varying milling time and sintering temperatures annealed a at 1000 °C, b at 1200 °C for 1 h, from Tavangarian et al.51.

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The combination of sol–gel and mechanical activation method was employed for the synthesis of single-phase nanostructured forsterite powder52. The obtained ball milled powder was sintered at low temperature starting from 600 up to 1000 °C for 1 h to get pure nano forsterite powder. With increase in temperature above 750 °C, the formed MgO reacts with the SiO2 at the surface level and forms enstatite (MgSiO3). Further, the produced MgO diffuses into enstatite particles, resulting in the formation of pure forsterite powder53. Research group of Bafrooei et al. synthesized nano forsterite powder through microwave-assisted high-energy ball milling technique. Magnesium hydroxide and silica gel were used as starting materials. The raw powders were milled up to 40 h; further the powders were calcined at 500–1200 °C using microwave heating. The results indicated that at 900 °C, pure single phase forsterite was formed without any other phases such as SiO2 and MgO. The obtained powders were compacted and sintered using conventional sintering (1150–1350 °C) and microwave sintering (1.1 kW, 2.45 GHz multimode microwave). From the studies, it was concluded that microwave sintered materials showed higher densification with smaller grain size and uniform grain growth54.

Using a novel cost-effective, low-temperature method, nanocrystalline forsterite was synthesized using magnesium nitrate and sucrose as a template material. The obtained powder was calcined in an electric furnace varying from 500 to 1000 °C for 3 h. Initial crystallization of forsterite was seen at 730 °C and with increase in temperature complete crystallization of the forsterite phase was seen at 800 °C with the average particle size around 200 nm50. Recently, Kheradmandfard et al. synthesized pure nano forsterite at low temperature (800 °C) using an ultrafast, green synthesis method via microwave irradiation technique55. Magnesium nitrate hexahydrate (Mg(NO3)2·6H2O) and tetraethyl orthosilicate (TEOS) were used as starting materials. The reaction mixture was refluxed in a modified microwave oven at 850 W and 2.45 GHz. Further, the mixture was subjected to microwave irradiation for 15 min. The obtained powders were dried at 90 °C for 1 h and sintered at 800 °C for 2 h. The main advantage of this technique is it increased the rate of the reaction and decreased the reaction time period. The average particle size of the prepared forsterite powder was found to be around 100 nm.

3.1 Addition of Dopants

The major advantage of forsterite bioceramics is its controlled degradation behaviour, enhanced mechanical properties and improved biological response. Further, these properties can be tuned by addition of essential cations. Several studies have reported that the incorporation of essential trace elements, such as; manganese (Mn), zinc (Zn), strontium (Sr), silver (Ag) and iron (Fe) in calcium phosphates played an important role in improving the mechanical and biological properties56,57,–58. With the scientific understanding of dopant-induced change in biocompatibility of CaPs, a new area of research on doped forsterite had started. Gheitanchi et al. prepared pure forsterite and Sr-doped forsterite nanopowder via the sol–gel route59. During the preparation process, the authors used PVA and sucrose to avoid the dissimilarity in hydrolysis and condensation process which bring out inhomogeneity of the reaction. They varied the percentage of Sr from 0, 0.05, 0.1, 0.2 and 0.4 at.% and sintered all the samples at 800 °C and 1000 °C for 2 h. Apart from the forsterite phase, other secondary phases such as MgSiO3, MgO and Sr2MgSi2O7 were also observed in the Sr-doped forsterite nanopowder. In a similar research, Devi et al. prepared pure ZnO- (0, 0.25, 0.5 wt%) and SrO (0, 1, 2, 3 wt%)-doped forsterite powder by the solid-state method. The prepared undoped and doped powders were compacted into circular discs and sintered at 1200 °C for 2 h using a muffle furnace. By increasing the dopant composition, the degradation and the porosity of the samples also increased10, 17, 46.

3.2 Composite Systems

Composite systems particularly in forsterite ceramics are still an underexplored area of research. Forsterite composite systems, particularly with bioglass, silk and polymers, are of particular interest, as the incorporation of these materials increases the potential application fields. Recently, Saqaei et al. fabricated 58S bioactive glass–forsterite nanocomposite using the sol–gel technique. They varied the forsterite powder from 0, 10, 20 and 30 wt% and studied the effect of addition of forsterite nanopowder on the antibacterial and bioactivity behaviour60. A novel forsterite/silk fibroin composite was fabricated via the freeze-drying method. Four different kinds of samples were prepared in the ratio of (forsterite/silk fibroin) 0: 100, 20:80, 30:70 and 40:60 and all the prepared samples were freeze dried for 3 days61. The new self-curing nano forsterite biocomposites were prepared by combining the forsterite powder (0, 5, 15, 30, 50, 70 wt%) with organic monomer 2,2-bis[4-(2-hydroxy-3-methacryloyloxypropoxy)-phenyl]propane{bis-GMA} and triethyleneglycol dimethacrylate {TEGDMA}. All the fabricated samples showed hydroxyapatite growth with improved mechanical and biological properties62.

3.3 Influence of Pore Forming Agents

Porosity plays an important role in the overall degradation and biocompatibility of bioceramics. The pore of specific size and distribution helps in implant integration, improves the degradation rate and decreases the implant rejection63, 64. The porous nanocomposite material consisting of forsterite nanopowder/polycaprolactone (PCL) was prepared through solvent-casting/particle-leaching method using sodium chloride particles as porogen. The PCL composites with varying (10–50 wt%) forsterite nanopowders were prepared. All the prepared samples were air dried for 48 h to evaporate the solvents completely. The synthesized composites exhibited well-interconnected porous structure with pore size varying from 100 to 400 µm and with the porosities around 90–92%14. Mesoporous forsterite with pore size of 5 nm was fabricated using magnesium nitrate hexahydrate and tetraethyl orthosilicate as starting material and P123 (EO20PO70EO20) as template. The prepared disc-shaped samples were calcined at 600 °C for 1 h in air to remove the template. The synthesized mesoporous forsterite showed enhanced degradation, good cytocompatibility, and higher surface area with pore volume of 0.41 cm3/g65. In a similar way, Bigham et al. fabricated ordered mesoporous magnesium silicate (OMMS) via the sol–gel route using P123 (EO20PO70EO20) as a surfactant. The synthesized product was air dried at 50 °C for 24 h and calcined at 350–750 °C for 4 h to remove all the surfactant. The samples calcined at 750 °C showed wider pore size distribution with around 20 nm pore size, making the material suitable for drug delivery systems66. In another study, the authors have prepared mesoporous magnesium silicate via the hydrothermal method using polyethylene glycol (PEG) as surfactant. The resulting mixture was washed several times to remove the inorganic ions and the samples were dried overnight at 90 °C. Finally, all the samples were calcined from 500 to 700 °C for 4 h. On increasing the calcination temperature, the pore size of the samples increased with decrease in surface area67.

4 In Vitro Behaviour

4.1 Dissolution and Biodegradation Properties

For any clinically adoptable implant, the degradation rate of the material should match the rate of new bone formation. This degradation takes place via two processes: (1) chemical dissolution and (2) cell-mediated resorption. The degradation or dissolution study in the artificial media is a well-known method in which pH, weight loss and ionic release profile are measured with respect to time. Various researchers have suggested that forsterite bioceramic can be employed in bone regeneration applications because of its enhanced degradation rate22, 62, 65. Choudhary et al. examined the degradation behaviour of forsterite powders prepared using two different fuels [glycine (FG) and urea (FU)]. They reported that forsterite ceramics when immersed in SBF solution showed weight loss of around 2.8% (FG) and 0.78% (FU) respectively. FG exhibited enhanced weight loss due to smaller particle size and higher surface area compared to FU45. Further, Kharaziha and Fathi synthesized forsterite nanopowder with a particle size of 25–45 nm. The authors evaluated the bioactivity of the samples by soaking the forsterite compacts in SBF for up to 28 days and found clusters of agglomerated hydroxyapatite that increased with increase in time [Fig. 2a (reproduced with permission)]. FT-IR spectra showed the characteristic bands of apatite crystals at 1030–1090 cm−1, 574 cm−1 and 471 cm−1. The CO32−groups of apatite were observed at 1418 cm−1, 1462 cm−1 and 872 cm−1. The bands at 3477 cm−1 and 1619 cm−1 are assigned to be hydroxyl groups in apatite [Fig. 2b (reproduced with permission)]. The XRD pattern confirmed the formation of apatite covering the surface of the forsterite samples [Fig. 2c (reproduced with permission)]. Further, the released concentration of Mg ions in SBF continued to increase, whereas the Ca and P ions gradually decreased confirming the formation of hydroxyapatite layer over the forsterite compacts41.

Figure 2:
figure 2

a SEM and EDX images. b FT-IR spectra. c XRD pattern of forsterite nanopowders immersed in SBF for various time points, from Kharaziha et al.41.

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In a similar study, Naghiu et al. evaluated the bioactivity of the synthesized nanopowder for 7 and 14 days in SBF solution and found that the forsterite nanoceramics were highly bioactive and biocompatible. After immersion in SBF, the formation of hydroxyapatite was confirmed using FT-IR, SEM–EDX and XRD18. The AAS analysis by Tavangarian and Emadi showed that the nanostructure forsterite bioceramics released Mg ions into the SBF solution, indicating that forsterite bioceramics were biodegradable and bioresorbable, hence making it an excellent material for bone tissue engineering applications49. The surface morphology of undoped and forsterite (0, 5, 15, 30, 50, 70 wt%)-doped polymer biocomposites (2,2-bis[4-(2-hydroxy-3-methacryloyloxypropoxy)-phenyl]propane (bis-GMA)/triethyleneglycol dimethacrylate (TEGDMA), before and after immersion in SBF for 0, 14 and 28 days, was analysed using AFM and the images are shown in Fig. 3 (reproduced with permission). From the AFM images, it can be seen that there is an increase in surface roughness for all the samples after 14 days immersion in SBF. The surface roughness (root mean square: RMS) value for undoped and forsterite-doped biocomposites before and after immersion in SBF are shown in Table 2 (reproduced with permission). The pure polymer sample did not show any characteristic morphology even after 28 days of immersion, whereas after 14 days of immersion forsterite/polymer composites showed new fibrous structure because of the presence of forsterite interaction with the SBF solution. After 28 days, there was a significant increase in surface roughness. Some imperfections and slight erosion are seen on the surface of the doped samples because of the deep penetration of SBF62.

Figure 3:
figure 3

2D AFM topographies of pure and forsterite/polymer biocomposites before and after immersion in SBF for 14 and 28 days, from Furtos et al.62.

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Table 2: Surface roughness values (RMS) (root mean square) for pure polymer and forsterite/polymer composites before and after immersion in SBF for various days, from Furtos et al.62.
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Devi et al. compared the biodegradation property of forsterite with that of commonly used HA and β-TCP ceramics (Table 3). All the sintered samples were immersed in SBF for 8 weeks to evaluate the degradation behaviour. It was found that forsterite bioceramics lost 9 wt% compared to that of 1wt% for β-TCP ceramics, whereas HA samples gained 2.5 wt% after 8 weeks of immersion. These results were further supported by increase in pH and ionic concentration of released Mg ions in the dissolution media10 [Fig. 4 (reproduced with permission)]. Further, addition of dopants such as Zn and Sr enhanced the degradation property of the ceramics. Upon addition of 0.5 wt% of Zn enhanced the weight loss to 22 wt% when compared to pure forsterite (9%) (Fig. 5). The pH of the dissolution media increased for Zn-doped ceramics because of the higher dissolution rate that was supported by the ionic release profile of Mg ions, which showed threefold release for Zn-doped forsterite ceramics17. In a similar way, addition of 3 wt% Sr into forsterite degraded ceramics much faster with a weight loss of 12% compared with pure forsterite (9%) (Fig. 5). The cumulative release profile of Mg2+ ions (1300 ppm) and Si4+ ions (300 ppm) increased rapidly for 3 wt% Sr after 8 weeks of immersion10.

Table 3: Summary of degradation behaviour of HA, TCP, undoped and doped forsterite.
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Figure 4:
figure 4

Comparison of the degradation behaviour of forsterite, HA and β-TCP a weight loss, b change in pH and c dissolution of released (Mg2+) (Ca2+) ion concentration at various time points, from Devi et al.10.

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Figure 5:
figure 5

Weight loss/gain behaviour of released HA, TCP, doped Zn and Sr ions at different time points, from Devi et al.10, 17.

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4.2 Mechanical properties

Mechanical properties of the bioceramics play an important role in determining the stability of the material in vivo for the bone regeneration application. The natural bone has an average compressive strength of around 130–190 MPa of cortical bone and 3.6–9.3 MPa for cancellous bone68. When implanted inside the human body, the strength of the designed implant should be similar to that of natural bone. Ni et al. synthesized forsterite powder using the sol–gel method and studied the mechanical properties of the ceramics by varying the sintering temperature and time. At 1350 °C, the bending strength and fracture toughness of forsterite ceramics was found to be 150 ± 8 MPa and 1.8 ± 0.4 Mpa m1/2. With increase in the sintering temperature from 1350 to 1450 °C, the bending strength (181 ± 9 MPa) and fracture toughness (2.3 ± 0.1 Mpa m1/2) also increased for these samples. On further increase in temperature from 1450 to 1550 °C, the bending strength (145 ± 8 MPa) and fracture toughness (1.6 ± 0.2 Mpa m1/2) of the ceramics decreased due to the formation of flaw structure and grain coarsening which was further confirmed by fractographical analysis. The fracture surface of forsterite ceramics sintered at 1450 °C and 1550 °C is shown in Fig. 6 (reproduced with permission). The samples sintered at 1450 °C showed sharp-edged pores with an average grain size of 10 µm. Further, increasing the temperature up to 1550 °C causes grain growth which traps the pores inside, resulting in decreased mechanical properties. The authors have shown significant improvement in fracture toughness for forsterite samples (2.3 ± 0.0 Mpa m1/2) when compared to hydroxyapatite (0.75–1.2 Mpa m1/2) ceramics19.

Figure 6:
figure 6

SEM micrographs of fracture surface of forsterite ceramics after sintering at two different temperatures at a 1450 °C and b 1550 °C for 8 h, from Ni et al.19.

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The forsterite scaffolds prepared using two different fuels, namely glycine (FG) and urea (FU), showed variations in the compressive strength and Young’s modulus. Both the samples had superior compressive strength in comparison to natural cortical bone. The FG showed a higher compressive strength (201 MPa) and Young’s modulus (4.8 GPa) compared to FU, which has compressive strength of 124 MPa and Young’s modulus 4.6 GPa (Fig. 7). This difference in mechanical property was due to difference in particle size. FG had smaller particle with higher surface area compared to FU. These results proved that the compressive strength of the synthesized forsterite scaffolds had better strength compared to cortical bone (130–200 MPa) even after 1 month of soaking time in SBF45. Ghomi et al. fabricated porous forsterite scaffolds with interconnected pore size from 50 to 200 µm using the gelcasting method69. The compressive strength and the elastic modulus of the prepared forsterite scaffold sintered at 1200 °C for 4 h was found to be 2.43 ± 0.11 MPa and 182 ± 19, respectively, which is close to the compressive strength (2–12 MPa) of cancellous bone70.

Figure 7:
figure 7

Compressive strength and Young’s modulus of forsterite prepared using glycine and urea as different fuels, from Choudhary et al.45.

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In another research, forsterite powder combined with silk fibroin composite was fabricated using the freeze-drying method. The compressive strength (1.25–4.6 MPa) and modulus (1.3–4.7 MPa) of the composite increased and the porosity of the samples (92–83%) decreased with addition of forsterite powder from 0 to 40 wt% [Fig. 8a–c (reproduced with permission)]. The authors found that the decrease in pore size and increase in the particle wall thickness were the major reasons for improvement in the mechanical properties of the composite scaffolds compared to pure silk fibron61. The dense two-step sintered (TSS) nanostructured forsterite powder with the crystallite size of 30–45 nm was prepared via the sol–gel method. The sintering behaviour and the mechanical properties (hardness and fracture toughness) are well studied. The results showed that with increase in soaking temperature from 750 to 850 °C for 5 h, both the fracture toughness (1.10 ± 0.5 MPa m1/2 to 4.3 ± 0.19 MPa m1/2) and hardness (520 ± 45 Hv to 1102 ± 25 Hv) values increased. This increase in fracture toughness could be due to dense forsterite powder and sintering temperature of the powder via the two-step sintering method71. These results were best in comparison with Ni et al., who prepared forsterite samples from coarse grain powder and obtained lower fracture toughness of 2.4 MPa m1/2. Overall, it was found that the fracture toughness of the dense forsterite ceramics (KIC = 4.3 MPa m1/2) is better than that of hydroxyapatite ceramics (KIC = 0.75–1.2 MPa m1/2)19.

Figure 8:
figure 8

Mechanical properties of silk fibroin and silk fibroin/forsterite composite scaffolds. a Compressive modulus, b compressive strength and c porosity values, from Teimouri et al.61.

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The nanoporous scaffolds comprising nano forsterite with polycaprolactone (PCL) were synthesized by the solvent-casting/particle-leaching method. The effects of the addition of nano forsterite powder varying from 0, 10, 20, 30, 40 and 50 wt% in the PCL were investigated. SEM micrographs of pure PCL and PCL–forsterite composite are shown in Fig. 9 (reproduced with permission). Pure PCL showed macroporous structure with the pores varying from 100 to 400 µm with 92.65% of porosity. With increase in forsterite powder (0–30 wt%), the porosity decreased to 91.38% with the open pores of about 100–300 µm. In addition to macropores, the composite scaffolds possess plenty of micropores (1–10 µm) with rough uneven surface walls. This open and porous interconnected structure enables the transportation and proliferation of osteoblasts cells, making it a suitable candidate for bone tissue engineering applications14. In a related studies, Furtos et al. investigated the novel forsterite biocomposite by mixing forsterite powder (5, 15, 30, 50 and 70 wt%) with 2,2-bis[4-(2-hydroxy-3-methacryloyloxypropoxy)-phenyl]propane (bis-GMA) and triethyleneglycol dimethacrylate (TEGDMA) monomers. Upon addition of forsterite powder from 0 to 50 wt%, the compressive strength (128.70 ± 12.90–167.49 ± 10.15 MPa), flexural strength (80.55 ± 12.51–83.20 ± 6.55 MPa) and diametral tensile strength (29.72 ± 2.70–31.55 ± 2.75 MPa) increased, whereas the compressive modulus (1.49 ± 0.20–2.75 ± 0.23 MPa) and flexural modulus (1.94 ± 0.60–7.37 ± 1.85 MPa) increased up to 70% addition of forsterite. Above 50 wt% addition, the compressive strength (147.49 ± 20.84 MPa), flexural strength (59.47 ± 9.81 MPa) and diametral tensile strength (25.45 ± 2.54 MPa) decreased as the excessive powder became a rigid filler, causing phase segregation and bringing more stress inside the polymer matrix62. These results are also in line with other published results72, 73.

Figure 9:
figure 9

SEM micrographs of surface morphology of a pure PCL, b PCL-10 wt% forsterite, c PCL-20 wt% forsterite, d PCL-30 wt% forsterite, e PCL-40 wt% forsterite and f PCL-50 wt% forsterite, from Diba et al.14.

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In a recent research, Devi et al. prepared pure forsterite, Zn- and Sr-doped forsterite using the solid-state method. The porosity, pore and materials distribution of all the samples before and after immersion in SBF for 8 weeks were analysed using 3D-micro-CT and are represented in Fig. 10 (reproduced with permission). Addition of dopants (Sr and Zn) increased the porosity from 1.15% (pure forsterite) to 1.80% (0.25 wt% Zn), 2.29% (0.5 wt% Zn), 1.70% (1 wt% Sr), 1.79% (2 wt% Sr) and 3.06% (3 wt% Sr), respectively. Similarly, because of enhanced dissolution and degradation behaviour of the forstertie samples, the porosity increased after 8 weeks of immersion for all the samples [(2.03% pure forsterite) (2.63% 0.25 wt% Zn) (2.53% − 0.5 wt% Zn) (2.63% − 1 wt% Sr), (3.00% − 2 wt% Sr) and (4.53% 3 wt% Sr)]. The interconnected pore size was found to be in the range of 510–580 µm for 0 week samples and 730–790 µm for subsequent 8 weeks immersion in SBF10, 17, 46.

Figure 10:
figure 10

3D Micro-CT images of the porosity, pore and materials distribution of pure, Zn- and Sr-doped forsterite before (a–f) and after immersion (g–l) in SBF, from Devi et al.10, 17.

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4.3 Cytocompatibility

In vitro cytocompatibility of magnesium silicate bioceramics, for bone tissue engineering applications, has been widely studied. The most commonly used cells were G292, MG-63 and MC3T3-E1 cells. Gheitanchi et al. evaluated the bioactivity of nano forsterite ceramics using MG63 osteoblast cells. The MTT assay showed that both pure forsterite and Sr forsterite samples gradually increased the MG63 proliferation from day 1 to day 759. Similar study of Furtos et al. showed that the incorporation of nano forsterite powder (0, 5, 15, 30, 50 wt%) in the bis-GMA/TEGDMA polymer matrix increased the Saos-2 viable cells from 57.8% ± 24.4% (0% forsterite only polymer, P) to 82.6% ± 6.6% for 50% forsterite powder after day 1. After 14 days, all the samples were examined using fluorescence microscopy (Fig. 11 (reproduced with permission) after live/dead staining which showed the complete coverage of viable and well attached cells for all the forsterite biocomposites compared to the undoped polymer matrix62. The MTT assay on forsterite powder with varying concentrations (6.25%, 12.5 and 25%) and days (1, 2 and 5 days) showed that the proliferation rate of U20s-type osteoblasts cells increased confirming the absence of cytotoxicity on all the samples18. Krishnamurithy et al. synthesized nanocrystalline forsterite powder via the sol–gel combustion route using urea as fuel (FU)74. The authors evaluated the in vitro biocompatibility and osteogenic differentiation properties. The forsterite samples seeded with human bone marrow-derived mesenchymal stromal cells (hBMSCs) were compared with commercial bone substitutes (cBS). The hBMSCs had completely colonized the surface of the forsterite scaffold compared to cBS confirming that the degraded products of forsterite were biocompatible and showed the positive way for cell attachment and proliferation. The F-actin analysis depicted that hBMSCs seeded on forsterite samples have underwent actin cytoskeleton remodelling to interact with the forsterite samples. From this phenomenon, we can confirm that the surface chemistry of forsterite was favourable for cell attachment75. Further, the authors analysed the osteogenic intra- and extracellular protein expression. Increased BMP2 protein secretion was noticed for hBMSCs seeded onto forsterite (day 1 and 14) when compared to cBS. This analysis showed that the hBMSCs seeded onto forsterite were committed to the pre-osteoblastic lineage. A significant expression of Col1 (both day 1 and 14) in hBMSCs seeded onto forsterite was observed compared to cBS. The substantial enhancement of cells secreting Osterix (OSX) was seen from day 1 to day 14 for forsterite scaffold compared to cBS. Further, the OPN matrix secretion was two times and eight times higher on day 14 and day 1 when hBMSCs were seeded on forsterite scaffold and cBS, respectively. These results supported that forsterite scaffold had greater potential to induce hBMSCs differentiation into osteogenic lineage compared to cBS. Similar research findings were seen when BMP2 expression on hBMSCs was higher when seeded on magnesium phosphate ceramics76. The released magnesium ions helped in promoting hMSCs differentiation via the Wnt signalling pathway77.

Figure 11:
figure 11

Fluorescence microscopy of Saos-2 cells cultured on discs for day 1 (a–f, a´–f´), day 7 (g´–l´) and day 14 (g–l). P (a, a´, g, g´), C5F-(b, b´, h, h´), C15F-(c, c´, i, i´), C30F-(d, d´, j, j´), C50F-(e, e´, k, k´), C70F-(f, f´, l, l´), from Furtos et al.62.

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In a similar way, the proliferation of MC3T3-E1 cells on TCP, forsterite and mesoporous forsterite samples was evaluated using the MTT assay. The optical density (OD) [Fig. 12a(a) (reproduced with permission)] values of both mesoporous forsterite and forsterite were found to be increased after 7 days of culture compared to TCP showing good cytocompatibility of the samples. In the same study, ALP activity of MC3T3-E1 cells (Fig. 12a(b) on mesoporous forsterite sample after 7 days showed significantly higher expression compared to TCP. The cell morphology of MC3T3-E1 cells [Fig. 12b (reproduced with permission)] on mesoporous forsterite exhibited well-flattened structure indicating no negative effect on cell viability or morphology. Thus, mesoporous forsterites are excellent candidates for bone tissue engineering applications65.

Figure 12:
figure 12

a (a) OD values of MC3T3-E1 cells on mesoporous forsterite, forsterite and TCP at 1,4 and 7 days (b) ALP activity of MC3T3-E1 cells on mesoporous forsterite, forsterite and TCP at 4 and 7 days. b Images of cytoskeletal morphology and spreading of MC3T3-E1 cells, (a) mesoporous forsterite and (b) forsterite after 4 days using confocal scanning microscopy (scale bar 50 µm), from Wu et al.65.

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Bigham et al. fabricated ordered mesoporous magnesium silicate (OMMS) using the sol–gel route and investigated the effect of calcination temperature on drug delivery property. The cell viability of all the samples with different calcination temperatures (350 °C, 550 °C, 750 °C) was investigated by osteosarcoma cell line (MG63) using MTT assay and it was found that all the samples showed no toxicity. The results indicated possible usage of the OMMS as drug carriers66. The biocompatibility of the forsterite ceramics was evaluated by Ni et al. who showed that at 7 days, a higher proliferation rate of G292 osteoblast cells was noticed19. An ultrafast, green synthesis method of nano forsterite showed no toxicity with improved cell proliferation. The ALP activity (Fig. 13a) (reproduced with permission) at day 7 and 14 improved the proliferation and differentiation of MG63 osteoblast cells compared to the control. SEM micrographs of cell morphology (Fig. 13b) (reproduced with permission) showed well-attached polygonal flattened and completely covered MG63 osteoblast cells on the forsterite samples. The dissoluted Mg and Si ions from forsterite powder induced the osteogenic differentiation of MG63 cells through up-regulating the expression of collagen and extracellular matrix proteins78,79,–80. The antibacterial activity of forsterite synthesized using two different fuels [glycine (FG) and urea (FU)] was analysed (Fig. 14) (reproduced with permission) against biofilm forming bacteria, such as S. aureus and E.coli. The authors reported that S. aureus was inhibited more than E.coli and the bacterial activity was higher for FG because of the enhanced surface properties such as surface area and particle size. The surface area and particle size of FG were found to be higher (65.1 m2/g) (28 nm) than that of FU (0.93 m2/g) (1.9 µm), respectively. Thus these results confirmed that forsterite can be used to prevent bacterial infection during surgeries and inhibit biofilm formation in medical implants81.

Figure 13:
figure 13

a ALP activity of MG63 cells cultured on forsterite and control after 7 and 14 days [*significant difference among the groups (p < 0.05)]. b SEM micrographs of MG63 cells cultured on forsterite sample after 2 days of culture, from Kheradmandfard et al.55.

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Figure 14:
figure 14

Antibacterial activities of forsterite ceramics prepared using two different fuels glycine (FG) and urea (FU) against S. aureus and E. coli. FG with S. aureusa 100 mg, b 200 mg, c 300 mg, FG with E. colid 200 mg, e 300 mg and FU against S. aureusf 300 mg, g 400 mg, h 500 mg, and FU with E. colii 300 mg, from Choudhary et al.81.

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In our recent research, the cytocompatibility of pure Zn (0.25 wt% and 0.5 wt%) and Sr (1 wt%, 2 wt%, 3 wt%)-doped forsterite was analysed using MC3T3-E1 cells at day 1 and day 3 and the results are shown in Fig. 15a (reproduced with permission). The fluorescence images of MC3T3-E1 after day 1 and day 3 presented similar cell attachment for pure and Zn-doped forsterite when compared with the control plate. The DNA quantification after day 3 showed significant increase in cell proliferation for forsterite and Zn-doped forsterite samples17 [Fig. 15b (reproduced with permission)]. In a similar way, we carried out the MC3T3-E1 cell proliferation of undoped and Sr-doped forsterite samples via DNA quantification [Fig. 15c (reproduced with permission)]. The cell proliferation rate of pure and Sr-doped forsterite samples was found to be increased with increase in culture time, in accordance with live/dead images which showed enhanced live cells on all the samples compared to the control10 These results confirmed that enhanced degradation behaviour of forsterite with and without any dopant showed good cell attachment and proliferation, rate making it a material suitable for biodegradable bone replacement.

Figure 15:
figure 15

a Live/dead images of MC3T3-E1 cells of undoped, Zn- and Sr-doped forsterite after day 1 (a–g) and day 3 (h–n) (scale bar = 300 µm). b, c DNA assay normalized to day 1 of control (#statistical significance), from Devi et al.10, 17.

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5 In Vivo Biocompatibility

Although one of the most important characterizations, in vivo biocompatibility of magnesium silicate bioceramics has rarely been evaluated. Very few literature exists on in vivo degradability and bone regeneration capability of these new class of bioceramics. In their recent work, Devi et al. for the first time investigated in vivo osteogenesis of pure forsterite, Zn (0.25 wt% and 0.5 wt%)46 and Sr (1 wt%, 2 wt% and 3 wt%)10 -doped forsterite samples by implanting the ceramics for 30, 60 and 90 days in the distal femur of white New Zealand rabbits. After the particular interval of time, the postoperated bone samples were studied for bone–implant interface. X-ray radiographs showed the loss in radio-opacity of all the ceramics indicating the enhanced degradation of the ceramics in vivo. After 90 days of implantation, the gradual bone regeneration was noticed in pure, Zn- and Sr-doped forsterite samples. Further, SEM [Fig. 16a (reproduced with permission)] was utilized to understand the bone–implant interfacial bridging. The 90 days images of Zn-doped forsterite implants showed thicker bone formation surrounding the implant material. The 90 days Sr-doped forsterite series indicated a strong fissure gap between the implant and bone. The thick interfacial new bone growth confirmed the binding properties of forsterite ceramics. The 3D micro-CT was used to study the detailed bone growth on the implanted ceramics. After 30 days, degradation on all the samples was noticed and bone started to regenerate on all the ceramics, confirming the osteoconductive property of the ceramics. Within 90 days, the new bone formed had undergone remodelling, which was revealed from the structural arrangement of bone [Fig. 16b (reproduced with permission)]. At 90 days, most of the ceramic material was found to be replaced by new trabecular bone. The enhanced dissolution behaviour of forsterite released Mg and Si ion in the vicinity of the implants. These released ions played a very important role in improving the bone formation in vivo26, 31, 82, 83.

Figure 16:
figure 16

a SEM images of bone implant interface of pure, Zn- and Sr-doped forsterite (scale bar = 30 µm) (red arrow showing the interface of bone and implant). b 3D micrographs showing new bone regeneration around the implant material for undoped, Zn- and Sr-doped forsterite for 30 (a–f) and 90 (g–l) days (red arrow showing the new bone formation and white arrow showing the implant material), from Devi et al.10, 46.

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The percentage of new bone formation was analysed using oxytetracycline (OTC) labelling and is shown in Figs. 17a, 18 and Table 4 (reproduced with permission). At 30 days, golden yellow fluorescence was seen in all the samples. At 90 days, an abundance of new patches of golden yellow fluorescence was noticed over all the samples and the percentage of new bone growth was calculated to be 42 ± 3% (pure forsterite), 51 ± 2% (Zn-0.25), 72 ± 3% (Zn-0.5), 48 ± 3% (Sr-1), 80 ± 2% (Sr-2), 75 ± 2% (Sr-3) for pure and doped forsterite samples, respectively. The detailed H&E stained histological micrographs [Fig. 17b (reproduced with permission)] of undoped, Zn- and Sr-doped forsterite showed enormous amount of new bone regeneration with proliferating osteoblast cells. To study the toxicity of the samples, we carried out the histological studies [Fig. 19 (reproduced with permission)] of three major organs, namely, the heart, kidney and liver. The H&E staining of heart depicted normal musculature confirming that forsterite ceramics have no harmful side effects. The kidney and liver structure do not show any major abnormal changes, indicating no toxic effects on all the samples. Thus, our research group, for the first time, evaluated the complete in vivo biocompatibility of forsterite bioceramics and their ability in promoting new bone formation.

Figure 17:
figure 17

a Fluorescence images of bone–implant interface of pure, Zn- and Sr-doped forsterite after 30 and 90 days (scale bar = 500 µm). b Histological H&E stained micrographs of pure, Zn and Sr forsterite implanted bone at 90 days (scale bar = 100 µm), from Devi et al.10, 46.

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Figure 18:
figure 18

Percentage of new bone formation in the defect site of Zn- and Sr-doped MgS with varying composition and time period, from Devi et al.10, 46.

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Table 4 Summary of mechanical and biological properties of forsterite
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Fig. 19
figure 19

Histological analysis of toxicity of forsterite. a Heart, b kidney, c liver after 90 days of implantation using H&E staining (scale bar = 100 µm), from Devi et al.10, 46

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6 Other Magnesium-Based Silicate Ceramics for Biomedical Applications

The widespread application of magnesium-containing silicate ceramics is an emerging area of research in the biomedical field. Apart from forsterite, other magnesium-containing silicate ceramics such as akermanite (Ca2MgSi2O7), bredigite (Ca7Mg(SiO4)4), diopside (CaMgSi2O6), merwinite (Ca3Mg(SiO4)2), monticellite (CaMgSiO4) and proto-enstatite (MgSiO3) have also been reported as potential bioceramics for bone tissue engineering applications84. In particular, akermanite, diopside and merwinite have recently drawn special attention because of their superior in vitro and in vivo biological properties22, 85, 86. The mechanical properties of merwinite and monticellite are similar to that of cortical bone85, 87. The dissoluted products of these materials further increased the osteoblast cell adhesion, proliferation and differentiation85. The incorporation of 2–10 wt% of merwinite nanoparticles in β-tricalcium phosphate (β-TCP) increased the mechanical strength and bioactivity compared to undoped β-TCP88. The in vivo studies in rat femoral defect model proved that the merwinite promoted osteogenesis and the rate of new bone regeneration was much faster compared to control (β-TCP), thereby making the material suitable for bone replacement89. Further, Jin et al. investigated the biocompatibility of enstatite up to 14 days in SBF and showed that enstatite could not induce hydroxyapatite growth. It is reported that the Mg ions are effective inhibitors of apatite nucleation and growth90, 91; however, the degraded products promoted the proliferation of mouse fibroblasts (L929 cell) than on traditional hydroxyapatite ceramics92.

7 Conclusions and Future Directions

The present review summarizes the synthesis methods and mechanical, in vitro and in vivo properties of magnesium silicate degradable bioceramics. With the considerable number of research findings, use of magnesium-based silicate bioceramics for bone tissue applications has been promoted. However, complete understanding of forsterite for orthopaedic application is still in a young stage. Several studies have reported that the degraded Mg and Si ions showed a positive effect on in vitro studies. In spite of promising aspects, there is a knowledge gap of detailed in vivo studies and clinical trials. In-depth research is needed in understanding the mechanism how various parameters influence the healing process, replacing the material with the new bone. The presence of porosity in the material plays an essential role in the choice of bone formation. Therefore, it is very important to study the nature and size of pore which helps in quality and quantity of bone formation.

The current research has been focussed on the development of materials which support osteogenesis and angiogenesis. Because of the advanced technologies in the synthesis methods, the implant materials can be tuned up with the desired porosity and other essential properties that support bone regeneration. The natural bone does not have uniform pore distribution and porosity throughout. Therefore, it is not an essential property of the implant to be uniformly porous with desired pore size and porosity. The distribution of interconnected porous structure can be tailored through manufacturing techniques that will govern both the mechanical strength and interconnected porous structure. Significant studies have been carried on calcium phosphate- and silicate-based ceramics. However, these ceramics do not satisfy the requirements of an ideal implant, mainly as a degradable material. Studies have showed that magnesium-based silicate ceramics had excellent mechanical properties with good cytocompatibility. Nano form of forsterite was highly bioactive with enhanced degradation rate, suggesting the tunable in vitro and in vivo biocompatibility. Further, combining these magnesium-based silicate ceramics with biopolymers and drugs will create a promising path to explore. This review shows that magnesium-based silicate bioceramics are promising materials for the development of various orthopaedic implant applications.