For the past decades, the incidences of bone defects caused by different musculoskeletal conditions (e.g., congenital, degenerative, post-traumatic, neoplastic, metabolic and infectious) are continuously increasing (Katagiri et al., 2017). Bone regeneration or substitution has been proved to be valid approaches for the current therapy (Bhattacharjee et al., 2017; Zhang et al., 2019a). In this context, bone transplantation has become the second frequent tissue transplantation after blood transfusion, with over two million cases worldwide per year (Liu et al., 2017). With the advent of an aging society, the clinical requirement for effective bone regeneration therapy will continue to increase. Despite autogenous bone transplantation has undoubtedly become the gold standard for bone regeneration, the applications are still limited because of the insufficient supply, loss of function as well as secondary defects at the bone donor site (Shukla et al., 2017; Benlidayi et al., 2018; Zhang et al., 2019b). Allogenous bone transplantation is an alternative approach, but this therapeutic method has to face risks such as disease transmission, infection, and immunogenicity (Calori et al., 2011; Campana et al., 2014). Although some existing synthetic biomaterials have achieved favorable clinical efficacy, there is still a lack of outstanding biomaterials that can provide excellent load-bearing, complete biodegradability, osteogenesis and osteoconductivity simultaneously (Fayyazbakhsh et al., 2017; Li et al., 2018a; Ju et al., 2021). As a result, there has been an urgent need for the development and application of synthetic bone graft substitutes.

Tissue engineering has been considered as a viable solution to the aforementioned challenges, which has led to significant advances in cell and organ transplantation over the past decades, as well as greatly stimulating innovation in new materials, application models, preparation techniques, and performance evaluation (Rose and Oreffo, 2002; O'Keefe and Mao, 2011). Tissue engineered bone grafts have great potential to alleviate the need arising from the lack of suitable autograft and allograft materials for bone repair (Yadav et al., 2021). Up to now, a great deal of effort has been made in the design, fabrication, characterization, and application of emerging materials, such as scaffolds, coatings, membranes. (Zwingenberger et al., 2012; Eivazzadeh-Keihan et al., 2019; Du et al., 2021). Among these promising solutions, scaffolds play a central role and exhibit several unique advantages in bone tissue engineering (Shadjou and Hasanzadeh, 2016). Scaffolds can not only provide mechanical support for local load-bearing, but also offer structural support for specific cells, and thus guide new tissue growth and promote bone regeneration (Sill and von Recum, 2008). Coatings for bone repair implants are perceived as another promising approach in bone tissue engineering. Coated materials can enhance the mechanical properties of the implants and improve the interfacial reaction physiologically (Pereira et al., 2020). Membranes for bone tissue engineering can provide an independent space for bone regeneration and act as a barrier against soft tissue ingrowth (Du et al., 2020).

The success of tissue engineering heavily depends on the performance of the materials. For the desired bone regeneration materials, excellent biocompatibility, controlled biodegradability, appropriate mechanical strength, and suitable porosity to support cell differentiation, growth, and proliferation all should be emphasized (Kim et al., 2013; Freeman et al., 2021; Hajiali et al., 2021). Research on graphene-based nanomaterials has boomed in biomaterial applications over the past few years. With outstanding physical, chemical and biological properties, graphene is considered to be revolutionary material and shows great potential for applications in tissue regeneration, drug delivery and other biomedical areas (Hu et al., 2020; Fu et al., 2021; Unal et al., 2021). Therefore, the purpose of this review is to highlight the scientific progress over the years and further summarize the physical and chemical properties, family members and applications in bone tissue engineering of this graphene-based nanomaterial.

Graphene, a single-atom thick and two-dimensional sheet of sp²-hybridized carbon atoms, was isolated from highly oriented pyrolytic graphite by two British physicists (Feng and Liu, 2011). This revolutionary discovery quickly attracted great attention in the fields of materials science, chemistry, physics and biotechnology, and the 2010 Nobel Prize in Physics attests to its extraordinary significance (Dresselhaus and Araujo, 2010). Since its discovery in 2004, graphene has received increasing interest for its remarkable properties, including high fracture strength, outstanding Young’s modulus, excellent thermal and electrical conductivity, large specific surface area, atomic structure stability and biocompatibility (Kumar et al., 2016; Sayyar et al., 2017; Zhao et al., 2017; Nezakati et al., 2018). There have been great expectations for the application of this “future material”.

Graphene is known as one of the strongest materials in existence (Lee et al., 2008). With exceptional mechanical properties, graphene stands out as the most promising candidate to be a major filling agent for bone repair composite. Nevertheless, it should be emphasized that the mechanical reinforcement effect from graphene is closely related to its distribution in the composite (Young et al., 2012). Homogeneous distribution leads to effective mechanical property enhancement, but the cohesion between graphene molecules can hinder the distribution and therefore needs to be overcome (Gao et al., 2018). Electrical conductivity can confer better osteogenic activity to bone repair materials (Huang et al., 2019). Owing to its unique molecular structure, graphene can be used to formulate three-dimensional composites with good electrical conductivity. Moreover, the large specific surface area of graphene can greatly improve cell adhesion, which likewise benefits the osteogenic activity (Gao et al., 2017). Large specific surface area also facilitates the further functionalization of graphene, thus being able to impart better chemical activity and improve its hydrophilicity and dispersibility (Trusek et al., 2020). The molecular size, content and uniformity of graphene will significantly affect the mechanical and electrical properties of the bone repair composite, and it is therefore important to determine the appropriate graphene content and ratio.

Biocompatibility is a prerequisite for the in vivo application of bone repair materials (Mehrali et al., 2014; Qi et al., 2020; Jyoti et al., 2021). A number of studies have demonstrated the biocompatibility of graphene through in vitro cell co-culture and in vivo metabolic analysis, however, this observation is accompanied by qualifying conditions (Alaghmandfard et al., 2021). The physical and chemical properties have been proved to greatly affect the interaction of graphene with living cells, and that the dose and concentration of graphene in the matrix are also related to its cytotoxicity (Cao et al., 2021; Pulingam et al., 2021). The resulting cytotoxic effects occur mainly at the cellular and molecular levels and may be attributed to increased oxidative stress (Sasidharan et al., 2016). Therefore, as with other nanomaterials, smaller particle sizes and higher concentrations are more likely to induce cytotoxic effects, while concentrations below 5–10 μg/ml are relatively safe (Chang et al., 2011). The long-term safety of biomaterials for in vivo applications also depends on their biodegradability. The degradation products from graphene have not been shown to cause substantial cell damage, however, the in vivo retention period may be associated with several pathological changes (Mukherjee et al., 2018; Daneshmandi et al., 2021). The enzymatic environment as well as specific chemical modifications are available to regulate the rate of degradation and thus reduce the potential of graphene-induced cytotoxicity (Palmieri et al., 2019; Ma et al., 2020; Peng et al., 2020).

The growing demand for solving clinical challenges is forcing the studies on tissue engineering to progress (Wang et al., 2021). As one of the most promising emerging materials, graphene has been widely explored in the fields of regenerative medicine (Kolanthai et al., 2018; Liu et al., 2019; Cojocaru et al., 2021). Nevertheless, there are still much room for improvement to overcome existing challenges for its application in bone regeneration (Cernat et al., 2020; Nimbalkar and Kim, 2020; Bruschi et al., 2021). In spite of remarkable properties, it is difficult for pristine graphene to form three-dimensional scaffolds on its own. In general, graphene needs to be compounded with other materials for bone repair (Raslan et al., 2020), but the strong van der Waals forces between particles lead to the poor dispersibility in aqueous media and physiological fluids (Li et al., 2017a). Functionalization of pristine graphene is an effective method to improve the solubility and dispersibility, so that a growing number of studies have been focusing on its derivatives for regenerative applications in recent years (Driscoll et al., 2021).

Graphene oxide (GO) is the oxidized form of graphene, which has been the most widely used graphene family member in biomedical applications today (Reina et al., 2017; Han et al., 2018a; Martín et al., 2019). After oxidation, GO retains its laminar structure. There are various oxygen functional groups distributed on the carbon atom sheet, mainly including hydroxyl and epoxy groups on the basal plane, carboxyl and carbonyl groups attached to the edge (Park et al., 2009). The introduction of functional groups not only offer the hydrophilicity and dispersibility (Chen et al., 2012), but also provide more opportunities to manipulate and customize the properties of GO (Kim et al., 2010). On the other hand, the presence of these functional groups produces high defect density in the perfect planar structure of graphene, which leads to the reduction of its mechanical, electrical, and thermal properties (Zhao et al., 2021). Therefore, GO is considered as an attractive and cost-effective alternative for graphene due to its accessibility, hydrophilicity, dispersibility, chemical tunability and processability.

Reduced graphene oxide (rGO) is obtained through chemical or physical methods to eliminate the oxygen functional groups (Dash et al., 2021). Reduction of GO is done to restore physical properties to some extent, but this also leads to a decrease in hydrophilicity and dispersibility, as well as a weakening of the chemical tunability (Bagri et al., 2010). rGO can be used as an alternative for large-scale production of graphene-based materials. Based on number of layers in the sheet, graphene can be classified into single layer graphene, few-layers graphene and multi-layer graphene (Daneshmandi et al., 2021). Ultimately, the application of graphene and its various functionalized derivatives can be selected according to specific clinical needs in bone regenerative medicine (Zhao et al., 2017). Currently, graphene and its derivatives (GDs) have been used to prepare various scaffolds, coatings, membrane materials and injectable hydrogels for bone tissue engineering by compounding with various matrices such as metals, polymers and inorganic substances (Park et al., 2016; Li et al., 2018b; Saravanan et al., 2018).

It is necessary to have a thorough understanding of the basic structure before discussing bone tissue engineering. In addition to bone cells, bone tissue contains a large number of matrices, mainly including collagen, non-collagenous proteins and calcium phosphate deposits (Katz and Meunier, 1987). Macroscopically, bone tissue can be divided into dense cortical bone and porous cancellous bone or trabecular bone. In general, cortical bone acts as a shell to encase cancellous bone. Tissue regions with higher mechanical stress contain a higher percentage of cortical bone (Rho et al., 1998). It can be assumed from the composition contained in bone tissue, the mineral component provides stiffness while the collagen assemblies provide viscoelasticity and toughness (Kadler et al., 1996; Iyo et al., 2004; Shah et al., 2019). Microscopically, pristine fibers, formed from collagen and minerals, can aggregate and assemble into nanoscale fibers. Subsequently, the nanoscale fibers gathering as lamellar structures that are arranged in cylinders parallel to the long axis in cortical bone and irregularly woven arrays in cancellous bone (Rho et al., 1998). Bone tissue is also a highly dynamic system. Continuous remodeling occurs through osteolysis by osteoclasts and osteogenesis by osteoblasts (Hadjidakis and Androulakis, 2006). The process of bone remodeling can be induced by mechanical stress, with areas of higher mechanical stress producing stronger tissue and higher overturn rates (Herring, 1968).

The ultimate goal of bone tissue engineering is to achieve restoration and reconstruction of human bone tissue to meet clinical treatment requirements, including but not limited to bone defects caused by trauma, infections, sarcomas and metabolic diseases (Holzapfel et al., 2013; Jakob et al., 2013; Fernandez-Yague et al., 2015). Bone regeneration is a complex and dynamic physiological process which macroscopically involves local mechanical stability, environmental matrix, and blood supply, and microscopically involves the interaction of multiple cells, signaling molecules, and effectors in a spatiotemporal sequence (Winkler et al., 2018). Therefore, an ideal bone graft substitute should exhibit the following characteristics: 1) good biocompatibility and non-immunogenicity to ensure safe clinical application; 2) suitable mechanical properties, including strength, viscoelasticity, toughness, and wear resistance to match the properties of host bone and ensure adequate longevity; 3) porous structure or rough surface to facilitate the ingrowth of cells and tissues; 4) favorable osteoconductivity and osteoinductivity to promote new bone production; 5) controlled degradability to match the rate of new bone ingrowth; 6) facile modifiability to meet the specific functional requirements of different application scenarios.

GD’s outstanding properties make it as one of the most anticipated materials for bone tissue engineering (Guo and Dong, 2011; Ding et al., 2015). Mechanical properties including strength, stiffness, and flexibility can best reflect the irreplaceability of GDs in bone repair. The mechanical reinforcement of GDs on the bone repair composites can be adjusted by the selection of different derivatives or by changing their content. GDs-based composites possess favorable viscoelasticity that can be molded as needed to better adapt to the physical characteristics of bone tissue. GDs also has strong structural stability and is less likely to be destroyed during complex preparations as well as in the physiological environment of the implant site. In addition, good electrical conductivity of GDs can not only directly stimulate cellular osteogenic activity, but also indirectly adsorb active factors to promote bone formation (Turk and Deliormanlı, 2017; Dalgic et al., 2018). Conductivity can also be used for signal control or magneto-thermal therapy in special cases, such as focal clearance for osteosarcoma or local infections. Moreover, GDs can promote biomineralization and bone-like apatite formation on the implant surface, thus enhancing osseointegration and osteoconductivity. Planar structure and large specific surface area impart GDs with excellent ability to immobilize various biomolecules, cells, drugs and other desired substances (Wang et al., 2011; Shen et al., 2012; Feng et al., 2013; Yang et al., 2013). On the other hand, such structure makes it easier for GDs to be modified with multifunctional groups, which in turn significantly improve the dispersibility and hydrophilicity of GDs in the composites. Based on the above advantages, GDs have become an indispensable component of bone repair composites in many studies.

Although promising progress has been made in current research of GDs-based materials, there are still many challenges to be faced before clinical application. Firstly, taking into account the cost-effectiveness and accessibility, the production and processing techniques of GDs are yet to be broken through. At present, the broad application of GDs is restricted due to the difficulty of large-scale synthesis. Cost may become a constraint for further research. Furthermore, the problem of GD’s aggregation in solution during the fabrication of composites also remains to be solved (Xu et al., 2016; Syama et al., 2017). How to promote the homogeneous distribution of GDs in the matrix will also be a focus of future research. Secondly, the long-term safety of GDs-based materials for in-vivo use is still unclear. It is imperative to gain insight into the interactions of GDs with biological systems. There have been concerns regarding its biocompatibility and toxicity, but convincing, high-quality studies are still insufficient. The observation period of existing studies is not long enough, so chronic toxicity studies longer than 6 months should be encouraged; on the other hand, the most of adopted animal models were rats and rabbits, lacking large animal models such as pigs, goats and monkeys. What’s more, the mechanism of interaction between GDs and the in vivo environment after implantation needs to be further clarified. Exploration of osteogenic mechanisms may point out the right direction for further study; analysis of cellular uptake and response mechanisms facilitates the prediction of acute and chronic adverse reactions; then the clarification of degradation and metabolism mechanisms will help to understand the spatiotemporal distribution of GDs in vivo. In summary, in spite of various challenges, GDs are likely to be a real breakthrough for future research in regenerative medicine, and advances in related technologies will pave the way for earlier clinical use of GDs.