Engineering vascularized and innervated bone biomaterials for improved skeletal tissue regeneration

Abstract

Blood vessels and nerve fibers are distributed throughout the entirety of skeletal tissue, and play important roles during bone development and fracture healing by supplying oxygen, nutrients, and cells. However, despite the successful development of bone mimetic materials that can replace damaged bone from a structural point of view, most of the available bone biomaterials often do not induce sufficient formation of blood vessels and nerves. In part, this is due to the difficulty of integrating and regulating multiple tissue types within artificial materials, which causes a gap between native skeletal tissue. Therefore, understanding the anatomy and underlying interaction mechanisms of blood vessels and nerve fibers in skeletal tissue is important to develop biomaterials that can recapitulate its complex microenvironment. In this perspective, we highlight the structure and osteogenic functions of the vascular and nervous system in bone, in a coupled manner. In addition, we discuss important design criteria for engineering vascularized, innervated, and neurovascularized bone implant materials, as well as recent advances in the development of such biomaterials. We expect that bone implant materials with neurovascularized networks can more accurately mimic native skeletal tissue and improve the regeneration of bone tissue.

Graphical abstract

1. Introduction

Despite significant progress in the field of bone tissue engineering, the treatment of large bone defects caused by extremity traumas, degenerative pathological conditions, or tumor resection, still remains a challenge. When the size of the defect is small and in a range where self-repair is possible, an endogenous and scarless healing process occurs through the synergistic actions of different types of cells, including mesenchymal stem cells (MSCs), osteogenic cells, and inflammatory cells [1, 2]. As a result, the newly formed bone has similar physicochemical and mechanical properties compared to the surrounding skeletal tissue [3]. However, when the size of the bone defect becomes larger than the critical size, it cannot heal spontaneously, so clinical intervention and artificial bone materials are necessary to repair damaged skeletal tissues. This type of permanent fracture is called a non-union bone defect, and occurs in 5–10% of all fracture cases [4]. The main challenge in treating large non-union bone defect is filling this gap, and allowing for vascularization and innervation to link with surrounding tissues to provide oxygen, nutrients, and cells [5].

Currently, more than two million bone graft surgeries are performed globally every year to fill the large defect sites in bone, with more than five hundred thousand implantations conducted in the United States alone [6]. In fact, bone implants are the second most frequent tissue transplanted worldwide, after blood transfusion [7]. Bone substitutes are defined as “synthetic, inorganic or biologically organic combinations, which can be inserted for the treatment of a bone defect instead of autogenous or allogenous bone” [8]. To replicate the function of bone tissue, an ideal bone implant material should provide an osteoconductive matrix for inducing vascularization, as well as supplying osteoinductive factors to direct the osteogenic differentiation of cells [9]. In the past half century, various types of bone materials in different forms, ranging from powders, cements, 3-dimensional scaffolds, and injectable composite hydrogels have been developed to improve bone healing [10, 11], with several bone implants having been approved by the US Food and Drug Administration as medical devices. Many of these bone grafts are composed of bone-mimetic ceramic components such as hydroxyapatite (HAP: Ca₁₀(PO₄)₆(OH)₂), beta-tricalcium phosphate (P-TCP: Ca₃(PO₄)₂), calcium phosphate cements (CPC), combined (or not) with natural or synthetic polymers, such as collagen, gelatin, hyaluronic acid, poly(lactic-co-glycolic acid) (PLGA), and poly(glycolic acid) (PGA) [12]. While these biomaterials can successfully recapitulate the macroscopic structural features and mechanical properties of bone tissues, current bone biomaterials are still not sufficient for generating vascularization or innervation, and often exhibit impaired healing. This limitation exists because these materials are restricted to a single tissue type, and mainly focus on inorganic bone minerals. As such, they ignore other important tissues in the skeletal organ system, such as vascular and nerve tissues.

Indeed, human bone is fully vascularized and innervated, where blood vessels and nerve fibers closely interact and enhance the development and function of each other, eventually supporting bone development and fracture healing [13–15]. Blood vessels carry cells to the bone surface, and also supply oxygen, nutrients, and hormones [16]. The nervous system in bone tissue also influences the function of bone cells, and actively participates in bone resorption/formation [17–21]. Notably, nerves enter bone alongside the arteries and remain closely associated with blood vessels inside bone, sharing common genetic pathways and stimuli [22–24]. In this respect, the lack of attention to the vascular and nervous networks within bone implant materials could be one of the main reasons for the delayed or impaired recovery of bone fractures.

For these reasons, the majority of clinical practice today is still based on the use of autografts, in which bone tissue is harvested from one part and subsequently grafted to another position in the same patient. Autografts are considered as the gold standard, since they can preserve blood vessels and nerve fibers, and can provide the proper osteogenic signals to guide new bone growth [25, 26]. Autografts also possess a cocktail of growth factors and cells necessary to recapitulate the bone developmental process [25]. However, there are drawbacks in its use, such as the risk of donor site pain, increased blood loss, donor site infection, and the limited amount of graft material available from a single patient [27–29].

To overcome the drawbacks of using autografts and to develop an ideal bone implant material, it is essential to more accurately recapitulate the microenvironment of bone by considering blood vessels and nerve fibers. It is expected that a neurovascularized bone graft will generate more efficient and effective bone regeneration by interconnecting surrounding tissues through neurovascular networks, and by providing a biomimetic niche for osteogenic cells. To develop such a neurovascularized bone implant material, a more comprehensive understanding of the anatomy and interaction mechanisms of blood vessels, nerve fibers, and bone tissue is required. Here, we discuss the anatomy of blood vessels and the peripheral nervous system (PNS) in bone, and their biological mechanisms for enhancing bone formation. Based on this biological background, we introduce a design criteria for developing vascularized and innervated bone graft materials, and the recent progress for fabricating such biomaterials. Lastly, we discuss the developmental mechanisms of coupled neurovascular networks in bone, and introduce current and potential strategies for engineering neurovascularized bone implant materials.

2. Engineering vascularized bone implant materials

2.1 Understanding development mechanism of coupled vascularization and osteogenesis in bone

The vascular system is pivotal in maintaining healthy bone tissue, since it continuously provides oxygen, nutrients and osteoprogenitor cells that are necessary for skeletal growth and remodelling [30]. When damage occurs in bone, adult skeletal tissue regenerates in a similar way to the growth of embryonic bone tissue, where blood vessels play an important role in supporting the growth and activity of osteogenic cells [31]. On the other hand, alterations to bone vasculature can reduce bone growth and lead to skeletal diseases [32–35]. In this section, we describe the blood vessel anatomy in bone and the role of vascularization in embryos and adult bone healing, with particular attention to the vasculogenic factors that are directly involved in osteogenesis.

2.1.1 Blood vascular system in bone

Long bone can be divided into three parts based on their anatomical shape and composition: tubular masts that extend along the ends of the bone (diaphysis), wider section extremities (epiphysis), and the cone-shaped metaphysis between the epiphysis and the diaphysis. Diaphysis are mainly composed of cortical bone, whereas metaphysis and epiphysis are mostly composed of trabecular bone, and are covered by a shell of cortical bone. The external surface of bone is covered by periosteum (peri– = “around” or “surrounding”), while the inner layer is covered by endosteum (end- = “inside”; oste- = “bone”). The periosteum is a connective tissue that contains a high density of blood vessels, nerve fibers, osteoblasts and osteoclasts. The inner surface of bone and the vascular canals are covered by a membranous structure of endosteum, which contains a protective lining of cells that covers the bone surface. The endosteum also contains connective tissue fibers and blood vessels that can supply nutrients [36].

Adult long bone includes four different types of vascular networks: diaphyseal, periosteal, metaphyseal and epiphyseal [37]. After the arteries enter bone through various foramina from the surrounding tissues, these blood vessels penetrate through Volkmann’s and Haversian canals and branch throughout the cortical bone (Fig. 1) [37, 38]. Among these vessels, the diaphyseal artery is the largest vessel, and supplies more than 50% of total blood that enters the long bones [39, 40]. This diaphyseal artery enters the diaphysis obliquely through the nutrient foramina of bone, and splits into branches along the medullary cavity. These branches reach the epiphysis and divide into smaller ramifications. Metaphyseal arteries enter the long bones along metaphysis, while the epiphyseal arteries compose the periarticular vascular arcades [41, 42]. Venous blood is drained through the venous central sinus, and exits the bone through veins of the cortical bone collar [42]. This vascular system provides nutrients, oxygen, and osteoprogenitor cells to the terminal regions of the skeletal tissue [43].

Figure 1.

The vascular system in bone. (a) Schematic diagram of hierarchically structured blood vessels in bone. Blood supply is derived from medullary arteries, which are connected to periosteal arterial blood vessels. Medullary blood then exists via veins that penetrate the bone cortex. (b) Light micrograph image showing blood vessels penetrating cortical human bone (image width: 1.5 mm). (c) Transverse sectional view of human bone showing blood vessel penetration at the center of Haversian canals. (d) MicroCT image of mouse tibia showing the medullary vessels using a radio-opaque contrast agent (image width: 3.5 mm). (e) 3D rendered image of the blood vessel network of the same mouse tibia from panel (d), after a decalcification process (Images reproduced with permission from [38]).

2.1.2 Role of vascularization on fetal bone development

The process of vascularization is essential to induce osteogenesis during embryonic bone formation and bone fracture repair. There are two different mechanisms in which the skeleton is formed during embryonic development: intramembranous ossification and endochondral ossification. Intramembranous ossification is an essential process that occurs during the fetal development of flat bones such as facial bones, cranium flat bones, mandible and clavicle bones [44], while endochondral ossification occurs during the development of long bones [45].

During intramembranous ossification, blood vessel capillaries invade a mesenchymal zone, inducing MSCs to differentiate towards osteoblasts [46]. The continuous bone matrix deposition leads to the formation of bone spicules, which grow into trabeculae, become woven bone, and are then replaced by lamellar bone [47].

On the other hand, long bones are developed through endochondral formation, in which hypertrophic cartilage is remodeled into bone tissue. Endochondral ossification occurs on a cartilaginous template as an intermediate step, which then develops into skeletal tissue through blood vessel invasion [48]. In particular, vasculature grows exponentially and recruits osteogenic and vasculogenic cells, while long bone expands both radially and longitudinally (Fig. 2 [49]). By this time, blood vessel capillaries become visible as a column-like structure in the metaphysis, and as a dense sinusoidal network in the diaphysis.

Figure 2.

Developmental angiogenesis in murine femur bone. (a) Confocal images showing blood vessel formation during the early stages of bone development, based on endomucin (Emcn)-stained endothelial cells. E: embryonic day, P: postnatal day. mp: metaphysis, dp: diaphysis. (b) Osterix immunostaining images illustrating skeletal development (Image reproduced with permission from [49]).

2.1.3 Vascularization role in adult bone healing

Vascularization is also essential in adult bone repair, where angiogenesis and osteogene sis occur in a coupled manner that resembles the developmental mechanisms of embryonic bone tissue [50]. After a fracture, an inflammatory response activates growth factors and cytokines that recruit osteoprogenitor cells to the damaged site. The initial clot is a temporary matrix for the invasion of vascular networks, and is replaced by fibrocartilage and remodelled into bone [51, 52]. During bone remodelling, osteoclasts remove necrotic bone tissue, while angiogenesis actively occurs to restore normal blood circulation. Based on this bone repairing mechanism, regenerated skeletal tissue has physicochemical properties that are comparable to the original tissues, with minimal formation of a fibrous scar. In particular, fibrous scar formation can affect the mechanical properties of the newly repaired tissue. As well, when the musculoskeletal system is damaged and vasculature is disrupted, acute necrosis and hypoxia of the surrounding tissue can occur [53].

Coupled angiogenesis and osteogenesis are induced based on signalling mechanisms that involve various types of cells, including MSCs, endothelial cells, osteoprogenitor cells, osteoblasts, and osteoclasts that cooperate for bone healing [50]. For example, MSCs and adipose-derived stromal cells (ASCs) interact with endothelial cells by secreting growth factors, as well as through physical conjunctions [54, 55]. In particular, connexin 43 (C×43) is a gap junction protein expressed by stem cells and endothelial cells, which plays an important role in cell signalling and the functional maintenance of bone tissue [56, 57]. A lack of C×43 expression affects the process of bone healing by reducing the angiogenic potential of progenitor endothelial cells [56, 58]. In particular, it was shown that a lack of C×43 in osteoblasts and osteocytes in mice affected bone formation and the remodeling processes, resulting in reduced bone density and cortical thickness [59].

Intercellular communication during coupled osteogenesis and angiogenesis also occur based on osteogenic growth factors and angiogenic growth factors produced by different kinds of cells. For example, MSCs secrete vasculogenic growth factors such as vascular endothelial growth factor (VEGF), platelet-derived growth factor (PDGF), fibroblast growth factor (FGF), insulin-like growth factor-1 (IGF-1), and angiopoietin-1 (Ang-1), which stimulate the migration, proliferation and differentiation of vascular cells in neighbouring tissues [60, 61].

Among various vasculogenic growth factors, VEGF is the principal mediator of angiogenesis that binds to endothelial cells through the vascular endothelial growth factor receptor 1 (VEGFR1 (Flt-1)) and VEGFR2 (KDR/Flk-1) [62]. Notably, VEGF plays a dual role by signalling to endothelial cells to induce vascularization, while also binding to osteoblast receptors to promote bone repair [63, 64]. Interestingly, secretion of these vasculogenic growth factors by MSCs is proportional to their exposure to bone morphogenetic protein-2 (BMP-2) [65]. Therefore, BMP-2 also has a dual role in bone healing by directly inducing osteoblast differentiation [66] and promoting angiogenesis by inducing VEGF secretion from osteoblasts [67]. Based on these observations, one of the developmental mechanisms of coupled osteogenesis and angiogenesis is the cellular crosstalk between MSCs and endothelial progenitor cells via secreted proteins such as VEGF or BMP-2 [65].

2.2 Material properties affecting vascularized bone regeneration

Engineered bone materials should have sufficient levels of mechanical stability to maintain their physical shape, and to support tissue regeneration at the implanted site until they are replaced by native tissue [68]. At the same time, developing blood vessel networks in bone materials is important for inducing mass transfer to cells embedded in engineered tissues. In the body, blood vessels are distributed in every distance within 100–300 μm to supply sufficient levels of oxygen and nutrients, as well as to remove metabolic wastes from cells [69–72]. Indeed, when the distance between cells and capillary vessels increases above this distance, the diffusion of nutrients and oxygen becomes impaired, and cell viability and proliferation decreases. In this respect, insufficient blood circulation in the implanted bone material results in a lack of integration with the host body, and inner graft necrosis. Therefore, bone implant designs should consider material properties that are effective for inducing vascularization (Fig. 3a). Since material properties for developing bone tissues and vascular tissues are different, it is important to find an optimal property of biomaterials that can simultaneously induce osteogenesis and vascularization [73].

Figure 3.

Schematic illustration of material properties that influence osteogenic differentiation and vascularization. (a) Osteogenic cells (blue) and vasculogenic cells (red) residing in the vascularized bone niche. (b) Material stiffness influences cellular adhesion, shape, and differentiation behaviour, based on mechanosensing of cells. (c) Materials that have nanoscale roughness can promote the attachment and differentiation of osteogenic and vasculogenic cells. (d) Generating an optimal level of porosity in bone materials is important to induce blood vessel growth, and maintain high mechanical strength of biomaterials.

2.2.1. Stiffness of vascularized bone materials

While most cells are attached to the ECM, different ECM types exhibit different levels of elastic moduli. Cells can sense these differences in mechanical properties of matrix materials and respond by altering their adhesion shape and differentiating into different cell lineages (Fig. 3b). This mechanosensing by cells occurs, at least in part, by a change in integrin/adhesion-ligand binding on matrix materials [74]. In the case of MSCs, their osteogenic differentiation occurs optimally when the rigidity of the matrix material is around 25–40 kPa [74, 75]. For example, when MSCs are cultured in decellularized cancellous bone scaffolds coated with a collagen and HAP mixture, their osteogenic differentiation was enhanced as the stiffness of the scaffolds increased from 6.7 kPa to 23.6 kPa [76]. Interestingly, when these scaffolds were implanted in rat subcutaneous tissues, an increase in vascularization (with higher expression of CD34) was observed as the stiffness of bone materials increased. In this respect, material stiffness influences not only bone formation, but also vascular development in a coupled manner. Indeed, vascular cells such as endothelial cells also show stiffness-dependent spreading behaviour [77]. For example, while endothelial cells exhibit a round morphology on soft matrix materials (~180 Pa), their adhesion area increases as the stiffness of material increases (~28.6 kPa). In addition, while actin acts as a major stress bearing component during cytoskeleton tension, actin stress fibers of endothelial cells are only developed when these cells are cultured on material surfaces with a stiffness greater than 2 kPa [78, 79].

2.2.2. Roughness of vascularized bone materials

Surface properties such as the topography of biomaterials are an important factor in regulating cellular growth and differentiation, since the material surface is the earliest point of contact of the cells to the scaffold materials [80]. In fact, natural bone and blood vessel walls have a nanostructured surface, as they are composed of nanometer-sized ECM proteins [81, 82]. Nanostructured surfaces have increased surface energy compared to a flat surface, which improves the hydrophilicity of the material and enhances the adhesion of proteins, further promoting the attachment of cells [83]. In addition, even small changes in the surface roughness of biomaterials at 10–100 nm scales can induce significant increases in surface energy [84]. In the case of titanium (Ti), a 0.7% change in its surface area can induce a 10% increase in its surface energy [85]. While Ti-based materials are widely used for orthopaedic and vascular implants due to their excellent mechanical properties, bare Ti surfaces have a poor level of cellular adhesion or tissue integration, often showing limited bioactivity [83, 85]. To overcome this problem, the surface of the Ti material can be modified to have nanostructured properties and mimic the natural tissue surface, by using various nanotechnologies such as anodization and electron beam evaporation. In fact, the spreading and adhesion density of both osteoblasts and endothelial cells can be significantly promoted when these cells are cultured on nanoscale Ti surfaces, compared to flat Ti surfaces [85]. Similarly, an increase in surface roughness of polymer substrates, such as polyurethane or poly(lactic-co-glycolic-acid), can enhance the attachment and growth of vascular cells including endothelial cells and blood vessel smooth muscle cells (Fig. 3c) [84, 86].

2.2.3. Porosity of vascularized bone materials

Bone materials with a high porosity can diffuse nutrients and metabolic wastes, and can facilitate the ingrowth of blood vessels by allowing cell migration [87, 88]. However, the mechanical strength of bone materials can become compromised if the porosity of the bone material is too high, which can be inappropriate for implantation at a load-bearing site [87]. Therefore, designing an optimal structure of porous bone materials is important (Fig. 3d). Pore size can also influence the formation of new bone and blood vessel tissue. The minimal pore size that can induce significant ingrowth of natural bone tissue is approximately 75–100 μm [89]. In addition, when the pore size is sufficiently large (>200 μm), bone ingrowth is improved based on the enhanced penetration of osteoblasts [89, 90]. On the other hand, smaller pores (<100 μm) can induce the formation of unmineralized osteoid tissue or fibrous connective tissue [89]. In addition, when different sized pores were generated in calcium phosphate-based bone implants, the density of the functional blood vessel capillaries was higher in bone scaffolds that had pore sizes larger than 140 μm. In part, this can be explained by a faster initial penetration of blood vessels in bone ceramic materials. Even for soft materials such as porous poly(ethylene glycol) (PEG) hydrogels, the invasion of cells and blood vessels are limited to the surface of materials when their pore size was below 50 μm, whereas PEG scaffolds with larger pores (>50 μm) allowed for the formation of mature vascularized tissue throughout the entire hydrogel material after implantation in vivo. While a larger pore size is generally advantageous for vascularization, there seems to be no difference in vascularization for pore sizes above 400 μm [91, 92]. Furthermore, when the pore size becomes excessively large, the cell-to-cell contact ratio decreases, and cells will recognize their surfaces as 2D substrates rather than 3D porous structures.

2.3 Current vascularized bone implant types and status

Vascularized bone implant materials that can anastomose with host tissues after implantation are expected to increase the long term survival rate of implanted materials by supplying nutrients and oxygen, which can promote the growth and osteogenic activity of cells [93]. The diffusion distance of oxygen and nutrients in large-sized engineered scaffolds are limited to 100–200 μm from the interface region, which often results in cell death and tissue necrosis at the center of the scaffold [94]. To overcome this critical limitation in diffusion by inducing vascularization, simple porous structured bone implant materials were fabricated based on methods such as: sacrificial template method, replication method using porous templates, and direct forming method using bubbles [95]. While these previous material designs have passively assisted vascularization from a structural perspective, nowadays, more advanced strategies have been developed that can actively accelerate vascularization in bone graft materials. These strategies include incorporating vasculogenic substances, such as: endothelial cells, vascular growth factors or angiogenic gene carrying vectors. In this section, we discuss recent representative strategies for inducing coupled vascularization and osteogenesis in bone implant materials. We first introduce engineering methodologies for inducing a peripheral vascular system in bone materials. We then highlight novel strategies for generating hierarchical blood vessel networks by incorporating large blood vessels throughout bone scaffolds.

2.3.1 Development of bone scaffold materials with peripheral vascular system

Vascular cells that can produce vasculogenic growth factors, such as endothelial cells, blood vessel smooth muscle cells, and other types of cells, can be directly injected into the defect site of bone to induce coupled angiogenesis and bone regeneration. In addition, delivering vasculogenic cytokines, (such as VEGF), or cocktails of growth factors (such as platelet-rich plasma (PRP) that includes VEGF, TGF, and PDGF), to the bone defect site can enhance alkaline phosphatase (ALP) activity, nodule formation of human osteoblasts, angiogenesis, and neo-vascularization [64, 96, 97]. While direct cell injection and in situ release of growth factors can significantly enhance vascularization and osteogenesis, these approaches are limited in their low targeting efficiency. Furthermore, uncontrolled release of vasculogenic growth factors can induce the formation of hemangioma, leakage of blood vessels, and growth of unstable vessels that resembles the chaotic structure of tumour vasculature [98–100].

To overcome the limitations of direct cell or growth factor injection methods, vascular cells and vasculogenic growth factors can be encapsulated into hydrogel-based bone scaffolds to enhance their delivery efficiency. Hydrogel biomaterials have the advantage of simple control over their physicochemical properties, mechanical properties, degradability, and cell adhesion, to promote cellular proliferation and differentiation [101]. In addition, hydrogel biomaterials can serve as a stabilizing matrix for vascularization, as it recapitulates the extracellular matrix (ECM) of human tissues and enables 3D migration of endothelial cells to form vascular structures [101, 102]. These hydrogel-based bone scaffolds can be used to develop peripheral vascular systems in vitro before implantation in vivo, to induce rapid anastomosis with host blood vessels and promote bone tissue healing.

Fuchs et al. co-cultured endothelial cells and osteoblasts on starch and poly(caprolactone) (SPCL) based 3D fiber meshes, since SPCL can increase the proliferation and differentiation of osteogenic cells [103, 104]. These composite fiber meshes were then embedded into Matrigel hydrogels to support the development of microvessel-like structures in vitro to induce fast anastomosis with the host vessels in vivo. In a mouse subcutaneous model, cellular proliferation and vessel number was enhanced when endothelial cells and osteoblasts were co-cultured compared with the case of monocultured endothelial cells, due to the close interaction between endothelial cells and osteoblasts [105]. In addition, these engineered microvessel-like structures in SPCL scaffold materials were functionally anastomosed with the host blood vessels and induced blood supply, by presenting erythrocytes in its luminal structure.

Since hypoxic conditions can enhance the expression of angiogenic factors (e.g. VEGF) and improve coupled vascularization and osteogenesis, Fan et al. induced vascularized bone tissue by treating bone marrow stromal cells (BMSCs) with the hypoxia mimicking agent cobalt chloride (CoCl₂) [106, 107]. CoCl₂ can upregulate VEGF by activating the hypoxia inducible factor-1 (HIF- 1) signalling pathway [108]. The authors prepared two types of BMSCs: osteogenic differentiated BMSCs and CoCl₂-treated BMSCs that had upregulated angiogenic gene expression. These cells were encapsulated in collagen scaffolds and implanted into an ectopic subcutaneous and orthotopic skull defect. From the von Willebrand Factor (vWF) staining analysis, the density of the blood vessels was significantly increased compared to implants that only contained osteogenic BMSCs or BMSCs. In addition, from the image analysis based on the 3D reconstruction of mineralized region after micro-CT scanning, the area of the newly formed bone tissue was significantly higher within the implant that encapsulated both BMSCs and CoCl₂-treated BMSCs than other implants with osteogenic BMSCs or BMSCs.

Although the vascular peripheral system in bone grafts can significantly improve osteogenesis and tissue integration, the formation of capillary vessel structures in bone grafts requires extensive in vitro tissue cultures before this implant can be used in the clinic. In this respect, there is a need to develop microchannel-structured bone implants that can allow for blood perfusion throughout the construct, and then into capillary vessels to mimic hierarchically structured blood vessels.

2.3.2 Development of bone scaffold materials with hierarchical vascular system

To develop biomimetic blood vascular structures and enable efficient blood flow throughout the bone implant in vivo, native blood vessels or hierarchically-structured blood vessels can be incorporated into bone scaffold materials.

To accelerate vascularization in bone graft materials, Boos et al. surgically created an arteriovenous loop in a sheep and integrated it into a bone substitute, which was then placed in a perforated chamber [109]. They demonstrated improved vascularization using a porous chamber instead of a closed one, because the arteriovenous loop and the capillaries sprouted from this loop can anastomose with extrinsic blood vessels ingrown through pores, and can recruit a higher number of bone cells. As a result, the newly formed bone area was enhanced in the vascularized bone graft that was placed in a perforated chamber, compared to bone grafts that were placed in a nonporous chamber. When BMP-2 was incorporated into this bone substitute with an arteriovenous loop, the number of vessels, length of vessels, and amount of bone area were further increased, based on histology analysis and micro-CT scans [109].

Recently, bioprinting has emerged as a powerful tool to develop cell-laden tissue constructs with any size and shape. Kang et al. used 3D bioprinting to develop calvarial bone scaffolds with highly interconnected microchannels to support the diffusion of nutrients and oxygen for enhancing both vascularization and bone regeneration [110]. The tissue construct was printed with two types of bioinks: a cell-laden bioink for tissue regeneration, and a poly(caprolactone) (PCL) polymer based ink to provide mechanical structural support (Fig. 4a–b). Human amniotic fluid-derived stem cells (AFSCs) were encapsulated in a hydrogel-based bioink composed of gelatin, hyaluronic acid, fibrinogen, and glycerol. After printing, the viability (>90%) and proliferation rate of cells were high, confirming the biocompatibility of the printing process. After implantation in a rat calvarial bone defect, these bioprinted tissue constructs regenerated vascularized bone tissue, with large blood vessel formation throughout the whole implants as visualized by vWF immunostaining (Fig. 4c). On the other hand, control groups without any treatment, or implantation of the scaffold without cells, induced fibrotic tissue formation in which a minimal amount of bone formation was restricted to the boundary of the implant.

Figure 4.

(a) An integrated tissue-organ printer used to develop microchannels for inducing bone vascularization. (b) A 3D architecture with basic patterning composed of multiple types of cell-laden hydrogels and PCL polymers. (c) (i) Motion program used for printing 3D scaffolds as a calvarial bone substitute. Green colors represent the PCL/TCP dispensing paths, while red colors indicate the cell-laden hydrogel paths. (ii) Photograph (bottom) and SEM image (top) of the printed calvarial bone scaffold. (iii) Images of the printed scaffold at day zero (top) and 5 months after implantation (bottom). (iv–vi) H&E staining representing bone formation of non-treated (iv), scaffolds without cells (v) and hAFSCs-printed constructs (vi) 5 months after implantation. (vii–Modified tetrachrome staining of non-treated (vii), scaffolds without cells (viii) and hAFSCs-printed constructs (ix): red represents mature bone, and blue represents osteoid and lining of lacunae. (x–xii) vWF immunostaining of non-treated (x), scaffolds without cells (xi) and hAFSCs-printed constructs (xii): red parts show blood vessels, NB indicates new bone, while PCL/TCP indicates remaining scaffold (Images reproduced with permission from [110]).

3. Engineering innervated bone implant materials

3.1 Understanding development mechanism of coupled innervation and osteogenesis in bone

Human bone is innervated with nerve fibers that are connected to the dorsal root ganglion (DRG) and central nerve system (CNS) (Fig. 5a) [111, 112]. The PNS is composed of a network of motor (signals from CNS) and sensory (signals to CNS) nerves that connect the CNS to the entire body, transmitting signals generated from internal or external stimuli to CNS via propagation of action potentials. The PNS in bone is also directly involved in osteogenesis through secretion of neuropeptides, which can stimulate the neuropeptide receptors of bone cells to modulate osteogenic differentiation. Therefore, the normal functioning of PNS in bone is essential for maintaining skeletal homeostasis and fracture healing [113]. It is reported that patients with insufficient peripheral innervation have delayed fracture healing or nonunion, and suffer from the recurrence of fractures [114–117]. For example, damage to proprioceptive receptors during bone fracture often results in the failure to unite [118]. This denervation process is known to decrease bone growth by reducing the activity of osteoblasts, while increasing the number of osteoclasts and their resorption activity [119–121]. Therefore, understanding the structure and function of the nervous system in skeletal tissue is important to maintain the healthy state of bone tissue, and to develop novel innervated bone materials for treating bone diseases and defects. In this section, we discuss the nerve anatomy in bone and major neuropeptides that regulate osteogenesis in bone.

Figure 5.

Nervous system in bone. (a) The human peripheral nervous network in bone is connected to the central nervous system (CNS). Peripheral nervous system (PNS) in bone tissue transmits signals from external stimuli to CNS through the spinal cord (Image reproduced with permission from [112]). (b) Schematic image showing distribution of sensory neurons that innervate bone with different receptor types of sensory neurons.

3.1.1 Nerve anatomy in bone

The cell bodies of afferent nerves are generally positioned in the DRG, which are connected to the CNS through the dorsal spinal roots or the cranial nerves. The sensory neurons terminate at the bone tissue, either in free nerve endings or in encapsulated endings (Fig. 5b) [122]. Free nerve endings are unmyelinated, branched throughout tissues, and can recognize and carry signals related to pain, temperature, and mechanical stimuli. On the other hand, encapsulated nerve endings can detect low frequency vibrations, and are enveloped in non-neural fibrous connective tissues that separate a single afferent nerve ending in gelatinous material.

Nerve fibers exist throughout the whole bone, including the periosteum, bone marrow, and mineralized parts of bone. These nerve fibers are predominantly located at metabolically active regions of bone, such as the epiphysis and metaphysis [123]. The total number of sensory fibers and sympathetic fibers is the highest in bone marrow, followed by the mineralized part of the bone and periosteum (Table 1) [123]. However, since the periosteum consists of fibrous connective tissue with approximately 30–60 μm of thickness that envelops the surface of mineralized bone, the density of nerve fibers is the highest in this region. Nerve fibers in bone marrow originate from the periosteum and the nutrient foramen, and then branch to cover the whole bone marrow. While most nerve endings are arranged around the arteries and capillary blood vessels, a few are scattered in the bone marrow [124].

Table 1.

Density and total number of nerve fibers at mineralized bone, bone marrow, and periosteum [123].

	Mineralized bone			Bone marrow			Periosteum
	proximal head	diaphysis	distal head	proximal head	diaphysis	distal head	proximal head	diaphysis	distal head
Innervation density (fibers/mm2)	20.3	15.9	2.1	21.4	10.4	9.9	117.0	73.1	177.4
Total number of nerve cells (mm3)	2.86	4.62	3.37	4.97	10.27	4.62	0.06	0.09	0.07

Nerve fibers also exist throughout the rigid, mineralized parts of bone, including cortical bone and cancellous bone [36, 125, 126]. Cortical bone is composed of osteons, which are cylindrically shaped structures composed of multiple types of cells and bone matrix proteins, and are directed parallel to the longitudinal axis of bone. The hollow tubes that exist at the center of the osteon are the Harversian canals, which contain blood vessels and lymphatic ducts [127, 128]. In addition, perpendicularly directed channels called Volkmann’s canals interconnect with the Harversian canals. Bone can be innervated with nerve fibers that penetrate through these Harversian canals and Volkmann’s canals, especially in association with blood vessels [129, 130]. The cancellous bone has a porous structure with a higher surface area to mass ratio compared to cortical bone, and thus has a softer and more flexible structure compared to cortical bone. This cancellous bone is innervated by nerve fibers that are densely distributed in the bone marrow, where these fibers penetrate through the porous structure of the cancellous bone along with blood vessels [123, 131–133].

Based on immunostaining image analysis which labels CGRP in thinly myelinated sensory fibers or unmyelinated sensory fibers, the density of CGRP-expressing nerve fibers in periosteum is the greatest compared to bone marrow and mineralized bone [134]. From immunostaining image analysis which labels neurofilament H (RT-97) structural proteins in myelinated sensory fibers, the periosteum has the highest density of RT-97 expressing nerve fibers, followed by bone marrow and mineralized bone [123].

Although the periosteum has the highest density of nerve networks, the total number of nerve cells in bone marrow or mineralized bone is greater than in the periosteum (Table 1). In all sites, bone marrow has the highest number of nerve cells, followed by the mineralized part of bone, and then the periosteum [123].

3.1.2 Neuropeptides for regulating bone cells

In human bone, neuropeptides are synthesized in the dorsal root, or localized in sympathetic/parasympathetic ganglia. These neuropeptides are then delivered by nerve fibers to bone by axonal transport [28, 135], to induce osteogenesis and neurogenesis during fracture healing and bone remodelling [113]. These nerve fibers are spatially closely associated with bone cells, while bone cells express receptors for these neuropeptides. In fact, the expression of neuropeptides is concurrent with mineralization during the ontogeny of nerves in the limbs of rodent models [136, 137]. Here, we introduce some of the most representative neuropeptides that are involved in bone formation, remodelling, and fracture healing, including: CGRP, vasoactive intestinal polypeptide (VIP), substance P (SP), and catecholamine.

Calcitonin Gene Related Peptide (CGRP)

CGRP is one of the calcitonin family peptides that is synthesized by both central and peripheral neurons [135]. In bone, CGRPs exist in the periosteum, epiphyseal trabecular bone, and bone marrow. It signals via the complex calcitonin receptor proteins, including receptor activitymodifying protein (RAMP-1) and calcitonin receptor-like receptor (CRLR), on the surface of cells [138]. CGRPs improve bone formation by enhancing the proliferation and osteogenic differentiation of MSCs and promoting cyclic adenosine monophosphate (cAMP) production in osteoblasts [139–144], while directly inhibiting osteoclast resorption activity and decreasing mineral dissolution [145–147].

Vasoactive Intestinal polypeptide (VIP)

VIP is a neuronal polypeptide that is expressed by parasympathetic nerve fibers and ganglionic sympathetic nerve fibers, and are localized to the epiphysis and periosteum [148, 149]. It is one of VIP/secretin/glucagon family neuronal polypeptides that are structurally related, and has a highly conserved N terminal end [150]. VIP stimulates VIP-binding functional receptors of osteoblasts and promotes cAMP production. VIPs also downregulate the production of osteoclasts by increasing the expression of osteoprotegerin (a receptor activator of nuclear factor κB (RANK) antagonist) in osteoblasts, thereby decreasing expression of RANK in osteoclasts [151].

Substance P (SP)

SP is one of nociceptive signaling molecules that are associated with sensory nerves in bone [152]. SP increases cAMP production and enhances the differentiation of osteoblasts, promoting bone formation [153]. Furthermore, the osteoclast-expressed receptor, neurokinin-1 (NK1-receptor) stimulates bone resorption activity when osteoclasts are activated by SP [153].

Catecholamine

Catecholamines are a family of neurotransmitters produced by postganglionic fibers of the sympathetic nervous system. The majority of these neurotransmitters are associated with blood vessels in bone marrow, with some being distributed as free nerve endings in periosteum [5]. Catecholamine hormones such as epinephrine and norepinephrine activate adrenergic receptors, increase the production of cAMP, prostaglandin E2 (PGE2), and enhance the proliferation of osteoblasts [154–157]

3.2 Material properties affecting innervated bone regeneration

In this section, we highlight the material properties of the bone matrix or substrate (e.g. stiffness, roughness, and porosity) and their impact on the proliferation and neuronal differentiation of stem cells (Fig. 6a).

Figure 6.

Material properties that influence growth and differentiation of osteogenic cells and neurogenic cells. (a) Schematic illustration of innervated bone niche. (b) Nerve scaffolds and bone scaffolds require different stiffness levels to mimic native tissues. (c) Neurogenic differentiation of MSCs occurred in soft gel scaffolds (< ~1kPa) as determined by the expression of ??3 tubulin neuronal cytoskeletal marker. On the other hand, osteogenic differentiation of MSCs was induced when cells were cultured in stiff gel scaffolds (> ~25 kPa) as determined by expression of CBFa1 osteoblast transcription factor (arrow). Scale bar: 5μm. (d) From the immunostaining images and Western blotting images, neuronal markers were only expressed when MSCs were cultured in soft gel materials (< ~1 kPa). (e) Secretion of osteogenic proteins by MSCs was enhanced when cells were cultured in relatively stiff gel materials (> ~25 kPa). (Images reproduced with the permission of [74, 75]). (f) Roughness change at the nanoscale can significantly decrease the adhesion level of neuroblastoma cells. Left: AFM images of gold substrate with flat surfaces (R_a=0.46 nm) and nanorough surfaces (R_a=99.8 nm). R_a is the mean surface roughness. Right: Neuroblastoma cells grown on flat surfaces had improved adhesion and spreading morphology compared to nanorough surfaces. (Image reproduced with the permission of [161]). (g) Porosity is another critical material property affecting the adhesion behaviour of neuronal cells. Left: AFM images of macroporous and mesoporous silicon substrates. Middle and right: From immunostaining images and SEM images, neuroblastoma cells grown on mesoporous substrates showed better spreading morphology than macroporous substrates. (Image reproduced with the permission of [166]).

3.2.1. Stiffness of innervated bone materials

While osteogenic differentiation of stem cells predominantly occurs at a matrix stiffness of 25–40 kPa, neural differentiation of stem cells is favoured in softer matrices (0.1–1 kPa) (Fig. 6b–e) [74, 75]. Furthermore, subtle differences in the stiffness of the microenvironment can direct the differentiation of neural stem cell into different neuronal lineages, such as neurons, oligodendrocytes, and astrocytes. For example, when the stiffness of the matrix material is below 500 Pa, neuronal stem cells differentiate into neurons, whereas stiffer materials induce astrocyte or oligodendrocyte differentiation [158]. However, when the stiffness of substrate is extremely soft (E’ ~10 Pa), the morphology of neural cells becomes round, and differentiation is inhibited [158]. In this respect, to enhance the growth and activity of both bone cells and neural cells in engineered bone biomaterials, finding an optimal stiffness level of bone biomaterials is important. Alternatively, a patterned structured bone biomaterial composed of osteogenic and neurogenic biomaterials can be constructed by using 3D bioprinting technology.

3.2.2. Roughness of innervated bone materials

In living systems, molecular and cellular behaviours occur at the nanometer scale. For example, human bone tissue is composed of calcium phosphate-based nanocrystallites in the range of tens of nanometers, which are assembled on collagen nanofibers that have a periodicity of about 67 nm [159]. In the case of neural tissue, its extracellular matrix is based on laminin that has a size of approximately 70 nm [160]. Neural cells can sense surface topography and respond differently depending on different levels of roughness at the nanoscale [161]. For example, when human neuroblastoma cells were cultured on gold substrates with different levels of roughness, cellular spreading behavior and viability were significantly decreased as surface roughness increased from flat surface to approximately 100 nm of mean surface roughness (Fig. 6f) [161]. This is because the increased hydrophilicity of the rough surface may decrease the adsorption of hydrophobic neuronal cell adhesive serum proteins (e.g. laminin), which can impede sensing by focal adhesion complexes. Furthermore, neurons grown on a rough substrate show disrupted polarity, nuclear condensation, and a functionally impaired actin cytoskeleton. In this respect, precise control of the nanoscale roughness of bone biomaterials is important to modulate neuronal cell growth into the bone matrix.

3.2.3. Porosity of innervated bone materials

Generating porous structured biomaterials is another important factor to consider for designing innervated bone materials to induce the axonal growth of neuronal cells. Porous substrates with increased surface area can bind greater amounts of neurogenic proteins, and improve the axonal outgrowth of neuronal cells. To evaluate the effect of pore size of bone materials on the growth of neurons, silicon materials are often used because of the ease of controlling pore size in a precise manner [162]. For neurons, porous structured silicon substrates are known to increase their adhesion and viability compared to flat silicon substrates (Fig. 6g) [163–166]. For example, when dorsal root ganglia (DRG) neurons were cultured on a mesoporous substrate with an approximate pore size of 300 nm, a greater number of axons were observed compared to planar areas [163]. In addition, when immortalized human cortical neuronal cells and mouse neuroblastoma cells were grown on porous silicon substrates, the density of adhered cells was higher on the nanoporous substrate than on the flat substrate, with improved spreading morphology [165]. The proliferation rate of neuronal cells was further increased as the pore size of the nanoporous substrate decreased from 20 nm to 5 nm. On the other hand, as the size of pores was increased to the micrometer range, no significant differences in the outgrowth of DRG axons were observed compared to flat surfaces [163]. In the case of neuroblastoma cells, when the pore size of the silicate substrate becomes larger than the microscale (>1 μm), these cells are known to stabilize themselves by increasing cell-to-cell contact and forming spherical cell aggregates with minimized area of filopodia at the surface [166]. Considering that a lower level of roughness is better for enhancing growth of neuronal cells, it seems that neuronal cells favour smooth material surfaces with a small pore size. Since osteogenic cells and neurogenic cells require different microenvironments for their growth and differentiation, there is a need to design composite materials that are composed of both hard materials and soft materials in a spatially controlled manner.

3.3 Current innervated bone implant types and status

Although it is known that the nervous system has important biological functions for regulating bone formation, only few studies have utilized the function of the nervous system in bone materials for improving osteogenesis and regulating skeletal homeostasis. To develop an implant material that can promote the formation of nerve tissue and bone tissue, nerve autografts can be integrated into bone implants or biomaterials.

Implanting sensory nerve tracts into tissue-engineered bone grafts can promote osteogenesis by neurotization [14, 20, 143]. Wu et al. integrated sensory nerve tracts in porous β-TCP bioceramic-based scaffolds by dissociating the femoral saphenous nerve and fixing it into the side groove of the implant. This allowed the neurons to grow through the porous structure of the scaffold [143]. The sensory nerve integrated system resulted in enhanced expression of CGRP and growth-associated protein 43 (GAP43), which is relevant to nerve healing after implantation into a rabbit critical-sized femur bone defect. Also, new bone formation was improved compared to the implanted group without integration of the sensory nerve.

Bone implant materials containing magnesium ions (Mg²⁺) can also increase CGRP after implantation in the rat femur defect site, particularly around the peripheral cortex region and the DRG. In particular, CGRP secreted from stem cells derived from periosteum (PDSCs) enhances bone fracture healing, inducing osteogenic differentiation (Fig. 7) [20]. Mg²⁺ ions enter DRG neurons through the transient receptor potential (TRP) cation channel, subfamily M, member 7 (TRPM) and magnesium transporter 1 (MAGT1), increasing production of adenosine triphosphate (ATP) and accumulating CGRP-encapsulating terminal synaptic vesicles. The released CGRP further activates receptors on PDSCs (CALCRL-RMP1), which induces phosphorylation of cAMPresponsive element binding protein 1 (CREB1) and upregulates the expression of osteogenic differentiation related genes, including runt-related transcription factor 2 (RUNX2) and Sp7 (Osterix). These results indicate that bone materials can directly regulate neurogenesis in skeletal tissue in a coupled manner with osteogenesis, through intercellular signaling mechanisms. Therefore, it is important to consider the neurogenic effects of biomaterials during bone implant design to mimic the innervated bone niche and induce better bone regeneration.

Figure 7.

Magnesium ions (Mg²⁺) can enhance CGRP neuronal production to promote bone regeneration. (a) Photo image of Mg rod implanted in a rat cortical bone, from a cross-sectional view. (b–c) H&E staining images (b) and calcein-green labeling images (c) showing increase of new bone formation after implantation of Mg rod for 2 weeks (BM: bone marrow, NB: new bone, OB: old bone, and P: periosteum. Scale bar: 200 μm). (d) CGRP immunofluorescence staining images in DRGs in rat lumbar, 2 weeks post Mg rod implantation. Cellular nuclei are stained with DAPI (Scale bar: 50 μm). (e) Radiograph images of the fractured rat femur bone, 4 weeks post implantation of intramedullary nail with/without Mg component. To confirm the effect of CGRP, Ramp1 was knocked down or overexpressed in vivo (IMN: innovative intramedullary nail, AdV-NC: a negative control treated with scrambled adenoviruses, AdV-Ramp1: a group with Ramp1 overexpression, and AdV-shRamp1: a group with knocked down of Ramp1). (f) Schematic diagram explaining the effect of (Mg²⁺) on coupled osteogenesis and neurogenesis. Implant derived Mg²⁺ ions can enter DRG neurons through Mg²⁺ channels or transporters, enhancing accumulation and exocytosis of CGRP-vesicles. The CGRPs released from DRGs then trigger the CGRP receptor on PDSCs and induce osteogenic differentiation (Images reproduced with permission from [20]).

4. Designing neurovascularized bone implants

4.1 Understanding development mechanism of coupled vascularization and innervation in bone

Human bone is fully vascularized and innervated with coupled blood vessels and nerve fibers, which are spread through Harversian canals and Volkmann’s canals that exist throughout the entire bone. These blood vessels and nerve fibers play fundamental roles in bone development by supplying nutrients and oxygen [47, 167]. Notably, there exists a synergistic effect between vascularization and innervation in bone, demonstrating that these processes cooperate during bone developmental stages [14]. In this respect, this section will firstly introduce the anatomy of the neurovascular network in skeletal tissue, and then describe its coupled effects on bone healing.In the human body, the distribution of blood vessels and peripheral nerves often follow each other, and reciprocal signals influence their growth and activity (Fig. 8a) [24]. Vascularization and innervation are also anatomically and functionally coupled in bone tissue: nerve fibers are present along the blood vessel walls and, at the same time, nerve fibers contain blood capillary networks (Fig. 8b–c) [167]. For example, long bones in neonatal rats contain nerve fibers around the periosteum and connective tissues, which are mostly located along blood vessels. In particular, a dense innervation in the diaphysis and deep metaphysis was observed adjacent to blood vessels (Fig. 8b–c) [167].

Figure 8.

(a) Vessels (red) and nerves (green) aligned together (reproduced with permission from ref. [24]). (b–c) Immunohistochemistry of nerve markers in metaphysis (b) and deep metaphysis (c) of neonatal rat femur bone. Positive nerve marker staining (arrows), local bone trabaeculae (t) and blood vessels (V), osteoclasts (Oc), osteoblasts (Ob) and hematopoietic cells (H) are shown. Original magnification: 1000× for panel (b); 1500× for panel (c). (Images reproduced with permission from [167]). (d) Role of β-NGF on angiogenesis. β-NGF binds the TRk-A receptor on endothelial cell surface and sensory nerves, inducing SP and CGRP-I release from sensory nerves, which target endothelial cells through the NK1 and CGRPR receptors. β-NGF also binds to macrophage surface to induce VEGF release. (Image reproduced with permission from [23]).

Based on their anatomical similarity, blood vessels and nerves are also functionally coupled and enhance the growth and development of each other. In particular, the integration of blood vessels within a tissue-engineered bone can improve innervation, and vice versa [14, 168, 169]. While blood vessels provide oxygen and nutrients that promote the formation of a nerve network, sensorial neuropeptides from nerve fibers influence angiogenesis, eventually promoting bone healing [14, 19]. Vasculogenic neuropeptides that are secreted from peripheral nerve fibers include neuropeptide Y (NPY), calcitonin gene-related peptide-I (CGRP-I), and SP. While NPY originates from autonomic nerves, CGRP and SP are mostly generated from sensory nerves [169–171].

NPY is a neuropeptide that is highly expressed in the PNS and CNS. It is mainly localized in nerve fibers in the periosteum, especially in epiphyseal marrow and the osteochondral junctions [172]. Bone tissue has a large supply of NPY from nervous networks, which are anatomically associated with blood vessels [148, 149]. NPY directly regulates osteoblast function [173, 174], and plays a key role during bone healing [175]. In particular, temporal changes in NPY activity have been observed during the initial and late phases of fracture healing. The early up-regulation of NPY has a direct response to the injury, leading to an enhancement of angiogenesis, eventually improving bone formation [176, 177]. On the contrary, a later increase in NPY secretion coinciding with the reduction of the initial callus thickness illustrates the role of NPY in bone resorption during the remodelling phase [175].

CGRP-I and SP are peptides that bind to specific vascular receptors, such as the calcitonin gene-related peptide-I receptor (CGRPRP) and neurokinin-1 (NKY) receptors, respectively. They can promote endothelial cell proliferation, migration, and improve angiogenesis [178]. In particular, CGRP-I is a neurotransmitter produced in primary sensory neurons. It is part of the calcitonin (CT) family of peptides, which targets cells through receptors such as the CGRP-I receptor, and receptor activity modifying proteins (RAMPs) that are present on endothelial cells and osteoblasts [179]. CGRP-I has a direct effect on vascularization and the regulation of bone regeneration. In particular, it binds the receptors of osteoblasts such as the CGRP-I receptor complex, enhancing their proliferation [144]. It also plays a key role in the dilatation of blood vessels by targeting receptors present on blood vessel smooth muscle cells and endothelial cells to modulate their activity [180–182].

In addition, SP enhances bone formation by directly promoting the differentiation of osteoblasts and enhancing BMP-2 secretion [179]. SP also influences angiogenesis by inducing the migration of MSCs to the lesion and promoting their VEGF secretion, which is essential for angiogenesis and consequently for bone regeneration [183–186].

β-nerve growth factor (β-NGF) binds to the β-NGF receptor and Trk-A receptor on the surface of endothelial cells, directly stimulating their proliferation and migration (Fig. 8d) [23]. At the same time, it binds to the surface of sensory nerves, inducing the secretion of the angiogenic neuropeptides, SP and CGRP-I [187]. β-NGF also binds β-NGF receptors and Trk-A receptors of macrophages, inducing VEGF release and subsequent angiogenesis [23, 188].

4.2 Current neurovascularized bone implant types and status

Bone materials that incorporate and link blood vessels and nerves are expected to more effectively recapitulate the microenvironment of bone tissue, resulting in better integration with the surrounding skeletal system. In this section, we will introduce the current research status and future possibilities for combining innervation and vascularization within bone materials.

Since blood vessels and nerve fibers are in close proximity in the human body, implanting autologous blood vessels can generally accompany parts of nerve tissue to develop neurovascularized tissue [149, 168]. In fact, vascular tissues can provide an ideal niche for nerve regeneration by supplying oxygen and nutrients. As a result, the majority of neurogenesis occurs in the vicinity of blood vessels [189, 190]. For example, Chen et al. evaluated the osteogenic capability of vascularized and innervated cell-laden bone scaffolds, which were composed of bone-marrow derived MSC-incorporated porous β-TCP bioceramic material [168]. This cell-laden bone scaffold was then implanted into a rabbit femur defect model, while the branch of the femoral blood vessels from the same hind limb were separated and implanted at the side of the bone scaffold. Histological results showed that the presence of vascular bundles or sensory nerves promoted osteogenesis with respect to the control. Interestingly, after up to 12 weeks from surgery, the expression of neural marker (CGRP1R and neuropeptide Y1 receptor (NPY1R)), was enhanced in the regenerated bone tissue when the scaffold was integrated with vessels, compared to the scaffold that was combined with sensory nerves. This is probably because the vascular bundles are able to supply oxygen and nutrients that are essential for nerve regeneration.

While vascularized tissue induces innervation by providing a vascular niche and innervated tissue, bone implant materials that can integrate both nerves and vessels have not been actively investigated, due of the difficulty of regulating multiple tissue types. However, there have been attempts to utilize neurovascularized bone implants by using autografts. For example, Feng et al, evaluated the effect of autologous neurovascular bundles for inducing vascularization and osteogenesis in a dog mandible defect model [169]. In particular, a Ti scaffold filled with a β-TCP bone material was implanted, and an alveolar neurovascular bundle was truncated and placed through the groove of the bone scaffold. The results showed that vascularization was induced throughout the bone scaffold, and bone regeneration successfully occurred. As a potential future concept, Fan et al. suggested integrating both blood vessel bundles and nerve fibers into the bone graft by microsurgery techniques during bone implant implantation in vivo [191].

Conclusions and future perspective

In this review, we have discussed the overall anatomy and developmental mechanisms of the vascular and nervous system in skeletal tissue, especially from the perspective of improving bone growth and fracture healing. We then elaborated on important design criteria for developing vascularized bone materials and innervated bone materials. We also introduced recent achievements in developing vascularized bone implants and innervated bone implants for promoting bone tissue regeneration. Finally, we highlighted the anatomically and functionally coupled distribution of blood vessels and nerve fibers in bone tissue, and their synergistic interactions for promoting osteogenesis based on intercellular communication. In summary, blood vessels promote neurogenesis by supplying oxygen, nutrients, and neurogenic growth factors, while nerve fibers enhance vascularization by providing vasculogenic neuropeptides. This process, in turn, facilitates bone growth and remodeling.

Despite successful generation of organic and inorganic materials that have similar physicochemical properties with bone, engineering neurovascularized bone implant materials are still in the infancy stage. In particular, most bone implant materials require a certain period of time for establishing a linking network with surrounding tissues to prevent necrosis.

Although there have been several encouraging studies for developing vascularized bone implant materials, most of them are still not applicable in the clinic. In particular, this is due to the difficulty of regulating multiple tissue types, and the complexity of pre-establishing peripheral vascular networks in bone grafts before implantation. Moreover, there are only limited studies that have developed innervated bone implant materials, which is due to the difficulty of developing a highly connective PNS inside the hard tissue and regulating its physiology and metabolism. Consequently, until now, engineered neurovascularized bone implants have not been utilized in the clinic.

Since human bone is highly vascularized and innervated, we expect that neurovascularized bone implant materials will recapitulate the microenvironment of bone tissue, and induce fracture healing based on improved integration with surrounding tissue. In this regard, a better understanding of the interaction mechanisms between blood vessel and nerve fibers in skeletal tissue is required. In addition, the development of a facile methodology for constructing neurovascularized bone implant platforms based on a biomimetic approach is needed. In this respect, predictive technologies such as numerical and molecular dynamic simulations to predict vessel formation and maturation [192, 193], as well as modelling the adsorption of proteins and growth factors [194–196], are expected to advance the design of biomimetic bone grafts. Taken together, we expect that this review paper will provide a holistic view of the anatomy and function of the neurovascular network in bone tissue, and aid in the development of neurovascularized bone materials for clinical applications.