Introduction

Bone defect caused by specific diseases or medications is very common. Autologous bone, allogeneic bone or xenogeneic bone transplantation is commonly used in clinical practice. However, autologous bone sources are limited. Xenogeneic bone cannot participate in metabolism. Because of the development of bone tissue engineering, the transplantation of new scaffold materials and autologous cells has opened up new treatment options for bone defects. The bone tissue engineering principle is applied to construct a degradable porous bone scaffold, which is implanted into the human body after loading bone cells, growth factors, etc. [1]. These scaffolds can be decomposed and absorbed by bone cells to form new bone. As a template for cell planting and tissue regeneration, the preparation of bone scaffolds has become an important part of bone tissue engineering, which is also a bottleneck restricting the clinical application of bone tissue engineering. Ideal scaffold needs to have good biocompatibility, biodegradability and suitable mechanical strength [2]. It is difficult for traditional bone grafting techniques to prepare individualized scaffold according to the structural characteristics of specific lesions, and the internal pore structure is also difficult to effectively control. 3D printing technology meets these needs very well, which can accurately produce personalized tissue engineering bone scaffolds based on bone defect and CT imaging data, and match the scaffold with the defect areas perfectly, imitate the normal nature tissue microstructure of human. With the continuous updating of 3D printers, seed cells and scaffold materials can be printed simultaneously, and the differentiation of osteoblasts is accompanied by the biodegradation of biomaterial. Some scholars have used the 3D printing technology to prepare human maxillary and mandibular scaffold models based on CT scan data and induced intravascular regeneration [3].

Principles of bone tissue engineering

In order to fundamentally understand the changes in pathological tissue structure and function, and to develop biological substitutes for repairing damaged tissues, tissue engineering has emerged [4]. The repair of bone defects caused by lesions or trauma is an important part of tissue engineering. Due to the different sites of bone defects, the constructed artificial bones have different structures, sizes, and biomechanical strengths. The traditional method of constructing is to inoculate osteogenic functional cells on a three-dimensional scaffold material in vitro and implant them into the human body after a period of culture to repair the bone defect tissue. Another way is to repair bone defect in vivo, which is using active growth factors or osteogenic functional materials to induce the osteogenesis of autologous tissue (periosteal or mesenchyme cells). The whole repair or reconstruct process is demonstrated in Fig. 1 [5, 6].

Fig. 1
figure 1

Bone tissue regeneration process

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Function of bone scaffold

In general, bone tissue engineering consists of three basic elements, which is cells, growth factors and bone scaffolds [7, 8]. As a framework for bone tissue regeneration, scaffold materials directly affect the biological properties of seed cells, such as its survival, migration, proliferation and metabolism. This scaffold plays the following roles in vivo:

  1. (1)

    As a scaffold for inoculated of cells in vitro, directing tissue grow into specific morphology.

  2. (2)

    Provide support for the defect site and hinder the growth of other surrounding tissues.

  3. (3)

    As a carrier of the signal molecules, transport those single molecules to the defect site and act as a sustained release body to slowly functioning.

  4. (4)

    As a place for the differentiation and metabolism of bone tissues, transporting nutrients and metabolic waste for cell growth.

  5. (5)

    Binding to specific cells, screening and selective adhesion to different types of cells.

Performance of materials for bone engineering

The ideal bone scaffold should be able to repair the bone defect and restore the bone tissue function to the greatest extent, and the system inside the body is extremely complicated, so the stent should meet the following requirements:

Biocompatibility

Biocompatibility refers to the performance of non-biological materials implanted in the body to cause autologous biological tissue reactions, which is related to the safety of the stent during clinical use [1]. Good biocompatibility requires that the scaffold material has good surface physicochemical properties to ensure normal adhesion and growth of the cells; the scaffold does not cause inflammatory reaction, any immunogenicity and cytotoxicity, and the degradation rate should be related to tissue growth. The speed is consistent, achieving a smooth transition from the stent to the ontogenesis [9, 10].

In addition, during the degradation of the scaffold, the tissue cells are provided with a constantly changing interface for adhesion and growth, which contributes to the firm adhesion of the cells to the material.

Osteoinductivity and osteoconductivity

In tissue engineering, the scaffold should not only serve as a carrier for cell proliferation, but also combine certain substances to induce the natural growth of bone tissue. There are two aspects: Firstly, the stent should have osteoconductivity. After being implanted into the human body, it can realize osteoblast transplantation, proliferating host bone tissue cells into the internal pores of the scaffold, and gradually dissolve the scaffold to form a new extracellular matrix. Secondly, the scaffold usually combines bone growth factors, which will be slowly released after implantation, and induces mesenchyme cells to differentiate into bone at non-skeletal tissue sites [11].

The three-dimensional (3D) microenvironment and porous matrix of the bone scaffold with good osteoconductivity enable the immigration of the bone cell along the surface or internal pores of the fabricated construct. The newly formed osteoblasts in bone scaffolds with good osteogenesis have the potential for further mineralization and formation of new external bone construct. Therefore, many scientists and technician have improved the osteoinductivity of bone scaffolds, by changing their physical and chemical properties or by incorporating biological constitute into the scaffolds [12]. Moreover, the porosity and pore topology structure of bone scaffolds can be modified, to enhance the osteoconductivity [13, 14].

Mechanical performance

The ideal bone scaffold needs to have mechanical properties that match the bone of the defect site, and the scaffold can always perform a perfect function from the beginning of implantation to the completion of the body repair [15,16,17,18]. Excessive hardness of the bone scaffold produces stress shielding. The load cannot transfer from the implant to the adjacent bone tissue, resulting in insufficient mechanical stimulation of the filled autologous bone and primitive cancellous bone. As the stent structure degrades, the implant gradually loosens, eventually leading to bone repair and reconstruction failure [19, 20]. Conversely, if the stiffness is too low, the load carrying capacity will be reduced and the bone will be more easily broken. Therefore, the scaffold cannot meet the mechanical strength requirements [21].

The scaffold material generally has a high porosity to facilitate the invasion and transport of host cells, and the corresponding mechanical strength is inevitably decreased. At the same time, the differences between different individuals are relatively large. Young individuals usually start weight bearing after 6 weeks of fracture. After about 1 year, the mechanical properties of the fracture site can be restored to the pre-injury level, while the recovery process of elderly individuals is relatively slow.

Among the existing scaffold materials, the organic material has a high elastic modulus, but the mechanical strength is low; while the bioceramic is brittle, the body is easily crushed. The combination of the two can obtain better mechanical properties, and the addition of chemical surfactant can increase the bonding energy between the composite interfaces, improve the dispersion of bioceramic particles in the organic phase and effectively improve the biomaterial of the scaffold material [22, 23].

Microstructural features

In addition to macroscopically connected pores (pore size > 50 μm) for excellent vessel growth and material transport [6], suitable microscopic pores (pore size < 50 μm) are also key to optimal osteogenesis [24]. Pore structure parameters, including pore size, porosity, connectivity between pores, degree of distortion of connected pores, and surface area, affect osteogenesis. The high-dimensional three-dimensional pore structure and high porosity significantly increase the specific surface area of the stent, which can ensure cell attachment and nutrient transport, and facilitate the smooth discharge of cellular metabolic waste and scaffold degradation products. The study found scaffold can achieve cell attachment, angiogenesis, nutrient transport, and metabolic waste discharge should have porosity more than 90% [25, 26].

Also, the extent of bone ingrowth depends on the pore size [27, 28]. The ideal pore diameter of scaffold materials should be similar to the size of the normal bone unit. (The average size of the human bone unit is about 223 μm.) Generally, a scaffold pore diameter ranging between 200 and 400 μm is considered adequate [29]. At the same time, some scholars have pointed out that it is not the size of the pores that restricts the invasion of bone tissue into the pores, but the degree of connectivity between the pores [30].

Bionic design of scaffolds for bone reconstruction

Biologically, bone tissue is a tough connective tissue composed primarily of cells, fibers, and stroma. The fiber is bone rubber fiber. There is a large amount of calcium deposits in the matrix. The cells mainly include osteoblasts, bone cells and osteoclasts. As shown in Fig. 2, the bone structure is divided into dense bone and cancellous bone. The dense bone consists of bone units composed of bone plate layer arranged by Harvard tube and its surrounding concentric layer. The cancellous bone is three-dimensional porous trabecular framework.

Fig. 2
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Anatomical structure of biological bone

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From the perspective of materialogy, biological bones are mainly composed of mineral salts and biological proteins. Inorganic mineral salts in bone tissue account for 60–70% of dry bone weight, most of which are hydroxyapatite [31]. Organic proteins account for about 30–40% of the dry bone weight, of which about 95% are type I collagen, which is ordered by the triple-stranded polypeptide chain, arranged in an order around the hydroxyapatite. In summary, the biological bone is formed by depositing calcium phosphate minerals on an ordered array of collagen substrates.

In tissue engineering, it is generally accepted that bionic design determines whether the biological structures prepared in vitro can survive in vivo, as well as whether the surviving bone tissue has proper biological functions. The bionic design of bone scaffolds includes the bionic structure design and bionic performance design. The existing bionic designed scaffold has the dense bone and cancellous bone structure of the human bone. Bionic designed sheep spine is shown in Fig. 3c, the problem of low strength of single porous structure is overcome, and the nutrient supply and shape matching of the bone scaffold can be ensured. At the same time, the mechanical properties of the compact bone are very high, and the mechanical properties of the cancellous bone vary greatly from person to person, and the individual bones have different lengths, complex surfaces and different joint angles. Therefore, it is still difficult to fabricate bone scaffold by taking both the spatial structure of natural bone and mechanical properties into account [21, 24, 32].

Fig. 3
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Bionic designed bone scaffolds before 3D fabrication [32]

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Formation methods of bone scaffolds

Numerous conventional techniques including solvent casting with particulate leaching, fiber bonding, membrane lamination, freeze drying, and gas foaming were developed to fabricate porous scaffolds [33, 34]. Although these techniques applicable to the preparation of porous scaffolds, but incapable of precisely controlling pore size, geometry and pore interconnectivity, the appearance is also unable to fully fit the defected bone tissues and therefore cannot be adapted to fabricate personalized scaffolds [35].

In addition, the size of the scaffold is often limited due to difficulties in removing pores and some of these conventional techniques use toxic organic solvents that may partially remain in the scaffold post-processing.

Rapid prototyping is well suited to the needs of tissue engineering for its advantages of high precision, high construction speed and capable of formation of complex structures to meet personalized medical needs and rejection reactions [36]. Additionally, additive manufacturing enables the incorporation of drugs/proteins as well as cells during scaffold manufacturing to produce very complex architecture similar to bone [37,38,39]. Therefore, additive manufacturing methodology is widely used for the formation of biological bone scaffolds.

In the process of personalized preparation of the porous structure and shape of the tissue engineering scaffold, the structure and shape of the damaged part are usually reconstructed in combination with medical CT scanning technology, and the rapid and accurate preparation of the scaffold is realized by means of rapid prototyping technology. The commonly used molding methods for preparing bone scaffolds can be classified into three types: laser-based molding systems such as photo-curable molding (SLA), selective laser sintering (SLS); nozzle-based extrusion molding systems such as fused deposition molding (FDM), Pneumatically assisted extrusion deposition (PDM); injection molding systems based on inkjet printing, such as three-dimensional forming (3DP), pneumatic micro-valve injection molding (PAM).

Fused deposition molding

The smelting deposition molding process is shown in Fig. 4. The filamentous thermoplastic material such as ABS, metal fuse or wax is used as raw material, and the raw material is fed into the nozzle by a pinch roller or a screw feeding mechanism and then heated under computer control. The nozzles are sequentially pressed to the printing platform according to the path of the model layered processing and then stacked layer by layer to complete the model construction. The structure of the model is determined by factors such as nozzle diameter, deposition rate, path spacing of the same layer, layer thickness and deposition angle.

Fig. 4
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Schematic diagram for FDM molding system

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FDM technology molding process is flexible and easy to control, and the application is more common. In addition to the early use of collagen, ceramics, PLGA and printing consumables, researchers have continued to experiment with composite materials and nanomaterials in recent years to obtain superior printing performance. Li et al. [40] demonstrated that PCL-TCP scaffolds can better promote bone growth and distribution by comparing the effects of titanium alloy and FDM molded PCL-TCP scaffolds as spine cages. Zein et al. [41] used FDM technology to prepare degradable porous scaffolds using filamentous polycaprolactone (PCL). Dong et al. [42] fabricated three-dimensional (3D) porous β-tricalcium phosphate (β-TCP)/calcium silicate (CS) composite scaffolds with different ratios by this kind of 3D printing technique. The scaffolds were composed of oriented filaments with an internal pore size of 160–200 μm and a porosity of 48–77%, and the strength of the bracket is highly related to the grid structure.

At the same time, there are some defects in FDM molding, such as the need for heating in the molding process, which can easily lead to degradation of polymer materials, inactivation of biological macromolecules, and is not conducive to compound biological factors. Long molding time, difficult to form the microporous structure of bionic bones and the smooth surface of the model, is not conducive to cell adhesion.

Stereolithography

Stereolithography is a kind of photopolymerization molding. Using the principle of photopolymerization of photosensitive resin, a specific wavelength and intensity of the laser are used to track the single-layer motion trajectory of the model so that the photosensitive material is solidified and formed, and then the movement of the lift table is completed to perform layer-by-layer scanning into a complete model (Fig. 5).

Fig. 5
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Schematic diagram for SLA molding system fabricating scaffolds

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The advantage of SLA is high resolution. The appearance of micro-stereolithography (MSLA) allows it to achieve an accuracy of 20 μm, allowing for the formation of more precise parts. Lee et al. completed polypropylene fumarate (PPF) with polylactic acid microspheres (PLGA) loaded with bone morphogenetic protein (BMP-2) and prepared highly connected porous scaffolds using MSLA technology [43]. Growth factors are gradually released during in vitro culture, and osteoblasts differentiate well, which provides an idea for the preparation of bone scaffolds. The main drawback of SLA is the lack of available biocompatible materials, and researchers are also experimenting with new composites. Ronca et al. used polylactide and nano-hydroxyapatite as substrates to obtain higher strength and better biocompatibility through SLA synthesis of composite scaffolds [44].

Selective laser sintering

The difference between selective laser sintering and SLA is that the substrate is in powder form. In the molding process, the powder material is first laid on a printing platform, and the cross-sectional shape of the model is scanned with a laser to perform selective sintering. Then, the lifting platform is adjusted and a new layer of powder is spread on the surface of the shaped part and is leveled off. The upper section of the sintered part is selectively spliced so that it is bonded to the molded part, and the stent model is stacked layer by layer as shown in Fig. 6. Since unbound solid particles can support any cantilever structure during sintering, SLS does not require a temporary support structure, simplifying the post-treatment process.

Fig. 6
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Schematic diagram for SLS molding system for fabricate scaffold models

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During the processing of bone scaffolds, polycaprolactone (PCL), polylactic acid (PLA), inorganic ceramics and composite materials thereof can be used as raw materials, and the processing precision is determined by the size of the powder, the diameter of the laser beam, and the pass of the powder layer. Roskies et al. demonstrated that porous polyetheretherketone (PEEK) scaffold can be prepared using SLS technology, which is capable of maintain the biological activity of cells and is conducive to the differentiation of osteoblasts [45].

Three-dimensional powder printing

The principle of 3D printing technology is similar to that of SLS. The difference is in the selective bonding of powder by spraying adhesive instead of laser sintering. The molding principle is shown in Fig. 7. In the printing process, the powder material is first laid on the printing platform, the inkjet nozzle sprays the adhesive according to the shape of the single-layer cross section of the model, and then the heater heats and dries the adhesive, then adjusts the printing platform, relays the powder and scrapes the flat, and then bonds the upper cross section to finally form the desired solid structure. Drying is required to ensure bond strength before further processing.

Fig. 7
figure 7

Schematic diagram for 3DP molding system

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Properties like powder bed thickness [46, 47], binder and powder composition [48, 49], surface roughness and flowability of powder [46], interaction of powder and binder [50] may affect the quality of the final product. Although 3DP does not require an additional support structure, it is difficult to completely remove the free powder when printing a bone support containing a microporous structure. In addition, due to the mild conditions in the printing process, extreme conditions such as laser heating are not required, which is beneficial to the addition of bioactive materials and is more suitable for the preparation of bone tissue engineering scaffolds. Hermann et al. used 3DP-prepared HA bone scaffolds with a pore size of 10–30 μm and a mechanical strength of up to 22 MPa, which can be used as a bone tissue engineering scaffold [51].

Challenges and future perspectives

Nowadays, various experimental verifications of scaffold materials and preparing methodology have been proven greatly contributing to the development of bone tissue engineering. In the meantime, still some issues need to be addressed.

Highly vascularization of bone scaffold

The effective of bone scaffold depends on the vascular system to transport nutrients, oxygen and carry away metabolic waste. Failed to establish a blood supply system will ultimately result in failure in scaffold implantation and may lead to immunological rejection. Vascularization of scaffold has become the bottleneck of clinical apply of tissue engineering. Traditional methods use biomaterial surface to conglutinate growth factors and peptides to promote the establishment of vascular system, but with little success. Fortunately in recent years, some scholars have proposed to establish a functional vascular system before implantation, that is to build a functional capillary system preferentially through 3D printing cells in vitro, and establish blood circulation through surgical techniques and receptors later [52]. Bose et al. [53] treated 3D-printed tricalcium phosphate bone scaffold with hydrochloric acid buffer containing Mg2+ and Sr4+ ions for 4–12 weeks, and 14–20% of the scaffold material surface has obvious vascular establishment. However, how to achieve the mutual signal transduction between cells, the invasion of vessels and how to make sure cells and bioactive factors keep active during the printing process to realize the final functionalization of the printed organs still remains a challenge.

Applicable powder material property and binder formula

When 3DP printing technology is applied to the manufacture of bone scaffolds, the material and structure determine the biological and mechanical properties of the bone scaffold, which directly affects the role of bone scaffold in the repair of bone defects. Firstly, the fluidity of the powder enables high-resolution printing, which mainly depends on the particle size and shape. The smaller the particle size, and the closer the shape is to the sphere, the better the fluidity. However, too small of the powder size tends to agglomerate due to internal special forces, resulting in reduced fluidity [54]. Secondly, binders must not only be able to effectively bond the substrate to meet the mechanical properties, porosity, and pore connectivity requirement, but also need to be non-cytotoxic, good biocompatibility, reasonable degradation rate, etc. Finally, the ratio of the binder needs to be adjusted, considering that the reactivity of the powder and the binder, the viscosity, density, surface tension and other parameters of the binder will affect the quality of the bond process.

Precision and quality control of three-dimensional forming of scaffold

Errors may occur due to various reasons during the preparation of the scaffold using different rapid prototyping methods. In the process of stereolithography, linear shrinkage and volume shrinkage occur during the polymerization of the resin from liquid to solid. Linear shrinkage will result in interlaminar stresses when layer-by-layer stacking causes the molded part to warp and deform, and volume shrinkage will cause changes in the size of the formed part, resulting in low precision. Three-dimensional powder printing experiences the stages of binder spray, impacting on powder bed, expansion, fusion and curing. Each stage may affect the forming quality [6], such as the satellite droplets generated during the injection, the compressive load generated when impacting the powder bed that will cause the longitudinal displacement of the printed thin layer, the excessive diffusion of the binder on the powder bed and the strength of the bond. To obtain scaffold with the desired accuracy and strength, it is necessary to study the influencing factors and find specific solution for each molding method.