Bone tissue engineering using 3D printing by Bose et. al



With the advent of additive manufacturing technologies in the mid 1980s, many applications benefited from the faster processing of products without the need for specific tooling or dies. However, the application of such techniques in the area of biomedical devices has been slow due to the stringent performance criteria and concerns related to reproducibility and part quality, when new technologies are in their infancy. However, the use of additive manufacturing technologies in bone tissue engineering has been growing in recent years. Among the different technology options, three dimensional printing (3DP) is becoming popular due to the ability to directly print porous scaffolds with designed shape, controlled chemistry and interconnected porosity. Some of these inorganic scaffolds are biodegradable and have proven ideal for bone tissue engineering, sometimes even with site specific growth factor/drug delivery abilities. This review article focuses on recent advances in 3D printed bone tissue engineering scaffolds along with current challenges and future directions.


Osseous tissue, known as bone, is made of two different structures; cancellous and cortical bone. Cancellous, or the inner part of bone, is spongy in nature having 50–90 vol% porosity. However, cortical bone is the dense outer layer of bone with less than 10 vol% porosity. Both types of bone undergo dynamic remodeling, maturation, differentiation, and resorption that are controlled via interactions among osteocyte, osteoblast, and osteoclast cells [1]. Osteoblasts are primarily responsible for new bone formation while osteoclasts are responsible for the resorption of old bone. Such a dynamic process involving osteoclasts and osteoblasts is known as bone remodeling, and is responsible for maintaining a healthy bone. Bone is well known for its self-healing abilities [2]; however, large-scale bone defects cannot be healed completely by the body [3,4], and in most cases, external intervention is needed to restore normal operations. Among different treatment options such as autografts (bone taken from the same person's body) and allografts (bone tissue from a deceased donor), bone tissue engineering that is focused on methods to synthesize and/or regenerate bone to restore, maintain or improve its functions in vivo[5,6] is becoming popular. Successful application of bone tissue engineering can avoid challenges related to other treatment options involving different materials such as autografts or allografts. Apart from material issues, a clear understanding of biology involving cells, extracellular matrix (ECM) and growth factors are pivotal in bone tissue engineering [7].

Scaffolds are an integral part of bone tissue engineering. Scaffolds are three dimensional (3D) biocompatible structures which can mimic the ECM properties (such as mechanical support, cellular activity and protein production through biochemical and mechanical interactions), and provide a template for cell attachment and stimulate bone tissue formation in vivo[3,5–7]. Besides chemistry, pore size, pore volume and mechanical strength are critical parameters which define a scaffold's performance. At an early stage, bone ingrowth happens at the periphery of scaffolds with a negative gradient in mineralization toward the inner parts [4]. For continuous ingrowth of bone tissue, interconnected porosity is important. Open and interconnected pores allow nutrients and molecules to transport to inner parts of a scaffold to facilitate cell ingrowth, vascularization, as well as waste material removal [4,6,8]. Since higher porosity increases surface area per unit volume, the biodegradation kinetics of scaffolds can be influenced by varying pore parameters. Biodegradation through a cell-mediated process or chemical dissolution are both important to ascertain stabilized repair and scaffold replacement with new bone without any remnant [8]. A minimum pore size between 100 and 150 μm is needed for bone formation [4,9]; however, enhanced bone formation and vascularization are reported for scaffolds with pore size larger than 300 μm [9–11]. Pore size also plays an important role in ECM production and organization. Poly(d,l-lactic acid) (PDLLA) scaffolds with pore size 325 and 420 μm led to well-organized collagen I network; whereas, smaller pore size of 275 μm prevented the human osteosarcoma-derived osteoblasts to proliferate, differentiate and produce functional ECM [12]. Pore volume also controls the permeability of nutrients to the scaffold and their mechanical properties. Permeability in poly-ɛ-caprolactone (PCL) increased with higher pore volume and resulted in better bone regeneration, blood vessel infiltration, and compressive strength in vivo, when other pore parameters were kept the same [13]. Apart from biological performance, the initial mechanical properties and strength degradation rate should match that of the host tissue for optimum bone healing [14]. The strength degradation kinetics of porous scaffolds are highly affected by pore size, geometry, and strut orientation with respect to the loading direction [15,16]. Finally, surface properties such as chemistry, surface charge and topography also influence hydrophilicity and in turn cell–material interactions for bone tissue ingrowth [17–19].

Porous bone scaffolds can be made by a variety of methods. Chemical/gas foaming [20], solvent casting, particle/salt leaching [12,21], freeze drying [22], thermally induced phase separation [23], and foam-gel [24] are some of those that have been used extensively. However, pore size, shape, and its interconnectivity cannot be fully controlled in these approaches. Moreover scaffolds with tailored porosity for specific defects are difficult to manufacture with most of these approaches [21–24]. Such scaffolds can be designed and fabricated using additive manufacturing (AM) approaches. Different AM approaches, for example, 3D printing (3DP), solid freeform fabrication (SFF), rapid prototyping (RP), are approaches that allow complex shapes for scaffolds’ fabrication directly from a computer aided design (CAD) file [25–27]. The concept of AM was first introduced by Chuck Hull in 1986 via a process known as ‘stereolithography (SLA)’ [28,29].

Some of the commercially available AM techniques are 3DP (ExOne, PA), fused deposition modeling (FDM, Stratasys, MN), selective laser sintering (SLS, 3D Systems, CA), stereolithography (3D Systems, CA), 3D plotting (Fraunhofer Institute for Materials Research and Beam Technology, Germany), as well as various forms of direct writing [27]. In all these AM approaches, 3D scaffolds are created layer-by-layer without any part specific tooling or dies [30,31]. These AM techniques can be classified as – (a) extrusion (deformation + solidification), (b) polymerization, (c) laser-assisted sintering, and (d) direct writing-based processes. Table 1[32–59] summarizes some of the AM techniques toward bone tissue engineering applications including their advantages and disadvantages.

Table 1. RP techniques for bone scaffold fabrication.

TechniqueProcess detailsProcessed materials for bone tissue engineeringAdvantages (+) and disadvantages (−)Reference
3D Plotting/direct ink writing→ Strands of paste/viscous material (in solution form) extrusion based on the predesigned structure
→ Layer by layer deposition of strands at constant rate, under specific pressure
→ Disruption of strands according to the tear of speed
→ Hydroxyapatite (HA)
→ Bioactive glasses
→ Mesoporous bioactive glass/alginate composite
→ Polylactic acid (PLA)/polyethylene glycol (PEG)
→ PLA/(PEG)/G5 glass
→ Poly(hydroxymethylglycolide-co-ɛ-caprolactone) (PHMGCL)
→ Bioactive 6P53B glass
→ Mild condition of process allows drug and biomolecules (proteins and living cells) plotting
→ Heating/post-processing needed for some materials restricts the biomolecule incorporation
Laser-assisted bioprinting (LAB)→ Coating the desired material on transparent quartz disk (ribbon)
→ Deposition control by laser pulse energy
→ Resolution control by distance between ribbon/substrate, spot size and stage movement
→ HA
→ Zirconia
→ HA/MG63 osteoblast-like cell
→ Nano HA
→ Human osteoprogenitor cell
→ Human umbilical vein endothelial cell
→ Ambient condition
→ Applicable for organic, inorganic materials and cells
→ Quantitatively controlled
→ 3D stage movement
→ Homogeneous ribbons needed
SLS→ Preparing the powder bed
→ Layer by layer addition of powder
→ Sintering each layer according to the CAD file, using laser source
→ Nano HA
→ Calcium phosphate (CaP)/poly(hydroxybutyrate–co-hydroxyvalerate) (PHBV)
→ Carbonated hydroxyapatite (CHAp)/poly(L-lactic acid) (PLLA)
→ β-Tricalcium phosphate (β-TCP)
→ No need for support
→ No post processing is needed
→ Feature resolution depends on laser beam diameter
SLA→ Immersion of platform in a photopolymer liquid
→ Exposure to focused light according to desired design
→ Polymer solidifying at focal point, non-exposed polymer remains liquid,
→ Layer by layer fabrication by platform moving downward
→ Poly(propylene fumarate) (PPF)/diethyl fumarate (DEF)
→ β-TCP
→ Complex internal features can be obtained
→ Growth factors, proteins and cell patterning is possible
→ Only applicable for photopolymers
FDM→ Strands of heated polymer/ceramics extrusion through nozzle→ Tricalcium phosphate (TCP)
→ TCP/polypropylene (PP)
→ Alumina (Al2O3)
No need for platform/support
→ Material restriction due to need for molten phase
Robotic assisted deposition/robocasting→ Direct writing of liquid using a nozzle
→ Consolidation through liquid-to-gel transition
→ 6P53B glass/PCL
→ Independent 3D nozzle movement
→ Precise control on thickness
→ No need for platform/support
→ Material restriction

3D printing (3DP) – history and methodology

3DP, a technology developed in the early 1990s at MIT (Cambridge, MA) by Sachs et al.[60], is a powder-based freeform fabrication method in which using a regular ink-jet print-head, binders are printed on to loose powders in a powder bed. Early research in this area was focused on rapid tooling using metals and ceramics [61].

Fig. 1 shows a schematic representation of the 3DP process [62]. For bone tissue engineering, 3DP is useful for the direct fabrication of scaffolds with tailored porosity from a CAD file. Before printing, essential parameters such as powder packing density, powder flowability, layer thickness, binder drop volume, binder saturation and powder wettability need to be optimized to improve the quality of the resultant part. Packing density is the relative density of the powder bed after uniform spreading. To start a build, enough powder should be packed homogeneously in a feed bed. A set of rollers spread a layer of powder to a predetermined thickness to create a powder bed. Powder flowability is critical in this process as it determines the spreading ability. Flowability is primarily determined by particle size, size distribution, surface roughness and shape. The desired layer thickness is in part determined by geometry and powder characteristics. Thinner layers cause binder penetration and excess spreading to other sites resulting in poor resolution and tolerance. However, thick layers need high saturation for the powders to bind [62].

Figure 1. (a) 3D printing schematic using an inkjet printing system. (b) 3D printed CaP sintered structures fabricated at WSU.

The printhead sprays the binder across the build layer in several passes, based on the instructions in the tool path file created according to the CAD file. The binder, which can be organic or water-based, locally binds the particles and hardens the wetted area, or results in a reaction similar to the hydraulic setting reaction in cements [63–65]. The binder drop volume and saturation play crucial roles. The binder drop volume is the amount of binder released from each nozzle per drop during printing, which depends on the binder density and viscosity. By coordinating the powder packing density and the drop volume, the binder saturation data required for printing is obtained. For a constant packing density, a higher drop volume demands a lower binder saturation [62]. Low saturation can cause layer displacement during processing. The binder saturation also depends on the powder wettability. The powder wettability, which is related to the powder particle chemistry and surface energy, determines the printing accuracy and the achievable tolerance [66]. While high wettability results in extensive binder spreading, low wettability causes week powder-binder integration [67].

After printing, the printed layer is moved under a strip heater to allow the binder to dry out and prevent spreading between layers [62]. This process is repeated until the printing of the designed part is complete. Heat treatment is needed to complete the binder reaction and increase the part green strength. Next step is depowdering, that is, the removal of loose powder from the printed body. This is one of the major challenges for porous scaffolds in 3DP due to the low green density of the part. Loose powder removal from fine pores can easily crack a green part [68].

In general, a large variety of ceramic, metallic, polymeric, and composite materials can be processed using 3DP; however, binder selection and process parameter optimization are the keys to successful part fabrication. In bone tissue engineering the advantages of this method arise through the control of fine features including interconnected porosity, no contamination issues related to any second material for support structures and the direct printing ability with both metallic and ceramic biomaterials [65,69]Fig. 2a shows some examples of 3D printed scaffolds with different pore sizes. It is important to note that extensive optimization is needed to process good quality parts with 3DP for any new material, a fundamental drawback for this approach.

Figure 2. (a) Photograph of the sintered 3D printed TCP scaffolds for mechanical strength and in vivotesting (small samples) [62]. (b) Compressive strength comparison of the scaffolds sintered at 1250 °C in conventional and microwave furnaces (**p < 0.05, *p > 0.05, n = 10) [62]; (c) SEM micrographs of hFOB cells showing the cell adhesion and proliferation on the scaffold surface and inside the 3D interconnected macro pores after 3 days of culture (white arrows indicate cells): 500 μm (i) & (ii), and 750 μm (iii) & (iv) [62]; (d) SEM image of the pure TCP scaffold showing the surface morphology and designed macro pore distribution [84]; (e) photomicrograph of 3DP pure (TCP) implants (i and iii), and Sr–Mg doped TCP implants (ii and iv) showing the new bone formation inside the interconnected macro and intrinsic micro pores of the 3DP scaffolds after 4 and 8 weeks in rat distal femur model. Modified Masson Goldner's trichrome staining of transverse section. OB: old bone, NB: new bone and BM: bone marrow. Color description: Dark gray/black = scaffold; orange/red = osteoid; green/bluish = new mineralized bone (NMB)/old bone [84]; and (f) histomorphometric analysis of osteoid area fraction (osteoid area/total area, %) from 800 μm width and 800 μm height tissue sections (**p < 0.05, *p > 0.05, n = 8). Completely mineralized bone formation was observed in presence of SrO and MgO in TCP after 12 weeks, hence no osteoid area was observed. All osteoid like bone was transformed into mineralized bone after 12 weeks in doped TCP due to the presence of strontium and magnesium. Hence, there was no osteoid like bone left after 12 weeks in Sr–Mg-doped TCP [84].

3D printed bone scaffolds

Table 2 summarizes a few selected material-binder system combinations for bone scaffolds using 3DP. Starch-based binders are one of the candidates for bone replacement applications. These binders are biocompatible and produce structures that have a mechanical strength close to trabecular bone [71,72]. Structural designs and post processing conditions both can influence the mechanical properties of 3D-printed starch-based scaffolds [73]. 3D-printed polyethylene (PE) scaffolds with 22.3–49.7% porosity have shown a tensile strength up to 4 MPa, and no toxicity to human osteoblasts [74].

Table 2. 3D printed materials for bone tissue engineering.

MaterialLayer thicknessBinderReference
TCP20 μmAqueous based[62]
α/β-TCP modified with 5 wt% hydroxypropymethylcellulose100 μmWater[64]
CaP mixture with Ca/P ratio of 1.7100 μm10% phosphoric acid[64]
Tetracalcium phosphate (TTCP), dicalcium phosphate and TCP100 μm25% citric acid[64]
HA300 μmSchelofix (water soluble polymeric compound)[65,75]
TTCP/β-TCP100 μm25 wt% of citric acid[68]
TTCP/calcium sulfate dihydrate100 μm25 wt% of citric acid[68]
HA100 μmNo information[69]
TCP100 μmNo information[69]
Biphasic calcium phosphate (BCP)100 μmNo information[69]
α/β-TCP (final product: dicalcium phosphate dihydrate (DCPD))No information20% phosphoric acid[70]
Starch/PLLA + PCLNo informationDistilled water + blue dye[73]
High density PE (HDPE)0.175 mmMaltodextrin + polyvinyl alcohol (PVA)[74]
SiO2–ZnO-doped TCP20 μmAqueous based[76]
PE or HDPE0.175 mmWater based binder[77,78]
PLANo informationChloroform[79]
TCP (final product:DCPD)0.1 mm20% phosphoric acid[80]
HA/maltodextrin0.175Water based binder[81]
TTCP (final product: HA)100 μm0.5 mol/l Ca(H2PO4)2 + 10% H3PO4[82]
TCP (final product: brushite)100 μm0.5 mol/l Ca(H2PO4)2 + 10% H3PO4[8,82]
HA/A-W glass0.1 mmWater based[83]

CaP ceramics are widely used in bone tissue engineering due to their excellent bioactivity, osteoconductivity, and similarities in composition to bone. Capillaries and vessel formation, and homogeneous osteoconduction from central channels, have previously been observed in 3D-printed HA blocks [75]. The effect of pore size on human fetal osteoblasts (hFOB) was studied with 3D-printed TCP scaffolds [62]. The decrease in designed pore size from 1000 to 750 and 500 μm resulted in an increase in proliferated cell density. 3D printed and microwave sintered β-TCP scaffolds fabricated using 3DP are shown in Fig. 2a, showing interconnected macro porosity across the sample. Fig. 2c (i–iv) presents the morphologies of hFOB cells on scaffold surfaces and pore walls after 3 days of culture showing good cell adherence and cell ingrowth into the pores, suggesting that the scaffolds were non-toxic. A secondary electron microscopy (SEM) image of the surface morphology and the designed macro pore distribution in a pure TCP scaffold is shown in Fig. 2d. New bone formation was observed at the implant/host bone interface as well as inside the interconnected macro and intrinsic micro pores after 4 and 8 weeks in both pure and doped TCP as shown in Fig. 2e. However, more osteoid like new bone formation was observed in SrO–MgO doped TCP scaffold as shown in Fig. 2f. Histological evaluation and histomorphometric analysis reveal that the treatment group (doped TCP scaffolds) facilitated higher osteoid like bone at an early stage, and completely mineralized bone later, which could be essential for fast bone healing and mineralization in vivo[62,84].

Further studies have shown that the addition of SiO2–ZnO dopants to TCP scaffolds increases cell viability in different pore size ranges [76]. The biocompatibility of 3D printed CaP ceramics has also been studied using osteoclasts. Tartrate resistant acid phosphatase (TRAP) staining, lacunae formation and microscopic images confirmed the monocyte differentiation to multinuclear osteoclast-like cells on a wide range of compositions [69]. It has been shown that the use of phosphoric acid instead of polymeric binders can improve both resolution and compressive strength [64]. HA scaffolds with high surface areas showed no cytotoxicity and adequate cell adhesion with MC3T3-E1 fibroblast cells [65]. In addition to in vitro experiments, in vivo biocompatibility and osteoconductivity of 3D-printed scaffolds showed that the 3D-printed brushite and monetite cements with controlled open porosity increased osteoconduction in vivo in a goat model [8]. 3D-printed TCP samples with micro and macro-porosity also facilitated osteogenesis in a rat femur model [53]. Cytotoxicity results of MC3T3-E1 cells on two different bone cement based compositions of TTCP/β-TCP and TTCP/calcium sulfate dihydrate have been reported for bone tissue engineering. A wide range of binders were used. It has been reported that the shortest hardening time can be obtained between 20–40% of citric acid, and 30–40% of lactic acid; however, a lower range of those binders and a different concentration of sodium hydrogenphosphate with sulfuric and phosphoric acids can be used to increase the hardening time for the cements [68]Fig. 3a and bshow patient specific 3D printed CaP implants. These results point to the application of 3DP in a large variety of materials and structures for bone tissue engineering scaffolds.

Figure 3. (a) 3D printed cranial segment [68], (b) general view of the implant bearing skull. Implants are fixed with miniplate osteosynthesis respectively bicortical osteosynthesis (mandibular defect). The drill holes for screw insertion were made after the positioning of the implants using a common bone drill [70], (c) representative macroscopic views of one half of bioceramic implant at retrieval, loaded with 56 ng copper [80].

Mechanical properties of 3D printed scaffolds

Low mechanical strength is a major challenge in porous scaffolds, and is primarily controlled by pore volume. This is also true for 3D printed ceramic scaffolds and limits their use only in non-load bearing and low-load bearing applications. Optimized post processing approaches and compositional modifications can improve mechanical properties of ceramic scaffolds. The compressive strength of 3D printed TCP sintered scaffolds is shown in Fig. 2b. In agreement with observed shrinkage and increased density, microwave sintering results in a higher compressive strength. The strength of the scaffold increases with decreasing pore size or volume, and a maximum strength of 10.95 ± 1.28 MPa has been observed for scaffolds with 500 μm pores, with 42% total open porosity, when sintered at 1250 °C for 1 h in a microwave furnace [62]. In another study, when a mixture of TTCP/β-TCP was sintered at 1400 °C, it increased the strength of the 3D printed scaffold. However, sintering a TTCP/calcium sulfate dihydrate composite caused a decrease in the strength due to water release [68]. Tarafder et al. reported an effective densification approach, using microwave sintering compared to conventional heating, and improved the mechanical properties of 3D-printed TCP scaffolds [62]. Bioactive liquid phase sintering aids have also been reported to increase strength. 3D printed HA/A-W glass, where the glassy phase is added as a liquid phase sintering aid, showed an increase in strength from 1.27 MPa to 76.82 MPa when sintered at 1300 °C for 3 h [83]. The enhancement of tensile properties was also found in PE scaffolds as a result of thermally induced densification and binder degradation [77]. To increase the strength of ceramic scaffolds without impairing biological properties of scaffolds, another approach is monomer or polymer infiltration. A mixture of bismethacrylated oligolactide macromer (DLM-1), containing 10 wt% of 2-hydroxyethyl methacrylate has been used to increase the strength of scaffolds before and after sintering [68]. The immersion of HA scaffolds in triethylene glycol dimethacrylate (TEGDMA), 2,2-bis[4 (2-hydroxy-3thacryloyloxypropyloxy)-henyl] propane (bis-GMA) resulted in an increase of the flexural strength by at least 20 times [85]Table 3 summarizes the mechanical properties of 3D printed scaffolds tailored for bone tissue engineering.

Table 3. Mechanical properties of 3D printed scaffolds.

MaterialCompressive strength (MPa)Compressive stiffness (MPa)Compressive yield strength (MPa)Bending modulus (GPa)Bending strength (MPa)Reference
TCP-sintered conventionally at 1250 °C6.4    [62]
TCP-sintered using microwave at 1250 °C10.9    [62]
TTCP/β-TCP0.7    [68]
DLM infiltrated TTCP/β-TCP76.1    [68]
Brushite    5.2[70]
Monetite    3.9[70]
Starch 11.151.12  [73]
PLLA/PCL infiltrated starch 55.191.77  [73]
TCP-sintered conventionally at 1250 °C5.5    [76]
SiO2–ZnO doped TCP-sintered conventionally at 1250 °C10.2    [76]
HA/A-W glass   0.351.27[83]
HA/A-W glass-sintered at 1300 °C   34.176.82[83]
HA   0.40.69[85]
HA/bis-GMA   6.250[85]

Bioprinting of tissue engineering scaffolds

Apart from inorganic scaffold manufacturing, AM approaches are also used to explore possibilities in fabricating scaffolds with live cells and tissues. Organogenesis of liver tissue using 3D printed PLLA/poly(lactic-co-glycolic acid) (PLGA) scaffolds has been investigated in vitro. It was shown that culturing a mixture of hepatocytes and endothelial cells on a channeled biodegradable scaffold results in the desired tissue structure intrinsically [86]. In 3D fiber deposition, a cell-laden viscous polymer paste was prepared and printed using a syringe dispenser. Alginate hydrogel-embedded multipotent stromal cells (MSCs)/chondrocytes were printed with a high cell viability using this method. The incorporation of MSCs and chondrocytes resulted in distinctive ECM formation both in vitro and in vivo. In addition, an increase in strand distan