Is it possible to 3D bioprint load-bearing bone implants? A critical review

Rehabilitative capabilities of any tissue engineered scaffold rely primarily on the triad of (i) biomechanical properties such as mechanical properties and architecture, (ii) chemical behavior such as regulation of cytokine expression, and (iii) cellular response modulation (including their recruitment and differentiation). The closer the implant can mimic the native tissue, the better it can rehabilitate the damage therein. Among the available fabrication techniques, only 3D bioprinting (3DBP) can satisfactorily replicate the inherent heterogeneity of the host tissue. However, 3DBP scaffolds typically suffer from poor mechanical properties, thereby, driving the increased research interest in development of load-bearing 3DBP orthopedic scaffolds in recent years. Typically, these scaffolds involve multi-material 3D printing, comprising of at-least one bioink and a load-bearing ink; such that mechanical and biological requirements of the biomaterials are decoupled. Ensuring high cellular survivability and good mechanical properties are of key concerns in all these studies. 3DBP of such scaffolds is in early developmental stages, and research data from only a handful of preliminary animal studies are available, owing to limitations in print-capabilities and restrictive materials library. This article presents a topically focused review of the state-of-the-art, while highlighting aspects like available 3DBP techniques; biomaterials’ printability; mechanical and degradation behavior; and their overall bone-tissue rehabilitative efficacy. This collection amalgamates and critically analyses the research aimed at 3DBP of load-bearing scaffolds for fulfilling demands of personalized-medicine. We highlight the recent-advances in 3DBP techniques employing thermoplastics and phosphate-cements for load-bearing applications. Finally, we provide an outlook for possible future perspectives of 3DBP for load-bearing orthopedic applications. Overall, the article creates ample foundation for future research, as it gathers the latest and ongoing research that scientists could utilize.


Introduction
A recent study on the global burden of bone fractures suggested that, as many as 178 million instances of new bone fractures were reported in 2019 [1], with a higher prevalence (of new instances per million population) in relatively developed regions (such as Australia and Europe) [1,2].Nevertheless, there is an ever-increasing trend of bone fractures across all regions of the world.In a majority of the cases the cause of bone fracture has been linked to traumatic injuries such as road accidents, occupational accidents and falls rather than pathological conditions [3,4].Risk of bone fracture gets further aggravated due to poor bone health stemming from poor nutrition and lifestyle [5][6][7].Globally, it has been estimated that as many as 200 million individuals are osteoporotic and are at a high risk of bone fracture [8,9].
Given the bleak statistics and worse projections, it is not too much of a stretch to comprehend the need to develop effective medical treatments for rehabilitation of bone injuries.This is especially true for fractures that require surgical intervention [10].Conventional treatments for such patients often involve use of bone fixation plates and screws, typically made of non-biodegradable material such as titanium, stainless steel and zirconia [11][12][13][14][15].Despite their clinical success and high mechanical strength, such materials are typically used as passive mechanical supports at the defect site.Not only these materials prevent further growth of bone, but can also result in complications such as problems associated with implant removal, inflammatory response and stress shielding [16][17][18][19].Moreover, preparing such materials into anatomical shapes matching the defect site poses an incredible engineering challenge [20,21].Modern biomaterials (such as those based on natural as well as synthetic biopolymers and bioceramics) have been extensively researched over the recent years, owing to their high customizability and favorable chemical, mechanical behavior [22][23][24][25][26][27][28][29].
Consequently, additively manufactured scaffolds have emerged as a viable solution to prepare patientspecific scaffolds with superior control of overall shape, internal microarchitecture, and in certain instances the distribution of multiple materials [53][54][55][56][57]. Furthermore, 3D printing (3DP) enables high customizability, while reducing processing times and minimizing waste.Chen et al categorized techniques used for 3DP of ceramics into three major categories (based on the form factor of the feedstock): (a) slurry-based; (b) powder-based; and (c) bulk solidbased [58].The fusion of ceramic particles in all such techniques is typically done either using a carrier phase (such as polymers in fused deposition modelling (FDM), binder jetting) or through application of elevated temperatures (such as through laser beam in selective laser sintering) [59].Furthermore, the 3D printed parts involving polymeric excipients are in a 'green state' and necessitate the sintering of printed parts at high temperatures (typically exceeding 1000 • C) [53].
Conventional ceramic 3DP methods (in particular vat polymerization and binder jetting), have also been extended to enable 3DP of bioceramics, such as HA, and β tricalcium phosphate (β-TCP) [25,53,60].Once again, the harsh processing conditions (high temperature, chemical environment etc) associated with these techniques, severely limiting their applicability for preparation of bone reparative scaffolds.This is since, living tissue and most biomolecules (viz.cytokines) cannot survive such temperatures.
Nevertheless, cell-free techniques while commonly used, fail to show enough sufficient repair capabilities in-vivo.This is especially true for large complex parts, since cells cannot penetrate deep into the implant [68,69].As such, it has been generally accepted by the scientific community that complete functional replication of native tissue (i.e.chemical, mechanical and biological) is crucial towards efficient rehabilitation of damaged tissue [70][71][72].This in turn has prompted the boom in the development of 'bioinks' (i.e.3D printable materials used to carry living tissue) [73][74][75][76][77]. Bioinks can be tailored to deposit living tissue in a spatially controlled manner as well as to mimic the composition of native extra cellular matrix [78,79].Nevertheless, hydrogel based bioinks are mechanically weak and cannot be used as-is for load-bearing applications (such as in orthopedics), besides having poor stability in physiological conditions owing rapid degradation rate.Moreover, cellladen bioinks require very specific processing conditions (in terms of pH, mechanical force, temperature etc), which limits the type of processes and materials that can be used.
As a work around, a few recent studies have reported the development of so-called 'biphasic scaffolds' , which comprise a mechanically strong (or load-bearing) phase printed along with the conventional bioink [80,81], thereby replicating the native environment of bone tissue to a certain degree.The stiff phase in virtually all such scaffolds comprises either low temperature melting thermoplastic biopolymers or phosphate cements, formulated as extrudable pastes.Unlike traditional bioceramics, these cements can be set or cured in mild conditions.Figure 1(a) qualitatively compares the properties of the biphasic scaffolds with bioprinted and acellular 3DP scaffolds.Biphasic scaffolds represent an improved balance between chemical and biological mimicry possible via 3D bioprinting (3DBP); and the mechanical stiffness, print resolution, and scaffold stability observed with acellular 3DP scaffolds.The compressive modulus of biphasic scaffolds can be approximately 1000 times that of the conventional 3DBP scaffolds (as depicted in figure 1(b)).
The present article discusses the current state-ofthe-art of 3D printable load-bearing inks that can be set/cured in mild conditions, thereby supporting 3D bioprinting (3DBP or bioprinting) of stiff bone tissue engineering (BTE) constructs.The authors introduce the different classes of materials used as loadbearing phases and the 3DP technologies that have been employed for bioprinting with these materials.Printing with multiple materials, known as multimaterial 3DP, which offers their own advantages and challenges have also been discussed.Furthermore, the spectrum of reports on load-bearing 3DBP for orthopaedic applications, along with the peculiarities of the load-bearing phases such as calcium or magnesium phosphate inks, and PCL, have been discussed.Next, a comparative analysis highlighting their mechanical properties, curing/setting mechanism, degradation behavior, and osteoinductive behavior has been performed.Finally, based on an extensive review of the literature, the authors present their opinion on the future direction research on 3DBP for load-bearing applications that may take towards clinical translation.

3DBP techniques for load-bearing application
Additive manufacturing (or 3DP) is an umbrella term used for the process of manufacturing 3D object via automated deposition of material in a layer-bylayer manner.Bioprinting then, is a subset of 3DP which typically involves fabrication of 3D cellularized constructs for diverse applications.In this context, bioprinting was defined as, 'the use of material transfer processes for patterning and assembling biologically relevant materials-molecules, cells, tissues, and biodegradable biomaterials-with a prescribed organization to accomplish one or more biological functions.'[83] A key advantage of 3DBP, in comparison to conventional biofabrication techniques, is the ability to fabricate patient-specific implants matching the geometrical complexity of the native tissue, while using patient specific bioinks that may even be derived directly from the host patient [84][85][86][87][88][89][90].In addition, it can replicate the spatial (pore architecture and interconnectivity) and chemical heterogeneity of the native tissue to a certain degree [91].Conventionally, bioprinting techniques involve extrusion, inkjet printing, laser assisted printing and SLA based methods, with each method having their unique set of advantages and challenges.These have been reviewed in detail elsewhere and are beyond the scope of the current article [92][93][94][95][96][97][98][99][100].
Despite their advantages, the conventional bioprinters are typically designed to handle only one type of cell-line and material [27,101].As such, they fail to mimic the native tissue environment, which often comprises of multiple cell types and extracellular components [102].For example, bioprinted orthopaedic constructs (especially for load-bearing applications) require high strength and stiffness, in addition to controlled distribution of cell-lines, growth factors (GFs), biomaterials, etc.
A key tenet in all these techniques is to decouple the contrasting requirements of high strength and cell viability, into two different phases: (i) a bioink phase, responsible for carrying cell-lines, GFs etc, and (ii) a stiff phase, responsible for providing mechanical stiffness to the scaffold.
Among the different techniques, extrusion-based multimaterial bioprinting (as shown in figures 2(a) and (b)) is the most straightforward to implement and can be found in relatively cheap desktop bioprinter models, resulting in their widespread adoption [98,107,108].Extrusion-based bioprinters, however, exhibit slow print speeds and poor print resolution (usually greater than a few hundred micrometers).In contrast, SLA and inkjet-based bioprinters can theoretically exhibit higher bioink print resolutions (usually < 50 µm).Thus far, inkjet or SLA-based multimaterial hybrid bioprinters have been reported to employ melt-extrusion for printing thermoplastic load-bearing inks in conjunction to bioinks printed via inkjet print head (as shown in figures 2(d)-(f)) [105] or SLA photopolymerization (as shown in figure 2(c)) [106], respectively.Table 1 depicts the key features, advantages, and limitations of these multimaterial bioprinting techniques.
Based on the foregoing discussion, 3DBP for loadbearing application may be understood as 'multimaterial 3D printing, with at least one of the printing inks being a load-bearing phase (like ceramic paste ink or thermoplastic), 3D printed alongside other bioink(s), such that the process of printing and setting/curing of the load-bearing phase occurs in mild conditions, without adversely affecting the viability of the printed living tissue and biomolecules.'

3DBP with thermoplastic
Thermoplastic biomaterials, such as poly(lactic-coglycolic acid) (PLGA), PLA and PCL, have long been accepted as promising materials for biomedical application.Their reputation has been further bolstered by their 3D printability and biodegradability.Furthermore, these biomaterials exhibit decent mechanical properties and hence have been employed for engineering load-bearing tissue constructs in the past [63,109].Even so, pristine thermoplastic biopolymers lack adequate bioactivity and the temperatures required for their 3DP can potentially induce cytotoxic effects [63,110].Nevertheless, successful attempts of 3DBP load-bearing implants using thermoplastic polymers as load-bearing phase have been reported in recent literatures [111][112][113]; these have been discussed in the subsequent sections.In essence, 3DBP of load-bearing implants with thermoplastic polymers entails multi-material 3DP of thermoplastic polymer (via melt extrusion), in conjunction with at least one bioink.

Bioprinting thermoplastics via extrusion 3D printing
Among the 3D printable biomaterials, PCL exhibits the lowest melting temperature (∼60 • C) [63].The printing temperature reported in most studies ranges between 80 • C and 100 • C (for PCL with M n 50 000-80 000).Lower print temperature is crucial for 3DBP, to minimize the effect of thermal shock on printed cells.Besides, at low 3DP temperature, the by-product of PCL degradation (caproic acid), is generally considered benign, making it the most preferred material for 3DBP application, by a considerable margin [63,114].Next, low printing temperature (∼120 • C) is exhibited by PLGA [115].
Studies concerning 3DBP for load-bearing application with PCL, started surfacing during the initial years of the previous decade.These early studies involved co-extrusion of PCL with cell-laden bioinks such as those of sodium alginate [116].A key concern in all these studies was to prevent the necrosis of the printed cell-lines from thermal shock, while providing a stiff framework (from PCL) towards efficient tissue rehabilitation.For example, Shim et al reported ∼91% viability of osteoblasts post printing with PCL, which was comparable to the cellular viability in unprinted gels [117].In a separate study, the authors employed an ultrasonic humidifier releasing a stream of aqueous CaCl 2 solution directly on to the scaffold during printing [118].This, the authors claimed, helped in improving shape fidelity of the scaffold, while alleviating the effect of thermal stress.Another study reported the issue of heat build-up in multi-layer scaffolds, such that the viability of chondrocytes (in 4%-6% alginate bioink) was reduced to 85% in multilayer prints, compared to 95% viability observed in single layer scaffold [114].We anticipate that the issue of heat build-up will intensify with increase in complexity and size of the scaffolds.Besides sodium alginate, reports employing bioink of atelocollagen, hyaluronic acid [119], as well as thermo-gelling bioinks based on chitosan and decellularized extra-cellular matrix (dECM) with promising biological response have also been published [120,121].dECM based bioinks are particularly advantageous, since they stimulate specific transduction pathways, thereby improving the in-vivo response of the printed cell-line [120].
Despite the appreciable engineering efforts, thermoplastic polymers like PCL and PLGA are hydrophobic in nature and exhibit poor bioactivity [111,122], as seen from the virtual absence of the adhesion of osteoprogenitor cell-lines on PCL filaments, even after 25 d of incubation (figure 6(c), (i)-(viii)).Furthermore, degradation time for PCL is very slow  (typically 2-4 years) [110].As a workaround, hydrophilic and/or bioactive ingredients have been incorporated in PCL filaments [123][124][125].For example, Kim et al printed a mixture of alginate and PCL [111].The resultant filaments exhibited improved hydrophilicity and porosity (from leached out alginate) without compromising its mechanical properties.
In other studies, incorporation of HA or tri-calcium phosphate (TCP) in PCL demonstrated improved bioactivity and faster degradation rate of the 3DP filaments [98,126].HA has been well established for its bioactivity.For example, use of HA in PLGA ink and alginate bioink based biphasic scaffolds resulted in distinct upregulation of osteogenic biomarkers (such as osteopontin and alkaline phosphatase (ALP)) during in vitro studies [115].
Although, load-bearing 3DBP implants mimic the native bone tissue to a certain extent, complete reproduction of the complex cellular and extracellular environment is not possible from current 3DBP technologies [127].As a stopgap solution, Daly et al reported an interesting strategy for engineering whole organ by bioprinting 'developmentally inspired templates' [128].This in case of bone involved 3DPB of stiff scaffold with chondrogenically primed MSCs (through action of RGDγ irradiated alginate), towards simulating hypertrophic cartilage, such that, these templates could inspire endochondral ossification in vivo.In fact, the samples harvested after 12 weeks of implantation in mice models exhibited significant mineralization and vascularization.
Furthermore, GFs, such as bone morphogenic protein-2 (BMP2) and vascular endothelial growth factor (VEGF), are often incorporated in bioinks to augment the natural bone tissue regeneration cascade [119][120][121].However, sustained release of these GFs is both of paramount importance and a considerable challenge, since burst release of GFs is associated with inflammation and impaired tissue regeneration [129,130].In an interesting attempt to overcome this, Cunniffe et al co-printed RGD-γ irradiated alginate (loaded with bone marrow-derived mesenchymal stem cells, BMSCs) with PCL for orthopaedic application [106].The authors, additionally, incorporated complexes of nano HA and plasmid DNA encoded for BMP2 and transforming GF beta-3 (TGF-β3), towards transfection of the printed MSCs.This resulted in sustained release of GFs at the implantation site, in-turn manifesting as enhanced osteogenesis in vitro, and upregulated neovascularization and mineralization in vivo.

Bioprinting thermoplastics via hybrid 3D printing
Extrusion-based 3DBP has significant limitations, including high shear stresses, low cellular densities, and poor print resolution [92,96,131].While 3DP of thermoplastic biopolymers may not be feasible through alternative methods, 3D printing bioink via laser-assisted or inkjet-based techniques may still be viable.This rationale has driven the emergence of 'hybrid bioprinters,' which enable multi-material 3DP by combining extrusion methods with SLA and/or inkjet-based 3DBP.
Daly et al developed a novel approach using inkjet bioprinting to rehabilitate osteochondral knee defects with high resolution and cellular density [105].They 3D printed PCL microchambers with non-porous side walls, followed by the deposition of MSC-laden gelatin methacrylate bioink using extrusion bioprinting, and MSC: chondrocyte co-culture via inkjet bioprinting (shown in figures 2(e) and (f)).The vertically arranged microchambers guided tissue growth and fusion.However, this method lacks full 3D print control and porosity in the z-axis, limiting nutrient exchange in anatomically sized implants.Despite this, the hybrid 3DBP and use of a dynamic bioreactor led to higher expression of collagenous biomarkers and improved tissue growth rate during 28 d of in vitro culture.
Similar to inkjet bioprinting, bioinks for SLA 3DBP typically have moderate to low viscosity.Shanjani et al developed the 'Hybprinter,' a hybrid printer that mechanically reinforced bioinks with PCL filaments [104].This printer combined the capabilities of a top-down resin 3D printer with a multiple head extrusion 3D printer (shown in figure 2(c)).The advantage of such printers is the expanded library of 3DP biomaterials and bioinks available for both SLA and extrusion bioprinting.In their study, the authors used PCL as a load-bearing phase and photocurable PEGDA (polyethylene glycol di-acrylate) based bioink cured using SLA, in addition to sacrificial Pluronic ink for printing complex features like horizontal cylindrical channels.The system achieved a z-axis resolution of ∼35 µm with a cellular viability of ∼90%.The lumen-like horizontal channel allowed rapid perfusion within the scaffold, essential for vascularization and fluid exchange in the implant.
Hybrid printers hold great potential for the future development of bioprinters, offering diverse capabilities.However, their current development is in early stages, with only proof-of-concept models reported thus far.Issues like complex printer design, compatibility between 3D printable materials, and slow print speed need addressing before hybrid 3DBP, especially for load-bearing applications, can gain broader acceptance.

3DBP with phosphate cements
Phosphate cements in their modern form were first reported in 1986, as water setting calcium phosphatebased cements [132].This cement was based on the solution precipitation of apatitic phases from water soluble calcium phosphates (CaP), such as tetracalcium phosphate (TTCP), and dicalcium phosphate (DCPA).The 'green' (i.e.uncured or partially cured) cement pastes can be injected in the wound site, where they harden into a ceramic solid through a hydraulic setting reaction, thereby eliminating the need for prefabrication of implants [133][134][135].Not only, calcium phosphate cements (CPCs) can set at body temperature, they also exhibit excellent injectability, mechanical properties, and bioactivity [29,[136][137][138].Given their early success and similarity (chemical and physical) to the mineral phase in native bone, CaPs have been the most studied type of phosphate cements.Nevertheless, other types of phosphate cements, namely: magnesium phosphate (MgP) [139,140] and zinc phosphate (ZnP) [141,142] based injectable cements have been developed towards targeting varied applications.It is to be clarified here that phosphate cements differ from acrylic bone cements (ABC), which are based on polymeric materials and cure via polymerization reaction [143][144][145][146]. Furthermore, ABCs are typically nonresorbable and inhibit further bone ingrowth.
According to the classical solubility data, depending upon the pH value of a cement paste, after setting, all CPCs can only form two major end-products: calcium deficient HA (CDHA, a poorly crystalline form of HA) at pH > 4.2 and dicalcium phosphate dihydrate (DCPD, also called 'brushite') at pH < 4.2 [147][148][149].Tan et al first applied a cement reaction between DCPA and H 3 PO 4 solution to 3D print ceramic scaffolds via binder jetting technique [150].Even though, bioceramic bone substitutes could be fabricated at room temperature, the printing process itself was carried out in a strongly acidic environment, adversely affecting the performance of the developed scaffold.In contrast, apatite-based CPC typically involve acid-base reaction between the reactants.For example, basic TTCP reacts with slightly acidic DCPA in aqueous suspension towards precipitation of CDHA as described in the equation below (equation ( 1)) [132,151], (1) In fact, such phosphate based injectable cements have been able to receive regulatory approval for clinical use in rehabilitation of skeletal and craniofacial tissues [152,153], with a few of them being available under the brand names of HydroSet™, Norian SRS and Biobon [154][155][156].
MgP is another class of cements that have emerged as a potential substitute to CPCs owing to their higher mechanical strength and relatively faster curing and degradation rate in vivo [140,157].Kanter et al reported a comparative study between struvite, brushite and CDHA based cements in the femoral condyle of adult sheep for a period of 10 months [158].It was observed that the MgP based struvite cement exhibited greater resorption and a correspondingly higher new bone growth than either of the two CaP cements, as shown in figure 3. MgP cements typically involve a solid magnesium precursor (such as MgO and tri-magnesium phosphate anhydrate, TMPA, Mg 3 (PO 4 ) 2 ) cured using a soluble phosphate salt (such as diammonium hydrogen phosphate, DAHP and dipotassium hydrogen phosphate, DKHP), to yield an insoluble MgP.For example, TMPA salt reacts with DAHP to form a combination of newberyite (MgHPO 4 • 3H 2 O), struvite (MgNH 4 PO 4 • 6H 2 O) as depicted in the equation below (equation ( 2)) [159], Clearly, phosphate cements have immense potential in BTE.However, their general acceptance is severely restricted by the lack of adequate porosity.Moreso, injected 'green' cement pastes are susceptible to wash-out and disintegration in vivo, which may lead to severe post operative complications [160,161].Although, 3D printed (preformed) scaffolds can address the issues of porosity and cement washout, they still cannot be used to carry living tissue and/or biomolecules (viz.GFs).This is since, the typical workflow for fabrication of such scaffolds involves a sintering/curing stage at temperatures exceeding 1000 • C. For further reading on 3D printing of bioceramics, we would refer to the reader other excellent reviews in the field [29,53,162].Another challenge with 3D printing such scaffolds is to enable sufficient 'printability window' .Water-based carriers typically used for 3DP of phosphate cement also initiate its crosslinking, thereby altering their rheological behavior, thereby adversely affecting the printability and part reproducibility.Some studies attempt to address these issues, by incorporating crosslink retarders, or by replacing aqueous carriers with lipophilic ones [108,163].
A few studies in the recent years have emerged, that focus on printing high stiffness phosphate cement-based scaffolds in mild conditions such that they (a) can carry biomolecules and (b) can be printed alongside living tissue [80,81].Such engineering solutions have the potential to replicate the native bone tissue in a wholistic manner and rehabilitate bone tissue efficiently.At this stage, the authors would like to clarify that owing to the high stiffness of cured phosphate cements, employing the cement paste as a carrier of cell lines may not be advisable.Figure 4 depicts the simplified workflow for fabrication and implantation of phosphate cementbased 3D printed scaffolds for BTE, along with the key feature of phosphate cement ink and cellladen bioinks used in such techniques.The subsequent sections discuss in detail two important classes of phosphate-based cements that are being developed to enable 3DBP of such high stiffness BTE scaffolds.

Bioprinting with MgP cements
Struvite phase is one of the end products of the MgP cement setting reaction (see equation ( 2)), and is the phase of primary interest owing to its significant bioactivity and biodegradability by both chemical dissolution and osteoclastic activity in the body, analogous to biodegradable CaP [164].Lee et al were the first to develop the procedure to 3D print and cure MgP cement-based inks at room temperature in 2014 [165,166].In their pilot study, the group used a farringtonite (or TMPA) paste in 1% w/v hydroxypropyl methyl cellulose (HPMC) and ethanol solution at a powder to liquid ratio (P/L) of 1.5.Ethanol in this study was used to improve the printability window of the paste.The prepared composition displayed adequate printability using both 19-and 23-gauge nozzles.The cementation reaction was carried out by immersion of green scaffold in DAHP solution post 3D printing.Furthermore, the authors could demonstrate that the scaffolds could be effectively used as carriers for drugs and other biomolecules owing to their room temperature cementation, such that the scaffolds maintained a stable drug release rate for over a period of 3 weeks.
As for their biological activity, the MgP scaffold displayed good cell affinity and consequently promoted good cell proliferation, growth, and migration, in addition to upregulation of biomarkers such as collagen-I, ALP, and osteocalcin (OC).In a follow up study, the group incorporated gelatin in the binder solution, to improve the scaffold's mechanical properties [167].At an optimized gelatin loading (i.e.6% w/v) the scaffolds exhibited a compressive strength of 16.7 MPa and a modulus of 56.1 MPa (indicating an improvement of 135% and 45%, respectively).However, higher gelatin content resulted in faster drug release, indicating a certain degree of tailor-ability in the overall properties of the bone scaffolds.
An alternative to conventional MgP cements is to use partially substituted magnesium phosphates [140].This is since, previous studies with calcium substituted tri-magnesium phosphates (CSTM, Ca x Mg (3−x) (PO 4 ) 2 ), have revealed that these magnesium/calcium phosphate (MgCaP) cements can exhibit better mechanical properties [168], in addition to being more osteogenic than either of the MgP and CaP cements when used alone [140,168,169].The equation (equation ( 3)) depicts the general cementing reaction of MgCaP cements, which results in the formation of a combination of both calcium (brushite) and magnesium phosphate (struvite and newberyite) phases, As such, Götz et al replaced the TMPA from previous studies with CSTM (Ca 0.75 Mg 2.25 (PO 4 ) 2 and MgO mixture along with 10%-14% w/v HPMC aqueous solution to prepare printable inks [170,171].It was observed that 10% HPMC samples resulted in good printability of cement pastes and a mechanical strength of ∼3 MPa in microporous samples (using 21 G needle).Furthermore, MgCaP cement demonstrated excellent printability, good extrudate stability (spans greater than 5 mm), and print window >2 h.While this is not explicitly discussed by the authors, the cement pastes seem to exhibit an inverse logarithmic relationship between injection pressure and aging time.The authors attribute the reduced printability with time to (a) structural changes in HPMC chain structure; and (b) hydration of some of MgO to Mg(OH) 2 .Even though the raw materials involved in the study are biocompatible, the authors fail to show any direct biocompatibility and/or bioactivity of the 3D printed paste.Nevertheless, this method demonstrated a viable alternative to the more conventional MgP-based 3D printable cements with improved mechanical and degradation characteristics.Owing to their room temperature cementation, the capabilities of the current MgP and MgCaP 3D printable cements can be extended in future to support printing alongside living tissue, as well as patient derived components, such as blood or its constituents.

Bioprinting with CaP cements
CaP cements or CPC rapidly set in the presence of water, posing challenges for 3D extrusion printing of aqueous CPC paste, resulting in a brief printability window and inhomogeneity [53,162].To address this, Lode et al reported 3D printing of CPC using non-aqueous yet water-miscible liquid, like shortchain triglycerides, at a powder-to-liquid ratio (P/L) of 2.5 mg ml −1 [172,173].The workflow involved 3DP of CPC-oil paste, followed by incubation in a water bath.The setting reaction of CPC involves the initial dissolution of α-TCP, CaCO 3 , and DCPA, followed by the formation/precipitation of carbonated CDHA [174] (equations ( 4) and ( 5)).The resulting fine-grained structure enhances the bioactivity of calcium phosphates [41,175], Room temperature processing avoids shrinkage issues associated with sintered bioceramics [58,[176][177][178].The scaffolds exhibited increased microporosity due to voids created by the removal of the oil phase during incubation in an aqueous environment.Complex surfaces resulting from this process can significantly improve cellular responses to 3DP scaffolds as cellular attachment to biomaterials is influenced primarily by surface topography [179][180][181].Subsequent studies using similar materials demonstrated their capability to act as carriers for biomolecules, such as GFs and anti-cancer drugs, while maintaining a stable release rate [82,182,183].

Multimaterial co-extrusion bioprinting
Ahlfeld et al advanced the technology for preparing stiff 3D bioprinted biphasic BTE scaffolds [80].They utilized multi-material co-extrusion printing of oil-based CPC inks and alginate bioink carrying mesenchymal stem cells (MSC), as depicted in figure 5(a).The presence of two phases during printing did not significantly affect the behavior of the individual inks.The scaffold was cured in two stages: (a) incubation in high humidity to cure the CPC surface, and (b) incubation in CaCl 2 solution to cure the alginate bioink and CPC bulk.This two-step regime was designed to prevent microcrack formation in the CPC phase.Directly immersing the CPC in a water bath for curing led to residual stress and crack formation during the setting reaction.On the other hand, initial exposure to humidity for more than 30 min resulted in gradual precipitation of the CaP.The localized curing on the exposed surface fixed the initial shape of the printed scaffold.Further incubation in an aqueous bath completed the curing process with increased moisture penetration, reducing residual stress and resulting in absence of surface microcracks and significantly increased compressive strength of the printed scaffolds.Overall, the dual-phase scaffolds exhibited a compressive modulus greater than 30 MPa and a compressive strength of 1.3 MPa, both considerably higher than that of conventional bioprinted scaffolds.The in-vitro cytocompatibility of the biphasic scaffolds revealed that the CPC ink did not cause toxicity (due to oil carrier release or setting reaction) in most of the bioink.However, a significant amount of cell death was observed at the interface of the two inks after one day (figure 6(a)).This was attributed to the hydrogen ion release during setting reaction (equation ( 4)).Long-term studies showed improved cytocompatibility with considerable cell growth and subsequent migration and proliferation on the CPC phase after 21 d of incubation (figures 6(a) and (b)).An optimized duration of initial scaffold incubation (<30 min) was found to strike a good balance between cell proliferation and scaffold strength.Due to the poor printability of alginate ink itself, pore closure of printed samples occurred due to the collapse of unsupported struts, potentially hindering cellular movement.Choosing a better bioink, such as a higher stiffness hydrogel, may enhance the in-vitro behavior of 3D bioprinted scaffolds by allowing printing of larger pores, minimizing interfaces between CPC and bioink, and correspondingly increasing cellular survivability during the initial phase of scaffold incubation.Moreover, the work was extended by developing patient-specific bioinks, incorporating blood plasma directly sourced from the patient [108].Delivery of immunoregulatory proteins to the defect site through this approach is expected to elicit a favorable immune response and enhance the healing capability of the damaged tissue.

Co-axial extrusion bioprinting
Despite the success of multi-material co-extrusion bioprinting, print speed remains slow due to repetitive print-head changes.To address this, Raja and Yun et al developed BTE scaffolds using a coaxial nozzle capable of depositing two distinct inks as a core and shell (shown in figure 5(b)) [81].The core was a blend of αTCP in 1% HPMC solution in ethanol (P/L = 1.67 g ml −1 ), while the shell consisted of partially crosslinked alginate bioink carrying preosteoblast bone cells with varying thickness.The bioink phase was first crosslinked in CaCl 2 solution and then immersed in PBS solution to set the ceramic phase (figure 5(b)).The use of hydrogel stabilized the CPC core, resulting in better mechanical properties compared to either brittle ceramics or the hydrogel bioink used alone.Although ethanol-based formulations are generally discouraged for bioprinting, a 30% ethanol solution was used as a cementation retarder in this study to allow sufficient printability.The samples maintained adequate cellular viability during the 5-week testing period.Coaxial printing offers advantages over multi-nozzle printing, including reduced printing time and stabilization of CPC filaments, resulting in higher mechanical properties.

COBICS bioprinting
Despite significant progress in 3D printed implants, many still involve stacking planar layers, limiting the printable shapes.Romanazzo et al introduced 'COBICS' (ceramic omnidirectional bioprinting in cell-suspension) for 3DBP freeform, biomimetic orthopedic constructs [184].Similar to FRESH bioprinting, they successfully printed α-TCP-based ceramic ink in a cell-laden gelatin slurry, enabling complex, load-bearing structures, as shown in figure 5(c).The printed CaP ink had minimal impact on cellular viability during setting and induced osteogenic differentiation of MSCs.Though not traditional bioprinting, the process showed favorable cellular distribution, suggesting 'COBICS' could be adapted for 3DBP of load-bearing scaffolds in the future.

Critical evaluation of load-bearing biomaterial for 3DBP
A key tenet in tissue engineering is the need for biodegradability of the engineered biomaterial.Biodegradation is critical as this allows for space into which neovascularization and subsequent bone development may occur [185][186][187].Although, there seems to be some contention regarding the ideal time required for degradation of biomaterial, it is generally accepted that rate of resorption/degradation should match the rate of bone growth [180,188,189].The bone growth rate in-turn can be controlled, to a certain degree, by the osteoinductive effect of the biomaterial [29,190].
CaP based 3DP inks, owing to their similarity to the mineral phase of native bone, have been studied relatively more than their magnesium-based counterparts for development of load-bearing 3DBP scaffolds.Biodegradation of CaP inks occur through both (a) solution driven dissolution and (b) osteoclast mediated resorption (as seen in figure 7(a)) [162,189,197].Among the two prominent phases (brushite and CDHA) found in cured CPC, apatitic phases are thermodynamically the most stable phases at physiological pH.As such, their resorption occurs primarily through osteoclastic activity [189,198].Even though, such resorption is relatively much slower, the osteoclastic activity results in formation of unique 'Howship's Lacunae' .Owing to their favorable topography, these resorption sites can act as hotspots for de novo bone formation [199,200].Furthermore, ionic substitution in CPC have been shown to control the osteoclastic activity, viz.Sr 2+ ion inhibits the activity of osteoclasts by suppressing the matrix metalloproteinase (MMP 1 and 2) secretion [190,[201][202][203].The authors claimed that the reduced resorption aided in enhancing the overall bone rehabilitation rate.Alternatively, CO 2− 3 ion substituted apatite show faster degradation potential, thereby accelerating the resorption process [48,204,205].Incidentally, 3D printable CaP ink reported by Innotere group, includes a finite amount of calcium carbonate, thereby alleviating the issue of slow degradation rate of CDHA [174].
Besides providing topographical cues, CaP have been shown to improve the crosstalk between osteoclasts and osteoblasts.For example, CaP substrates could upregulate the expression of coupling factors in osteoclasts (viz.C3A, Cthrc1, and Ephrin B2), these in-turn enable improved osteoblast functionalization and differentiation [206,207].Moreover, the Ca 2+ and (PO 4 ) 3− ions released from biodegradation of CaP cements modulate the cellular response of the neighboring tissue.For example, (PO 4 ) 3− ions promote osteoclastic differentiation by increasing RANK and RANKL binding via increased NF-Kβ signal transduction, thereby enhancing resorption rate and subsequent bone remodeling [40].
CaP have also been reported in some studies to favorably modulate the immune response.For example, Ca 2+ ions favor the anti-inflammatory M2 polarization of macrophages (via calcium-sensing receptor (CaSR) pathway), while simultaneously inhibiting the activity of M1 polarized (pro inflammatory) macrophages [208,209].Consequently, the CaP degrade via dissolution, leading to release of small fragments which may be phagocytosed by macrophages, while bulk particles are resorbed by osteoclasts.Also, dicalcium phosphate dihydrate (DCPD) may convert to more stable octa calcium phosphate (OCP).Reproduced from [189].CC BY 4.0.(b) shows the simplified degradation modes observed in PCL i.e. surface erosion, bulk degradation, and bulk degradation with autocatalysis.(c) Represents the in vivo degradation of PCL, showing hydrolysis intermediates 6-hydroxyl caproic acid and acetyl coenzyme A, these are then eliminated from the system via the citric acid cycle.Reprinted from [110], Copyright (2010), with permission from Elsevier.secretion of GFs (BMP2, VEGF, TFG-β1, PDGD etc) and anti-inflammatory genes (IL-10 and IL-1α) is upregulated and that of pro-inflammatory molecules like TNF-α, IL-6 and IL-1β is downregulated.These conditions favor recruitment of MSCs and neovascularization at the implantation site [210][211][212].
Ca 2+ ions (from CaP resorption) further aid osteoblastic activity in a dose-dependent manner.Not only optimum levels of Ca 2+ ions favor differentiation of MCSs in osteogenic lineage [193], it also upregulates the synthesis of collagen type 1 (COL-1) and other non-collagenous bone proteins like bone sialoprotein (BSP) and osteopontin [180,195].
Another promising candidate for load-bearing material in 3DBP are MgP and MgCaP based 3D printable inks.Even though, their suitability for 3DP along with cell-laden inks is yet to be established, these inks exhibit superior mechanical strength, osteoinductive behavior, and faster degradation rate compared to CaP cements [158,213].Much like CaP based inks, Mg based 3DP inks are bioresorbable and orchestrate the bone regeneration through a range of topographical and chemical cues.For example, 3DP MgP implants have a fair degree of porosity and consequently a topographically complex surface.These surfaces facilitate adsorption of non-collagenous bone proteins, thereby facilitating de novo bone formation over the ceramic implant and enhanced bone bonding [180,214].
In addition to topographical cues, MgP cements also stimulate osteogenic activity in the similar manner as CaP.For example, Mg 2+ ions induce specific signaling pathways that not only stimulate differentiation of stem cells [140,194], but also upregulates production of osteogenic proteins such as COL10A1 and VEGF [212]; in addition to modulating the immune response to the implanted ceramic scaffold (viz.M2 polarization of macrophages) [215].
The key difference, however, with MgP based implants is the absence of osteoclast mediated resorption.That is, the resorption in these scaffolds occurs exclusively via dissolution of ceramic in physiological media [214].In fact, Mg 2+ ions released from the degrading scaffolds inhibit osteoclastic differentiation and activation [216].Besides, Mg 2+ ions have been reported to facilitate adhesion and motility of osteoblasts, owing to a higher binding affinity with αintegrin subunits [217,218].
Overall, MgP cements have been reported to exhibit relatively higher degradation rate and faster bone growth than CaP cements.Towards comparing their relative rehabilitative potential, Kanter et al [158,213] implanted MgP and CaP cements (specifically struvite, brushite, and CDHA cements) in both: partially load-bearing and non-load-bearing defects in large animal models (merino sheep).At the end of 10 month implantation period, not only MgP scaffold was completely degraded; it was fully replaced with new bone.In contrast, CaP although demonstrated some resorbability in vivo, the degradation rate and therefore bone regeneration was much slower.Figure 3 compares the results obtained from different biomaterials during the study.
Despite the distinct advantages of MgP based cements, MgP chemistry-based inks are to be used successfully for 3DBP.The primary obstacle towards this is the use of DAHP as a curing agent towards formation of struvite phase [140,167].Nevertheless, a few studies have reported substitution of DAHP with KH 2 PO 4 to form 'K-struvite' (MgKPO 4 .6H 2 O) [219][220][221], while preserving the key properties of struvite (i.e. initial strength and setting rate).However, 3D printable inks based on K-struvite are yet to be established.Table 2 summarizes the key features of the load-bearing inks for 3DBP proposed thus far.
The only other class of biomaterial that has been reported in literature for 3DBP of load-bearing orthopaedic scaffolds are 3D printable thermoplastics; in particular PCL [63], with only a handful of exceptions where the authors attempted to use PLA or PLGA based inks with limited results [115,222].The technology for 3DBP with thermoplastics has been adopted from the conventional FDM based processes, wherein PCL (in the form of filament or pellets) is extruded as thin strands at temperatures ranging between 70 • C-100 • C depending on the polymer's weight average molecular weight (M w) .In general, the initial reports pertaining to bioprinting for loadbearing scaffolds involved extensive use of 3D printed PCL [114,[116][117][118].However, the high heat involved in the printing process is of significant concern, as it tends to adversely affect the cellular viability.The issue of heat built-up must be more significant for anatomically sized, larger scaffolds.Nevertheless, through optimization of printing parameters, cellular viability >90% freshly prepared 3DP implants has been reported [111,114,223].
Unlike ceramic biomaterials, PCL is not biodegradable in vertebrate animals.Nevertheless, it can be resorbed post degradation.However, degradation of PCL is among the slowest, occurring exclusively through hydrolytic degradation of poly(α-hydroxy) esters into caproic acid, at the surface and/or bulk of the implant (figure 7(c)) [110].The hydrolysis intermediates 6-hydroxyl caproic acid and acetyl coenzyme A, are then eliminated from the system via the citric acid cycle [234].Although, studies involving small specimens have claimed good cytocompatibility of PCL and its degradation by-products, the acidic nature of these degradation by-products can be of significant concern.Crucially, in case of implants having low perfusability, the released oligomers and acidic by-products have been shown to build-up locally, resulting in chronic inflammatory response in vivo [110,235].In fact, the local acidic gradients tend to accelerate degradation and result in autocatalysis.Figure 7(b) depicts the modes of degradation observed in PCL-based implants.Such implants typically degrade over a time span of 2-4 years [225].Even though, PCL based scaffolds provide good strength to 3DP scaffolds for load-bearing applications, neither PCL nor its by-products are osteogenic in nature.Furthermore, the absence of topographical complexity and inherent hydrophobicity of PCL, result in poor osteointegration of the scaffold [63,236].
For example, figure 6(c) depicts in vitro response of a biphasic scaffold (PCL ink and MC3T3-E1 cell laden alginate bioink) after 25 d of incubation.Most of the cellular activity is restricted in the vicinity of the alginate bioink phase, clearly indicating the aversion of the cellular tissue to the PCL phase.This contrasts with the in vitro response of biphasic (CaP based ink and MSC laden alginate bioink) scaffold (as shown in figures 6(a) and (b)).Herein, not only, the loadbearing CaP phase exhibits high biocompatibility, the MSC can be seen to actively migrate from bioink phase on to the ceramic phase after 21 d of incubation,  [140,165,168,232,233] in addition to exhibiting markers for osteogenic differentiation [80].Nevertheless, biological response of PCL phase can be improved by incorporation of bioactive particles (such as bioceramics) [98,126,235,237] and hydrophilic polymers (such as alginate, gelatin) [111,125,236,238] which aid in improving its bioactivity and degradation characteristics.Based on the foregoing discussion, figure 8 summarizes the various types of materials (load-bearing biomaterials ink and bioinks) and bioprinting techniques reported in literature towards development of load-bearing 3DBP scaffolds for orthopaedic applications.Overall, while individual phases such as phospahate cements (HydroSet™, Norian SRS and Biobon for dental fillings [154][155][156]) and PCL (for medical devices such as sutures) have received regulatory approval and are available for clinical use, we are yet to witness clinical translation of bioprinted scaffolds or organelles.Similarly, acellular 3D printed scaffolds for BTE application have completed pre-clinical trials successfully and are currently undergoing clinical trials [239,240].Yet, they are not available for clinical use to date.

Bioinspired biomineralization of collagen-based high stiffness hydrogel
Invertebrate bone is a highly optimized structure, exhibiting remarkably high specific fracture toughness and stiffness, despite the use of relatively brittle HA [241].The significant role played by the staggered arrangement of mineral phases within the collagen fibrils contributes to these properties [242].
Naturally, for development of effective BTE scaffolds, it is worthwhile to examine the mechanism involved in bone biomineralization.In native bone, the mineralization processes are dominated by endochondral ossification [243].While collagen matrix acts as a template for mineral deposition, it is widely accepted that non-collagenous proteins (NCPs), including glycoproteins and proteoglycans play a crucial role in regulating biomineralization process.In addition to directing the self-assembly of mineral phase at the gap region within the collagen fibrils [244].NCP are highly negatively charged, and form protein-mineral complex with amorphous calcium phosphate (ACP).These negatively charged complexes are advantageous since (i) it prevents HA precipitation within the interstitial sites, and (ii) provide the electrostatic potential for intrafibrillar infiltration of ACP towards the positively charged Nterminus in the gap region of collagen [245].Figure 9 represents graphically the mechanism, along with TEM images of unmineralized and fully mineralized collagen fibril.
While these models for collagen mineralization are widely accepted, only a handful studies have attempted to employ biomimetic biomineralization to develop BTE scaffolds with superior mechanical properties.For example, Thrivikraman et al reported the development of cell-laden bone scaffolds with biomimetic intrafibrillar collagen mineralization [246].Herein, the authors reported the formation of nanoscale HA in the interstices of collagen fibrils, using bone derived osteopontin as an NCP analog.Crucially, the mineralization process did not adversely affect the viability of encapsulated MSCs and resulted in a localized elastic modulus of approximately 20 GPa (as measured from atomic force microscopy).The high stiffness of mineralized matrix could stimulate the expression of osteogenic and preosteocytic markers.Despite these promising observations, the scaffold exhibits poor bulk mechanical properties, characteristic of hydrogel-based scaffolds.Although, these scaffolds may not be applicable for load-bearing BTE application in their current state, we believe biomimetic biomineralization can be a promising avenue for future development of loadbearing BTE constructs.

Future perspective
New bone formation in mammalian specimen is a dynamic process, which entails recruitment of MSCs to the target site followed by their differentiation to osteoblasts.These osteoblasts secrete soft polymeric matrix (primarily comprising of collagen-I) to form osteoid on top of old bone, followed by their calcification (with minerals such as HA) to form new bone [181,247,248].However, fabrication of synthetic BTE scaffolds of such complexity is not possible through existing fabrication technology.Nevertheless, functional rehabilitation of damaged tissue can be achieved by use of bone scaffolds that can mimic the native bone tissue in a wholistic manner (i.e.chemically, mechanically, and biologically).One possible solution that has emerged over the years is the use of multiphasic 3D bioprinted scaffolds that comprise of at-least one mineral phase printed along with an extracellular matrix (ECM) like, cellladen bioink.
Early attempts in developing such load-bearing scaffolds, almost exclusively employed thermoplastics such as PCL, perhaps because 3DP with these materials is reasonably well established.Primary aim of most such studies was to prevent the adverse effects of thermal shock on the bioprinted cell-lines.Indeed, through engineering ingenuity cell viability >90% was reported in most instances.Even so, PCL is inherently hydrophobic and relatively bioinert, as such, it does not contribute significantly to the tissue regeneration, apart from providing mechanical integrity to the scaffold.Furthermore, the resorption rate of PCL is extremely slow, which obstructs further bone growth in the long run.Although, incorporation of bioceramics or hydrophilic polymers improved bioactivity and degradation rate, in the absence of long-term studies with such scaffolds conclusive evidence is missing.Nevertheless, a few interesting studies using PCL based 3DBP scaffolds were reported, these could be used to inspire further research in BTE.For example, successful transfection of MSCs were done towards sustained release of GFs, thereby, overcoming side-effects arising from their burst release.Elsewhere, 'developmentally inspired templates' were developed to overcome the limitations of current 3DP techniques.
In contrast, bioceramic based inks show favorable degradation behavior and superior bioactivity.However, 3DBP with bioceramics posits unique challenges, such as need for setting/curing of high stiffness bioceramic phase in mild conditions, without adversely affecting the nearby cellular tissues.Additionally, the phosphate cement inks must allow sufficient printability window.The current article discussed in some detail the various research undertakings targeting this issue.Yet, the technology is in early developmental stages and certain challenges remain.These will need to be resolved before clinical trials can be attempted.
For example, MgP and MgCaP based 3D printable inks show immense promise as bioceramic phase owing to their superior mechanical and osteoinductive properties post-cementation.These cements can be set at room temperature, exhibit adequate mechanical strength and have been reported as effective carriers for biomolecules.However, virtually all the reports involving 3D printing of such scaffolds, require incubation in diammonium hydrogen phosphate (DAHP).This can be problematic, as DAHP adversely affects cell survivability.As a workaround, we hypothesize (a) substitution of DAHP with KH 2 PO 4 solution; or (b) shielding the printed cell lines (from the adverse effect of cementation reaction) by use of a secondary carrier vehicle, such as fibrin gel.Although, such techniques have been never been employed in 3D printing of bioceramics, Song et al successfully demonstrated that human bone marrow stromal cell (hBMSC) laden fibrin microfibers could shield the cell lines from the harsh processing conditions and setting reaction pertaining to phosphate cements [250].
3D printing with CaP cements, in contrast, have different challenges, for example, aqueous media cannot be reliably used as a carrier medium owing to their short print window.As such, a few studies involving oil-based CPC inks have been reported with excellent printability alongside cell-laden bioinks.While these studies demonstrate good cytocompatibility and excellent cellular proliferation during in-vitro cultures, no study, to date, has been able to demonstrate the regenerative capability of such biphasic scaffolds.Ahlfeld et al in a recent study attempted to test the in-vivo repair capabilities of biphasic CPC scaffold in craniofacial defects in Lewis rats with promising results (see figure 10) [80].However, the authors failed to 3D print parts small enough using multichannel printing owing to the resolution limits of the printing technique.A possible alternate to such issue may be the use of multiaxial core/shell printing, as reported elsewhere by Raja and Yun [81].Nevertheless, issue pertaining to low print resolution with multi-material extrusion 3D printing remains at large and posits a significant challenge in 3DBP of complex, anatomically shaped parts.
Finally, there is the issue of mechanical properties, that is, even though the biphasic 3D bioprinted scaffolds exhibit compressive strength much higher than any conventional bioprinted samples, they fall short in matching the mechanical properties of bone.Furthermore, most of the studies thus far, have solely focused on developing material parameters towards ensuring 3D bioprintability and cell survivability.Future studies with such materials may benefit significantly by employing bio-mimicked scaffold microarchitecture towards enhancing cellular response and optimizing the load-distribution within the scaffold.

Conclusion
3DBP is widely acknowledged within the biomedical engineering community for its superior regenerative capabilities.However, poor mechanical properties and stability of 3DBP scaffolds are significant drawbacks, representing the primary challenges in their translational research.Among the available fabrication techniques, load-bearing 3DBP scaffolds have emerged as a viable solution towards biomimicking the heterogeneity of native tissue, while improving the mechanical properties of tissue engineering constructs for orthopaedic applications.3DBP for loadbearing application may be understood as, 'multimaterial 3D printing, with at least one of the printing inks being a load-bearing phase, 3D printed alongside other bioink(s), such that the process of printing and setting/curing of the load-bearing phase occurs in mild conditions, without adversely affecting the viability of the printed living tissue and biomolecules.'The strict material and processing requirements severely limit the type of load-bearing biomaterial inks that can be used.Thus far, only thermoplastic and phosphate cement-based inks have been reported as viable candidates, with latter being preferred in the recent years due to their osteogenicity, high stiffness and superior degradation behavior.Nevertheless, the field is in its early developmental stages and challenges, such as poor print resolution; limited materials library; and sub-par mechanical strength need to be addressed before clinical trials can be attempted.Laboratories with highly multidisciplinary skills and establishment of regulatory protocols will be crucial for translation of such 3D bioprinted load-bearing scaffolds into clinical practice.By means of the current review, the authors aim to consolidate the pertinent information available in diverse domains, such that it can be used as a foundation to support future research in the field.Overall, we believe that load-bearing 3DBP has enough potential to address the unmet challenges of personalized medicine towards efficient bone rehabilitation.With time, it is anticipated that the technology will mature enough to enable repair of bone tissue defects as well as whole bone replacement in a clinical setting.

Figure 2 .
Figure 2. (a) Depicts graphically the construction of the multimaterial 3D bioprinter used for to print PCL via melt-extrusion along with three bioinks.(b) Shows a representative image of the 3D bioprinted scaffold architecture with microchannels for nutrient exchange.[107] John Wiley & Sons.© 2017 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim.(c) Depicts the construction of 'hybprinter' and the typical workflow involved in printing SLA/extrusion based 3DBP scaffold.Reproduced from [104].© IOP Publishing Ltd.All rights reserved.(d) Represents the typical workflow employed by Daly and Kelly for bioprinting inkjet/extrusion based hybrid scaffolds, (e) image of the hybrid 3D bioprinter in action, and (f) the design of implant along with macroscopic image of tibial implant after 28 d of in vitro culture.Reprinted from [105], Copyright (2019), with permission from Elsevier.

Figure 4 .
Figure 4. Schematic representation of the typical workflow involved in synthesis and implantation of phosphate cement-based 3D printed scaffolds for bone tissue engineering, along with the key feature of 'phosphate cement ink' and 'cell-laden bioinks' used in such techniques.

Figure 5 .
Figure 5. (a) Graphically depicts the typical workflow involved in 3DBP via multi-material co-extrusion bioprinting, the post-processing was optimized by Ahlfeld et al to improve scaffold's mechanical properties.Reproduced from [80].© IOP Publishing Ltd.All rights reserved.(b) (left) represents the simplified nozzle design used for coaxial bioprinting with ceramic as core and bioink as shell, (right) post-processing for the 3DBP scaffold involves immersion in CaCl2 and PBS.Reproduced from [81] with permission from the Royal Society of Chemistry.(c) Depicts the direct-ink-writing reported by Romanazzo et al towards 3D printing CaP ink in cell-laden support bath, resulting in high cell-adhesion with 3DP ink and tailored cellular distribution.[184] John Wiley & Sons.© 2021 Wiley-VCH GmbH.

Figure 6 .
Figure 6.In-vitro results of biphasic 3D printed scaffold based on CPC and alginate bioink, showing significant amount of necrosis near the interface of bioink and CPC after 1 d of incubation (see red arrows); this is reversed upon longer incubation, in fact marked increase in cellular adhesion and proliferation on both CPC and bioink phase is seen after 1 and 3 weeks.(b) Staining of actin cytoskeleton (green) and nuclei (blue) of cells, attached to the CPC strand surface at day 21, further indicating high cellular adhesion to CPC ink.Reproduced from [80].© IOP Publishing Ltd.All rights reserved.(c) Shows in vitro results of PCL/alginate scaffold (i-v) shows fluorescence images with live (green), dead (red) in plan and cross-sectional views of scaffold.(vi) optical and (vii, viii) fluorescence images of implant after 25 d of incubation with actin cytoskeleton (green) and nuclei (blue).[118] John Wiley & Sons.© 2013 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim.

Figure 7 .
Figure 7. (a) Schematically depicts the fate of CaP biomaterials post implantation.CaP degrade via dissolution, leading to release of small fragments which may be phagocytosed by macrophages, while bulk particles are resorbed by osteoclasts.Also, dicalcium phosphate dihydrate (DCPD) may convert to more stable octa calcium phosphate (OCP).Reproduced from [189].CC BY 4.0.(b) shows the simplified degradation modes observed in PCL i.e. surface erosion, bulk degradation, and bulk degradation with autocatalysis.(c) Represents the in vivo degradation of PCL, showing hydrolysis intermediates 6-hydroxyl caproic acid and acetyl coenzyme A, these are then eliminated from the system via the citric acid cycle.Reprinted from [110], Copyright (2010), with permission from Elsevier.

Figure 8 .
Figure 8. Overview of the materials and bioprinting methods reported for load-bearing 3D bioprinting.

Figure 9 .
Figure 9. (a) Graphical representation of collagen mineralization mediated by NCP.Herein NCP-mineral complex gradually infiltrates collagen matrix, followed by precipitation of HA crystals in the gap region.(b) TEM image of unmineralized collagen fibril with ACP complexes, and (c) TEM image of fully mineralized fibril having oriented apatite crystals.Inset: electron diffraction showing increased mineral phase.Reprinted from [245], Copyright (2013), with permission from Elsevier.

Table 1 .
Techniques used for 3D bioprinting of multimaterial orthopaedic scaffolds for load-bearing application.

Table 2 .
Summary of the biomaterials and their properties, used for 3D bioprinting of load-bearing orthopaedic scaffolds.