Co-axial wet-spinning in 3D bioprinting: state of the art and future perspective of microfluidic integration

Recently, a bioprinting paradigm that exploits the simultaneous and distinct delivery of the printing and crosslinking solutions in a core–shell fashion has been developed [37]. This system can be considered an unconventional wet-spinning process, in which the bioink and the coagulation bath are delivered simultaneously to the extruder, triggering the gelification of the bioink at the ending tip of the dispensing head. As the dispensing head moves, following the printing code, a hydrogel fibre is spun out and deposited in 3D. By adjusting the flows of the two solutions and speed of deposition, fibres of tuneable dimension can be deposited. The basic configuration of a co-axial extrusion system is depicted in figure 2. The two solutions of bioink and crosslinker are individually pumped through distinct needles assembled in a co-axial arrangement, permitting the deposition of either a bulk (bioink in the inner needle, figure 2(a)) or hollow fibre (bioink in the external needle, figure 2(b)). Co-axial extrusion/wet-pinning can be considered as an evolution of coagulation baths, in which the bioink and the crosslinking solution are co-ejected from a single depositing head in laminar flow conditions, ensuring a high repeatability of the deposition process.

Zoom In Zoom Out Reset image size

The successful creation of living tissues with 3D bioprinting can be pursued through two different routes. The first one, based on a bottom-up approach, is based on the spontaneous self-organisation of cells, and relies on formulating a bioink that besides being biocompatible is also bioactive. In this case, the bioink is designed to be conductive of different types of stimuli (e.g. chemical, electrical, mechanical), capable for example of guiding stem cell differentiation towards a desired cell phenotype, favouring cell spreading, motility and proliferation within the 3D hydrogel thus promoting cellular intrinsic capacity of organising into functional tissue-like structures. The second route, based on a top-down approach, deals with mimicking a priori the complexity of human tissue histo-architecture and relies on the constant evolution of bioprinting techniques to design living constructs of complex structures with high accuracy and resolution. However, it must be borne in mind that self-organised or predetermined cellular architectures are not separated options in tissue engineering, and the target of recapitulating the complexity of human tissues and organs will most probably involve both strategies.

In the next sections, we will highlight how the adoption of co-axial extrusion in 3D bioprinting presents important technological benefits with respect to these two approaches: (i) the achieved decoupling of printing accuracy from bioink rheological behaviour allow for the ease tuning of its composition (in terms of macromolecular composition or cellular density) to design specific and instructive cellular environments; (ii) the possibility to create heterogeneous structures that mimic defined histo-architectures can be obtained using a single extrusion head with potential integration with microfluidic platforms; (iii) the intrinsic capability of this system in creating hollow fibres can be used to overcome the great limitations associated with the vascularisation of thick tissue constructs.

In co-axial wet-spinning there is the necessity of achieving a fast and cell compatible gelation of the bioink during the deposition; nowadays, such a fast and mild gelation mechanisms has been demonstrated using mainly alginate-based bioinks crosslinked with calcium-chloride solutions [38–40, 43]. However, it must be considered that, beside its technological advantages and biocompatibility, alginate lacks binding sites for supporting cells adhesion and spreading which results in suboptimal functionality of the bioprinted constructs. At the scope of creating bioactive inks suitable for co-axial wet-spinning, alginate has been blended with more bioactive biopolymers. For instance, Colosi et al embedded human umbilical vein endothelial cells (HUVEC) in a blend of alginate and gelatin methacrylate (GelMA) to exploit the presence in gelatin of integrin binding sites that favour cell adhesion, spreading and motility (figures 3(a), (b)) [38]. Bio-constructs printed using alginate ionic gelation were then exposed to UV light to promote chemical crosslinking among methacrylic substituents of GelMA making the structure of the scaffold permanent. It was observed that the spreading behaviour of HUVEC was determined by the degree of crosslinking of gelatin-methacrylate network which could be tuned by the time of exposure to UV light, independently from the printing step. When the network density was correctly tuned, HUVEC migrated towards the periphery of the deposited fibres and organised into vascular-like structures. Cells at the borders developed tight junctions (figure 3(b)), forming a tubular-shaped endothelium templated on the hydrogel cylindrical fibres. At the junction areas among fibres belonging to adjacent layers, lumen-like structures emerged, revealing the capacity of endothelial cells to spontaneously organise in networks (figure 3(a)).

Zoom In Zoom Out Reset image size

The same kind of bioink was used by Zhang et al in the engineering of endothelialized-myocardial tissues. In this study the bioink containing endothelial cells was printed according to an anisotropic fibres arrangement (high aspect ratio) in order to induce post-seeded primary cardiomyocytes to align and undergo spontaneous and synchronous contraction [43]. These bioprinted constructs were embedded in a microfluidic perfusion bioreactor, and the obtained endothelialized-myocardium-on-a-chip model might represent a useful platform for cardiovascular drug screening.

An example of how the formulation of the bioink permits to direct the differentiation of stem cells towards the desired phenotype is reported in the work of Costantini et al [39]. In this work, the Authors targeted articular cartilage as the tissue to be engineered. The formulation of the bioink was driven by the attempt to mimic the natural ECM of cartilage, a tissue rich of proteoglycans and collagen type II (figures 3(c), (d)). A series of photocurable bioinks of increasing complexity constituted by gelatin methacrylate (GelMA), gelatin methacrylate and chondroitin sulfate amino ethyl methacrylate (GelMA + CS-AEMA), and gelatin methacrylate, chondroitin sulfate amino ethyl methacrylate and hyaluronic acid methacrylate (GelMA + CS-AEMA + HAMA) were blended with alginate and human bone marrow-derived human mesenchymal stem cells (hBM-MSCs) to disclose the influence that each component exerted on the differentiation of hBM-MSCs. Interestingly, the Authors demonstrated that the aforementioned bioinks can be used not only to print complex geometrical structures but also to 3D bioprint with high accuracy anatomical parts such as neonatal-size ear (figure 3(d)). Among the formulated bioinks, the one composed of alginate, GelMA and CS-AEMA turned out to be the best candidate in neocartilage formation with the highest production of collagen type II/collagen type I and collagen type II/collagen type X ratios.

A great advantage of bioprinting using the co-axial extrusion method is that once a certain hydrogel has been proven particularly effective in inducing the differentiation of stem cells and their organisation into relevant tissue-like structures, the precursor components of the hydrogel can be readily bioprinted with alginate and cells according to a pattern mimicking the histo-architecture of the target tissue. This approach has been pursued with skeletal muscle tissue. In a previous report [46], Fuoco et al demonstrated that vessel-associated muscle progenitors, termed mesoangioblasts (Mabs), underwent myogenic differentiation in a hydrogel derived from photocurable PEG-Fibrinogen. Hence, C2C12 cells—a gold standard cell line for skeletal muscle tissue engineering—were suspended in a blend of alginate and PEG-Fibrinogen methacrylate and bioprinted in a compact array of highly aligned fibres (figure 3(e)) [40]. After 21 d of culture in vitro, C2C12 properly spread and fused forming highly aligned long-range multinucleated myotubes, with abundant and functional expression of myosin heavy chain and laminin (figures 3(f), (g)). The obtained myotubes showed high degree of alignment along the direction of hydrogel fibre deposition, further revealing maturation, sarcomerogenesis, and functionality.

Co-axial wet-spinning has been also employed in cardiac tissue engineering. A major problem commonly encountered in this field is the poor electrical conductivity of the biomaterial that delay efficient electrical coupling between adjacent cardiac cells. In a recent article, Zhu et al formulated a bioink that beside alginate, comprised gelatin methacrylate and gold nanorods (GNR) [44]. The concentration of nanorods was optimized in order to guarantee low viscosity of the bioink, thus permitting the loading of high cell concentration and at the same time minimising the effect of shear stress field on cell viability. It was shown that cardiac cells in the printed GNR constructs performed better in term of adhesion and organisation when compared to the constructs without GNRs. Furthermore, the incorporated GNRs improved the electrical propagation between cardiac cells and promoted their functional improvement in the printed cardiac construct (figures 3(h), (i)).

Bioprinted 3D cellular constructs represent an important result in the realisation of functional tissues in vitro. The control over cellular disposition and the presence of hydrogel ECM equivalents that provide a 3D environment for cell growth and organisation can lead to the replication of essential structural and functional features of living tissues. However, a further improvement is represented by the simultaneous deposition of different types of ECM and cell populations so as to produce heterogeneous bio-constructs on demand, replicating the architecture of complex organs.

ECM components, collagen in the first instance, represent the ideal bioink for embedding cells, thanks to its bioactive property that favours cellular processes. Unfortunately, printability of collagen is challenging because of its low solubility, inadequate viscoelastic properties and mechanical weakness of the ensuing scaffolds. To circumvent such a problem, Yeo et al—by adapting an approach developed by Onoe et al [47]—3D printed core-sheath microfibers consisting of a collagen core embedding cells and an alginate shell crosslinked with calcium [45]. The use of alginate guarantees printability with good fidelity and protect cells from shear stress during deposition. As a result, when viability, spreading and hepatogenic differentiation of human adipose stem cells was assessed against a control scaffold constituted by alginate alone, the composite fibre-based scaffolds outperformed the alginate one (figure 4(a)).

Zoom In Zoom Out Reset image size

Following the same approach, Liu et al confined a core of GelMA inside a shell of alginate serving as a fibre template [41]. In this way the polymeric network density of the cell embedding phase could be tuned by regulating GelMA concentration. It was observed that among the cell lines cultured in the laid scaffold, each one has its optimal network density. The core–shell approach permitted the fabrication of cell-laden GelMA constructs at extremely low concentrations (<2.0%), a result not achievable using conventional bioprinting strategies (figure 4(b)). The major drawback of such core–shell approach is that alginate fibres are hold together only by ionic crosslinked junctions that have scarce endurance in cell culture medium greatly limiting cell culture times.

Scaffolds among other requirements, must also be capable of withstanding the mechanical environment of the native tissue to be replaced. This is particularly true for load bearing tissues such as bone and articular cartilage. To satisfy such a criterion, one strategy consists in the development of specialised scaffolds formed by combining two or more dissimilar materials. The result is a composite structure that possesses unique properties that cannot be replicated by a single material. Also in this respect, co-axial printing in combination with alginate, offers the opportunity to create composite structures in which the load bearing and the hydrophilic and bioactive phases are spatially segregated in a core–shell fashion.

Following this approach, Cornock et al manufactured a steel co-axial extruder using selective laser melting and refined it by electropolishing [48]. This extruder was used for the simultaneous deposition of a composite fibre composed of an alginate core and a poly(caprolactone) (PCL) shell (figure 4(c)). Process parameters were optimised to minimise the offset of the core fibre from the centre of the co-axial fibre. At the end of the scaffold deposition process, the alginate core was removed leaving a hollow fibre of PCL. Such scaffolds displayed excellent mechanical properties attributable almost exclusively to the PCL shell. Fibroblasts were either post injected inside the fibres of the scaffold or encapsulated in the alginate matrix during deposition. While in the first approach, cell viability was preserved at high level, in the second one, the heat stress caused the death of all cells. This co-extrusion process of both a synthetic, biocompatible and biodegradable polymer and a hydrogel while interesting does not hold great promises in the field of bioprinting. Even though the problem associated to heat stress could be overcome, the hydrogel core supporting cells is surrounded by impermeable hydrophobic shell that hampers the diffusion of nutrients and oxygen which can proceed only longitudinally to the fibres, starting from the ends. A similar solution completely based on the use of hydrogel materials was presented in the work of Mistry et al in which two different bioinks were delivered through the outer and inner needles of a co-axial extruder [49]. In particular, the core bioink containing collagen or Matrigel was loaded with HUVECs, while the shell fluid was composed by a mixture of alginate and poly(ethylene glycol) diacrylate (PEGDA) embedding hepatocytes. The interpenetrating network consisting of a dual crosslinking networks, namely covalently bonded PEGDA chains and a Ca²⁺ mediated alginate network, guaranteed good mechanical properties in particular with reference to gel shape and dimension recovery after deformation.

Co-axially bioprinted structures represent a useful platform on which investigate cellular mechanisms involved in tumours. Cellular spreading and cell–cell interactions are influenced by the surrounding matrix which should be at the same time supportive for cells and not hinder cell migration. Such a construct was developed by Dai et al by co-extruding an alginate solution sheet and a core of brain tumour cells and stroma MSCs suspended in a fibrinogen solution [50]. Fibres were deposited in a coagulating bath of CaCl₂ forming a 3D scaffold. The alginate sheet provided confinement and support to the cells that interacted with each other and self-assembled into multicellular heterogeneous brain tumour fibres. The alginate sheet was subsequently removed by complexing calcium ions with ETDA. The so obtained self-assembled heterogeneous tumour tissue-like fibres provided valuable 3D models for studying tumour microenvironment in vitro.

Human tissues are assemblies of multiple cell types arranged into definite histo-architecture. This implies that the cell-laden printing strategy has to provide the possibility of simultaneous bioprinting of different bioinks endowed with tailored mechanical, chemical properties and embedding different cell types, disposed in specific positions in the 3D space. The standard approach in extrusion-based bioprinters involve the use of multiple printing heads for the fabrication of complex and heterogeneous biological structures. However, an increased number of printing heads requires a more complex printing system and the deposition of the construct is rather slow. Co-axial needles printing heads in conjunction with alginate fibre templating offers new and straightforward opportunities in this respect. Two recent articles showed that, by exploiting microfluidic platforms, the bioinks delivered to the co-axial needle system can be changed on demand using a single printing head [38, 40]. The key of the system was based on a Y configured microfluidic chip mounted upstream the co-axial dispenser (figures 4(d), (e)). The inlets of the top channels are supplied with different bioinks that can be delivered to the co-axial printing head for instance in an alternate or simultaneous fashion, thus realising heterogeneous structures made up by differently composed layers or composite fibres (figure 4(d)). This strategy was implemented in the co-culturing of myoblasts with fibroblasts segregated in the two longitudinal half of the extruded fibres of alginate and PEG-MA-fibrinogen. It has been demonstrated that fibroblasts support myoblasts differentiation by secreting extracellular matrix components and growth factors (figure 4(e)).

The creation of a vascular network within engineered tissue constructs to promote the transport of oxygen, nutrients, and waste products, is a hot topic in tissue engineering and great efforts are being spent to achieve this goal. Co-axial extrusion is particularly promising for the creation of vascular-like structures. Given the configuration of the basic extruder composed of two co-axial needles, a quite intuitive development regarded the modification of the extruder for the deposition of hollow fibres (figure 2(b)) [51]. This development opens up the possibility of printing a blood vessel network that can guarantee the uniform supply of nutrients and oxygen to cells present in all compartments of the scaffold.

One of the first attempt to create a perfusable scaffold composed of hollow fibres using a shell/core nozzle tip was proposed by Luo et al [28]. These authors extruded a self-standing alginate/PVA paste through the hollow space of the extruding system. After deposition, the scaffold was soaked in a CaCl₂ solution. Young's moduli were of the order of few MPa, one or two order of magnitude higher than that characterizing conventional gels. At the same time, fibres walls were permeable to a dye simulating nutrients and oxygen (figure 5(a)). This article paved the way to the manufacturing of hollow fibres based scaffolds even though its applications in bioprinting are ruled out by the extremely high viscosity and network density of the extruded blend.

Zoom In Zoom Out Reset image size

A completely perfusable scaffold was created by printing hollow alginate fibres using co-axial printing head. The concentrations of alginate and calcium chloride solutions were adjusted to allow the deposited fibres to fuse together [52]. The Z stage of the printer moved down across a CaCl₂ solution to keep the printed structure immersed in for complete crosslinking (figure 5(b)). The authors demonstrated that the crosslinked alginate walls were permeable to a model protein and as a consequence the embedded cells could be supplied with nutrient and oxygen delivered by the solution flowing through the hollow fibres network.

A design that allows producing and depositing on a collecting plane either solid or hollow fibres was described in the work of Li et al [56]. The key element for the obtainment of one or the other kind of fibre is the relative position of the tips of the core and external needle. When solid fibres are desired, the inner needle protrudes 500 μm out the outer one and the solution of alginate and calcium chloride are delivered through the core and external needles, respectively. This means that gelation starts from the exterior of the alginate flow and proceeds by diffusion of calcium ions towards the centre. When hollow fibres are the target of the extrusion process, the converse needle arrangement is exploited. In this case, gelation of the shell alginate solution starts from the interior and proceeds outwards. The locus of the beginning of the physical gelation has important implications as far as adhesion among laid fibres is concerned. When gelation starts from the outer shell of the printed fibres, as in the case of solid gel fibres, the alginate solution is completely gelled and fibres belonging to adjacent layers, glue to each other only weakly at the points of mutual contact. As a consequence, scaffolds suffer of poor stability especially in the presence of interfering ions (e.g. in cell culture medium). On the contrary, if gelation started from the internal layers of the printed hollow fibres, the surface layers being still in a sol-gel transition state could effectively glue one to each other at the points of mutual contact. This concept was exploited by Gao et al in the fabrication of hollow vessels-like structures [53]. By using in tandem two spinnerets, hollow fibres embedding either smooth muscle cells or fibroblast were collected around a rotating mandrel (figure 5(c)). Since the outer shell of the alginate hollow fibres was still gelling after deposition, contacting fibres fused to each other creating an unicum construct. This approach permitted to arrange according to a multilevel order, cells making up blood vessels. Thus, smooth muscle cells were deposited within the first fibre layer, on top of which a second layer of fibre supporting fibroblasts was laid down. The inner walls of the vessel-like structure were coated with collagen to favour endothelial cell adhesion. The printed structures with multilevel fluidic channels have structural similarity to blood vessels, have sufficient mechanical strength to support loading, and exhibit biocompatibility. A foreseeable drawback of this approach is related to the creation of groves perpendicular to flow direction that may cause local turbulence on vessels walls. This, in vivo applications can bring about to the formation of deposits on walls which on long terms may cause the clogging of the vessel.

The regeneration of heart tissue damaged by ischaemic events is a goal of great impact in medicine due to the high rate of incidence of this pathology worldwide. The treatment of ischaemic tissues involves the replacement of the tract of the blood vessel clogged. The achievement of this goal is not trivial since the hostile ischaemic conditions (e.g., low nutrition, oxidative stress, inflammation, reactive oxygen species) limit survival and differentiation of transplanted endothelial progenitor cells (EPCs). In a recent publication Gao et al used as bioink vascular-tissue derived decellularized extracellular matrix (VdECM) blended with alginate, EPCs and PLGA microparticles loaded with atorvastatin, a molecule that enhances angiogenesis by reducing subdural haematoma, and promotes endothelial cells function [54]. As stressed previously, the co-axial extrusion system is particularly well-suited for the straightway formation of hollow structures. The gel inducing solution is constituted by Pluronic F127 dissolved in a solution of CaCl₂ that helps avoiding the deformation of the tubular structure before the thermal induced gelation of VdECM (figure 5(d)). These devices once implanted in a mouse ischaemic model, increased the rate of neovascularization and salvage of ischaemic limbs.

A major drawback of most blood vessel bioprinting strategies, is the weak mechanical strength of the constructs that hampers its surgical anastomosis with the host blood vessel in the implantation process. The use of co-axial-needles extruders offers the opportunity of printing mechanically robust and perfusable constructs. In the work by Jia et al [55], tri-axial nozzle was used to create perfusable hollow fibres of various diameters and wall thickness in a single step. In this approach, the bioink was delivered through one or both of the two external co-axial needles while the calcium chloride solution was pumped through the innermost needle (figure 5(e)). In this way either a single or double walled concentric hollow fibre were produced permitting to modulate its mechanical property. The bioink was a blend of gelatin methacrylate, alginate and 4-arm poly(ethylene glycol) acrylate PEGTA that conferred to the bioprinted hollow constructs improved mechanical properties by increasing the crosslinking density. The co-axial needles printing method possesses distinct advantages over conventional complicated microfabrication and sacrificial-templating approaches where a templating bioink is firstly deposited and embedded in a hydrogel matrix, followed by removal of the bioink to obtain the hollow vessel-like structures.