Emerging Technologies in Multi-Material Bioprinting

Abstract

Bioprinting, within the emerging field of biofabrication, aims at the fabrication of functional biomimetic constructs. Different three-dimensional bioprinting techniques have been adapted to bioprint cell-laden bioinks. However, single-material bioprinting techniques oftentimes fail to reproduce the complex compositions and diversity of native tissues. Multi-material bioprinting as an emerging approach enables the fabrication of heterogeneous multi-cellular constructs that replicate their host microenvironments better than single-material approaches. Here, we briefly review bioprinting modalities, discuss how they are being adapted to multi-material bioprinting, as well as analyze their advantages and challenges, encompassing both custom-designed and commercially available technologies. The review offers a perspective of how multi-material bioprinting opens up new opportunities for tissue engineering, tissue model engineering, therapeutics development, and personalized medicine.

Graphical Abstract

A comprehensive overview of the available 3D bioprinting technologies for fabricating multi-material constructs is presented. The technologies are classified into four major categories, and their advantages and shortcomings are compared. Available multi-material commercial bioprinters are reviewed, concluding with a perspective of the future path for developing multi-material technologies.

1. Introduction to bioprinting modalities

Three-dimensional (3D) printing refers to the fabrication of constructs from a digital 3D model in a layer-wise^[1] or volumetric^[2] programmed manner. The flexibility, versatility, and functionality of 3D printing enable the fabrication of exquisite and intricate structures^[3–5] with details as small as hundreds of nanometers^[6]. Here we focus on one sub-category of 3D printing, designated 3D bioprinting^[7,8], in which a combination of cells, growth factors, or biomaterials (i.e., bioink^[9]) may be used as the printing material for additive manufacturing of biological constructs^[10]. As part of the rapidly evolving field of biofabrication, 3D bioprinting is being explored for a broad range of applications within tissue engineering^[11,12], regenerative medicine^[13], organ-specific tissues^[14], patient-specific grafts^[15], tissue model engineering^[16], and drug screening^[17]. The most frequently used technologies for 3D bioprinting include nozzle-based and laser/light-based techniques. Extrusion^[18] and inkjet^[19] are arguably the most common modalities of nozzle-based bioprinting at the moment, while laser-induced forward transfer (LIFT)^[20] and vat-photopolymerization^[21] are the two frequently used for laser/light-based 3D bioprinting^[22].

Extrusion bioprinting involves fabrication on a bioprinting platform using bioinks extruded from one or several nozzle(s) (Figure 1A). The extrusion process may be pressure-controlled, with the bioink entrained by means of pneumatic actuation, or flow rate-controlled, with the bioink forced by mechanical impulses through syringes^[23]. Solidification of the bioprinted structures as they are delivered is obtained through physical, chemical, or photo-crosslinking^[24]. Extrusion bioprinting is relatively inexpensive, straightforward, and convenient. It has been embodied, for example, within handheld and portable devices^[25–29]. But such convenience must be traded off against significant challenges. Nozzle extrusion necessarily entails a high level of shear stress near the fluidic channel walls. Excessive shear, particularly in high-viscosity bioinks, jeopardizes cell viability^[30]. High-resolution bioprinting often requires small-diameter nozzles. The greater shear required imposes a limit on bioink flow rate and throughput. This problem may be addressed through the use of printable shear-thinning biomaterials^[31], for which the viscosity decreases under shear stress. But shear-thinning bioink materials that meet all design requirements are often difficult to find. Further notable challenges of extrusion methods include difficulties in finding stable in situ crosslinking methods for non-shear-thinning bioinks, as well as low printing resolution^[32] in relation to other methods, and complications in the fabrication of free-standing constructs^[33,34]. These shortcomings have spurred many enhancements, notably co-axial/core-shell bioprinting^[35], described in Section 3.1.3, and embedded bioprinting^[36], in Section 3.1.4.

Figure 1-.

Schematic showing the different bioprinting modalities. A) Extrusion. B) Inkjet. C) LIFT. D) Vat-photopolymerization.

Inkjet bioprinting (Figure 1B), similar to home/office inkjet printing, delivers small droplets of bioink to a substrate and can produce high-resolution voxelated constructs^[37]. Unlike the extrusion method where shear-thinning bioinks with a broad range of viscosities can be used, inkjet bioprinters are mainly designed to work with low-viscosity bioinks^[38]. The deposition of droplets/voxels is controlled either by thermal, piezoelectric, or electromagnetic actuation^[39,40]. Short thermal pulses within the printhead result in gasification of the bioink, and the subsequent pressure increase causes the ejection of droplets. Piezoelectric actuators impart acceleration to the bioink to trigger controlled droplet ejection. Electromagnetic inkjet bioprinters use solenoid-actuated valves within the nozzle to regulate the extrusion of droplets^[41]. Inkjet approaches are faster than extrusion. Yet, they are typically not suited for the fabrication of thick structures, and not used as broadly for multi-material bioprinting. The printability of bioinks for inkjet bioprinters is usually determined by the Ohnesorge number, $Oh=\mu /\sqrt{\rho \sigma R}$, where µ is the viscosity, ρ is the density, σ is the surface tension, and R is the characteristic length scale for the flow, which can be taken as the radius of the orifice of the printing nozzle^[42]. This dimensionless number represents the ratio of viscous forces to inertia and surface tension. A material with an Ohnesorge number between 0.1 to 1 is usually deemed printable via the inkjet method^[42]. For a given bioink composition, the ejected droplet size is mainly determined by nozzle dimensions. The printing speed mainly depends on the frequency of the pulsated heating or piezoelectric motion. Inkjet printing speed and droplet size are therefore controlled independently.

In LIFT bioprinting (Figure 1C), the bioink is initially deposited as a thin layer on the bottom of a substrate, termed the donor layer. Upon irradiation of the donor layer surface with a pulsed laser, droplets are created and transferred onto the receiver plate in a voxelated fashion^[20]. The LIFT method is both highly accurate and fast. A range of bioinks is available for LIFT^[20], despite restrictions in viscosity (values should fall between 1 to 300 mPa s) and crosslinking mechanisms^[22,43]. At the time of this review, LIFT is not suitable for the high-throughput fabrication of multi-material or multi-cellular constructs^[44]. It is not often used for high-aspect-ratio constructs.

In vat-photopolymerization bioprinting (Figure 1D), the fabrication process involves the selective exposure of liquid photocurable bioink in a vat to ultraviolet (UV), visible, or near-infrared (NIR) light^[45–48]. Three light-patterning methods have mainly been used for vat-photopolymerization: (i) one single programmed laser beam, such as in conventional stereolithography (SLA); (ii) digital light processing (DLP); or (iii) two-photon polymerization (TPP). The programmed pattern in the precursor vat is illuminated by the laser beam and solidifies. Once each layer is crosslinked, the build platform moves to a neighboring position to allow the fabrication of the next layer.

Other commonly used vat-photopolymerization bioprinting methods include DLP and conventional SLA. These methods are able to fabricate constructs with details only of a few tens of micrometers^[49]. In the DLP method, the projected light is usually masked using an array of mirrors termed the digital micromirror device (DMD). This allows the simultaneous illumination of one entire layer. Pixels are either exposed or blocked to obtain the desired shape outline. Other devices scan a laser beam along the horizontal axis to rapidly solidify the bioink across each plane^[50,51]. The speed of DLP bioprinting is greater than that of conventional SLA or TPP as it is not limited by layer geometrical complexity (XY area), but rather by layer thickness and exposure time. However, the accuracy of DLP is limited by the pixel size of the projected image. Thus, the printed surface area must be traded off against the level of detail along the horizontal axes (XY). Light intensities in DLP-based 3D bioprinting are often lower than that in other vat-photopolymerization printers. Nevertheless, some high-resolution DLP light engines deliver light energy densities within the same range as SLA laser beams (>100 mW/cm²) and solidify bioinks within hundreds of milliseconds. In these two methods, the XY resolution is primarily determined by the projected pixel or laser beam spot size, bioink reaction kinetics, and the diffusion of free radicals^[52]. The vertical resolution is mainly dictated by the light penetration depth, itself a function of the absorbance/scattering of the ink, and the layer thickness^[53]. Both SLA and DLP technologies can generate bioconstructs with high cell densities (>2×10⁷ cell/mL)^[54–56]. Limitations include phototoxicity, UV-triggered mutations, photoinitiator toxicity, and insufficient choice of photocurable bioinks.

The use of TPP for 3D bioprinting has been less common due to lower cell viability, insufficient throughput for bioprinting of large constructs, and a limited number of efficient biocompatible water-soluble photoinitiators^[57]. Most TPP 3D bioprinters are based on femtosecond lasers operating in the NIR wavelength range. They enable the fabrication of 3D microstructures with submicron details^[58]. Due to the operational mechanism of TPP, high-throughput 3D bioprinting of cell-laden hydrogels requires a high-speed scanning system^[59]. In TPP, the spatial resolution is largely dictated by the laser irradiation intensity (~TW/cm²) and exposure time, which primarily depends on the initiating efficiency of TPP photoinitiators^[60,61].

The performances of these four primary bioprinting modalities according to feature size, cell viability, and printing speed are shown in Figure 2. As mentioned before, nozzle-based techniques, especially the extrusion-based method, deliver a relatively lower average cell viability rate. Viscosity of a bioink, which depends on the molecular weight and concentration of the dissolved polymers, also plays an important role in the efficacy of the bioprinting methods. Particularly for extrusion bioprinting, structures with high fidelity can be bioprinted when viscous bioinks with high yield strengths are utilized, although oftentimes they would still go through another crosslinking step post-bioprinting^[42]. However, higher viscosity values usually induce higher shear stresses and lower cell viability^[62]. This trade-off needs to be considered when live cells are used in a biofabrication process. If viscous bioinks are adopted in conventional extrusion bioprinting, the cell viability declines due to the high shear stresses. In the inkjet technique, low-viscosity bioinks are used with high cell viability, but additional polymerization is almost always needed to yield a crosslinked construct with a high yield strength that resists the deformation due to gravity^[63]. Recently, an acoustic droplet ejection approach has been employed in conjunction with the inkjet bioprinting to reduce shear stress on the bioink^[64]. In embedded bioprinting, where low-viscosity bioinks are also employable, higher cell viability and smaller feature sizes can be achieved. As a result, the ranges of cell viability and feature sizes for the extrusion method are considerably larger than those of the other methods (Figure 2A). Laser/light-based methods, i.e., LIFT and vat-photopolymerization, generally yield high cell viability values as well as better bioprinting resolutions and faster bioprinting speeds^[47,49].

Figure 2-.

Comparison of different bioprinting methods. A) Cell viability versus minimum feature size. In nozzle-based techniques, the excessive shear stress applied to the cells significantly decreases the cell viability, whereas in light/lased-based methods, the overall cell viability is higher. The light/laser-based methods are also more capable of generating well-defined constructs with higher resolutions. B) Printing speed versus minimum feature size. Despite its simplicity, the extrusion method is generally the slowest modality among the four primary bioprinting techniques. Inkjet and vat-photopolymerization are the two fastest methods, and LIFT is considered a medium-speed method^{[32,38,43,49,217,300–303]}. The comparisons are general and may not hold true for all specific cases.

Bioprinter selection must consider the limitations of the intrinsic printing mechanism. Printing speed and detail resolution are shown for various techniques in Figure 2B. No modality is absolutely preferable over other methods in terms of printing speed. In general, extrusion bioprinting is slower, while inkjet and vat-photopolymerization technologies provide faster fabrication paces^[43]. The range of bioink viscosities available for extrusion bioprinters is greater than that for inkjet bioprinters^[65]. The lower viscosities in inkjet bioprinting yield higher flow rates and thus increased printing speeds. For nozzle-based techniques, the printing speed considerably affects detail resolution^[66], unless it is increased through the addition of multiple nozzles. According to the Hagen–Poiseuille law^[67], nozzle diameter, nozzle length, bioink viscosity, and bioink flow rate affect detail resolution in extrusion bioprinting. In newly developed volumetric 3D bioprinting^[2] and xolography^[68], the resolution does not necessarily affect the printing speed. Inspired by cell patterning^[69] and cell packing approaches^[70], recent efforts have been devoted to exploiting acoustic and ultra-sound impulses to organize and pattern cells within printed layers^[71,72]. Overall, the selection of bioprinting technology depends on the targeted application and the required detail resolution. Although printing speed, cell viability, and resolution are important, other key features, such as flexibility, accessibility, operability, and cost-effectiveness, should also all be considered. These aspects are discussed in Section 6.

The capability of fabricating multi-cellular/multi-material constructs is of paramount significance as the native human tissues and organs possess heterogenous cellular and extracellular structures. Recently, tremendous attempts have been made to design multi-material bioprinters, encompassing a broad range of technologies, from open-source desktop platforms^[73] to standalone commercial bioprinters^[74]. Several surveys have been published with discussions on multi-material additive manufacturing^[75,76]. However, these reviews have devoted scant attention to the multi-material printing of cell-laden bioinks, i.e., bioprinting. Other reviews with a focus on biomedical applications of 3D printing^[3,77,78] have brief discussions on multi-material methods, principally emphasizing the material design rather than its technology. Other recently published reviews describe microfluidics-assisted bioprinting^[44]. Extrusion-based multi-material bioprinting has been hailed as one of the main applications of microfluidic systems^[79]. But the salient evolution of multi-material bioprinting both in nozzle-based and laser/light-enabled technologies over the last few years, along with the upsurge of available multi-material commercial bioprinters, entails the need to recapitulate the recent advances.

Herein, we classify (Figure 3) and summarize the state-of-the-art multi-material bioprinting approaches mostly developed over the past 5 years. Recent advances in multi-material bioprinting are critically discussed, broken down along the four dominant technologies presented above. A section on applications of multi-material bioprinting is included. Commercial multi-material bioprinters are presented as they are key for translation to clinical applications. Perspectives on opportunities for future discoveries conclude the review.

Figure 3-.

Classification of multi-material bioprinting technologies.

2. Multi-material bioprinting concept

Multi-material bioprinters are generally better-suited than conventional bioprinters for the fabrication of constructs that mimic the heterocellular structures of native tissues^[80–82], enabling for example the incorporation of graded composition and properties or environmental adaptations^[83–85]. The versatility added to biofabrication methods via the development of multi-material bioprinting has already led to improved sacrificial supports^[86], multi-functional systems^[87], vascularized structures^[88], customizable tissues and organs^[89], and advanced spatiotemporal control^[90].

At the onset, it is useful to clarify some important definitions. Multi-component bioink refers to mixtures of two or more different biomaterials making up one single-phase homogenous bioink^[91,92]. Once properly crosslinked, such bioinks may yield multi-network or single-network structures. Composite bioinks are, specifically, multi-phase materials composed of two or more immiscible components^[93]. The most common types of composite bioinks are fabricated through the incorporation of nanoparticles/nanofibers in hydrogel matrices^[94–96]. Composite bioinks may have isotropic or anisotropic properties, depending on the spatial distribution and orientation of the additives^[92]. Multi-material 3D bioprinting designates the sequential/simultaneous bioprinting of two or more (bio)inks in a programmed manner to achieve region-specific features and performances. Accordingly, each (bio)ink may be single-/multi-component, single-phase, or composite.

As for conventional bioprinting, various advanced materials have been developed for multi-material modalities to achieve either superior properties, proper crosslinking, or better biomimicry^[97]. Many are hydrogel polymers in view of their intrinsic characteristics, such as hydrated microenvironments, facilitated crosslinking mechanisms, cytocompatibility, printability, and vast rheological properties over different temperatures. Chemically modified bioinks, e.g., methacrylated polymers^[98,99], are widely employed since they offer precisely controllable crosslinking regimes^[100]. The polymer molecular weight and concentration significantly affect bioink performance^[101]. The gelation mechanism of a biomaterial is a critical factor in multi-material bioprinting, where multiple bioinks are to be crosslinked in a compliant manner. Physically and chemically crosslinked hydrogels follow different gelation regimes^[91]. Special attention should be, therefore, devoted to bioinks with different crosslinking mechanisms when concomitantly employed in multi-material bioprinting. Composite biomaterials, as described above, are another category of advanced materials that have been developed and used in multi-material bioprinting. One of the main advantages of composite bioinks is their enhanced printability and functionality. The reader is referred to our previous review for more details regarding composite bioinks^[92].

Due to the employment of multiple materials, interactions at the interfaces of materials must be considered. A strong interfacial adhesion can improve the toughness and fatigue-resistance of bioprinted constructs^[102]. Different bioprinted parts are commonly bonded by using materials that can be crosslinked via the same covalent or ionic crosslinkers^[103]. Other methods employing supramolecular forces such as guest-host interactions have also been explored^[104]. But, it is challenging to bond hydrogels to materials from a different nature, especially elastomers and thermoplastics, in a multi-material (bio)printing process. Recent advances in hydrogel adhesives shed light on new ways to design material interfaces in bioprinting. For example, hydrogels and elastomers can be strongly bonded by using free-radical polymerizations^[105], topological adhesion^[106], bridging polymers^[107], or catechol chemistry^[108]. Thermoplastics can also be functionalized with certain specific functional groups to allow hydrogel bonding^[109]. However, it is to be emphasized that most existing hydrogel adhesion strategies suffer from cytocompatibility issue, and the cytotoxicity are dose-dependent. Cell-friendly strategies, such as bio-orthogonal click chemistry^[110], could be a potential remedy. While this review is focused on multi-material bioprinting technology, the reader is referred to reviews covering the design and applications of hydrogel adhesion elsewhere^[111–113].

3. Multi-material bioprinting technologies

Over recent years, four types of bioprinting technologies have been exploited to fabricate multi-material products. Multi-material bioprinting has grown from the use of rather cumbersome tools to a well-integrated and automated process. In this section, the latest efforts for advancing the field of multi-material bioprinting are reported with a focus on technology. Broader studies about 3D printing are at times included when they have the potential to be adapted for 3D bioprinting.

3.1. Extrusion bioprinting

Extrusion bioprinting is the most popular method for multi-material bioprinting^[3]. Many innovative concepts have been demonstrated for the fabrication of multi-material constructs. In multi-material extrusion bioprinters, the printhead comprises cartridges (reservoirs), mixers, tubing, and nozzles. Nozzle delivery may be divided into single-nozzle^[114] and multi-nozzle^[115] technologies. Conventional single- and multi-nozzle technologies are among the most prevalent, accessible, and easy-to-implement technologies. In single-nozzle bioprinters, there is essentially one printhead consisting of only one nozzle. Different concentrations of bioinks may be sequentially extruded or mixed within the printhead. The use of only one nozzle for multiple materials requires the same working temperature for all the (bio)inks, which is not always feasible. More importantly, single-nozzle delivery increases the risk for cross-contamination. Multi-nozzle bioprinters have one or multiple printheads, each of which is equipped with one or several nozzles. Simultaneous delivery through several nozzles tends to increase fabrication speed. Co-axial nozzles^[116–118] enable the fabrication of multi-layer/core-shell constructs through the simultaneous and concentrically collocated extrusion of different bioinks. These are often needed for bioinks with rapid crosslinking mechanisms, e.g. those based on or containing alginate. Embedded bioprinting^[119] is helpful for making freeform structures that are difficult to fabricate using classic extrusion methods, or to use low-viscosity bioinks that are not compatible with other extrusion-based methods. Co-axial and embedded bioprinting techniques can be potentially used with either single-nozzle and multi-nozzle bioprinters.

3.1.1. Single-nozzle bioprinting technology

Extrusion of different bioinks through one single nozzle remains the baseline for multi-material bioprinting. Multi-reservoir systems and mixers are the two most common related modalities. In the former, several reservoirs of bioinks are connected to the nozzle within the printhead (Figure 4A-D). The multi-material construct is fabricated following sequential or simultaneous activation of multiple reservoirs, resulting in the alteration of the feed. This method is well-illustrated by a pneumatically controlled bioprinter consisting of seven reservoirs^[120]. This bioprinter may be used with a broad range of bioinks, including shear-thinning and conductive biomaterials. Its solenoid valve technology enables precise bioink flow control and rapid switch between bioinks. As for most extrusion techniques, with the exception of embedded bioprinting, the smallest feature size is limited to around 100–200 μm. The nozzles’ inner surface properties may also affect the switching rate between bioinks. For example, Cameron et al. used a readily available single-nozzle printhead mounted on a commercial bioprinter (Figure 4E, F)^[121] to show that switching between bioinks is faster when a hydrophobic coating, e.g., silicone, is used for the nozzles’ inner channels. Much work is needed to further explore the influence of nozzle coatings on the effectiveness of multi-material bioprinting and the viability of encapsulated cells.

Figure 4-.

Single-nozzle multi-material bioprinting technologies. A, B) Schematic illustration of the multi-reservoir technique. Each cartridge is actuated using a separate pneumatic valve. C, D) Optical image of the valves and the printhead setup. Reproduced with permission.^[120] Copyright 2017, Wiley-VCH. E, F) Schematic configuration of a multi-material single-nozzle printhead assembled on a commercial 3D bioprinter, and its photograph. Adapted under the terms of the CC-BY Creative Commons Attribution 4.0 International license (https://creativecommons.org/licenses/by/4.0).^[121] Copyright 2020, The Authors, published by Multidisciplinary Digital Publishing Institute.

Microfluidic devices facilitate switching between bioinks when following a multi-reservoir approach^{[79,117,122,123]}. Since the flow of bioinks in microchannels is often laminar (i.e., with a low Reynolds number), microfluidic devices work as switches that deliver bioinks sequentially with minimal mixing. Switching rate, however, is limited by the system’s compressibility ^[124]. Longer transient periods tend to reduce the sharpness of the edges in the printed structures.

While implementing the multi-reservoir system is simple, this method is not functional for systematically fabricating constructs with continuous gradient properties. Various mixers have been used to overcome this shortcoming by blending two or more bioinks in different concentrations to facilitate multi-material printing^[124]/bioprinting^[125] jobs. The mixers are divided into two general categories; active and passive^[126]. Active mixers, such as the so-called “on-the-fly” designs^[127], consist of either a motor-driven impeller or an acoustic source (Figure 5A, B)^[128], in which the mixing capacity can be finely controlled through changing the input power. The active mixers have been studied using computational models to anticipate the mixing capabilities^[129,130]. Passive mixers operate by introducing turbulent cross-stream flows using geometrical discontinuities or sharp edges within the microchannels (Figure 5C)^[131]. An important problem with passive mixers is the limited capability of mixing bioinks on small scales. The solution to this problem is employing microfluidic mixers, a popular category of passive mixers commonly used in multi-material bioprinters. Bioink volumes as small as 10⁻⁹ to 10⁻¹⁸ L can be properly mixed in a controlled manner using microfluidic mixers^[44]. The mechanism of mixing in such devices is based on transitioning the flow of bioink from laminar to turbulent (i.e., high Reynolds numbers)^{[123,131,132]}. The reader is referred to another recent review for a more comprehensive insight into microfluidics-based 3D bioprinting^[44].

Figure 5-.

Different types of mixers used for multi-material bioprinting. A, B) Photograph and schematic of an impeller-based active mixer used for blending two bioinks. Reproduced with permission.^[127] Copyright 2015, The Authors, published by National Academy of Sciences. C) Schematic design of a passive microfluidic mixer. Reproduced with permission.^[131] Copyright 2019, Institute of Physics. D) Continuous chaotic printing experimental setup. E) Cross-section of the multi-lamellar printed fiber. Scale bar: 250 µm. F) Schematic of the experimental design for the chaotic printing of two inks using a KSM and a syringe pump. G) Side views of KSM at two different angles. H) Illustration of flow splitting action in a six-element KSM with a diameter of D. The distance between lamellae is shown by δ. Reproduced under the terms of the CC-BY Creative Commons Attribution 4.0 International license (https://creativecommons.org/licenses/by/4.0).^[134] Copyright 2020, The Authors, published by Institute of Physics.

Passive mixers are generally easier to integrate and more biocompatible as they induce less shear stress to the encapsulated cells^[133]. However, it is impossible to tune the mixing capacity of passive mixers without changing the input bioinks’ flow rates^[127]. Changing the flow rates, on the other hand, may adversely affect the bioprinting quality. Furthermore, the mixing capacity of microfluidic devices highly depends on the viscosity of the bioink, as well as the length, diameter, and geometry of the microchannels. The small-scale mixing issue, mentioned above, becomes more serious when viscous bioinks are employed. In such situations, the mixing capacity significantly declines. Kenics static mixers (KSM) were recently used to facilitate the continuous chaotic printing of multi-material structures (Figure 5D-H)^[134]. Using this technology, multi-lamellar fibers with a well-defined internal microarchitecture were created by exploiting the mixing capacity of chaotic flows (Figure 5E), which importantly, were highly predictable. These fibers can be used to bioprint defined multi-material structures featuring large interfacial areas. This KSM technology is growing fast^[135,136] to overcome the limitations of mixing viscous bioinks in passive microfluidic mixers.

3.1.2. Multi-nozzle bioprinting technology

One drawback of the single-nozzle technology is the risk of cross-contamination as the (bio)inks flow in one nozzle. Multi-nozzle multi-material bioprinters are utilized to resolve this shortcoming and increase the throughput. In such technology, one or more printheads with an array of nozzles are implemented for delivering various bioinks^[137]. Compared to single-nozzle bioprinters, more complex features at a faster pace and a larger build volume can be fabricated when multi-nozzle technology is employed. Depending on the design complexity, each of the bioinks can flow through one nozzle or every nozzle. In the latter case, a more meticulous design of microfluidic channels is essential to ensure synchronized delivery of the bioinks to the bioprinting platform. A basic method, in which each nozzle is dedicated to only one bioink, was demonstrated in 2014^[138]. In this work, four nozzles were mounted on separate printheads, each independently controlled along the Z-axis. Heterogenous 3D constructs with interpenetrating vasculature were successfully bioprinted taking advantage of the fugitive properties of Pluronic F-127. In 2016, a multi-nozzle printer, named integrated tissue–organ printer (ITOP), with the capability of printing four different materials, including two bioinks, was proposed^[139]. As shown in Figure 6A, this single-printhead multi-dispensing module consisted of four microscale nozzles connected to separate repositories and air pressure controllers. The researchers could successfully bioprint different organ models such as ear cartilage, skeletal muscle, and mandible bone with the aid of polycaprolactone (PCL) as the supporting material (Figure 6B) and Pluronic F-127 as the sacrificial compartment. While elegant, the designed bioprinter is not able to fast-switch between the bioinks. The prolonged transient period results in a lower sharpness in the edges and hinders the construction of intricate multi-material tissue models.

Figure 6-.

Multi-nozzle multi-material 3D printing technologies. A) The Integrated tissue-organ printer setup consisting of four separate nozzles. B) Schematic of 3D-printed basic pattern with multiple bioinks and the supporting PCL ink. Reproduced with permission.^[139] Copyright 2016, Springer Nature. C) Optical images of top and side views of multi-nozzle printheads with various types of inks. D) Generating voxelated multi-material filaments at an increasing switching frequency, where only one of the nozzles is shown. E) Comparing the effect of subcritical and supercritical pressures in the occurrence of backflow. F) The effect of using asymmetric microfluidic channels on the maximum flow of the active channel (Q_f^max), where only one of the nozzles is shown. Reproduced with permission.^[140] Copyright 2019, Springer Nature.

Pneumatic actuators are reliable instruments to achieve fast switching between the inks^[120,140]. Recently, Lewis and colleagues built a multi-material multi-nozzle 3D printer (Figure 6C), which was actuated through a series of pneumatic solenoids, enabling fast switching between the inks (up to 50 Hz)^[140]. In this design, a maximum of eight different materials could flow in the nozzles by means of a microfluidic system. They used pressure-driven flows in merging microfluidic channels to achieve seamless switching between the inks (Figure 6D). One common concern with the advanced multi-nozzle designs that may negatively influence the operation of the 3D printer is the backflow from the active channel to the static channels due to the higher pressure at the junction of the channel. Pressures below the maximum critical pressure values for the active channel (P_cr) were used to successfully prevent the backflow into the static channels (Figure 6E). Also, it was reported that increasing the channel length for lower-viscosity inks and using asymmetric configurations could effectively mitigate the backflow (Figure 6F). Although these conclusions are derived for general multi-material 3D printing, the same concept can be potentially expanded to multi-material 3D bioprinters, with consideration of the presence of the cells.

3.1.3. Co-axial bioprinting technology

Co-axial nozzle (Figure 7A) is a mechanism for bioprinting multi-material core-shell structures^{[117,118,141,142]}, such as vascular constructs^[143], heterogenous microfibers^[144], and tumor models^[145]. This technique enables the user to fabricate hollow structures with compositional and geometrical complexities. Co-axial nozzles are specifically suitable for bioinks that rapidly crosslink upon mixing with the crosslinker. Alginate is the most popular hydrogel used in co-axial bioprinters in conjunction with calcium chloride (CaCl₂) as the physical crosslinker. For example, a handheld co-axial system of extrusion was developed to print multi-material structures, which resembled the brain’s cortical tissue^[26]. Gellan gum and the proper crosslinkers (calcium and magnesium ions) were used to print the multi-material brain-like structures through this co-axial configuration. While convenient, this method has certain limitations in resolution, accuracy, reproducibility, and production rate, which are attributed to the lack of full automation. By reconfiguring the nozzles in the co-axial design, multi-layered tubular tissues can be fabricated. This configuration was achieved by concentrically placing three nozzles with different diameters inside each other and simultaneously bioprinting a mixture of alginate, gelatin methacryloyl (GelMA), and eight-arm poly(ethylene glycol) (PEG)-acrylate with a tripentaerythritol core (PEGOA)^[146]. As shown in Figure 7B, CaCl₂ was used in the inner nozzle to ionically crosslink the bioink. The double-layer tubular construct was subsequently crosslinked again by UV light to form a stable structure. Using this design, the users can continuously alter the shape, size, and the number of layers in a single step without changing the nozzle. These circumferentially multi-layered constructs can be used as human cannular tissue models.

Figure 7-.

Co-axial multi-material bioprinting technologies. A) Custom-designed multi-layered co-axial nozzles with various diameters. Reproduced under the terms of the CC-BY Creative Commons Attribution 4.0 International license (https://creativecommons.org/licenses/by/4.0).^[264] Copyright 2017, The Authors, published by Springer Nature. B) Schematic illustration of multi-material bioprinting of a tubular tissue using a blend of PEG and PEGOA in GelMA/alginate as the bioink. Reproduced with permission.^[146] Copyright 2018, Wiley-VCH. C) Using microfluidic systems to achieve a crosslinked Janus flow pattern of multiple cell-laden bioinks via a co-axial nozzle. Reproduced with permission.^[80] Copyright 2017, Elsevier. D) A multi-scale fluidic system used for biofabricating vessel-like structures made of alginate. E) Multi-scale perfusable vessel-like constructs, (a) Samples of single-layer and double-layer constructs, (b, c) Photographs of the single-layer construct, (d, e) Photographs of the double-layer construct, (f) Scanning electron microscopy of the longitudinal section. F) Bioprinted multi-cellular vessel-like structures (red: L929, green: MOVAS, and orange: HUVEC). Reproduced with permission.^[148] Copyright 2017, American Chemical Society.

Microfluidic printheads can be coupled with co-axial extruders to add more versatility to the bioprinters. Costantini et al. used this technology to achieve a Janus flow pattern of mouse myoblast (C2C12) and fibroblast (BALB/3T3) cell-laden bioinks in the printhead (Figure 7C)^[80]. The compartmentalized bioprinted structure was used to study how fibroblasts could expedite the myogenic differentiation. Microfluidic printheads in conjunction with co-axial nozzles are also beneficial in fabricating more complicated configurations, such as multi-compartmental fibers^[147]. Furthermore, by implementing an innovative approach with the aid of co-axial nozzles, a configuration of multi-scale fluidic systems (i.e., macrochannels and microchannels) for fabricating multi-level vascularized tissue constructs was proposed^[148]. As depicted in Figure 7D, two co-axial nozzles were used to bioprint two types of cell-laden alginate bioinks (with fibroblasts and smooth muscle cells) along a rotating rod. The inner nozzles contained CaCl₂ as the crosslinker. After bioprinting, the rod was removed, and the double-layer spiral construct was soaked in the CaCl₂ bath so that the outer surface was fully crosslinked. Collagen solution was then injected in the macrochannel to enhance the adhesion of endothelial cells, which were seeded to resemble the vascular network. The method was used to fabricate single- and multi-layer vessel-like constructs (Figure 7E). Also, multi-cellular vessel-like structures containing mouse fibroblasts (L-929), mouse vascular smooth muscle cells (MOVSMCs), and human umbilical vein endothelial cells (HUVECs) were successfully fabricated (Figure 7F). In the fabrication process, the constructs should be manually translocated between each step to receive proper crosslinking. As a result, the main drawback of this method is the increased risk of contamination. One possible improvement to such technology is automating the system and eliminating the manual translocation of the construct.

The liquid rope-coil effect has been recently exploited as an excellent method for creating multi-material constructs^[149,150]. Using this approach, Shao et al. employed co-axial multi-material bioprinting to fabricate GelMA microfibers with various morphologies^[151]. Non-viscous GelMA was surrounded by viscous alginate in the nozzle, forming a laminar co-axial flow. The nozzle was connected to a transparent tube to provide enough room for a sequential crosslinking process (Figure 8A, B). As the first step of crosslinking, the co-axial flow was exposed to the UV light, which fully solidified the GelMA compartment, resulting in the formation of GelMA microfibers flowing inside alginate. The second step of crosslinking was subjecting the sheath layer to CaCl₂. Once alginate rapidly solidified, the velocity of the GelMA compartment dominated, and the microfibers began to coil inside the alginate matrix. By adjusting the flow rates of GelMA and alginate solutions, the nozzle diameters, and the concentration of the GelMA phase, a variety of microfiber shapes was achieved. A similar methodology was used with different cell-laden hydrogels as the core bioinks to achieve multi-compartmental microfibers (Figure 8C-K). Using HUVEC-laden GelMA as the core bioink in the co-axial configuration proved that endothelial cells could migrate towards the border of GelMA coils after 12 days of culture. Other types of cells can be used through similar co-culture systems to study the effect of microfiber shape on the interaction between the cells.

Figure 8-.

Fabrication of multi-material microfibers. A) Schematic illustration of the co-axial nozzle for the fabrication of GelMA microfibers based on the co-flow rope-coil effect. B) Two-step sequential crosslinking of the microfibers. By changing the setup and adjusting the flow rates, microfibers with different morphologies were obtained: C, D, E) Janus structures, F, G, H) multi-layer patterns, and I, J, K) paractic configurations. Reproduced with permission.^[151] Copyright 2018, Wiley-VCH.

3.1.4. Embedded bioprinting technology

In the embedded printing approach, the hydrogel (bio)ink, which is not printable via conventional 3D printing techniques, is extruded into a liquid-like or gel-like bath that supports the fidelity of the printed structure. Besides the ability to fabricate freeform constructs, embedded printing is very effective for printing extremely low-viscosity inks through employing the aqueous two-phase system (ATPS)^[152,153]. One of the earliest works using the concept of embedded printing was reported in 2011^[154]. Therein, fugitive ink filaments were printed in a photocurable hydrogel bath to form omnidirectional patterns, e.g., vascularized constructs. Later on, in 2015, an upsurge in the field of embedded bioprinting occurred by three concurrent papers. The guest-host complexes were exploited to print structures within self-healing support hydrogels through supramolecular assembly^[155]. The other approach employed granular hydrogels as the supporting bath to create large-aspect-ratio 3D structures^[156]. A freeform reversible embedding of suspended hydrogels (FRESH) method, which relied on the supporting bath’s thermoreversible properties, was also reported in the same year^[157]. Very high resolutions, down to 20 µm, have been achieved for acellular inks using the upgraded versions of the embedded printing technique^[158]. This capacity is attributed to the compatibility of the embedded printing method with low-viscosity inks and the presence of a supporting matrix, which enhances the fidelity of the printed structures.

Since the embedded bioprinting job is performed inside a supporting bath, the rheological and mechanical properties of this substrate should be precisely controlled. In addition to the importance of the nozzle’s inner diameter, which is also critical in conventional extrusion-based bioprinting, the nozzle tip’s outer diameter is a crucial factor determining the bioprinting fidelity. A needle with a large outer diameter may disturb an excessive amount of hydrogel in the bath and preclude its recovery, which eventually results in bubble-formation or lower resolution. The nozzle’s rigidity is also critical considering that nozzles with low stiffness may bend or break while moving inside the viscous hydrogel bath^[119]. Embedded bioprinting can work with lower ranges of viscosity values for the bioinks comparing to other extrusion modalities. However, the existence of the supporting bath in embedded bioprinting sets practical limits on the bioink flow rate and the nozzle displacement. Particularly, since the nozzle’s fast movement may agitate the supporting bath and damage the bioprinted structure, the average printing speed is generally lower than other extrusion methods.

As for multi-material embedded bioprinting, we employed a multi-nozzle technology to perform embedded multi-material bioprinting in a Pluronic F-127 bath^[119]. A set of 27G needles were bundled together and placed in metal-tube supports in a telescopic fashion so that the needles do not fail during printing. This multi-material bioprinter could print cell-laden structures with an acceptable switching rate between the materials. One impediment in this design is the limitation in the number of clustered needles, which restricts the versatility of the multi-material bioprinting. Therein, alginate with different colors was used to illustrate the feasibility of multi-material bioprinting. However, using materials with diverse rheological properties entails a custom design for each of the nozzles, which makes the multi-material embedded bioprinting a cumbersome job. The thermosensitive rheological properties of the supporting bath should also be precisely designed so that different bioinks can be properly embedded. These complications also made the embedded bioprinting technology a challenging technique for fabricating multi-material constructs. Although significant advancements have been reported in the embedded bioprinting method^[158], this technique is in the early stages of development, and more effort is needed to foster its maturation for multi-material bioprinting.

3.2. Inkjet bioprinting

Inkjet bioprinters are perhaps the most affordable type of bioprinter and can be easily modified from commercially available two-dimensional (2D) ink-based paper-printers at a low cost^[159]. However, the pricings for commercially available inkjet bioprinters are generally more expensive than extrusion or vat-photopolymerization bioprinters. One advantage of inkjet bioprinting is that it can obtain higher printing resolutions (up to 50–75 µm) compared to extrusion-based bioprinters^[38], as already mentioned. While the droplet size is usually predetermined by the printing nozzle and the viscosity of the bioink used, the resolution can be changed by the movement speed between the nozzle and the stage, bioink surface energy (contact angle between the droplet and the substrate), and the polymerization speed post-printing. For bioinks with higher viscosities (greater than 1,000 mPa s), stronger jetting mechanisms, such as highly localized acoustic pressure, can be utilized to reach ~100-µm resolution^[160]. The multi-material inkjet bioprinters can rapidly and precisely deposit multiple bioinks to form heterogeneous constructs with gradient properties by spatially varying the droplet sizes of biomaterials, cells, and growth factors^[161]. Another feature of the inkjet technology is that it works with bioinks with low viscosities on the order of 10 mPa s ^[65]. Unlike single-nozzle bioprinters, multi-material inkjet bioprinters usually do not share printheads (i.e., each material has its own nozzle)^[162]. Therefore, there is limited chance for cross-contamination during the bioink switches.

Multi-material inkjet bioprinting technology has been long-explored for building tissues and their models. Some pioneer works can be traced back to as early as 2003, where organic molecules and aggregates have been shown to be able to patten onto solid supports and form stable and functional cellular assemblies^[163,164]. A fairly complex printhead design, such as a 3-by-3 parallel nozzle array, was demonstrated to be a feasible way to create heterogeneous engineered tissues. The high-precision dispensing ability is especially beneficial for drug screening and development^[40]. It has been demonstrated that this technology can dispense different compounds into 384-well plates for high-throughput tests, such as biochemical assay, cell-based reporter-gene assay, and cytotoxicity assay, with clean and reproducible results^[165].

One main limitation of this technology is the difficulty in fabricating cellular constructs with clinically relevant sizes. This is caused by the fact that the mechanical strengths of compatible bioinks are generally low prior to necessary crosslinking, which limits the maximum building heights of the printed constructs^[38]. Due to this reason, constructs with high aspect ratios are also difficult to fabricate. Efforts have been made to overcome those challenges. For example, an inkjet bioprinting platform printed a silk fibrin/alginate bioink on a motorized stage^[166]. When the bioink was collected on the stage, it gradually moved the bioprinted construct into a cell-friendly crosslinker bath while the fabrication continued for the part which was still out of the crosslinker solution. This method offers high cell viability (evidenced by greater than 6-fold increase in metabolic activity) and high-aspect-ratio shapes (height-to-width ratio greater than 1). Other limitations of inkjet bioprinters include the incompatibility of bioinks with high cell densities (greater than 1×10⁶ cells/mL) due to viscosity issues and relatively low printing fidelity compared to the vat-photopolymerization method^[38]. For more technical and translational aspects of inkjet bioprinting, we refer the readers to other excellent reviews^[41,63,167].

3.3. LIFT bioprinting

Although the LIFT technology was invented in the 1980s^[168], it has rarely been applied for the fabrication of complex multi-cellular and multi-material constructs. Specific printing parameters associated with each cell-laden hydrogel comprise one of the greatest challenges of the LIFT method that limit its use in multi-material bioprinting^[169]. Koch et al. designed a multi-material LIFT bioprinter using a carrier for transferring different cells coated on the donor slides (Figure 9A)^[170,171]. They deposited human skin cell lines and mesenchymal stem cells (MSCs) with high survival rates^[170]. Fabien Guillemot, the founder of Poietis (France), proposed a high-resolution positioning platform, allowing the exchange of up to six different donors using a carousel holder^[172]. Therein, different cell types could be switched on demand using a motorized system to fabricate multi-cellular structures (Figure 9B). This bioprinter was applied for 3D assembly and patterning of multiple cells and biomaterials. However, due to the complexities of the LIFT method regarding the required high precision and calibration, the exchange process can affect critical parameters associated with jet-initiation in the deposition procedure. A method that allows well-controlled spatial micropatterning of cell-laden beads have recently been developed. Using this method, multi-cellular embryoid bodies and tumor spheroids were produced with precise control over the size and shape of the beads (Figure 9C-E)^[173]. This method is only limited to self-aggregating and self-assembling cells. Further improvement is needed in the development of affordable and accessible LIFT bioprinters for the fabrication of heterogeneous structures.

Figure 9-.

LIFT multi-material bioprinting. A) Multi-component LIFT bioprinting using rapid replacement of donor-slides by a carrier (scale bars = 500 mm). Reproduced with permission.^[171] Copyright 2012, Wiley-VCH. B) High-resolution ribbon switching system based on a carousel holder with a loading capacity of five different ribbons. Reproduced with permission.^[304] Copyright 2010, Elsevier. C-E) Illustration of multi-cellular LIFT mechanism. Reproduced with permission.^[173] Copyright 2019, Elsevier. C) alginate bead-deposition; D) conversion of microbeads into core-shelled structure; E) aggregate-formation.

3.4. Vat-photopolymerization bioprinting

The past decade has seen increasingly rapid advances in vat-photopolymerization 3D printing, including microstereolithography that enables micron-scale features^[174,175], continuous liquid interface production (CLIP) commercialized by Carbon enabling 100 times faster printing compared to conventional vat-photopolymerization 3D printing^[176], volumetric 3D printing inspired by computed tomography^[177,178], and more recently, xolography 3D printing based on using photoswitchable photoinitiators^[68]. However, the majority of existing studies have focused on single-material 3D bioprinting. When compared with other bioprinting methods, the development of multi-material vat-photopolymerization-based 3D bioprinters is still relatively challenging and limited. Different strategies have been applied to overcome this problem, including using multi-vat, sequential injection, sequential deposition, and multi-wavelength.

3.4.1. Multi-vat-photopolymerization

The most common multi-material vat-photopolymerization 3D-printing method is based on using a system that automatically alters the resin vat. Choi et al. developed a system with multiple resin containers on a top-down DLP exposure with an automated rotating vat-carousel mechanism^[179]. The use of cleaning solutions to wash the uncured resin turned out to be damaging for features smaller than ~300 μm^[179]. Similarly, composite structures as small as 30 µm with shape-memory hydrogels were 3D-printed using a downward exposure (Figure 10A, B)^[180]. However, top-down exposure is not considered an efficient and cost-effective method for bioprinting since the resin vat must be completely filled with cell-laden hydrogels. Zhou et al. reported a multi-material 3D printer based on the bottom-up exposure method with a rotating resin container system. To minimize cross-contamination, two different cleaning steps, including rough cleaning using two soft brushes and fine cleaning using ultrasonication along with a dryer, were implemented^[181,182]. Likewise, an osteochondral composite scaffold was fabricated using an efficient washing step with two sponges and consequently sub-merging the printed layer in a cleaning medium^[183]. In the same way, Grix et al. developed a liver model with vascular networks for organ-on-a-chip applications^[184]. Photopolymerizable degradable PEG-bis-(acryloyloxy acetate) (PEG-bis-AA) and cell-laden GelMA hydrogel-precursor (cell density of ~1×10⁸ cell/mL) were 3D-bioprinted and lumens, as small as 200 µm, were formed by degradation of the PEG-bis-AA hydrogel after 3 days in culture (Figure 10C-F). The 3D-bioprinted liver model showed a higher albumin and gene expressions compared to monolayer controls over 14 days of the post-bioprinting period. The multi-step washing process in these methods significantly increased the fabrication time and might damage the bioprinted parts especially when they are made of soft hydrogels. In addition, several labs fabricated multi-component structures by manually changing the resin vat for different applications, such as fabrication of a porous membrane within a microfluidic device and soft hydrogel-based actuators^[185–187].

Figure 10-.

Multi-vat photopolymerization. A) Automated material-exchange process based on a top-down exposure system enabled fabrication of shape memory structures. B) A multi-component thermo-sensitive construct. Reproduced under the terms of the CC-BY Creative Commons Attribution 4.0 International license (https://creativecommons.org/licenses/by/4.0).^[180] Copyright 2016, The Authors, published by Springer Nature. C) Schematic illustration of a multi-vat multi-material bioprinter based on the bottom-up DLP approach. D-F) Vasculature networks 3D-bioprinted with degradable PEG within cell-laden GelMA hydrogel. Reproduced under the terms of the CC-BY Creative Commons Attribution 4.0 International license (https://creativecommons.org/licenses/by/4.0).^[184] Copyright 2018, The Authors, published by Multidisciplinary Digital Publishing Institute.

3.4.2. Sequential injection

The sequential injection of different (bio)inks is another common multi-material photopolymerization approach to minimize bioink usage. This method was initially introduced by Chen and colleagues, who developed a method based on the delivery of pre-hydrogel solutions into the gap between a hydrophobic glass and a servo-stage that enabled fabrication of heterogeneous scaffolds with features as small as 50 µm. In this top-down DLP 3D printing mechanism, rapid biomaterial-exchange with minimal material consumption was achieved by injection of 10 µL of a pre-hydrogel solution. Since the inlet of the monomer was placed within the servo stage, this method could only be used for the fabrication of lattice structures such as honeycomb and woodpile microstructures that allow efficient cleaning with free transfer of solvent within the structure. Recently, Han et al. addressed this limitation by placing multiple ink inlets equipped with precise control valves on the side of the servo stage^[188]. Complex structures with features as small as 100 μm were 3D-printed using this method (Figure 11A-E). To achieve an efficient cleaning step with minimal cross-contamination, prolonged pumping time is required, which eventually increases swelling, material-consumption, and printing time that are not desirable for bioprinting.

Figure 11-.

Sequential injection multi-material vat-photopolymerization approaches. A) Schematic illustration of dynamic fluidic control-based multi-material 3D printer. B-E) Multi-component structures: (B) two-component tensegrity structure with high aspect ratio beams, (C) Taiji symbol, (D) two-component bilayer micro-capillary structure, and (E) three-component helix structure. Reproduced with permission.^[188] Copyright 2019, Elsevier. F) Schematic illustration of the DLP-based bioprinter based on the sequential injection of different bioinks into a microfluidic chamber. G) the four-step bioprinting process inside the microfluidic chip. H) multi-component constructs: a two-component GelMA structure (left), and a three-component star-shaped pyramid (right). I) A tendon-to-bone model bioprinted with this approach with different cell-laden GelMA hydrogels. Reproduced with permission.^[189] Copyright 2018, Wiley-VCH.

Alternatively, a multi-material DLP 3D bioprinter was developed based on the sequential delivery of up to four hydrogels into a custom-designed microfluidic chamber (Figure 11F-H)^[189]. The material-exchange was performed by flushing the microchannel with phosphate-buffered saline (PBS) to wash out the previous bioink. In addition to pneumatic valves for sequential delivery of different bioinks, we implemented an elastomeric polydimethylsiloxane (PDMS) membrane within the microfluidic chip allowing in-chip depth control for constructing 3D objects (Figure 11G-H). This method was exploited to bioprint multi-component structures using three hydrogels in 20 seconds, which is significantly faster than other multi-material vat-photopolymerization methods. The same technique was successfully employed for the fabrication of multi-cellular structures such as a tumor angiogenesis model and a tendon-to-bone insertion model (Figure 11I). Satisfactory proliferation and metabolic activities were achieved on day 1 and day 7 after bioprinting. The chance of dehydration during the bioprinting process is significantly lower for this method since the printing is happening in a closed chamber. The main drawback of this method is that the remaining hydrogel precursors in the chamber (~125 µL) will be discarded during the washing step. Furthermore, the printing area is limited to the closed microfluidic chamber space that restricts the height on the Z-axis. In addition to aforementioned methods, human induced pluripotent stem cell (iPSC)-derived cardiomyocytes and hepatic progenitor cells (hiPSC-HPCs) were 3D-bioprinted by manually pipetting of cell-material solutions between the gap^[51,190]. The bioprinting process consists of printing iPSC-derived cells followed by washing extra prepolymer solution with PBS and patterning of supporting cells that fill in the empty space of the pattern. After 7 days, high maturity marker expressions were observed in 3D-bioprinted multi-cellular liver and cardiac models.

3.4.3. Sequential deposition

Another approach in multi-material vat-photopolymerization is the sequential deposition of different precursors, in which no closed chamber is needed for delivering the materials. Wang et al. developed a bottom-up DLP exposure-based multi-material 3D printer by sequentially depositing different material droplets onto a rotating wheel^[191]. Using this technique, multi-material structures with negative thermal expansion were successfully fabricated. This method, however, suffers from two main drawbacks in multi-material 3D printing, including the increased risk of contamination and prolonged fabrication time (~6 hours for a structure with a volume of 220 mm³). More recently, the same group developed a novel minimal-waste material-exchange process using an air jet to remove residual resin attached to the previous layer^[192]. They concluded that this method is 58% faster than existing methods that are based on cleaning solutions. However, in the case of resin contamination, the hydrophobic layer must be washed. This step, which is only functional for resins with relatively low viscosities, may disrupt the printing process. In addition, the air jet might dehydrate the printed layer and lead to unexpected deformation of the layer. Recently, the same method has been applied for 3D printing of complex hybrid structures consisting of covalently bonded elastomer, rigid polymer, and soft hydrogels (Figure 12A-E)^[105]. Miller’s group also developed a 3D bioprinter using a very similar approach capable of utilizing up to four bioinks at a time in a semi-automated way (Figure 12F-H)^[193]. Using this approach, several multimaterial architectures, such as structures consisting of acellular and cellular domains, were bioprinted with a high cell density (2×10⁷ cells/mL) in a GelMA bioink. This method requires manual washing and drying of the printing platform during the material-exchange process, which significantly increases the fabrication time and could lead to reduced cell viability. Another limitation of their technology is the inability to fabricate large constructs since this method is based on deposition of material droplets. Hohnholz et al. used a top-down exposure system equipped with an aerosol jet mechanism to sequentially deposit different resins^[194]. In this approach, various compositions of different materials from one layer to another layer are attainable using the aerosol jet. Due to the complexity of the system and the use of a laser beam instead of a DLP projector, the fabrication time is longer than other multi-material vat-photopolymerization 3D printing methods. Furthermore, by employing the SLA and fused-deposition modeling (FDM) methods, a hybrid 3D printing method was proposed for the fabrication of multi-material scaffolds with soft hydrogels and rigid scaffolds^[195].

Figure 12-.

Sequential deposition in vat-photopolymerization 3D printing technology. A) High-Efficiency multi-material 3D printer based on the sequential deposition of different liquid puddles. B) Hydrogel composite reinforced by rigid polymer structure. C) Composite structure under tensile test. D) Printed multi-material structure consisting of rigid polymer, elastomer, and PEG-diacrylate (PEGDA) hydrogel. E) High stretchability of the printed structure. Reproduced under the terms of the CC-BY Creative Commons Attribution 4.0 International license (https://creativecommons.org/licenses/by/4.0).^[105] Copyright 2021, The Authors, published by American Association for the Advancement of Science. F) Schematic illustration of a multi-material stereolithography bioprinter. G) Printed structure with four different bioinks. H) Heterogeneous construct with an embedded vessel. Reproduced under the terms of the CC-BY Creative Commons Attribution 4.0 International license (https://creativecommons.org/licenses/by/4.0).^[193] Copyright 2021, The Authors, published by Springer Nature.

3.4.4. Multi-wavelength photopolymerization

In addition to the above-mentioned approaches that are based on different material-exchange mechanisms, more recently, a continuous multi-material technique was introduced based on selective polymerization of ink components using a multi-wavelength strategy^[196]. The photocurable ink consisted of acrylate and epoxy monomers with their corresponding photoinitiators with different absorption spectrums. This allowed independent polymerization of various photocurable monomers using different photoinitiators and wavelengths. A soft hydrogel could be achieved by exposing the ink to blue light, whereas a stiff material can be printed under short-wavelength (UV) irradiation (Figure 13A-B). This strategy enabled 3D printing of multi-material structures containing soft hydrogels along with hard solid epoxide networks (Figure 13C-D). Therefore, spatial heterogeneity in 3D-printed parts can be achieved by exposing the resin to different wavelengths (365 and 450 nm) to tailor the mechanical properties of the structure^[196,197]. However, due to the toxicity of epoxide monomers, the application of this method is still only limited to non-biological applications. Further investigation in developing biocompatible formulation is, therefore, needed to implement this method for bioprinting.

Figure 13-.

Multi-wavelength photopolymerization approach. A) Schematic illustration of the multi-wavelength 3D-printing setup. B) multi-component photoresin consisting of acrylate and epoxy monomers. C) Design of a two-component sea star construct. D) Swelling actuation of the 3D-printed part with soft hydrogels and stiff hydrophobic networks. Scale bars = 25 mm. Reproduced under the terms of the CC-BY Creative Commons Attribution 4.0 International license (https://creativecommons.org/licenses/by/4.0).^[196] Copyright 2019, the Authors, published by Springer Nature.

The printing rate in multi-material vat-photopolymerization is mainly dependent on the model complexity, material-switching, and washing process. 3D printing of structures in which each layer consists of multiple materials is more laborious. The fastest material-exchange rate that has been reported is based on sequential deposition of inks on a glass plate with the air-jet cleaning process that takes around 3 s^[198]. Using this method, structures with different materials side-by-side within the same layer could be 3D-printed with a fabrication time ~2.2 times longer than single-material 3D printing^[192]. In methods based on sequential injection, bioinks could be washed within 5–20 s by the subsequent flow of PBS^[189]. These methods are more suitable for 3D bioprinting due to using PBS for cleaning that avoids unwanted dehydration of the printed layers. Ink-replacement rate must be optimized based on the maximum air jet pressure and flow rate of the washing buffer that would not damage the printed layers.

Although outstanding advances in the field of multi-material vat-photopolymerization have been reported, several significant challenges for vat-photopolymerization bioprinting of biomimetic tissue constructs remain. The incapability for continuous printing of hydrogels with clinically relevant dimensions and the inability to fabricate multi-material complex constructs with high efficiency and accuracy are among the major concerns^{[180,182,186,195,199–201]}. The main challenge in using multiple polymers during the polymerization process is cross-contamination while switching between the different bioinks. The washing step might lead to deformation and dehydration of 3D-bioprinted constructs during the bioprinting process that could limit the application of this method in fabricating small-scale constructs.

In general, multi-vat photopolymerization is more suitable for fabricating organ-sized objects in comparison to other methods since the bioprinting area is not restricted by the material-exchange mechanisms. In contrast, methods based on sequential injections and deposition are more efficient in terms of material-switching, cleaning process, and bioink-consumption. Washing and material-exchange mechanisms are time-consuming and complex steps in all vat-polymerization methods other than multi-wavelength photopolymerization that would elongate the fabrication time for multi-material constructs. This could also affect the cell viability and lead to dehydration of 3D-bioprinted parts. Although many of the aforementioned techniques are used for 3D printing, they can be implemented readily in 3D bioprinting by using biocompatible photocurable bioinks. The reader may refer to other reviews on vat-photopolymerization 3D printing for more technical aspects^{[47,60,202,203]}.

4. Commercial multi-material bioprinters

The arrival of commercially available 3D bioprinters has enabled access to these technologies for many more research laboratories and industrial utilities. The bioprinting market is expected to reach a value of approximately 4.1 billion US dollars by 2026^[204]. There are numerous commercial bioprinters in the market based on different technologies. Commercial bioprinters were initially prohibitively expensive and not accessible to most academic research facilities, making their use largely exclusive to the industry. However, many companies recently have addressed the unmet need for accessible 3D bioprinters by producing modular and affordable bioprinters. Since commercial bioprinters play an important role in the advancement of multi-material bioprinting technologies, we devoted this section to introducing a number of commercial multi-material 3D bioprinters.

4.1. Extrusion bioprinters

Due to the affordability and simplicity of the extrusion-based bioprinting method, most commercial bioprinters are based on this technology with the highest market share^[204]. EnvisionTEC, a globally leading 3D printing company, developed the first commercial bioprinter in 2000. The 3D-Bioplotter was initially invented by Mulhapt’s group at the University of Freiburg and then commercialized by EnvisionTEC^[205]. Since then, several research laboratories around the world have employed the extrusion bioprinting approach for tissue engineering and other fields of biomedical engineering. Another extrusion-based multi-material bioprinter, NovoGen MMX, was developed by Organovo in 2009 and could deliver cells and hydrogels using two syringes^[74,206]. However, this company does not market its bioprinter anymore; instead, it only sells this bioprinter mostly to pharmaceutical companies for drug screening^[74]. Other than the two pioneering extrusion-based bioprinters, 3D-Bioplotter and NovoGen MMX, many companies have developed bioprinters based on different extrusion bioprinting methods. Several companies such as Allevi, Aether, CELLINK, and REGEMAT use value adds-on to provide more affordable and accessible 3D bioprinters. Allevi and CELLINK, arguably two of the largest bioprinter-providers, have developed several extrusion-based bioprinters with different specifications and price ranges. Allevi 3 bioprinter has three extruders with precise temperature control from 4 °C to 160 °C enabling bioprinting a wide range of bioinks. CELLINK BIO X6 bioprinter consists of six printheads that allow fast and versatile multi-material bioprinting. Aether, another hybrid multi-material bioprinter, is capable of depositing twenty-four materials at a time using eight syringe extruders, two heated nozzles and fourteen droplet jet extruders. Aspect Biosystems, a Canadian bioprinter company, has also developed a microfluidic bioprinter based on its patented Lab-on-a-Printer (LOP) microfluidic technology^[132]. Biological fibers can be deposited with varied diameters by changing the flow rate in the flow-focusing LOP co-axial method. Similar to Organovo, Aspect Biosystem only provides its bioprinters based on partnership programs. In addition to the conventional extrusion-based bioprinters, a scaffold-free bioprinter, based on the Kenzen method, was created by Nakayama at Saga University and commercialized by Cyfuse Biomedical. This approach is capable of 3D assembling and positioning of different spheroids as building blocks of biological constructs using an array of 160-µm-thick stainless-steel microneedles^[207,208]. Revotek is another scaffold-free bioprinter based on a proprietary technology named Biosynsphere which makes it suitable for fabrication of scaffold-free vascular structures. Blood vessels with iPSC-laden bioinks were successfully 3D-bioprinted using this technology and implanted into rhesus monkeys^[209]. Furthermore, the BioassemblyBot (Advanced Solutions), a six-axis bioprinter, employs a robotic arm for the deposition of different materials in 3D space, enabling multi-material bioprinting^[210].

4.2. Inkjet bioprinters

In comparison with commercial extrusion-based bioprinters, the number of available inkjet bioprinters on the market is very limited. The RASTRUM 3D inkjet desktop bioprinter by Inventia is equipped with an inbuilt laminar flow hood with a deposition capability of droplets as small as 5 nL. This multi-material bioprinter can rapidly print up to eight different droplets of cells and matrix components simultaneously in a reproducible way^[211]. Ricoh, a Japanese imaging and electronics company, has developed a new inkjet head for precise and gentle deposition of different cell types. Ricoh has also partnered with Elixirgen Scientific, a pioneer company in high-speed production of iPSCs, in drug screening and precise positioning of multiple disease-specific iPSC lines^[212]. The Jetlab printer, manufactured by MicroFab Technologies, is a high-precision industrial drop-on-demand piezoelectric inkjet printer with a large printing area. This 3D printer is capable of microdispensing up to four different materials simultaneously. Since this 3D printer is not specifically designed for bioprinting, prior optimization on the viscosities of bioinks and dispensing parameters is essential to eliminate nozzle clogging during cell-laden bioprinting. Another inkjet 3D bioprinter is CellJet, which was developed based on the Digilab inkjet technology. Using sixteen independent channels, this technology ensures high cell viabilities (~95%) through non-invasive dispensing of droplets ranging from 20 nL to 4 µL by preventing cells from the adverse impact of shear forces. In addition, the bioprinter software enables full control over dispensing parameters such as the inkjet speed and the height, which are quite critical for viscous bioinks and delicate cells. This 3D bioprinter can be placed in most biological safety cabinets^[213,214]. There is still an unmet need for commercially available, affordable, and on-the-bench inkjet 3D bioprinters.

4.3. LIFT bioprinters

Despite the complexity of the LIFT technology, it has been successfully commercialized. NGB-R™ is the first commercially available LIFT 3D bioprinter that has been developed. This 3D hybrid bioprinter uses the laser-assisted bioprinting (LAB) technology along with multiple dispensers, allowing fabrication of heterocellular patterns with a single-cell resolution^[215–217]. Besides, Poietis partnered with large cosmetic and pharmaceutical companies such as L’Oréal and BASF to assess the efficacy and toxicity of cosmetic products and drug candidates using 3D-bioprinted tissue models such as hair follicle and human skin^[218]. Precise Bio, a North Carolina-based company is another leading company in LIFT bioprinting. It achieved the first transplantation of a 3D-bioprinted cornea graft into an animal using its innovative laser-assisted four-dimensional (4D) biofabrication technology. This technology allows the fabrication of complex tissues with single-cell resolution and spatial accuracy with a high cell viability (>95%)^[219].

4.4. Vat-photopolymerization bioprinters

Even though vat-photopolymerization is among the high-resolution and cost-effective 3D bioprinting methods, there are only a few commercially available multi-material bioprinters. To the best of our knowledge, Cellbricks is the only commercially available multi-material bioprinter based on vat-photopolymerization. This bioprinter employs rapid material-exchange using multiple vats for bioprinting of heterogeneous constructs with 10-µm accuracy^[184,220]. Most of the SLA 3D printers can also be used for bioprinting cell-laden hydrogels. For example, the EnvisionTech 3D printer has been widely used in research labs for bioprinting jobs^[221,222]. Lumen X, another commercial SLA 3D bioprinter, was developed by Miller’s group and commercialized by CELLINK^[223]. Stemaker™ bioprinter (Allegro 3D) is also a DLP-based 3D bioprinter that is capable of direct printing in multi-well plates. In addition, EFL’s DLP bioprinter has been used for fabrication of nerve guidance conduits using GelMA hydrogels^[224]. Although these bioprinters are not capable of performing multi-material 3D bioprinting jobs, heterogeneous structures can be successfully 3D-bioprinted by manually altering the resin container. The readers may refer to Table 1 for a list of commercial most common multi-material bioprinters and selected applications of them that have been reported in the literature.

Table 1-.

The most common commercial multi-material bioprinters available in the market.

Bioprinter	Company	Technology	Selected Applications
3D NovoGen	Organovo (USA)	Extrusion	Liver model, kidney model, intestinal model, vascular constructs[305–309]
3D-Bioplotter	EnvisionTEC (Germany)	Extrusion	Tendon-to-bone model, adipose tissue model[310,311]
3D Discovery™ Evolution, BioFactory™	regenHU (Switzerland)	Extrusion (capable of using electrospinning and melt electrowriting technology)	Bone-regeneration, perfusable cardiac patches, meniscus, skin[312–315]
Aether 1	Aether (USA)	Extrusion (capable of depositing twenty-four materials at a time)	Superficial skin cancer therapy[316]
Allevi 1, 2, 3	3D Systems (USA)	Extrusion (with precise temperature control from 4 °C to 160 °C)	Angiogenesis and vascularization model[317,318], Liver model[319], 3D cell culture[319]
BAT Series	nScrypt (USA)	Extrusion	Spheroid-manufacturing[320], knee cartilage[321]
BIO X, BIO X6, INKREDIBLE	CELLINK (USA)	Extrusion (compatible with various modular printheads)	Vascularized skin graft[322], bone and nerve tissues[323], corneal stroma[324]
BioAssemblyBot	Advanced Solutions Life Sciences (USA)	Extrusion (six-axis printhead with a robotic arm)	Vocal fold tissue, tumor modeling[287]
Bio-Architect®-Pro, WS, x, Sparrow	REGENOVO (China)	Extrusion	Cartilage[325], sweat gland[325], skin[326], neuronal tissue[327]
BioScaffolder	GeSiM (Germany)	Extrusion	Multi-cellular tumor spheroid[261,262]
BRINTER 1	BRINTER (Finland)	Extrusion	-
Fabion 1, 2	3D Bioprinting Solutions (Russia)	Extrusion (compatible with various modular printheads)	Vascularized thyroid gland construct[328,329]
FELiX BIOPRINTER	FELIXprinters (The Netherlands)	Extrusion	Bioprinting of human iPSCs[330]
FLUX-1	Frontier Bio (USA)	Extrusion	-
Genesis™ I and II, Reactor™, Octopus™, BioFDM™	3D Biotechnologies Solutions (USA)	Extrusion	-
MedPrin	MedPrin (China)	Extrusion	Lung cancer model[331], brain tumor model[332], artificial dura mater[333]
PCPrinter BC	Particle Cloud (China)	Extrusion	Bone[334]
REG4LIFE, BIO V1	Regemat 3D (Spain)	Extrusion (compatible with various modular printheads)	Articular cartilage[335]
Revotek	Sichuan Revotek (China)	Extrusion (scaffold-free 3D bioprinting of vascular structures)	Blood vessels[336,337]
ROKIT INVIVO	Rokit (South Korea)	Extrusion (with built-in cell incubator)	Cartilage-regeneration[338], vascularized tumor model[143,339]
RX1	Aspect Biosystems (Canada)	Extrusion (Based on Lab-on-a-Printer (LOP) microfluidic technology)	Glioblastoma tumor model[340], neuronal tissue[341], brain tissue model[342], contractile smooth muscle tissue[343]
SUNP BIOMAKER	SunP Biotech International (USA)	Extrusion	Cancer model[344], liver[345]
T&R Biofab	T&R Biofab (South Korea)	Extrusion	Pre-vascularized cardiac patch[23], superficial knee cartilage and subchondral bone[346], organoid-manufacturing[347]
WeBio	WeBio (Argentina)	Extrusion	Cartilage regeneration[348]
Autodrop	Microdrop	Inkjet
RASTRUM™	Inventia (Australia)	Inkjet	Production of multi-cellular spheroids[349]
CellJet	Digilab (USA)	Inkjet	Bioprinting of human MSCs[350]
Hp D300e	Hewlett-Packard (USA)	Inkjet	Skin model[351]
Cellbricks	Cellbricks GmbH (Germany)	Vat-photopolymerization	Liver model[184]
NGB-R	Poietis (France)	LIFT	MSC-patterning[352], skin model
Precise Bio	Precise Bio (USA)	LIFT	Ophthalmology[353]
Regenova, S-PIKE	Cyfuse Biomedical (Japan)	Scaffold-free bioprinting	Liver tissue model[354], glioblastoma invasion model[355]

5. Applications of multi-material bioprinting

5.1. Tissue fabrication

The current accomplishment in the bioprinting field has opened new arenas for regenerative medicine, tissue modeling, and beyond. However, many existing developments, especially at the beginning of bioprinting technology, have been limited to the creation of monocellular and homogeneous environments. In contrast, human tissues are multi-cellular constructs with hierarchical structures. They are mechanically anisotropic and intrinsically heterogeneous. It is challenging, or even unrealistic, to use one single material to analog all human tissues. Therefore, advances in materials science and innovations in multi-material bioprinting technologies are critically needed towards the fabrication of more biomimetic tissue and organ constructs.

Multi-material dispensing systems, such as co-axial and microfluidic nozzles, can offer simultaneous deposition of different bioinks to build complex constructs. For example, the co-axial nozzle has been employed to engineer 3D brain-like structures using bioinks comprising peptide-modified gellan gum with primary cortical neurons^[103]. The co-extrusion of low-viscosity bioink and crosslinker enabled by the co-axial nozzle allowed the bioprinted filaments to have mechanical strength to fabricate laminated brain-like structures with a high aspect ratio. In such multi-material constructs, the precise arrangement of material concentration, cell distribution, and height gradients among different layers are important to resemble tissues with layered architectures.

Besides the ability to recreate the structural complexity of various tissues, another salient feature of multi-material technologies is their potential to improve the printing speed (mainly for extrusion-based bioprinters, see Section 3.1 for details). Because most existing extrusion bioprinting tasks happen at ambient environment, the viability of encapsulated cells may be decremented due to the exposure to room temperature and low humidity for a prolonged period. Such a consideration is critical for the fabrication of large-scale cellular constructs, especially when the direct ink-writing (DIW) methods are used. Therefore, it is preferred to complete the bioprinting tasks in a quick manner. A co-axial printhead has been designed to fabricate complex liver tissue models with compartmentalized arrangement similar to hepatic lobule structures through single-filament writing^[82]. The highly customized design of the nozzle allowed the co-extrusion of hepatocellular cells, endothelial cells, and a lumen at the same time (Figure 14A-C). Without employing this multi-material single-nozzle design, bioprinting such a construct would have been a cumbersome job as each compartment of the hepatic lobule would have to be bioprinted separately using a single-material printhead. The technology enabled the buildup of multi-scale functional, heterogeneous, and multi-cellular hepatic structures in a facile and rapid fashion. Other tissue-engineering applications are also achievable with adjustment to the co-axial nozzle design.

Figure 14-.

Extrusion-based multi-material bioprinting for tissue-fabrication. (A) Schematic illustration of multi-material extrusion printhead design for hepatic lobule printing. (B) Left: Immunostaining of CD31 (green) and f-actin (red) for the bioprinted epithelial cells. Nuclei were counterstained in blue; right: Live (green) and dead (red) staining showing morphological changes and viability. (C) Immunostaining of CD31 (red), albumin (green), and MRP2 (green) for the bioprinted hepatic lobule with hepatocelluar cells. Nuclei were counterstained in blue. Reproduced with permission.^[82] Copyright 2020, Wiley-VCH. (D) Schematic showing the multi-material bioprinting setup for dual-cell MTU constructs fabrication. (E) Tensile behavior of the bioprinted MTU with both soft and rigid regions. (F) Fluorescence image showing bioprinted MTU after 7-day culture (green: DiO-labeled C2C12 cells; red: DiI-labeled NIH/3T3 cells). (G) Differential expression between the two cell types at the interface region (depicted by the dotted line) is observed. Reproduced with permission.^[225] Copyright 2015, Institute of Physics. (H) Schematic of dual-material FRESH printing using a collagen ink and a high-concentration cell ink. (I) Micrograph of FRESH-printed multi-component ventricle. (J) Side view of FRESH-printed ventricle stained with calcium-sensitive dye showing uniform cell distribution. (K) Calcium mapping of the subregion [yellow box in (J)] showing spontaneous, directional calcium wave propagation. Reproduced with permission.^[158] Copyright 2015, American Association for the Advancement of Science.

So far, single-nozzle technologies have been demonstrated to accommodate the co-extrusion of several different bioinks at the same time^[120]. It has been shown that human dermal fibroblasts, human hepatocellular cells, MSCs, and HUVECs could be hierarchically assembled and co-cultured using such a system to form complex tissues. Microfluidic nozzles can also be used in conjunction with co-axial nozzles to fabricate filaments with compartments containing different bioinks. An example using such a hybrid nozzle is the printing of PEG-fibrinogen/alginate bioinks^[80]. The alginate content of this compartmented bioink was ionically crosslinked by a crosslinker solution from the co-axial nozzle, therefore enabling the DIW ability of the low-viscosity bioink. The printed constructs were further stabilized by UV irradiation to initiate free-radical polymerization of the PEG content before the removal of alginate with ethylenediaminetetraacetic acid (EDTA). The co-cultured constructs showed a proper spread of C2C12 and highly aligned long-range multi-nucleated myotubes, with abundant and functional expressions of myosin heavy chain and laminin^[80]. The better recapitulation of the whole muscle histoarchitecture in vitro and in vivo by bioprinted multi-cellular hydrogels compared to cast-hydrogels also revealed the advantages of the multi-material bioprinting technology.

Bioprinters equipped with multiple dispensing modules are helpful in fabricating multiple materials with different processing conditions. For instance, (bio)inks that require different printing temperatures and pressures can be loaded into separate extruders and be printed individually at their own desirable conditions. This is particularly useful when fabricating materials with dissimilar natures, such as thermoplastics and hydrogels. Printing such combinations are important to provide mechanical resemblance while maintaining cellular support. The muscle-tendon unit (MTU), for example, possess regional differences in cell types and mechanical properties^[225]. Muscle is predominately composed of elastic fibrous myofibers while the tendon is stiff and rich in collagen fibers. It is traditionally challenging to mimic such distinctly different environments with a single material system. However, bioprinters with multi-material dispensing capacities can integrate the two parts seamlessly. A study bioprinted an MTU construct with thermoplastic polyurethane (PU)/C2C12 myoblast-laden bioink and PCL/NIH-3T3 fibroblast-laden bioink to mimic the muscle and tendon parts, respectively. The bioprinted MTU comprised of both soft and stiff sides, similar to the real tissue. The excellent viability, highly aligned cell morphologies, and increased MTU-associated gene expressions showed the potential of this multi-material bioprinting strategy for the reconstructions of complex tissues (Figure 14D-G).

Multi-nozzle bioprinting can also enable the direct writing of bioinks that do not comply with suitable rheological properties for bioprinting. One example is decellularized ECM (dECM)-derived bioinks. They have been proved to provide an optimized microenvironment leading to the growth of cells for specific tissues^[226]. However, they are usually too weak (i.e., low yield stresses) to withstand their weight during fabrication, resulting in poor spatial resolution. An exemplary printing method for providing stable structural support to soft biomaterials is to dispense the low-viscosity bioinks within a printed stiff construct. The bioprinting of dECM bioinks within printed PCL structures has been demonstrated to greatly improve the printing fidelity^[226]. Similarly, a study employed cell-laden hydrogels, PCL, and Pluronic F-127 to fabricate human-scale tissue constructs, such as ear, bone, and muscle^[227]. PCL and sacrificial Pluronic F-127 ensured the stability of the bioprinted structures during and after the fabrication, respectively. The cell-laden hydrogels were found to secrete the corresponding ECM with demonstrated long-term functionalities.

Multi-material bioprinting has been implemented in embedded bioprinting as well. As mentioned before, embedded bioprinting allows the deposition of multiple bioinks within a supportive liquid reservoir. The freeform modeling ability and the compatibility with low-viscosity bioink-writing well-comply with the trend of multi-cellular whole-organ fabrication. Cells and microtissues can also be bioprinted directly through this method^[228]. A multi-material bioprinting system comprising a collagen ink, a cell-only bioink, and a gelatin-based supporting bath displayed the potential to recreate cardiac ventricles and tri-leaflet valves with high resolution^[158]. The bioprinted ventricles containing human cardiomyocytes exhibited synchronized contractions and directional action potential propagation, recapitulating the structural, mechanical, and biological properties of native cardiac tissues (Figure 14H-K). Another study employed embedded bioprinting to bioprint decellularized cardiac bioinks^[229]. The fabricated multi-material, thick, vascularized, and perfusable cardiac patches recapitulated the structural, immunological, and anatomical properties similar to those of human hearts. Cell viability could be well-preserved owing to the mild and all-liquid environment of the supporting reservoirs. This cell-friendly feature is important for time-consuming multi-material fabrication on a whole-organ scale. Other novel functions such as the erasing of existing filaments by withdrawing the extruded bioinks also offer flexibilities to correct errors during a complicated tissue-construction task^[36].

5.2. Vascularization

Most tissues in the human body require vessels and microvasculature systems to supply oxygen and nutrients and to remove wastes. Although biocompatible hydrogels have been shown to support cellular activities, their pore sizes are typically on the nanometer scale and incapable of enabling sufficient mass transfer when the thickness of the scaffolds is greater than ~800–1,000 µm^[159]. Vascularization is critical to ensure the survival and functions of thick scaffolds but is a major challenge in fabrication process at the same time. The precise positioning of biomaterials opens the opportunities for creating controllable, repeatable, and freeform modeling of engineered vascular channels. An intuitive strategy is to leave spacings between printed filaments as open pores or channels. Such spacings have been proven to effectively support cell survival in relatively large constructs^[230]. However, the open pores and channels generated using this strategy may not be reliable. For instance, the swelling, shrinkage, and degradation of printed hydrogels can be significantly affected, or the construct may become structurally unstable, which may clog the spacings. The ECM deposition inside scaffolds by cells can also deteriorate the channels^[231].

Multi-material bioprinting can be employed for fabricating physiologically relevant vasculature through co-printing of hydrogel and sacrificial materials. This strategy relies on the removal of sacrificial channels within the bioprinted hydrogels after the scaffold fabrication. Pluronic F-127 is a thermosensitive ink often used as the sacrificial material^[232]. This thixotropic material has an excellent shear-thinning property favoring the extrusion bioprinting process but resumes high yield stress quickly when shear stress is removed so that the printing fidelity is preserved. Cell-laden hydrogels can be bioprinted or cast surrounding the sacrificial channels. Pluronic F-127 is subjected to phase-transition and becomes hydrophilic when the temperature is lowered and therefore can be flushed away at 4 ℃ or lower, resulting in multi-cellular and heterogeneous vascularized tissue constructs^[232]. HUVECs can be seeded to the channels to provide long-term perfusion stability. Using this approach, multi-material bioprinting was employed to fabricate vascularized scaffolds with a large thickness (>1 cm)^[233]. These constructs have been demonstrated to support long-term cell survival for over 45 days^[233]. Programmable cellular heterogeneity and in situ development of MSCs within the bioprinted vascularized scaffolds have also been shown to recapitulate physiologically relevant features.

Other thermosensitive biopolymers have also been explored in conjunction with multi-material bioprinting technologies to create vascularized constructs. Studies on the co-printing of matrix bioinks and gelatin-based sacrificial inks have provided a vascularization solution^[234]. Both matrix and sacrificial inks can be printed adjacent with no spacing needed to establish void-free multi-material structures. The matrix phases can be crosslinked during fabrication while the sacrificial ink can be removed through mild triggers, such as temperature change. The embedded endothelial cells can then be released from the sacrificial bioink to functionalize the channels with a confluent endothelial layer. Such an approach avoids the structural collapse and instability during the fabrication of mechanically weak materials (Figure 15A-B). A similar idea has also been explored on the co-extrusion of collagen and gelatin bioinks, in which HUVEC-laden gelatin was used to print sacrificial vascular channels^[235,236]. The functional vascular channels with perfused open lumens were capable to support cell viability in scaffolds up to 5-mm-thick. Such a strategy enabled by multi-material bioprinting shows great potential to investigate fundamental mechanisms of vascular functions and maturation under physiological flow conditions.

Figure 15-.

Vascularization strategies with multi-material bioprinting. (A) Schematic of the void-free 3D bioprinting process, where a biocompatible templating bioink (green) and a matrix bioink (yellow) are printed side-by-side, followed by photo-crosslinking of the matrix phase and 37 °C incubation to release the templating phase. Preloading endothelial cells in the templating bioink allows in situ endothelialization of the channels. (B) The widespread formation of endothelialized channels after 7 days of culture. Reproduced under the terms of the CC-BY Creative Commons Attribution 4.0 International license (https://creativecommons.org/licenses/by/4.0).^[234] Copyright 2020, The Authors, published by Wiley-VCH. (C) Schematic showing the multi-material bioprinting of vascularized constructs with sacrificial materials. (D) A bioprinted vascularized cardiac patch. (E) Transplantation of the printed patch in between two layers of rat omentum. Dashed, white line highlights the borders of the patch. (F-G) Sarcomeric actinin (red) and nuclei (blue) staining of sections from the explanted patch. Reproduced under the terms of the CC-BY Creative Commons Attribution 4.0 International license (https://creativecommons.org/licenses/by/4.0).^[229] Copyright 2019, The Authors, published by Wiley-VCH. (H) A concentration-gradient model comprising PEGDA and GelMA hydrogels generated by multi-vat-photopolymerization bioprinting. (I) Photographs showing retrieved implants with evidence of vascularization within VEGF-laden constructs. Reproduced with permission.^[189] Copyright 2018, Wiley-VCH.

Embedded multi-material printing of sacrificial channels has been utilized to vascularize thick scaffolds with biomimetic vasculature structures^[154]. A study printed sacrificial gelatin channels inside a supporting matrix composed of patient-specific iPSC-derived spheroids^[14]. The rapid building of perfusable vascularized constructs with high cell density (~1×10⁸ cells/mL) enabled patient- and organ-specific tissues at the therapeutic scales. A similar approach was also adopted to bioprint vascularized cardiac patches^[229]. Endothelial cells maintained high viability after bioprinting and formed physically robust vessels after gelatin channel removal. The functionality of the vascularized patches was evaluated by transplanting to rat omentum. The cells inside the patches exhibited elongated and aligned morphology with massive striation, which indicated functional contractility potential (Figure 15C-G).

Multi-material bioprinting through co-axial nozzles is another prominent advancement towards the reconstructions of vasculature. This type of printing nozzle accommodates the interior flow of the crosslinker solution surrounded by the exterior flow of (bio)inks, leading to the formation of hollow microchannels that are perfusable. Sodium alginate and CaCl₂ ionic crosslinker are one of the most common combinations. Engineered vessels made of other materials, such as GelMA, hyaluronic acid (HA), cellulose, and dECM have been also explored using this method in conjunction with their associated crosslinking mechanisms^[237–239]. Circumferentially multi-layered tubular tissues can be created with multi-channel multi-material co-axial bioprinters. For example, cannular urothelial tissue constructs containing human urothelial cells and human bladder smooth muscle cells, or cannular blood vessels containing HUVECs and human vascular smooth muscle cells, have been bioprinted with such a system^[146]. The bioprinted cannular tissues were perfusable with cell culture medium to promote growth and proliferation of the embedded cells.

Multi-vat-photopolymerization, as a fast and high-resolution multi-material bioprinting modality, has been explored to (bio)print the vasculature. As a result of its superlative resolution, this method is favorable for bioprinting minuscule vasculatures. Through sequential changing of (bio)inks with and without endothelial cells, prevascularized tissues with complex 3D microarchitectures can be created. A study using anastomosis between grafted prevascularized bioprinted constructs and the host vasculature showed the formation of functional vasculature within the engineered tissues^[240]. Heterogeneous PEGDA and GelMA constructs printed by a microfluidics-enabled multi-material vat-photopolymerization bioprinter have also been tested for their neovascularization potential in a rat model. The presence of vascular endothelial growth factor (VEGF) in the bioprinted constructs led to more blood vessel-formation in the implants compared to those without VEGF (Figure 15H, I)^[189]. In terms of technology itself, the relative ease to scale up in multi-vat-photopolymerization can potentially lead to creation of large-scale vascularized constructs for tissue repair or even organ transplantations. Gene expression could also be patterned through tuning the channel network architecture, medium temperature, and flow directions, which expands the capacity to regulate cellular development and regeneration through vascularization^[241]. However, it should be noted that the generation of bioink wastes associated with vat-photopolymerization may increase the costs and therefore limit the scalability.

5.3. Organ-on-a-chip (OoC) models

OoC is a type of microfluidics-based cell culture platform recreating the key elements of biological tissues^[242–244]. In the past decade, various OoCs have been developed to mimic different human organs for fundamental biological studies, tissue/disease modeling, and drug screening^[245–249]. Traditional OoC fabrications involve a steep learning curve and time-consuming manual operations, which hinder their adoption to biomedical research labs. However, the flexible and automated bioprinting technology can bring more user-friendly fabrication approaches to researchers with rapid prototyping capability. To this end, multi-material bioprinting facilitates the construction of functional OoC models with hierarchical architectures and spatial heterogeneity.

OoC models often require the arrangement of multiple cell types and ECM-mimicking materials in a spatially defined fashion to resemble the physiological environment of individual organs. One advantage of OoC model fabrications using multi-material bioprinting is the design freedom offered by this technology. Cell types, amounts, and spatial arrangements can be controlled with relative ease. The 3D environments provided by the bioprinted constructs, usually composed of polymeric hydrogel networks, have been shown to yield more realistic cellular behavior such as spreading and migration compared to locally 2D cultures in most conventional OoCs. For example, a multi-cellular liver OoC model containing hepatocytes and endothelial cells has been developed (Figure 16A-C)^[250]. The co-culture of these two cell types showed significantly enhanced liver functions such as higher values of urea and albumin secretions. The OoC model can also be bioprinted in conjunction with conventional 3D printing to yield a complete OoC device to further improve the fabrication and research efficiencies^[250].

Figure 16-.

OOC and tumor modeling with multi-material bioprinting. (A) Schematic illustration of the 3D bioprinting technology for the OOC applications. (B) Digital image showing a bioprinted liver chip. (C) Various configurations of cells and biomaterials within the printed chips. Reproduced under the terms of the CC-BY Creative Commons Attribution 3.0 International license (https://creativecommons.org/licenses/by/3.0).^[250] Copyright 2016, Royal Society of Chemistry. (D) Schematics showing precisely controlled multiple-cell patterning in microfluidic chips by inkjet printing and the detection of drug metabolism and diffusion. Reproduced with permission.^[252]. Copyright 2016, Royal Society of Chemistry. (E) Multi-vat-photopolymerization bioprinting of hydrogel-based hepatic construct. (F) Gene expression profiles and albumin and urea secretion levels of HPCs from 2D, 3D single-cellular, and 3D multi-cellular cell cultures. Reproduced with permission.^[51]. Copyright 2016, The Authors, published by National Academy of Sciences. (G) A 3D-bioprinted in vitro tumor model mimicking metastatic dissemination. (H) Plots and micrographs of the population of disseminated A549 lung cancer cells detected in the collection chamber versus time. Reproduced with permission.^[268] Copyright 2019, Wiley-VCH.

Multi-material bioprinting can assist microfluidic devices with patterning cellular constructs for high(er)-throughput detection and screening^[251]. A study utilized a multi-material inkjet bioprinter to dispense cell-laden alginate hydrogels to form precisely distributed cell arrays in a microfluidic chip (Figure 16D)^[252]. Hepatocellular carcinoma (HepG2) and glioblastoma tumor (U251) cells were co-patterned for drug metabolism and diffusion tests under a biomimetic environment. The efficacy enabled by multi-cellular bioprinting of the cell arrays can significantly reduce the extent of laborious experimental work from fabricating multiple conventional microfluidic devices. In addition to cell-laden hydrogels, the array can also be composed of cell spheroids. The response of multi-cellular OoC model could be comparable to the corresponding animal model, which confirmed the utility of this technology for biological testing and potentially facilitate drug development^[253]. With the aid of multi-material bioprinting technologies, the ability to recapitulate human physiology and disease states can potentially exceed the animal models and offer more accurate therapeutic prediction^[254]. Multi-material bioprinters with hybrid modules have the potential to further streamline the fabrication steps and decrease costs (both material- and financial-wise). An exemplar is a structurally heterogeneous skin model/device bioprinted with a combination of extrusion and inkjet modules^[255]. Reductions of 50-fold in cost and 10-fold in culture medium-consumption were achieved compared to traditional cell culture platforms. Such migrations from traditional cell culture platforms to bioprinted microdevices has great socioeconomical implications and could potentially decrease the amount of time and cost for technological translations in a greener way.

Sophisticated miniature biological constructs have also been developed with the assistance of multi-material vat-photopolymerization bioprinting. For example, human-derived hiPSC-HPCs, adipose-derived stem cells, and HUVECs have been demonstrated to pattern into microscale hexagonal architecture, representing the building unit of the hepatic tissue (Figure 16E). Multiple fold of increases in morphological organization, higher liver-specific gene-expression levels, increased metabolic product-secretion, and enhanced cytochrome P450-induction were observed from constructs created by multi-material bioprinting (Figure 16F). The vascularized hepatic model showed both phenotypic and functional enhancements, which is essential for building personalized platforms for in vitro drug screening and disease studies^[256]. A similar liver spheroid model with perfusable intrinsic channels has been established with multi-material vat-photopolymerization bioprinting^[184]. The strengthened metabolism and gene expressions compared to 2D culture models support long-term drug screening for pharmaceutical developments.

5.4. Tumor modeling

Tumors are dynamic 3D ECM networks consisting of stromal, immune, and vascular cells^[257]. It is challenging for 2D tumor models to recapitulate these multi-cellular, highly heterogeneous microenvironments, which yield non-physiological cellular behavior, such as altered gene expressions, cell-cell interactions, and drug responses, among others^[258]. Multi-material bioprinting can mimic the complicated 3D microenvironments of tumors and therefore has the potential to create better biomimetic tumor models compared to traditional methods^[145,259]. For instance, an in vitro cervical tumor model comprising Hela cells and gelatin/alginate/fibrinogen promoted cell proliferation, matrix metallopeptidase (MMP)-expressions, and chemoresistance when compared with 2D planar models^[260]. Therefore, multi-material bioprinting is becoming a standard approach to recapitulating the multi-cellular, hierarchically architectured, and dynamic microenvironments of tumors.

Bioprinted heterogeneous tumor models enable the study of the dynamic tumor progression process. For instance, a few recent studies on triple-negative breast cancer cells are focused on the development of multi-cellular tumor spheroids composed of MDA-MB-231 triple-negative breast cancer cells and IMR-90 cancer-associated fibroblasts with the aid of multi-material bioprinting^[261,262]. The flexibility of multi-material bioprinting allowed convenient construction of different tumor microenvironments that involved different bioink compositions, matrix elasticities, and MDA-MB-231 initial cell densities. The developed 3D models were robust and suitable for long-term cultures and probed various parameters of initial cancer model environment in regulating breast tumor progression and metastasis. 3D-bioprinted mini-brains containing two types of bioinks have also been developed to study the crosstalk between glioblastoma cells and glioma-associated macrophages (GAMs)^[263]. The two types of cells were found to actively interact with each other, resulting in changes in macrophage phenotype and cancer progression and invasiveness. Such clinically relevant models are important to improve the understanding of tumor biology for evaluating novel cancer therapeutics. Other types of multi-material bioprinting, such as co-axial technology and DLP-based bioprinting, have also been used for tumor modeling. A study fabricated self-assembled multi-cellular heterogeneous brain tumor fibers with co-axially printed alginate/gelatin hydrogels as the external shell, and cell suspension containing fibrinogen as the core^[264]. This biomimetic design improved viability, proliferation, and tumor-stromal interactions, which was benefited from the recapitulation of the dynamic tumor microenvironment using multi-material bioprinters. Other configurations, such as GelMA/alginate (core/sheath) have also been investigated for tumor modeling^[116]. In another study, glioblastoma models consisting of distinct tumor, acellular ECM, and/or endothelial regions were built using a DLP-based bioprinter^[265]. The stiffness of different regions corresponding to glioblastoma stroma, healthy or pathological brain parenchyma, and brain capillaries was tuned to recapitulate the biochemical and biophysical heterogeneity of the glioblastoma microenvironment. While enhancement of hypoxia, stemness, endothelial protruding morphology, and angiogenic potentials related to malignant phenotypes was observed in the stiff models, rapid proliferation and expansion of cells with the classical phenotype occurred in the soft models. In a similar way, glioblastoma microenvironments containing patient-derived glioblastoma stem cells (GSCs), macrophages, astrocytes, and neural stem cells (NSCs) in an HA-based hydrogel were 3D-bioprinted to investigate the growth of GSCs with or without macrophages^[266]. The presence of macrophages yielded transcriptional profiles that were closer to those of patient-derived glioblastoma tissues. The collective findings from these studies suggest that building biomimetic and controllable 3D tumor models can be realized by leveraging the precise spatial programming of matrices and cells with multi-material bioprinting.

Creating biomimetic in vitro 3D tumor models is critical for anticancer drug screening and development^[267]. The spatiotemporal control of signaling molecular gradients enabled by multi-material bioprinting has been employed to modulate cellular behaviors of tumors at a local level (Figure 16G, H)^[268]. Vascularized tumor models were created to mimic cancer invasion, intravasation, and angiogenesis. The flexibility to add functional materials and precise placement of different combinations of tumor-relevant cells and hydrogels enabled programmable tumor microenvironments, which potentially could provide insights to identify the fundamental problems in preclinical settings^[268]. Another study proposed a bioprinted tumor-on-a-chip device with co-axially bioprinted blood/lymphatic vessel pair to recapitulate the different levels of drug transport profile^[269]. This unique design demonstrated the capacity of simulating the complex transport mechanisms of certain pharmaceutical compounds inside tumors and exhibited great potential to improve cancer drug screening accuracy.

5.5. Organoids

Organoids are miniature and simplified model organs derived from self-organization and differentiation of pluripotent stem cells or progenitor cells isolated from tissues. They recapitulate similar gene/protein expressions, metabolic functions, and microscale architectures of the native organs^[270]. Owing to the remarkable resemblance, they show significant advantages over some traditional 2D and 3D culture platforms to serve as model tissues and organs for cell therapy, drug screening, and disease modeling^[271,272]. Despite much progress in the field, organoids formed with manual techniques face problems such as low throughput and high variability, which hinder the translatability of organoids^[270,273]. Besides, the size of organoids is typically restricted on a millimeter-scale due to the lack of nutrient supply, which results in a necrotic inner core when the organoid grows large^[270]. This size-limitation precludes the culture of physiologically relevant scales to be used as implantable medicines.

Automated 3D bioprinting techniques have demonstrated outstanding promise to fabricate conformable, scalable, and reproducible organoids in a higher-throughput manner^[274]. In a typical procedure for bioprinting organoids, concentrated cell suspension solutions or pastes are dispensed to a location, either on a well plate or within a biochemical- and biophysical-mimicking niche (such as collagen and Matrigel). The cell numbers and the dispensed locations can be precisely controlled, which can drastically improve the reproducibility of bioprinted organoids. A pioneer work has shown that kidney organoids can be bioprinted within 6- and 96-well plates with high throughput and low individual variation^[273]. Such a reproducibility favors the accuracy of drug screening, such as the toxicity screening of doxorubicin and aminoglycoside, compared to the unrepeatable results for organoids from manual culture. Conformation of the organoids can also be controlled by spatially define the location of dispensed cells, which enables the easy programming of the organoid maturation. Furthermore, dispensing different cell densities and geometries can alter the spontaneous cellular self-organization into mesoscopic organoids. Spatiotemporal modulation of morphogenesis of mesenchymal aggregates, intestinal organoids, and vascular organoids was proven successful^[275]. Organoids with centimeter-scales were achieved by bioprinting the stem cells in their permissive environments, which hold promise in fabricating physiologically relevant organs for implantable regenerative medicines. Multi-material bioprinting, in principle, can in addition, expand the functionality and integration of various bioprintable organoids within a single system.

Despite several transformative studies mentioned above, the use of bioprinting for organoid-fabrication is still in its infancy. Currently, organoid-bioprinting heavily relies on mechanically actuated extrusion techniques, such as syringe pumps, due to their excellent volume control capability. The employed bioinks are also often comprised of low-viscosity cell-only suspensions or pastes. Pneumatically actuated extrusion bioprinters are intrinsically not suitable to promise repeatability due to their incapability in flow rate control, which is associated with the inconsistency in cell number. Another intrinsic limitation that hinders the adoption of bioprinting organoids is the high costs of stem cell expansion. Bioprinting organoids often requires the cell density to reach 1×10⁷ cells/mL. To compromise the limited number of cells in practice, specialized hardware, such as gas-tight low-volume syringes (~10–10² μL) and small nozzles (10–10² μm), are often needed to accommodate the limited volumes of the bioinks. Treatments for the prevention of drying and cell adhesion to the printing nozzle are also important to ensure successful bioprinting. To accommodate the process for organoid-fabrication, development in bioprinting hardware and accessories that are easy to control flow rate and with low ink-consumption is a critical mission. Meanwhile, other bioprinting techniques such as robotic microtissue assembly (the Kenzan method) show potentials in precisely placing spheroids to form complicated biological constructs, which could also be explored for the assembly of organoids^[276–278].

6. Conclusions and perspectives

Multi-material bioprinters outperform their single-material counterparts in fabricating heterogenous/multi-cellular tissue constructs. Developments in the technology of multi-material bioprinters were discussed, and different modalities were compared. Each of the four primary bioprinting technologies has its own advantages and disadvantages. In Figure 17, we further compared these technologies in three aspects: bioprinter basics, cost-effectiveness, and functionality. Due to their simplicity and prevalence, nozzle-based techniques provide the basics of a bioprinter in a more cost-effective manner. However, in terms of functionality, the laser/light-based technologies, i.e., LIFT and vat-photopolymerization, are likely more advantageous.

Figure 17-.

Comparison of the four primary bioprinting technologies in different aspects. Features are ranked in four levels with the outer level presenting the highest score. The comparisons are general and may not hold true for all specific cases.

Regarding the bioprinter basics, extrusion and inkjet bioprinters are very flexible to be upgraded for bioprinting of more than a single material. There are, however, complications for accommodating this feature in laser/light-based technologies, although not impossible. Another advantage of multi-material nozzle-based techniques is the relatively high material-switching speed, which is missing in laser/light-based methods. LIFT and vat-photopolymerization are excellent in terms of printing speed; however, their building volumes and user-friendliness cannot compete with the extrusion method, at least at this stage of development. In general, nozzle-based modalities are more commercially available, which results in a lower price for their hardware and consumable parts. Furthermore, the minimal bioink usage, due to the intrinsic nature of nozzle-based techniques, and the abundance of compatible bioinks make this technique a cost-efficient and accessible method. In the functionality spider chart of Figure 17, it is evident that there is still room for enhancing the resolution and cell viability of the extrusion technique. Laser/light-based methods deliver excellent resolutions and cell viabilities. Specifically, the most suitable approach for printing multi-scale vascular structures is perhaps vat-photopolymerization. A critical drawback of the current multi-material vat-photopolymerization techniques is the risk of cross-contamination as well as swelling or dehydration during the printing process, which could be a potential direction for future endeavors. The washing buffer during the material-exchange process could lead to deformation of the construct that would affect the 3D-bioprinting resolution especially for organ-sized constructs. In addition, vascular networks with a physiological dimension yet have not been 3D-bioprinted and studied using cell-laden hydrogels. Fabrication of such constructs is more challenging due to absorption of light by cells that could affect the penetration depth depending on the cell density and type as well as wavelength of the light source.

As illustrated in Figure 18, several improvements are envisioned to overcome the current limitations. One challenge faced by most existing multi-material bioprinting technologies is to scale up the printing volume to clinically relevant sizes within a reasonable fabrication time. For instance, with a typical extrusion-based bioprinter, it can take hours to even days to create an entire multi-cellular organ structure with a decent resolution. Such a long fabrication process keeps the cells away from physiological conditions and significantly decreases the cell viability in the fabricated cell-laden constructs. Although this issue can be potentially remedied with a 3D-bioprinting setup within a temperature-, humidity-, and CO₂-controlled environment, such attempts would inevitably increase the equipment and fabrication costs and hinder the accessibility of the technology. Maintaining the physiological temperature during a fabrication process might also be incompatible with some thermosensitive materials, such as gelatin and collagen. Speeding up the biofabrication process, therefore, still seems to be the most feasible and economical pathway at the current stage. Embedded bioprinting can possibly alleviate this concern for mid-sized organs to some extent since cells stay in a humidified environment. Yet, bioprinted full-sized organs with high cell viability have not been reported so far with embedded bioprinting. Other designs such as multi-nozzle printheads have also been explored^[140], but they are mainly suitable for creating repeating patterns.

Figure 18-.

Future perspective of potential improvements in multi-material bioprinting.

Large-scale bioprinting has higher chances to introduce defects, such as bubbles and voids. The involvement of multiple materials further impedes the quality control for layer-by-layer methods, leading to a low success rate for production. Such complications are mainly due to the relatively unstable properties of soft bioinks. Although the defects are difficult to avoid or predict, they are usually easy to detect during the printing process. Machine-learning approaches have been recently employed to enhance the printing quality^[279]. By introducing artificial intelligence (AI), bioprinting on dynamically moving substrates, such as breathing lungs, has been successfully demonstrated^[280–282]. With the assistance of AI in conjunction with computer vision, substrate height, printing pressure, or writing speed can be dynamically adjusted to correct the defects and errors, yielding better printing quality^[283].

Although several multi-material vat-photopolymerization methods have been developed, one of the remaining challenges is fabricating large constructs while not compromising printing resolution and speed. The advent of DLP light engines with a higher number of pixels and 3D printers with a dynamic projector could increase the throughput without reducing the print speed or resolution. The throughput of the TPP method can be enhanced by using multiple beams as well as developing highly efficient water-soluble TPP photoinitiators. In addition, further improvement is needed in the rapid bioink-exchange process with minimal bioink cross-contamination and high cleaning efficiency. The LIFT bioprinting method also requires further development to make it more accessible and affordable to fabricate multi-cellular constructs.

Hardware aside, the capacity of multi-material bioprinting can be further broadened through the use of advanced material concepts, such as self-assembling materials^[284]. For instance, a self-assembled collagen-fibrin hydrogel has been shown to improve mechanical strength and stretchability while providing a better ECM-mimetic structure for cell alignment^[285] or growth^[286]. Self-assembling chitosan bioinks have also been shown to generate cell-sized pores within the scaffolds. Their stiffness and viscoelasticity can be modulated over a wide range by controlling the strength of supramolecular interactions^[287]. The delivery of bioinks within a Bingham fluid material could enable the freeform fabrication of constructs with geometrical details that are not easily achievable using DIW^[155,288]. The combination of a top-down bioprinting approach and a bottom-up material self-assembly is desirable to maximize mechanical, structural, and biological functions.

Application-wise, multi-material bioprinting may be employed to further improve the current single-material designs through the use of sacrificial inks. For instance, in a co-axial configuration, where the researchers used HUVEC-laden GelMA as the core and cell-free alginate as the shell, a lumen structure formed around the core fiber^[151], as explained in Section 3.1.3. However, the microfibers were not hollow, and they cannot essentially resemble a vessel-like structure. Tubular vessel-like constructs may be fabricated using a sacrificial material as the core ink^[237]. Generally, the fabrication of vascularized constructs can be fostered by effectively exploiting multi-material bioprinting of bioinks and sacrificial inks. Another unique perspective in bioprinting is integrating multi-material technologies in 4D bioprinting, a method to use stimuli-responsive materials as the bioink^[289–291]. Multi-material bioprinters can accommodate the use of multiple smart bioinks to fabricate multi-functional stimuli-responsive constructs.

The accessibility of commercial multi-material bioprinters has motivated many researchers from different fields to leverage this technology for desirable tissue engineering applications without spending time and resources on the production of bioprinters. Previously, multi-material technologies were mostly limited to biomedical engineering labs with essential engineering backgrounds for its development. Most commercial bioprinters are based on the extrusion technique due to its accessibility, affordability, and easy-to-use platform. There is still an unmet need for commercial multi-material LIFT and vat-photopolymerization bioprinters to make these technologies accessible and available to more research labs, where they can be used for cutting-edge applications. The cost barrier and the complexity of the LIFT bioprinting method are the main challenges in commercializing this method.

Looking into the future, we envision hybrid bioprinters comprising multi-material modules to be developed. With the aid of various parallel modules, different tissue parts can be fabricated and assembled to form a complete and functional whole-size organ. An exploratory 3D printer combining an aerosol jet head, a photonic cure head, extrusion modules, and inkjet heads has been recently reported to create complex structures with high precision^[292]. Hybrid bioprinters comprising molten material-extrusion and DLP/SLA modules has also been used to print soft hydrogels and thermoplastics simultaneously^[293]. Such a combination is favorable to fabricate cell-laden constructs with gradients and high mechanical strength. Acoustic-based technologies have recently been utilized for the spatial assembly and patterning of cells or organoids into complex microtissues^[294–296]. Ultrasound acoustic waves could potentially be used as a contact-free and cell-friendly approach to align multi-cellular 3D constructs within layers to recapitulate the intrinsic microarchitectural organization of many native tissues. For instance, hybrid bioprinters capable of preferentially organize cellular arrays within bioprinted constructs have pioneered the integration of acoustophoresis with nozzle-based bioprinting methods^[297,298]. Other works such as manual^[299] or robotic^[278] mini-tissue assemblies also prove the feasibility of modular tissue fabrication.

Significant advances in multi-material bioprinting technologies have been achieved over the last two decades. However, synergetic efforts on developing bioprinting techniques, bioinks, and gantries are still needed to realize the full potential of these technologies. This review focused on the technologies developed for multi-material bioprinting and provided a comprehensive overview of their design, standard techniques, commercialization progress, and biomedical applications. Multi-disciplinary research and collaborations between academic research and industries will be crucial to bringing this technology into clinical use.

Author Photographs and Biographies

Hossein Ravanbakhsh is a Ph.D. candidate in the Department of Mechanical Engineering at McGill University (Canada), and a visiting researcher in the Division of Engineering in Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School (USA). He received his M.Sc. in Mechanical Engineering from Amirkabir University of Technology (Iran) and B.Sc. also in Mechanical Engineering from Isfahan University of Technology (Iran). His research interest focuses on 3D bioprinting technologies, composite hydrogels, and biomaterials for tissue engineering.

Vahid Karamzadeh is a Ph.D. candidate in the Department of Biomedical Engineering at McGill University (Canada). Previously, he worked at the Organ Manufacturing Group at United Therapeutics Corporation. He obtained his M.Sc. degree in Mechanical Engineering from Concordia University (Canada) and B.Sc. degree in Mechanical Engineering from Sharif University of Technology (Iran). His research interests mainly focus on 3D (bio)printing technologies, biofabrication, microfluidics, and biomaterials.

Guangyu Bao is currently a Ph.D. candidate in the Department of Mechanical Engineering at McGill University (Canada). He received his M.S. degree in mechanical engineering from Washington University in St. Louis (USA) and B.Eng. degree in mechanical engineering and automation from Beijing University of Posts and Telecommunications (China). His current research is focused on developing novel biomaterials and biofabrication approaches for tissue engineering and other biomedical applications.

Luc Mongeau is Distinguished James McGill Professor of Mechanical Engineering at McGill University. He holds joint appointments in Biomedical Engineering, Bioengineering, and Otolaryngology Head and Neck Surgery. His research program includes investigations of bioprinting, airway organoids, vocal fold tissue engineering for voice restoration, head and neck cancer therapies, and needleless drug injection. He is a professional engineer member of several professional organizations, including ASME, TERMIS, ESB, and CSB. A list of publications and accomplishments is available at ORCID: 0000–0003-0344–227X.

David Juncker is Professor in the Biomedical Engineering Department at McGill University, and has been serving as Chair of the department since 2018. David’s research interests include micro-bioengineering technologies for the development of 3D-printed and miniaturized organ-on-a-chip devices, ultrasensitive assays for the analysis of proteins and cellular secretions such as extracellular vesicles in liquid biopsies and single cells, as well as novel microfluidic circuits for point-of-care diagnostics.

Yu Shrike Zhang received a B.Eng. in Biomedical Engineering from Southeast University (2008), a M.S. in Biomedical Engineering from Washington University in St. Louis (2011), and a Ph.D. in Biomedical Engineering at Georgia Institute of Technology/Emory University (2013). Dr. Zhang is currently an Assistant Professor at Harvard Medical School and Associate Bioengineer at Brigham and Women’s Hospital. Dr. Zhang’s research is focused on innovating medical engineering technologies, including bioprinting, organs-on-chips, microfluidics, and bioanalysis, to recreate functional tissues and their biomimetic models towards applications in precision medicine.