Bioprinters for Organs-On-Chips

3.1. Inkjet bioprinting of vascularized systems

The first bioprinting technology introduced into lab-on-chip platforms is inkjet bioprinting [53], which directly deposits droplets of the ink onto a substrate by using a heating or piezoelectric head attached to the nozzle (Figure 1). Under pneumatic pressure or a mechanical load, the biomaterials and/or cells are dispensed in droplets at desired positions [32]. In one work, alginate-based bioink formulation was placed along a bifurcating microvasculature capable of supporting physiological flow rates [54]. They showed that cells can be patterned in the channels using inkjet bioprinting. Regarding the practical limitation of inkjet bioprinting, such as the requirement of low viscosity bioink and the inherent inability to perform continuous flow, constructing 3D architecture is very challenging while applying hydrogel bioinks. Therefore, this method has been less used for 3D bioprinting.

3.2. Extrusion-assisted bioprinting

Extrusion-assisted bioprinting is the most multipurpose process and enables the broadest range of bioink viscosities to be used for bioprinting for clinical applications [55]. Sometimes extrusion-assisted bioprinting is called pressure-assisted bioprinting (PAB) [56]. In such a case, the motion of air or plunger-based pressure is released in the form of a continuous filament through a micro-nozzle or a micro-needle on a motionless substrate. Li et al. believe that it benefits from homogenous distribution, room temperature processing and direct combination of cells [57] PAB has been applied to the printing of cell and organs with confirmed retention of activity in many applications. Bioprinted cells include but not limited to human mesenchymal stem cells (hMSCs), mouse pre-osteoblasts, endothelial cells, and osteogenic progenitors [58]. Kolesky et al. made-up a single tissue by co-printing three different cells and multiple inks including Pluronic F-127 filaments and hMSCs-laden gelatin/fibrin filaments and neonatal dermal fibroblasts inside a modified ECM together with embedded vasculature (Figure 2A), which is then coated by human umbilical vein endothelial cells (HUVECs). Towards having a dense osteogenic tissue, culture media containing growth factors for osteogenic differentiation was perfused [33]. A mixed tissue model was fabricated and transformed into a vascular network filled by HUVECs. Brigham et al., have fabricated semi-interpenetrating networks of hyaluronic acid and collagen to achieve adequate mechanical properties and to present a new class of hydrogels in 3D printing. They designed many synthetic materials, such as the biodegradable polycaprolactone, poly (lactic acid) (PLA), poly (glycolic acid) (PGA), and PLGA, PLA-PGA co-polymer as an alternative to natural polymers [59].

Recently researchers have started using decellularized ECM as a promising bioink to reproduce the natural microenvironment of the cells to mimic the native tissue. [60]. Homan et al. reported 3D human renal proximal tubules in vitro (Figure 2B) by placing the renal proximal tubules inside an ECM. The perfusable tissue microchips was then preserved for greater than eight weeks. Their proposed technique offers a new way for programmable fabrication of human kidney tissue models [34]. Toprakhisar et al. have stablished a novel approach in bioprinting of dECM hydrogel bioink for use as micro-capillary-based bioprinting (Figure 2C). The main advantage of this method is bioprinting without any need to use support structure or any cross-linker components.

Freeform reversible embedding of suspended hydrogel (FRESH) has been applied for creating vascular constructs, where bioink is extruded into a buoyant support from granular dense medium [35]. The presence of a sacrificial support bath allows direct printing of bioinks such as alginate and Matrigel^® with high fidelity [35]. FRESH has been utilized to fabricate vascular networks. In a recent effort, xanthan gum was deposited into a CaCh-infused gelatin to form sacrificial channels as rapid prototyping of microfluidic platforms [61]. Proper characterization of bioinks and the role of processing parameters such as pressure on the biological properties of printed constructs have been neglected [62].

In summary, extrusion-assisted has shown great potentials in fabricating mechanically-stable microfluidic constructs through both sacrificial and direct 3D printing. The wide range of materials and low costs associated with extrusion-assisted bioprinting have made it the first choice for many efforts on bioprinted microfluidic platforms. However, it suffers from a resolution limitation of around 100 μm for current bioink systems. Optical-based techniques have provided a better resolution, as explained below.

4.1. Laser-induced forward transfer (LIFT)

The basic configuration for LIFT includes a laser system which produces pulsed focused beams, an absorbent layer where a sheet of ink is coated, and a collecting surface on which the ink is deposited (see Figure 3A) [74]. The laser energy yields absorption of ink onto a transparent substrate and the light-matter interaction at the interface results in a strong pressure at the irradiation region. A small pixel of the biological material is ejected onto the collecting substrate. The properties of the incident laser spot controls the final shape and size of transferred ink. This method has a bioprinting resolution of around 50 μm [75] but it is costly due to the complexity of the laser source and the control of the laser.

The LIFT technique was employed to repair Pt-based heater elements with promising results on the ability of LIFT to write electrodes for the control of electro-osmotic flow in microfluidic platforms [76]. The LIFT technique has been used to pattern cellular compositions with almost any desired cell density and suspension viscosity [77]. In one study, LIFT was used to print thin layers of fibroblasts- and keratinocytes-laden collagen for skin tissue formation [36]. Wu and Ringeisen have fabricated tissue models loaded by human umbilical vein smooth muscle cells (HUVSMCs) and HUVECs using small droplets with diameters of approximately 50 μm [37]. Such work showed that the LIFT technique can be used for tissue biofabrication and it can be scalable to manufacturing macroscale constructs.

4.2. Stereolithography (SLA)

Conventional SLA is popular in manufacturing industries, while most SLA printers are configured so that the light source polymerizes a bath filled with a resin onto a moving 3D stage (Figure 3). Once one layer is polymerized, the stage descends a predefined distance for another pre-defined layer to be polymerized (Figure 3B). The stage is then elevated by a distance equivalent to the thickness of the layer, thus this method suffers from a lower resolution [80]. Despite these techniques, digital light processing (DLP) generates dynamic masks using computercontrolled array of micromirrors to polymerize a unit volume of the resin thus significantly expediting the curing process (Figure 1) [81]. During the exposure time, the DLP chip can be selected independently to deflect the light or to project it, which results in a light pattern. Each layer is solidified based on the pattern generated by the DLP chip and the light is modulated and transferred through a reduction lens onto the liquid resin. Compared to other 3D bioprinting technologies, light-assisted 3D bioprinters have the advantage of higher printing speed and better bioprinting resolution [31]. Due to the high degree of directionality in laser light and less light scattering perpendicular to laser light, the bioprinting resolution in SLA- and DLP-based bioprinting is predominantly determined by the thickness of the photocured resin. The thickness can be tuned by modifying laser characteristics such as pulse duration, wavelength and energy, repetition rate, and beam focus diameter, as well as the resin properties such as viscosity and surface tension [82]. Although the polymerization kinetics of the curing process can be very complex, the following relationship is often used to describe the thickness of the photo-cured resin [83]:

SLA bioprinting has been initially used for creating tissue constructs. In a study, the authors used a 100-μm-resolution SLA printer to solidify a biodegradable bioink made of diethyl fumarate, poly(propylene fumarate), and bisacylphosphine oxide towards healing critical-sized bone defects [84]. SLA has been recently employed to the rapid fabrication of microfluidic systems using transparent biocompatible polymers, with potential outlooks in organ-on-chip platforms [38]. Since SLA eliminates the need of expensive molds and it has the advantage of rapid prototyping, SLA-based microfluidic systems may lead to an efficient commercialization of organ-on-chip platforms.

The printing speed of light-based printing technology has been greatly enhanced by maskless DLP-based bioprinting. A manual exchange of two bioinks was used, including GelMA as the parenchymal tissue and GelMA-glycidal hyaluronic acid as the vasculature. The bioinks were mixed with human-induced pluripotent cell-derived hepatic progenitor cells and HUVECs. A recent study by Zhu et al. demonstrated DLP bioprinting of prevascularized tissue models with complex geometries of varying widths and heights of 50 μm using endothelial cells that were encapsulated in a mixture of glycidal methacrylate-hyaluronic acid and gelatin methacryloyl (GelMA) [39]. This study not only highlighted the versatility and the accuracy of the maskless method, but since all tissue models were bioprinted under 1 minute, it also emphasized the high speed of bioprinting associated with DLP. A recent effort on making multi-material DLP-based bioprinter has led to high-resolution fabrication of micro-tissue models, such as tumor model shown in Figure 3C [31]. They used a novel microfluidic chamber to switch different bioinks.

4.3. Two-photon lithography (TPL)

TPL does not rely on the use of complex optical systems or photomasks to print in a photosensitive bioink and it involves the use of a near-infrared ultrafast femtosecond laser. A multi-photon absorption process occurs in a medium that allows penetration of the emitted light. The freedom that laser systems offer in terms of scanning and tightly modulating the energies at the point of interest allows for submicron features. Applegate et al. prepared 3D multiscale patterns in protein hydrogels made of soft silk with low-energy femtosecond lasers without the use of exogenous or chemical crosslinkers [88]. Since then several improvements, such as multi-material bioprinting and patterned geometries, have been made in the direction of manufacturing of tissue models via TPL [87, 89]. Indirect bioprinting based on TPL has been applied to generate hollow models. In a recent work, multiscale channels were created within polyethylene glycol (PEG) tetrabi-cyclononyne hydrogels by photodegrading the bioink at the target pattern [40].

5.1. Current approaches

There are currently four different approaches for shape transformation of a material/biomaterial: 1) manual transformation 2) spontaneous shape transformation 3) cell contraction 4) stimuli-responsive [94]. Among these, the stimuli-responsive process may occur with different stimulus such as moisture, temperature, pH, magnetic field, electric field, light and so forth [92]. Use of stimuli-responsive (bio)materials offer several advantages over the other conventional methods, such as precise control on the moment when shape transformation is needed, simultaneously folding of multiple objects made of different materials, and the ability to fold micron-sized objects. These important characteristics may allow the biofabrication with cells being seeded onto the stimuli-responsive material or encapsulated within the smart material [94].

Temperature stimulation has been used to induce shape transformation in bioprinted constructs more than any other factor. Figure 4B shows a folding star-shaped thermos-responsive polymer bilayer. As it can be seen, both theoretical (simulation) investigations and real microscopy images indicate that this thermos-responsive hydrogel swells and deforms into a capsule-like structure at reduced temperatures, and it unfolds and releases the encapsulated cells by increasing temperature, offering a promising approach for drug delivery applications [95, 96].

Another frequently-used stimulus is water, based on the level of water sorption (i.e., swelling) in various hydrogel systems. For instance, the swelling of the bioprinted scaffolds occur through water sorption when they immersed in an aqueous solution. Figure 4C shows a doublelayered printed construct consisted of patterned PEG following self-folding controlled by the swelling differences of individual layers in water environment, leading to the deformation of the constructs [93]. When cells were embedded in the PEG bilayers, water exposure will induce the self-folding behavior of the construct into cylinders of various sizes, with minimal damage to the cells.

Another approach for shape transformation of a biomaterial is cell contraction. Cells, with the ability of exerting traction force, can adhere to a substrate/biomaterial surface and induce self-folding based shape transformation to fabricate complex cellular constructs (Figure 4D). In this approach, the angle θ (i.e., the angle between the folding microplate and glass substrate and can be measured by the cell amount adhered onto microplates) plays a vital role in making desired 3D cell-laden microstructure.

5.2. Proposed solutions for vascular chips

There has been a great effort on developing novel 4D biofabrication approaches for vascular systems. For instance, a group of researchers have employed shape-morphing hydrogel systems to fabricate very small internal diameter tubes, as low as 20 μm. This is comparable to blood capillaries in vivo and is not yet feasible by other bioprinting approaches (Figure 4E) [90]. The self-folded hydrogel-based tubes fabricated via their proposed approach supported cell growth and proliferation, as well as cell survival, while maintaining high viability of the cells by day 7.

Although the 4D bioprinting has been recently developed, several groups tried to provide the bioprinting scientific community with a better understanding of this approach through writing comprehensive review papers on this matter [93, 94]. But in a general overview, with the aim of creating in vitro models that resemble tissue structures found in nature, 4D bioprinting promotes dynamic, structural, and cellular changes of a tissue over time, overcoming the static nature of 3D bioprinting [97].

6.1. Concluding remarks

Three major factors for the design of microfluidic organ-on-chip models are chip function, degree of integration, and application [8]. To mimic complex architectures of micro-tissues or organoids in traditional organ-on-chip models, the fabrication of multiple PDMS layers and sequential integrations are required, which are expensive, time-consuming, labor-intensive, and difficult to get automated for continuous production. The emergence of 3D bioprinting techniques have led to novel opportunities for economic and rapid prototyping of micro-tissues, and they have shown potentials for constructing organ-on-chips in automated procedures for continuous production. Adapting 3D bioprinting in microfluidics platforms allows the construction of desired micro-tissue in organ-on-chip models using inexpensive desktop 3D bioprinter platforms, while the microfluidic set-up includes expandable electrochemical monitoring, sensing elements for electrochemical transductions, and enzymes and biomarkers. This can be achieved by the integration of supporting technologies with bioprinting platforms.

6.2. Future directions

There are still some challenges for creating functional microfluidic constructs towards commercialization. When applying 3D bioprinting to produce micro-tissue models, the post-integration of microfluidic chips and micro-tissues is important. Compared to traditional organ-on-chip models, bioprinted micro-tissues inside microfluidic chips will encounter much more shear stresses given by perfusion flows due to their spatial architecture. Optimum scaffold structure for better adherence and preserving structural integrity over long-term will be another challenge. For example, three-(trimethoxysilyl)-propyl methacrylate coating has been applied for better adherence onto glass substrate for bioprinted GelMA- and PEGDA-based bioinks [31]. Creation of customized supports and connecting scaffolds are also straightforward using 3D printing techniques. Furthermore, micro-tissue fabrication has become more automatic and robust, which will ultimately lead to single-step biofabrication of organ-on-chip for scaling-up from lab scale to mass production. Microfluidic chips can be also made in an automated manner since 3D printing has shown the capability to construct microfluidic platforms [41, 43], as well as some advances in microelectronic systems [99]. The combination of 3D printing of microfluidics fabrication and bioprinting techniques shows great promise in the direction of single-step organ-on-chip engineering. In addition, we can develop new capabilities, such as repeatable microelectrodes, which are unrealistic by traditional microfabrication methods. We envision further technologies that will enable online rapid fabrication of (electrically-) conductive materials and microchannel encasements along with integrated micro-tissue constructs. This way, the required labor works to make such assemblies will be significantly reduced during and after the printing process. In this perspective, we propose using multi-material, extrusion-assisted printers to deposit a mixture of electrically-conductive, nonconductive, and sacrificial materials as an integrated construct. This strategy would pave the path towards commercializing organ-on-chip platforms.