Breaking the resolution limits of 3D bioprinting: future opportunities and present challenges

• High-definition (HD) bioprinting enables spatial resolution on a cellular and subcellular level in 3D, allowing reproduction of key features of the cellular microenvironment at a scale not achievable with conventional bioprinting techniques, and allowing control of material properties, geometry, and chemical and physical properties of cell-containing constructs.

• Light-based, precision jetting, and electrohydrodynamic technologies can already achieve such resolution, and will be increasingly applied to engineer disease models, organ-on-a-chip devices, and implantable microscaffolds with high complexity.

• Standing challenges include preserving microscale and submicroscale resolution while enabling high-throughput and volumetric construct bioprinting, streamlined multimaterial processing, and the development of new functional (bio)inks and (bio)resins.

Bioprinting aims to produce 3D structures from which embedded cells can receive mechanical and chemical stimuli that influence their behavior, direct their organization and migration, and promote differentiation, in a similar way to what happens within the native extracellular matrix. However, limited spatial resolution has been a bottleneck for conventional 3D bioprinting approaches. Reproducing fine features at the cellular scale, while maintaining a reasonable printing volume, is necessary to enable the biofabrication of more complex and functional tissue and organ models. In this opinion article we recount the emergence of, and discuss the most promising, high-definition (HD) bioprinting techniques to achieve this goal, discussing which obstacles remain to be overcome, and which applications are envisioned in the tissue engineering field.

In their native environment, cells receive and decode a broad variety of highly specialized signals from the extracellular matrix (ECM) (see Glossary) and from neighboring cells, as well as a range of soluble biochemical cues. The combination and spatiotemporal presentation of these signals is intimately linked to cell behavior at all levels, and plays a key role, for instance, in stem- and progenitor-cell differentiation [

], cell migration [

], and tissue and organ development [

]. In particular, the ECM has been extensively demonstrated to steer cell and tissue function during tissue morphogenesis and development, and tissue regeneration, as well as in degenerative diseases and cancer, as cells sense local variations in mechanical properties, molecular composition, and 3D (micro)architecture [

] (Figure 1). For example, it was recently demonstrated that the composition of the hydrogel matrix – that mimics the mechanical properties and chemical composition of the ECM– plays a decisive role in replicating morphogenic events allowing intestinal organoid maturation in vitro [

]. Furthermore, creating ~30 μm wide regions of softer hydrogel next to an embedded intestinal stem-cell colony enabled precise control of its shape and the formation of characteristic crypt-like buds [

]. Such spatiotemporal control of intestinal organoid shape on a microscale gave rise to a characteristic distribution of cells (Figure 1B,C). Likewise, inspired by the notion that cells align and accommodate their shape in agreement with the orientation (or lack of thereof) of ECM fibrous structures in vivo [

], topographical elements and matrix mechanical properties at the micro and submicron scales have been extensively studied in the past two decades [

], resulting in the production of biomaterials and cell-laden constructs capable of giving directionality to tissue regeneration, mimicking the anisotropic architecture of tissues (i.e., in skeletal and cardiac muscle) [

], and boosting stem-cell differentiation [

]. Phenotypic transition of fibroblasts to myoblasts is another important phenomenon associated with mechanobiological mechanisms. A recent study demonstrated a predominant role for the directionality of the mechanical signals from the ECM rather than their intensity [

]. Cells interacted with the surrounding collagen fibers to develop tension anisotropy in a two-way cell–ECM feedback. Only cell pretreatment with protrusion-inhibiting nocodazole led to significantly reduced collagen alignment (Figure 1F,G). Most often, these phenomena are investigated on 2.5D patterned surfaces, that is, surfaces with low aspect ratio features, or whose features are in contact with only one side of the cell, produced using conventional replica molding and lithographic techniques [

]. While these methods are highly versatile, allow high resolution (down to hundreds of nanometers) and are compatible with high-throughput screening of broad arrays of topographical features [

], they do not provide sufficient freedom for realizing arbitrary 3D structures typical of native cell-laden environments.

In the past decade, most of the advances focused on variations of extrusion-based techniques, in which cell-laden hydrogel formulations (also termed bioinks) are plotted as filaments upon shearing through a nozzle, in a layer-by-layer fashion to form 3D objects. In order to limit the shear stresses imparted on cells, and therefore ensure their viability, extruded filaments typically have diameters larger than ~100 μm, and are thus unable to resolve smaller features of the native microenvironment. An exception is constituted by elements printed out of collagen in a support bath, where generation of filaments down to 20–30 μm has been reported; these have been used to produce hollow vessels, cardiac valves, and a neonatal-sized heart [

]. On the downside, printing fidelity for centimeter-sized structures is still not better than 500 μm.

Light-based techniques such as SLA and DLP are gaining popularity in the bioprinting community. Nevertheless, their resolution is usually still in the range of several tens of micrometers [

,

]. This is in part because the lateral resolution is often limited by the photochemistry of the crosslinking process rather than by the minimum laser spot or pixel size, while layer thickness is inherently bound to the light penetration depth. For instance, researchers working with a micrometer-resolution DLP found that, despite the theoretical optical capabilities of their setup, good-quality features were obtained only from a lateral dimension of 100 μm and vertical dimension of 300 μm [

], while many authors reported a layer spacing of between 40 and 50 μm on hydrogels [

,

,

]. In a recent work, Bhusal and colleagues showed lines of 15 μm thickness (Figure 2A ) using cell-free polyethylene glycol diacrylate (PEGDA) hydrogel [

], which is already close to cell size. In general, DLP resolution is also a trade-off between the desired feature size and the sample size, both bound to projector resolution and optics used. An interesting variation, based on the combination of lightsheet excitation and DLP with a dual-color photoinitiator system, could reduce the layer thickness and eliminate the need for layer deposition, reaching 25 μm horizontal and 50 μm vertical resolution in a cell-free resin [

].

A new class of light-based VP techniques has shown potential as future HD bioprinting candidates. The unprecedented speed, absence of layer-by-layer material deposition, scalability, and high-volume capabilities (tens of cubic centimeters) provide multiple advantages rarely found simultaneously with other techniques. 3D structuring is achieved by delivering an anisotropic 3D visible light dosage distribution within a photocurable resin using filtered tomographic back-projections [

,

]. This is of particular interest for biofabrication, where the absence of any shear stresses is beneficial for cell viability. To date, VP has demonstrated feature sizes down to ≈40 μm, and realization of cell-laden and even complex organoid-laden biomaterials with >90% viability [

]. Current challenges regarding the maximum attainable resolution are mainly due to intrinsic losses of reconstruction information, resulting from the back-projection algorithms employed, diffusion of reactive species in the material during crosslinking, and material-dependent fragility of the as-printed part. Minimizing optical attenuation, aberrations, and scattering becomes crucial for accurate results. In cell-laden resins, where cells cause scattering, techniques such as refractive index matching [

] and software-level corrections of the tomographic projections [

] can be effectively employed. Such mitigation strategies are still an active area of research.

Inkjet printing [

,

] and laser-induced forward transfer (LIFT) [

,

] are two patterning techniques that exploit the deposition of small cell-containing material droplets. Despite the nominally high lateral resolution, they generally struggle to produce thick and high-aspect-ratio structures with the same precision, because the bioinks used are usually unable to (i) crosslink quickly after deposition, and (ii) support multiple layers. A few strategies have been proposed to overcome this issue, such as combining the inks with crosslinkers [

], or print at the surface of a crosslinker bath [

].

Among the high-resolution techniques that were recently adopted by the biofabrication community, EW of polymer melts of solutions, together with electrospinning, offers the possibility of producing polymer structures by controlled deposition of thin fibers [

], even below 100 nm of lateral resolution with special near-field setups [

]. EW has been applied also to water-rich hydrogels [

,

], and recently it was demonstrated to print subcellular size fibers (≈5 μm) with cell-laden hydrogel-based bioinks (cell EW, Figure 2B) [

].

Currently, the most widespread technique for high-resolution processing of biocompatible materials, including cell-laden hydrogels, is MPL. It can work freely within a material volume, achieving resolution down to <1 μm in 3D. However, it is inherently slower than DLP, since (like SLA) it generally relies on scanning line by line the whole volume of interest. There have been many examples of producing scaffolds for cell culture [

,

,

,

], building simple organ models [

,

,

] (Figure 2D,F), and even in situ printing within living animals (rodents) [

]. Notably, besides conventional polymerization, MPL enables photodegradation and cleaving of sacrificial moieties [

], molecule grafting [

], and local densification of a partially crosslinked hydrogel matrix [

] (Figure 2C,E).

Together, these techniques will eventually form a versatile toolbox, able to take bioprinting capabilities to the cellular and subcellular scales; we refer to them in general as HD bioprinting techniques. An overview of the aforementioned techniques plotting throughput versus resolution is presented in Figure 3.

The convergence of high-resolution additive manufacturing approaches has recently accelerated the development of progressively more complex and clinically relevant applications in bioprinting, with highly accurate spatial control. This is imperative to achieve the detailed microarchitectures necessary for recapitulating – at least morphologically – the relevant physiology or anatomy in question [

,

]. Compared to photolithography and replica-based techniques, which typically generate 2.5 D geometries, HD bioprinting techniques produce fully 3D structures, providing a better mimicry of tissue architectures. Notably, light-based technologies can produce structures even across transparent tissue culture labware. Bioprinting directly within a microfluidic chip becomes possible [

,

,

,

], hence allowing the constructs to be fabricated noninvasively, avoiding sterility issues, and reducing the number of assembly steps. Faster development cycles can be enabled by HD bioprinting in the domain of organ-on-a-chip, by producing in a standard chip different internal 3D structures at each redesign and optimization step. Also, the biofabricated constructs can be manipulated in a contactless fashion at multiple time points, adding time as a relevant parameter to mimic the evolution of biological processes [

]. At present, as application fields or markets for HD bioprinting, we can envision the production of organ-on-a-chip for basic/fundamental research, personalized drug testing and diagnostics, and microstructured tissue engineering scaffolds for implantation.

Organ-on-a-chip platforms have been leveraged for the modeling of physiological and pathological processes, even as complementary and potentially alternative platforms to animal experimentation. They aim to recreate an organ’s elementary unit or portion (as depicted in Figure 4) in an observable and controllable environment to better understand the underlying biological mechanisms, and enable a faster drug development and testing cycle, within both academia and the pharmaceutical industry [

,

,

]. HD bioprinting allows recreation of a highly organized heterogeneous microenvironment with multiple cell types, where the fine features are produced directly instead of relying on microfabricated molds; examples include the study of microvasculature using direct embedding of cells via MPL [

], the creation of semipermeable barriers to mimic the placenta [

], and the use of micro-SLA for an HD approach towards cell patterning to study soluble cell niche factors [

]. Inkjet printing has also been demonstrated for the bioprinting of a heart-on-a-chip [

] to test hydrogel formulations in the context of cardiac tissue response to different drugs. In the future, organ-on-a-chip devices could lead to point-of-care diagnostics and even personalized in vitro drug testing, on clinically relevant models built with patient-specific cells.

Implantable microstructured tissue scaffolds are a key step towards functional tissue engineering in those cases where it comes to resolving micro and submicron features of the native ECM. Nerves and capillary alignment, for example, require not only sufficient resolution to print the microscale features, but also the fidelity to avoid mismatch between grafted elements [

,

]. A few examples of microstructured tissue scaffolds for implantation have been demonstrated through MPL, even though they could not be properly labeled as bioprinting, since cells were seeded or invaded them after printing [

,

,

]. Intravital bioprinting using MPL has also been demonstrated, in which a cell-laden hydrogel precursor is injected subcutaneously, and subsequently patterned and cross-linked to create a scaffold for cell growth directly inside a living animal, leveraging the deep tissue penetration typical of the infrared light used in MPL [

]. At larger scales, HD bioprinting techniques such as microscale continuous projection (based on DLP) have been demonstrated to produce cell-laden scaffolds for central nervous system repair, with promising in vivo results in rescuing spinal cord alignment and connectivity in a murine model [

].

The MPL community has been facing the throughput issue for years, and some recent advances enabled the production of meso-scale objects while retaining microscopic features, trying to increase the scanning speed, combine efficiently many fields of view, and adapting the shape and the number of laser foci (Box 1).

Besides tight geometrical control, another critical aspect is the combination of different materials; conventional and commercially available extrusion bioprinting hardware can easily use many printheads dispensing different bioinks, and cell EW could offer the same advantage. Vat polymerization techniques, which include DLP, VP, and MPL, may typically need multiple processing steps involving washing off the current printing material and substituting it with the next one, a very time-consuming operation. Elegant solutions involving microfluidic systems to exchange and rapidly mix multiple (bio)resins [

,

,

], as well as multivat DLP stations, have recently been proposed [

]. As these may still face limitations in terms of throughput and potential loss of valuable cells and resins, further development towards new approaches to enable rapid multimaterial processing in vat polymerization and HD bioprinting will be crucial in the future.

Particularly relevant for bioprinting, and not restricted to the HD aspect alone, is the topic of functional materials. Hydrogels are often chosen for their chemical and mechanical resemblance to the ECM, and many literature reports about synthesis and functionalization have been published to obtain well-processable materials that fit the requirements of the various techniques in terms of viscosity, biocompatibility, crosslinking speed, swelling, and degradability [

,

,

,

]. Processability and biocompatibility are necessary conditions for bioprinting, but they are not the only properties that could be controlled in a biomaterial; specific molecules and motifs should be added to the materials to elicit a response from the encapsulated cells. An interesting approach is the use of recombinant materials that replicate the structure of natural proteins such as elastin and collagen with high reproducibility, purity (and therefore improved safety), and strongly reducing batch-to-batch variations. Being bioengineered proteins, these can be custom-modified to include other functional components, for example arginine–glycine–aspartate (RGD) peptides to increase cell adhesion [

,

,

], and photoactive moieties to be processed through MPL [

]. Another strategy to improve biofunctionalization consists in the use of protein-adhesive resins (by grafting binding domains, that is, via MPL or molecular imprinting methods) that promote coating from specific biological molecules onto biofabricated scaffolds [

,

].

HD bioprinting is a growing family of techniques. Herein we have discussed the current limitations and presented some possible strategies to overcome them. While this field is still in its infancy, the publications and reports so far relate to studies on more elaborate, interconnected, and biologically representative structures. For example, the work on vascularization [

,

] and on artificial barriers [

,

] will enable the study of larger size multitissue organoids.

With the currently available HD bioprinting systems, key advances in a relatively short time span can be expected for organ-on-a-chip devices, in which the printing of more native-tissue mimetic architectures (see Outstanding questions) could contribute to creating reliable humanized in vitro models for pharmaceutical testing and precision medicine.