PLGA microspheres were used in this work, but other injectable microscaffolds/microcarriers e.g. microspheres and microfibrils30, fabricated from biodegradable polymeric materials such as polyesters23, polypeptides31, polysaccharides32, and their combinations33, could serve the same purpose in this context. The scaffolds provide 2D surfaces with pseudo-3D environment for the anchorage-dependent cells to adhere and proliferate. At this instant, scaffolds fabricated from biomaterials with a controllable degradation time frame are preferred. PLGA microspheres could degrade within 12 weeks hydrolytically in vitro at 37 °C (Supplementary Fig. S1a,b). Also, a highly porous microscaffold is favourable compared to a full solid scaffold in order to pack high density of cells into the scaffold and achieve expedite degradation after printing. As amorphous PLGA has a glass transition temperature (T g ) higher than 37 °C (Supplementary Fig. S1c), the microspheres could remain non-cohesive23 for smooth printing without clogging.

By using cell-laden microscaffolds in bioprinting of a 3D construct, a lower initial cell density is achievable as compared to that required by conventional cell-laden hydrogels or tissue spheroids/strands. For example, when using highly porous microspheres, initial density of ~2.7 × 104 cells/mm3 is sufficient. By contrast, cell-laden hydrogel printing requires a cell density of ~1.7 × 105 cells/mm3 while tissue spheroids printing entails ~1.8 × 106 cells/mm3 (calculated in Supplementary Data file S1). With the readily available biomaterials and the mature technology in fabrication of the microscaffolds, microscaffold-based bioprinting becomes easy to implement. Size distribution of the microscaffolds can be controlled well by varying the fabrication parameters. Furthermore, the polymeric microscaffolds were found to be stable over a relatively long period after fabrication if properly stored. With the support of these biodegradable microscaffolds, printed constructs can undergo a slower and more controllable process of tissue maturation as compared to the scaffold-free constructs.

The microspheres size range was selected with good reasoning. The chosen as-fabricated microspheres size range (before EtOH-NaOH treatment: sieved to ~90–150 μm) is ideal for this study as pores’ enlargement using EtOH-NaOH is limited by the microsphere size, especially when microsphere sizes are small. It is known that most of the mammalian cells have a diameter of ~10 μm when rounded up after detachment. Cell infiltration would be impossible if the pores on the microspheres are too small, which would reduce the cell density per microsphere. The optimized pore size range was found to be ~0.1–20 μm. Further pore enlargement treatment causes microsphere fragments as shown in Fig. 2a. Submicron pores assist in nutrient exchange while the macropores allow cells to infiltrate. After enlargement of pores, the microspheres size became smaller (60–130 μm) due to the “etching effect” of the EtOH-NaOH. Obviously, a higher printing resolution could be realized by using smaller microspheres. Hence the adopted microspheres size range was determined by a trade-off between printing resolution and cell density. Of note is that size of CLMs became bigger (60–150 μm) because of the slight swelling of PLGA microspheres after submerging in cell culture media at 37 °C34. Cells seeded on the microspheres also contribute to the larger size of the CLMs when compared to the acellular microspheres. The printing resolution achievable here is ~300–600 μm depending on the printing speed, which is of no difference to the current bio-printable resolution35,36,37.

Agarose hydrogel is not uncommon in TE scaffolds fabrication38,39. It has also been used in bioprinting but usually printed as molds14, thanks to its superb thermal properties, extrudability, and non-stickiness. Agarose exhibits pronounced hysteresis between the gelling and the melting40, gels at below 18–42 °C and re-melts at above 60–90 °C. The gelling and melting temperatures depend on its end groups. Low gelling point agarose was chosen in this work so that the hydrogel remains as liquid at a printing temperature of 37 °C. Despite its advantages, agarose is a polysaccharide that lacks the arginyl-glycyl-aspartic acid (RGD) for cell attachment and proliferation. Type I collagen is added as a favored adhesive substrate to make the AC composite hydrogel more cell-affinity. Cell spreading is promoted using the AC hydrogel as compared to the agarose. The AC hydrogel with optimized concentrations of agarose and collagen shows good printability and cell attachment compared to other concentrations of AC hydrogels. Here, the printability means good extrudability; enables immediate gelation upon extrusion onto cold platform; and allows formation of collagen fiber networks within the printed construct during subsequent culture at 37 °C. High concentration gelatin was chosen as a support material due to its attribute as a biocompatible hydrogel gelling at low temperatures. It is commonly used as support material in bioprinting41 as it melts at culture temperature of 37 °C, allowing easy removal of the temporary support material once the bioprinted construct can self-support.

The current bioprinting strategies include extrusion-based, microvalve-based, laser-based, and inkjet-based printing2,42. Microneedle, tapered tip, or nozzle based syringe extrusion methods were widely applied in the reported extrusion-based bioprinting approaches7,12,43,44. We herein found that the micropipette-based extrusion method is a good approach to print tightly packed PLGA microspheres. The tightly packed microspheres encapsulated in thin AC hydrogel samples could sustain a much higher stress and strain than both the AC hydrogel and the loosely packed samples. The mechanical strength of the bioink was dramatically improved by more than 100 times when compared to that of the AC hydrogel. Meanwhile, when the microspheres are loosely packed in AC hydrogel, the mechanical strength of the construct was increased only by 1.5 times compared to the AC hydrogel. The compressive moduli of the tightly packed microspheres samples were superior compared to the AC hydrogel. The improvement in mechanical properties was attributed to the packing of strong polymer microspheres.

This 3D bioprinting approach, when compared to pure hydrogel bioprinting, allows a better stacking ability for the fabrication of 3D constructs. A volumetric 3D tissue can be built up in a few layers. With cushioning or shielding effect from the microspheres, there will be less shear stress-induced cell damage during extrusion-based printing process. The printed construct has been proven to provide a suitable 3D environment for different types of cells to grow. For example, bioprinted rat smooth muscle A10 cells can be clearly seen under SEM (Supplementary Fig. S2a) after 3 days of culture. The F-actin architecture was observed from the printed A10 construct but it was random with no preferential alignment, as opposed to the control (Supplementary Fig. S2b,c). This can be attributed to the random distribution of the isotropic microspheres. It is suggested that cell-laden microfibrils or microfibers45 can be used in bioprinting to provide anisotropic alignment when an aligned tissue is needed such as with neural46 and gastrointestinal tract tissues47,48.

Multipotent mesenchymal and pluripotent stem cells can be expanded by culturing with the microscaffolds49,50. By using microscaffolds in 3D bioprinting, well-established surface treatment on scaffolds can be applied to enhance cell affinity51. Bioactive factors can be added into the scaffolds and thus effectively control the stem cell fate. Levato et al. used a collagen-functionalized solid microcarrier-based approach to print bilayered osteochondral models for the bone compartment24. They proved that the microcarriers facilitated cell adhesion and supported bone cells differentiation by mesenchymal stem cells. The incorporation of proteins and drugs such as growth factors, antibody, and cell adhesion peptides into the microscaffolds can be realized for a sustainable release system in the printed construct. Poldervaart et al. printed bone morphogenetic protein 2 (BMP-2) and vascular endothelial growth factor (VEGF) loaded gelatin microparticles for bone regeneration52,53. A graded concentration of these molecules can also be accomplished in printing. In addition, the microscaffolds can be functionalized, for example by encapsulation of magnets for in vivo imaging of the implanted organ54.

In this study, we have shown the ability of bioprinting CLMs by using the micropipette extrusion-based method. Tight packing of microspheres was achieved. However, both resolution and speed of extrusion-based printing method are still not in the desired range for large tissue printing. Indirect printing of the construct, e.g. printing of sacrificial moulds before extrusion of the bioink, can improve the printing resolution due to dimensional constriction whilst within the limitation of the microsphere sizes. We also suggest that 3D bioprinting of microscaffolds is not restricted to extrusion-mode. For instance, CLMs could be tightly packed in photocurable hydrogels, followed by photo-curing layer by layer for highly efficient printing of tissue constructs55. To further explore the possibility of the microscaffolds bioprinting approach, future strategies could be extended to utilize new material combinations, produce different forms of microscaffolds, and develop novel printing methods to assemble CLMs.