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Gu, et al.
properties of gelatin and GelMA were then demonstrated. a little bit larger than the outer diameter of the nozzle.
Shear-thinning behaviors of GelMA/gelatin were Along with the decrease in Q outer /Q inner , the inner diameter
observed, with excellent extrusion properties of the fiber increased gradually.
(Figure 3Aiii). When facing periodic amplitude sweep In consideration of a good nutrient supply
of shear strain, gelatin showed good response-to- and subsequent unobstructed perfusion, the tubular
deformation capability with fast switching: In the case constructs with outer/inner diameters of approximately
of small amplitude (1%), G′ > G″, representing a gel- 3 mm/1 mm were ideal. Hence, nozzle-1-3 was
like status; in the case of large amplitude (200%), G′ < selected for follow-up bioprinting. The relationship
G″, representing a liquid-like status. However, GelMA between the outer/inner flow rate and outer/inner fiber
had a relatively poor response capability (Figure 3Aiv). diameter could be calculated according to the formula
That is why a 2°C cooling platform was introduced for Q outer S outer D − d 2
2
printed fiber deposition, as shown in Figure 2Ai, to Q = S = d 2 , where S outer and S inner were
ensured better structural fidelity. inner inner
the cross-sectional area of the outer/inner channel of
3.2.2 Printing of unicursal patterns the fiber, and D and d were the outer/inner diameters
of the fiber. Putting D = 3 mm, and d = 1 mm into the
To mimic the complicated structural features of formula, Q /Q = 8, which was identical to the result
organoids, freeform biofabrication of coaxial bioprinting presented in Figure 3Eii. Therefore, Q of 0.3 ml/min
outer
inner
outer
GelMA/gelatin bioinks was desirable. The versatility of and Q of 0.0375 ml/min were employed in subsequent
inner
bioprinting was well exhibited through several unicursal experiments. Under these bioprinting parameters, the
patterns. When using a coaxial nozzle-18G/25G, all average outer/inner diameters of hydrogel tubes were
printed patterns that consisted of one continuous hollow 3116 μm and 1063 μm, respectively.
filament closely resembled the morphology of targets
(Figure 3B and Figure S3). 3.3. Comparison of hydrogel constructs with/
without a PCL stent
3.2.3. Flow rate analysis
3.3.1. Mechanical properties of the hydrogel bulk with/
The outer diameter of vessel-like structures biofabricated without a stent
through nozzle-18G/25G was approximately
1200 μm. To biofabricate large-scale tubular constructs To better understand the advantage of PCL stents
mimicking human blood vessels for perfusion culture, during perfusion culture, a mechanical property test
the use of nozzles with larger diameters was inevitable. was first applied to measure the improvement of the
Nozzle-1-2.5 (inner/outer diameters of the nozzle were structural deformation-resistance capacity that a stent
1 mm/2.5 mm) and nozzle-1-3 (inner/outer diameters would provide, as shown in Figure 4Ai. The hydrogel
of the nozzle were 1 mm/3 mm) were adopted for bulks containing tube with/without stent were exhibited
the experiments below. The printability windows of vividly in lateral and cross sections of SEM images in
nozzle-1-2.5 and nozzle-1-3 are shown in Figure 3D Figure S4. Figure 4Aii shows the compressive stress-
as the inner/outer flow rates varied. If the difference strain curves of the bulks. In the early stage of the curves
between the flow rate of the outer channel (Q outer ) (strain <20%), it was mainly the hydrogel part that
and the flow rate of the inner channel (Q inner ) was too resisted deformation by force, whereas the two curves
large, the tubular lumen could not be maintained; if were almost indistinguishable. As pressure increased
Q outer and Q inner were too close, the shell layer could not (strain >20%), the tubular lumen of the bulk began to
completely enfold the core layer. These two conditions be squeezed. The hydrogel bulk with a stent presented
were considered not printable since no clear hollow tube increasing superiority with ascending force. The higher
was achieved. the strain was, the greater the stress differed between
Apparently, the construction of tunable tubular bulks with/without a stent. When facing the same stress
structures could be achieved by regulating the flow of 30 kPa for instance, the hydrogel bulk with a stent
rate of GelMA/gelatin or changing the size of nozzles. sustained a strain of approximately 60% while the bulk
It was easy to understand that the fiber extrusion speed without a stent reached 70%. The stresses the hydrogel
was affected by the flow rate, and the inner/outer bulk experienced at the same strain (50% and 60%) were
diameters of the printed fibers were affected by the ratio also studied, as shown in Figure 4Aiii. At a strain of 50%,
of the outer/inner flow rate (Q outer /Q inner ). Figure 3E the stress the bulk with a stent sustained was 1.53 times
displays the variation of printed fiber diameters greater than that of the bulk without a stent. At the strain
corresponding to the modulation of Q outer /Q inner while of 60%, this value increased to 2.2 times. These results
using nozzle-1-2.5/nozzle-1-3. It is worth noting that the demonstrated that the PCL stent offered greater resistance
outer diameter of the fiber remained almost the same as to deformation.

International Journal of Bioprinting (2022)–Volume 8, Issue 4 299