The results of flow behavior analysis for the 40FAH solution at 4 and 22 °C are shown in Figures 6.7A, B. The hydrogel solution exhibits a non-Newtonian, shear thinning flow behavior because the resulting shear stress-shear rate curve has a nonlinear, concave pattern [32]. Figure 6.7A also verifies that the yield stress of the 40FAH solution at both temperatures approximately equals zero [33]. Therefore, a power law model without considering the yield stress can be used to represent the flow behavior of the hydrogel solution [34]. Figure 6.7C exhibits low viscosity values of 40FAH at both 4 °C (1.23-7.35 Pa∙s) and 22 °C (0.56-4.36 Pa∙s) in the tested range of shear rate. The viscosity of the control group (alginate and HA solution) at 22 °C shows a litter lower viscosity than the performance of 40FAH solution. Compared to previous study with alginate and HA solution, the 40FAH solution at 22 °C could be dispensed under controllable manner [20].
One important element for controlled production of a scaffold by bioprinting is the solution mass flow rate of the printed biomaterial, which can be regulated by changing the dispensing pressure. Figure 6.7D shows the mass flow rate of the 40FAH solution under a series of dispensing pressures ranging from 20 to 60 KPa. A higher pressure will extrude more hydrogel solution from the bioprinting needle. Based on the flow behavior pattern obtained from the rheometer test, the flow rate trend can be described and predicted [33], with the result shown in Figure 6.7D.
If the pressure is known, the mass flow rate can be obtained accordingly from the relationship described above. Based on the flow rate, the dispensing head speed is therefore determined in terms of scaffold structural requirements, such as the porosity and structural stability. Because our bioprinting method is combined with the submerge technique by which the hydrogel solution is deposited in a reservoir containing crosslinking solution for scaffold gelation, the concentration of crosslinkers becomes another crucial element to determine the success of scaffold bioprinting. Figure 6.7E shows how the printability is affected by the dispensing head speed and concentration of the calcium crosslinker. In detail, the influence of dispensing head speed on structure’s integrity shows that, in a fixed
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20 mM of calcium ions, a wide range of speeds (2 to 10 mm/s) can be applied to build scaffolds. In contrast, when the speed is too fast (over 11 mm/s) or too slow (less than 1 mm/s in this case), stacked scaffold with multiple layers is hard to produce. Meanwhile, when the dispensing head speed is preset at a given value, low concentrations of calcium ions (20 to 40 mM) are adequate for the success of scaffold bioprinting. However, three- layer scaffolds are hard to achieve at calcium concentrations >50 mM or <10 mM.
Figure 6. 7 Bioprinting control for building scaffolds: A, shear stress of 40FAH hydrogel at a low shear rate; B, flow behavior of 40FAH hydrogel over a wide range of shear rates;
C, mass flow rate under various dispensing pressures; and D, printability under various dispensing head speeds and calcium concentrations (check represents adequate
printability, × represents poor printability).
The speed of the dispensing head not only affects the printability of scaffolds but also determines the diameter of scaffold strands and their inner stresses if the dispensing pressure is known. Theoretically, the diameter of a printed strand equals the inner diameter of the bioprinting needle if the dispensing head speed is preset [35]:
𝑣 = 4𝑄
𝜋𝜌𝑑2, (6.3)
where v is the speed of the dispensing head, Q is the mass flow rate, ρ is the solution density (close to 1 g/mL herein), and d is the inner diameter of the bioprinting needle. Figure 6.8 shows the variation of strand diameters (top view) under different dispensing head speeds for a constant dispensing pressure and crosslinking solution concentration. A faster speed
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can lead to narrower strands under the same dispensing pressure; however, the strands all display larger diameters than predicted by theoretical calculations for the dispensing head speeds applied (theoretical values: 4 mm/s, bigger than needle diameter; 6 mm/s, close to needle diameter; 10 mm/s, less than needle diameter). This is probably due to surface tension of the hydrogel solution as it deposited on the preset surface, and the effect of crosslinking rate [36].
Figure 6. 8 Influence of dispensing head speed on the diameter of printed strands: strands formed at A, 4 mm/s; B, 6 mm/s; and C, 10 mm/s. D, Comparison of actual and
theoretical diameters.
Immunofluorescence staining of fibrin allowed for the observation of fibrin fibers inside the printed strands. Thus, the influence of dispensing head speed on the orientation of fibrin fibers within strands can be observed. Fibrin fibers appeared to be oriented parallel to the printing direction, with this orientation being more obvious when higher dispensing head speeds were applied (Figures 6.9A-C). For example, the strand printed at 2 mm/s had 74.67±8.34% of fibrin fibers distributed within ± 20° of the strand orientation, with this value increasing to 94.36±3.51% at 9 mm/s. However, fibrin fibers in droplets formed random mesh networks, with only 40.67±7.23% of fibers oriented within ±20° relative to the 0° orientation defined for image analysis (Figures 6.9D1, D2).
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Figure 6. 9 Orientation of fibrin fibers inside printed strands and dispensed droplets: fibrin fiber orientation at a dispensing head speed of A, 2 mm/s; B, 6 mm/s; and C, 9
mm/s. D, Fibrin fiber orientation inside a droplet.
After investigating elements including the flow behavior of the hydrogel solution, dispensing pressure, speed of the dispensing head, and crosslinking agent concentration, scaffolds were produced using low viscosity 40FAH and FRAH solutions, resulting in the printed structures. Figure 6.10 reveals the example of scaffold produced by 40FAH. The 3D multi-layer scaffold has integrated morphologies with predesigned pores and architecture (Figures 6.10A to 6.10C), indicating that porous tissue structures can be achieved using a low viscosity hydrogel solution following our bioprinting method. The sufficient mechanical support from subjacent layers of the scaffold made it stable to be handled by hands or forceps during the transition for image capturing. Both phase retrieval images and 3D reconstructed structure after SR-inline-PCI-CT imaging confirm that the scaffold with fully interconnected channels and pore networks can be fabricated (Figures 6.10D to 6.10F), with the quantitatively average porosities of 39.42±11.93% and 43.65±8.42% for 40FAH and FRAH hydrogel scaffolds, respectively.
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Figure 6. 10 Observations of printed 40FAH scaffolds. A, scaffold in the crosslinking medium after bioprinting; B, top view of scaffold shape and size; C, side view of scaffold; D, a phase retrieval slice of printed scaffold captured using our SR-inline-PCI-
CT imaging technique; E, top view of a scaffold after reconstruction; and F, reconstructed 3D scaffold