## 4 Discussion  
Assigning optimal levels of process parameters is critical to the micro-architecture and geometric fidelity of bioplotted scaffolds. A previous study had demonstrated how changes in nozzle diameter and extrusion temperature affect the bioplotted PCL strand width [11]. Phase-1 results of this study show that for a given nozzle diameter and extrusion temperature, changes in extrusion pressure and nozzle speed significantly affect the strand width, while the effects of the two geometry-specific factors (lay orientation and strand length) are not significant. The strand width is directly proportional to extrusion pressure. This phenomenon occurs due to the increase in flow rate of the extruded polymer through the nozzle with increasing pressure [17]. On the other hand, the strand width is inversely proportional to the nozzle speed. As the speed increases, there is a decrease in the dispensing time of the polymer per unit length traversed by the nozzle. With the extrusion flow rate remaining constant, this leads to a decrease in the volume of polymer deposited per unit length traversed by the nozzle. Figure 6: FE estimated equivalent (Von Mises) stress at a maximum strain of 0.59: (**a**) Stress distribution along X–Y plane, and (**b**) Stress distribution across the cross-section of the model, and (**c**) average stress across each layer oriented along or perpendicular to the direction of tensile loading (0\({}^{\circ}\) and 90\({}^{\circ}\), respectively). Figure 7: (**a**) Stress versus strain plots for three porous tensile measures represented by lines with different types of line style and (**b**) Comparison of experimental tensile properties of bulk and porous specimens. Error bars denote standard deviation (n = 3). This decrease in volume deposited per unit length, in turn, translates to a decrease in strand width [33]. The results from the aB reduction assay confirm that the strand width and corresponding inter-strand spacing, which govern the micro-architecture of the scaffold, play an important role in regulating cellular activity. Whereas the overall size and ratio of strand width to inter-strand spacing were the same in both 3 DB1 and 3 DB2 scaffold designs (\(\mathcal{O}\,20\times 2\) mm and 1:1, respectively), the smaller biplotted strand width and corresponding inter-strand spacing resulted in higher effective surface area of 3 DB2 (3462.85 mm\({}^{2}\)) compared to 3 DB1 (2367.61 mm\({}^{2}\)), as estimated from their CAD models. Higher cell proliferation observed in 3 DB2 can be primarily attributed to the higher surface area available for the adherent cells to attach to and proliferate. This observation is consistent with higher cellular migration and adhesion of pre-osteoblastic murine cells reported in scaffolds having higher effective surface area [34]. As such, having precise control over the strand width through mapping of process parameters can streamline the efforts in designing scaffolds for different tissue engineering applications. Tensile testing of bulk specimens in Phase-2 was performed in order to create the FE model. Any FE model requires a base line of the mechanical properties of the materials used in the model. Since, the mechanical properties of solid bioplotted PCL are not reported in the literature, the bulk samples of 0\({}^{\circ}\) and 90\({}^{\circ}\) configurations were fabricated and tensile tested to determine the appropriate solid 3 DB PCL's tensile properties. Table 3 presents a summary of PCL tensile properties reported in the literature. Similar to most other polymers, tensile properties of PCL vary with the molecular weight [35; 36] and the degree of mechanical compaction due to the pressure applied during primary and secondary processing operations [37; 38]. For example, the strength of structures manufactured via injection molding and extrusion is typically greater than that of structures processed using AM [22; 29].