## 4 Discussion
Our data show pore sizes of about 800 \(\upmu\)m and 200 \(\upmu\)m. The recommended ranges of pore size for bone tissue regeneration are 150-900 \(\upmu\)m [42, 43, 44, 45, 46]. Thus, we assumed the developed pore architecture to be suitable for BTE applications. As studies on PCL scaffolds have shown the evocation of osteogenesis in vivo after the seeding of osteoblasts [47] and the successful use in BTE applications [48, 49], we expect comparable behavior of the scaffolds of the present study. PCL is a semi crystalline thermoplastic that is widely used in BTE applications [6, 50] and shows sufficient biocompatibility [11]. The polymer induces only minor immune responses after implantation and is biodegradable [51, 52]. Thereby, degradation starts in situ after 4 to 6 months. Full loss of mass can take more than 48 months [53, 54]. Furthermore, PCL shows positive effects on bone healing processes and the growth of fetal bovine chondrocytes [55, 56]. On mechanical testing, PCL shows viscoelastic behavior [10]. Thus, the material properties depend on the testing velocities and testing temperature, and all testing in this study was performed under the same conditions. Our mechanical data show plastic behavior of the polymer, with a molar mass of 45.000 and a production temperature of 80 \({}^{\circ}\)C, in the range of the comparable literature [9]. Tensile testing of scaffolds shows a linear-elastic elastic-plastic behavior, as shown in other studies [22]. The behavior of PCL scaffolds under compression loads was studied by [PERSON] et al. [57]. Due to the small specimen size of the scaffolds, a perfect axial load application cannot be guaranteed; therefore, the occurrence of bending moments cannot be excluded. Fabricated scaffolds were used for the validation of the parametric numerical models. During testing of the real scaffolds, eccentric load application is not fully preventable. Thus, forces and stresses beyond the loading axis can arise. These effects are not represented in the numerical model and lead to a potential underestimation of the internal forces in the numerical model compared to the testing. Considering these effects, basic variations of the real scaffolds' geometries and material behaviors of PCL, the calculated force ratios demonstrate a sufficient numerical approximation of displacements and forces for strains up to 3% by the developed parametric numerical models. Thus, in this range, forces and strains can be validly determined for specific loads with the present numeral model. Due to the hydrophobic properties of PCL [13], and to improve cell-scaffold interaction [58], scaffolds were treated with plasma generators and were coated with rat collagen type I. Various comparative studies showed a positive effect of plasma treatment and collagen coating on cell seeding and growth [59, 60]. In the present study, collagen coating was visible in electron microscopy images. Studies with immunoassays on comparable scaffolds with the same coating technique show successful coatings [25]. We therefore assume a sufficient collagen coating on the scaffolds in this study. To improve the stability and durability, the crosslinking of collagen via EDC was performed [61]. Seeded cells (Osteoblasts, MSC, MYL-Y4, MG63) are widely used in BTE applications and research [62; 63; 64; 65; 66; 67]. Thus, they can be considered suitable for this study. All cells show growth on PCL strands after three days. Thereby, MG63 cells show pronounced colonization. As these cells are derived from osteosarcoma cells, a more rapid cell division could be expected [68]. The results from the Live-Dead Assay analysis underline the successful seeding of cells, as vital osteoblasts can be demonstrated after three, seven and ten days of growth. In the current study, we aimed to show a qualitative representation of growth. In a further study with the same material and coating, but different scaffold geometry, we were able to demonstrate cell growth. The data are provided in Supplementary Figure S2.