## 4 Discussion  
Hemolysis can be defined as damage to the cytoplasmic membrane of RBCs due to mechanical or chemical factors in the environment (in vivo or extracorporeal), resulting in lysis and content release into plasma. In the context of blood and medical device interactions, hemolysis may occur due to local mechanical forces (mechanically induced hemolysis) and biochemical factors (material-induced hemolysis). Particularly, mechanically induced hemolysis, caused by forced blood circulation in medical devices, is a risk factor that should be considered during the development of medical devices [1]. The present study confirmed the comparative indicators of RBC morphology and quantitative hemolysis level and the suitability of the evaluation system using in vitro MCL based on ASTM F1841-19 (a standardized method approved by the FDA). Furthermore, we designed a hemocompatibility evaluation model that could be used as an interpretation  
Figure 5: Hemolysis analysis. (**A**) Delta plasma-free hemoglobin (\(\Delta\)pfHb) and (**B**) normalized hemolytic index (NIH) over time. Roller pumps are depicted in dark gray and centrifugal pumps are depicted in light gray. (**C**) Light microscopy observations of untreated red blood cells (200\(\times\)). (**D**) The percentages of morphologically intact red blood cells over time. index for the repeatability and reproducibility limitations for the application of mechanically induced hemolysis results. The evaluation of mechanical-induced hemolysis in blood pumps requires loop systems to simulate the clinical environment. Effort has been made previously to minimize hemolytic variables. [PERSON] et al. [29] applied resistance to 150 mm tubing to minimize hemolysis caused by the local pressure of the tube when controlling the flow rate and pressure head of the pump. [PERSON] et al. [30] minimized the effect of fluid inertia by narrowing the distance between the atrium and ventricle to reduce hemolysis caused by blood flow in a simulated circulatory loop for pulsatile total artificial heart testing. We constructed a loop system according to ASTM F1841-19 to evaluate mechanically induced hemolysis. To minimize hemolytic variables, changes in the internal diameter were minimized, and an optimized reservoir was designed to prevent distortion caused by blood congestion. Although the blood temperature was controlled to exclude thermal factors generated in the system, the blood equilibrium temperature of the centrifugal pump was higher than that of the roller pump under the same temperature-control conditions. This implies that the heat generated in the centrifugal pump system, including the motor, is higher than that generated in the roller pump system. According to [PERSON] et al. [31], RBCs exposed to high temperatures experience increased hemolysis because of thermal denaturation. Nevertheless, the hemolysis level in the centrifugal pump in the present study was lower than that of the roller pump, which implies that the heat-related hemolysis effect was limited in our test system and that the evaluation system was properly deployed. We also observed the morphological changes of RBCs as auxiliary indicators for the evaluation of hemolysis in medical devices. The ASTM F1841-19 and F1830-19 standards specify that \"blood is generally used within 48 h of blood draw, including the time for transport. Best consistency may be expected when blood is used within 24 h; older blood may only provide reliable comparison in paired studies.\" [6; 7]. However, we attempted to confirm the test method by reflecting the contents of the standard through a small number of repetitions as a limitation section. Specific information on how to determine blood conditions for test application and the differences in blood preservation techniques between anticoagulants is not provided. We confirmed morphological changes in RBCs using porcine blood based on ASTM F1830-recommended anticoagulants ACD-A, CPDA-1, and heparin.