## 4 Discussion  
In conclusion, we have shown that rMM-OCT provides a higher magnetomotive signal at a lower power consumption compared to previous pulsed MM-OCT. This improvement was mostly achieved  
Figure 6: Quantification of the cellular magnetomotive signal using OCT and laser speckle analysis. In each OCT (**a**) and laser speckle image (**b**) a region of interest (ROI) representing the cell is chosen and the mean magnetomotive signal is determined; (**c**) Magnetomotive signals retrieved from OCT data (left column) and laser speckle data (right column) for five different cells imaged with our previous pulsed MM-OCT setup (dark bars) and the novel rMM-OCT setup (bright bars) using a capacitor charging voltage of 120 V and 30 V, respectively. The magnetomotive signals represent the normalized mean variance in the ROI, i.e., the mean pixel variance divided by the average pixel intensity across the magnetic pulse. Data represents mean values \(\pm\) standard deviation (color online). Figure 7: Volumetric rMM-OCT imaging of labeled ARPE-19 cells embedded in an agar scaffold. (**a**) Top view and (**b**) side view of a recorded 3D OCT image stack showing two cells next to each other. (**c**) 3D view of the two cells. All views represent maximum intensity projections of the 3D image data (color online). by adding a small permanent magnet in an off-axis configuration to the MM-OCT setup. The thickness of the permanent magnet is only 2 mm; hence, the reduction of working distance from 6 mm (relative to the surface of the solenoid) in the original setup to 4 mm (relative to the surface of the permanent magnet) in the new setup is still acceptable for retinal imaging in small animal models used in cell transplantation studies, such as mice (eye length 3 mm). Further, the dual-magnet configuration of rMM-OCT is a single-sided setup that circumvents the restrictions of previous dual-coil configurations to small samples. In this context, we have also shown that rMM-OCT allows robust and reliable contrast-enhancement in different cell types, such as MSCs and transplantable ARPE-19 cells. We note that magnetomotion in cells can also be induced by other MM-OCT methods, such as pulsed MM-OCT using our previous setup. However, our results shown in Figure 6 imply that the magnetomotive signal of rMM-OCT is significantly higher compared to the previous pulsed MM-OCT. Besides particle chaining and rotation, this signal increase might be additionally supported by particle clumping induced by the permanent magnet leading to enhanced magnetic susceptibility of the contrast agent and, thus, improved displacement in response to the applied magnetic field. As a consequence, rMM-OCT permits higher sensitivity, the use of low power magnetic fields or corresponding larger working distances, and single-sided access to biological specimens. This makes it a promising method for future 3D cell tracking in target tissues and organs, such as the retina in small animal models, under in vivo conditions. However, further investigation is required to adress several aspects of the method presented herein. Most important, as the induced magnetic particle changes to elongated chains (as shown in Figure 4) end up being very long in glycerin (~10 \(\upmu\)m) and approach the diameters of cells, cell viability experiments are necessary to investigate whether particle chain formation and rotation during rMM-OCT imaging may be destructive to cells. Additionally, it is not yet clear what the size these chains in cells may reach, how controllable this is in conjunction with the magnetization time (currently 30 min), and what the minimum detectable concentration of magnetic particles is or the minimum detectable size of particle aggregates. In this context, the current magnetization time of 30 min is relatively long, especially with regards to prospective in vivo applications.