## 2 Materials and Methods
Finally, the averaged flat-field (sample-free) signal at the detector was tabulated as a function of MGD for several experimental conditions, including X-ray energy, sample-to-detector distance, and for different detectors. After the MGD calibration, the ion chamber was removed from the X-ray beam and was not used during the tomographic scanning of samples (the width and height of the ion chamber acceptance are smaller than the corresponding dimensions of the X-ray beam and the samples). Breast tissue imaging at IMBL was conducted on an approximately weekly basis to correlate with the breast surgeries at Monash Health, Victoria, Australia. All the necessary set-up configurations were standardized using a suite of purpose-designed computer scripts to maintain the same experimental conditions for every sample. This included setting the slit positions, detector position, sample position, incident energy setting from the monochromator, etc. Each tomographic scan was performed by rotating the sample over a 180\({}^{\circ}\) range. The scan data consisted of 4800 projections with a step angle of 0.0375\({}^{\circ}\) with a corresponding MGD of 8 mGy. A total of 100 dark-current images and 100 flat-field images were collected before each scan to account for the nonuniformity of the imaging system. The dark-current images were collected without X-rays. Flat-field images were collected with the X-ray illumination on but without a sample within the field of view (FOV). From one set of collected data, we produced three sets of reconstruction results: the original 8 mGy MGD result (4800 projections, step angle = 0.0375\({}^{\circ}\)), a 4 mGy MGD result using half the number of projections (2400 projections, step angle = 0.075\({}^{\circ}\)), and a 2 mGy MGD result using a quarter of the projections (1200 projections, step angle = 0.15\({}^{\circ}\)). The distinctive properties of synchrotron X-ray sources, such as high coherence, energy tunability, and high brightness are very important for producing PCT scans with short exposure times. These source characteristics need to be complemented with a highly efficient detector having a high frame rate, high spatial resolution, high detection efficiency, large area, and low noise level in order to maximize the image quality at the minimal possible radiation dose to the sample. The first X-ray detector used at the IMBL for these scans included a large-area flat-panel imaging detector, the Hamamatsu complementary metal-oxide-semiconductor (CMOS) C10900D with a Cesium Iodide (CsI) scintillator deposited directly on 2D photodiode array with a pixel size of 100 \(\upmu\)m \(\times\) 100 \(\upmu\)m and a FOV of 12.16 cm (horizontal) \(\times\) 12.32 cm (vertical). A second detector was also used in this study in order to compare and contrast the optimal detector characteristics in the context of PCT for soft-tissue imaging. The second detector was a Xineos 3030 HR CMOS flat-panel detector with a medical-grade columnar CsI scintillator with a pixel size of 99 \(\upmu\)m \(\times\) 99 \(\upmu\)m.  
Both the Hamamatsu and Xineos detectors have a high quantum efficiency and very low noise level, which is important for low-dose PCT scans with short exposure times. The main specifications of both detectors can be seen in Table 1.  
Preprocessing of the projection data is required prior to 3D reconstructions. This includes the correction for the dark-current and flat-field images and geometrical correction of the images to account for the gaps between pixel rows or columns and to mask defects in the detector pixels. Both detectors had intermodular gaps or defect lines in both the horizontal and vertical directions that created intensity artefacts in the images. These intensity artefacts need to be corrected in the projection images to prevent them from producing strong ring artefacts in the reconstructed images.