In this chapter, a bench-top DCT system was constructed and used to investigate the feasibility of a stationary DCT using a spatially distributed CNT x-ray source array. One of the key questions investigated was whether the CNT x-ray source array is capable of generating sufficient x-ray output for chest tomosynthesis. Projection images in current commercial DCT systems are typically acquired at 120 kVp and 180 cm SID. The tube output (mAs) is usually derived using automatic exposure control (AEC) from a scout view image.[3, 9, 72, 27] For instance, the GE VolumRAD DCT system multiplies the mAs of the scout view by 10, and uniformly distribute the derived mAs into 60 projection views.[3] The scout view uses the standard PA CXR technique. Sabol calculated the mean technique of 1.9 mAs and 0.1 mGy incident air kerma for the standard PA CXR used in clinic from 294 adult cases.[3] Based on the method to derive tomosynthesis radiographic technique from the scout view, the mean tomosynthesis technique is 0.32 mAs and 16.7 µGy incident air kerma per projection, where the SID is 180 cm and the incident air kerma was measured 25 cm in front of the detector.[3] The incident air kerma is a more reliable reference since the tube output in mAs used also depends on the kVp, tube filtration, and SID. For the CNT source array used in this study, the incident air kerma per mAs at 100 cm was measured as 74.47 µGy/mAs at 80 kVp. The phantom images in this study were acquired at 0.6 mAs per projection. Using this technique, assuming an average-sized patient with a chest thickness of 25 cm and extending the s-DCT SID to 180 cm as current systems, the incident air kerma is scaled to 18.6 µGy at the patient entrance plane (155 cm from source), which is higher than the mean incident air kerma per projection in the tomosynthesis technique. With larger patient, the mAs per projection needs to be increased accordingly. Therefore, the x-ray output from CNT source array in this s-DCT system is comparable to current clinical tomosynthesis systems. All 75 CNT sources in the tube can be operated stably at this condition with reasonable source-to-source consistency in the cathode-gate voltage and focal spot size.
The x-ray source array was operated at 80 kVp in this study because that was the maxi- mum design specification for this specific tube, which is why this study was done at a lower energy than the energy used in commercial chest tomosynthesis systems (typically 120 kVp). However, CNT x-ray source arrays using the same technology with energies up to 160 kVp, with even higher power outputs, are now available for other imaging applications from XinRay Systems.[59] Future plans include designing and building a dedicated source array for chest tomosynthesis for higher x-ray energies, the feasibility of which has been demonstrated through the 160 kVp source array developed at XinRay Systems.
This particular model of CNT source array has an elongated focal spot size of 2.5mm× 0.5mm FWHM by design, as opposed to the isotropic shape found in most commercial DR systems (0.6 mm –1.2 mm nominal size).[1, 29, 28] The anisotropic focal spot leads to asymmetric system MTF values in this system. The shape of the focal spot is determined by the geometry of the CNT cathode and the focusing electron optics design. Smaller isotropic focal spots can be achieved using a different electron optics, as demonstrated in the CNT source array designed for breast tomosynthesis.[47] There is a tradeoff between the focal spot size and the maximum tube power, which needs to be considered when designing a dedicated source array for chest tomosynthesis. A detailed anode heat simulation is needed to determine the optimum focal spot size based on the anode heat capability.2[51] For instance, the tungsten anode can safely withstand 1.1 kW power in a 0.6mm×0.6mmeffective focal spot size during a 250 ms pulse according to simulation.[47, 51] If the tube operates at 120 kVp, the source array can be safely operated at 9 mA tube current. Given the performance of this tube, a slight reduction of the focal spot area is not expected to affect the ability to generate the x-ray output needed for chest tomosynthesis.
Due to the short length of the source array used in this feasibility study, the phantom and the detector were translated to mimic different imaging configurations. A longer x-ray source array covering the entire angular range can be manufactured with this technology. 2Anode heat load analysis and simulation will be covered in detail in Section 6.4
Preliminary results show that as the angular range of the image acquisition increases, the system in-plane resolution (system MTF) remains the same, while the in-depth resolution (FWHM of the ASF) improves greatly. These findings agree with previous studies on imaging parameters of tomosynthesis, using both stationary source array and conventional moving source commercial systems.[71, 5, 70] The system MTF, corresponding to the direction with shorter focal spot size, resulted an in-plane resolution up to 3.4 cycles/mm at 10% MTF, which is comparable to the 3.5 cycles/mm at 10% MTF in GE VolumeRAD system and 3 cycles/mm at 10% MTF in Shimadzu SONIALVISION tomosynthesis systems.[30, 29] Moreover, the flexibility in the spatial configuration of the individual CNT sources allows new tomosynthesis imaging geometries, beyond the linear scanning mode, which are mechanically challenging for conventional systems. These novel imaging geometries may improve image in- depth resolution, and reduce the image artifacts. A systematic investigation and evaluation of the effects of imaging acquisition geometry and parameters on image quality in the s-DCT system will be discussed in Chapter 7.
In this study, an anti-scatter grid was not used as is normally done in clinical settings, because the main purposes were to investigate the feasibility of s-DCT and system require- ments for CNT source arrays. Scatter reduction is an important issue in chest imaging. For the s-DCT system using linear source array, conventional focusing anti-scatter grid with matching focal length can be used to reduce the scatter signal in the images. Besides the anti-scatter grid, a novel method using primary beam sampling apparatus (PSA) has been proved to be effective.[73] Initial results of using PSA for scatter rejection of s-DCT will be covered in Chapter 9.