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We previously introduced four fiducial marker-based strategies to compensate for involuntary knee-joint motion during weight-bearing C-arm CT scanning of the lower body. 2D methods showed significant reduction of motion- related artifacts, but 3D methods worked best.
However, previous methods led to increased examination times and patient discomfort caused by the marker attachment process. Moreover, sub-optimal marker placement may lead to decreased marker detectability and therefore unstable motion estimates. In order to reduce overall patient discomfort, we developed a new image-based 2D projection shifting method.
A C-arm cone-beam CT system was used to acquire projection images of five healthy volunteers at various flexion angles. Projection matrices for the horizontal scanning trajectory were calibrated using the Siemens standard PDS-2 phantom. The initial reconstruction was forward projected using maximum-intensity projections (MIP), yielding an estimate of a static scan. This estimate was then used to obtain the 2D projection shifts via registration.
For the scan with the most motion, the proposed method reproduced the marker-based results with a mean error of 2.90 mm +/- 1.43 mm (compared to a mean error of 4.10 mm +/- 3.03 mm in the uncorrected case). Bone contour surrounding modeling clay layer was improved. The proposed method is a first step towards automatic image-based, marker-free motion-compensation.
We used a Monte Carlo (MC) model based on Geant4 (GEometry ANd Tracking) to generate dose profiles in the central plane of the CTDI phantom. MC simulations were carried out for three different sizes of z-collimator and different tube voltages (80, 100, or 120 kVp), a tube current of 80 mA, and an exposure time of 25 ms.
We derived optimal weighting coefficients by taking the integral of the radial dose profiles. The first-order linear equation and the quadratic equation were used to fit the dose profiles along the radial direction perpendicular to the central plane, and the fitted profiles were revolved about the Z-axis to compute the mean dose (i.e., total volume under the fitted profiles/the central plane area). The integral computed using the linear equation resulted in the same equation as conventional CTDIW, and the integral computed using the quadratic equation resulted in a new CTDIW (CTDIMW) that incorporates different weightings ("2/3 and 1/3") and the middle dose point instead of the central dose point.
Compared to the results of MC simulations, our new CTDIMW showed less error than the previous CTDIW methods by successfully incorporating the curvature of the dose profiles regardless of acquisition protocols. Our new CTDIMW will also be applicable to the AAPM-ICRU phantom, which has a middle dose point.
The requirements for patient imaging are low patient dose, fast imaging time, and high image quality. For GBI, these requirements can be met most successfully with a narrow energy width, high- ux spectrum. Additionally, to penetrate a human-sized object, the design energy of the system has to be well above 40 keV. To our knowledge, little research has been done so far to investigate optimal GBI filtration at such high x-ray energies.
In this paper, we study different filtration strategies and their impact on high-energy GBI. Specifically, we compare copper filtration at low peak voltage with equal-absorption, equal-imaging time K-edge filtration of spectra with higher peak voltage under clinically realistic boundary conditions. We specifically focus on a design energy of 59 keV and investigate combinations of tube current, peak voltage, and filtration that lead to equal patient absorption. Theoretical considerations suggest that the K edge of tantalum might provide a transmission pocket at around 59 keV, yielding a well-shaped spectrum. Although one can observe a slight visibility benefit when using tungsten or tantalum filtration, experimental results indicate that visibility benefits most from a low x-ray tube peak voltage.
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