A new dual-drum CT system architecture has been recently introduced with the potential to achieve significantly higher temporal resolution than is currently possible in medical imaging CT. The concept relies only on known technologies; in particular rotation speeds several times higher than what is possible today could be achieved leveraging typical x-ray tube designs and capabilities. However, the architecture lends itself to the development of a new arrangement of x-ray sources in a toroidal vacuum envelope containing a rotating cathode ring and a (optionally rotating) shared anode ring to potentially obtain increased individual beam power as well as increase total exposure per rotation. The new x-ray source sub-system design builds on previously described concepts and could make the provision of multiple conventional high-power cathodes in a CT system practical by distributing the anode target between the cathodes. In particular, relying on known magnetic-levitation technologies, it is in principle possible to more than double the relative speed of the electron-beam with respect to the target, thus potentially leading to significant individual beam power increases as compared to today’s state-of-the-art. In one embodiment, the proposed design can be naturally leveraged by the dual-drum CT concept previously described to alleviate the problem of arranging a number of conventional rotating anode-stem x-ray tubes and power conditioners on the limited space of a CT gantry. In another embodiment, a system with three cathodes is suggested leveraging the architecture previously proposed by Franke.
A new scalable CT system architecture is introduced with the potential to achieve much higher temporal resolution than is possible with current CT designs while maintaining the flux per rotation near today’s levels. Higher effective rotation speeds can be achieved leveraging today’s x-ray tube designs and capabilities. The new CT architecture comprises the following elements: (1) decoupling of the source rotation from the detector rotation through the provision of two independent, coaxial and coplanar rotating gantries (drums); (2) observation of a source at a range of azimuthal angles with respect to a given detector cell; (3) utilization of a multiplicity of x-ray sources; (4) use of a wide-angle iso-centered detector mounted on the independent detector drum; (5) the detector drum presents a wide angular aperture allowing x-rays from the various sources to pass through, with the active detector cells occupying about 240-degrees in one configuration, and the wide aperture the complementary 120-degrees; (6) anti-scatter grids with absorbing lamellas oriented substantially parallel to the main gantry plane; (7) optional sparse view acquisition in “bunches,” a unique sparse sampling pattern potentially enabling further data acquisition speed-up for specific applications. Temporal resolution gains are achieved when multiple sources are simultaneously in view of the extended detector. Accurate data acquisition then relies on multiplexing in space, time, or spectra. Thus the use of an energy-discriminating detector, such as a photon-counting detector, and of tube pulsing will be advantageous. Volume-based scatter correction methods have the potential to apply when space multiplexing is used.
KEYWORDS: Sensors, X-rays, Signal detection, Signal to noise ratio, Imaging systems, Chest, Charge-coupled devices, Radiography, X-ray detectors, Lung cancer
This paper describes a system for multi-spectral single-scan lung imaging. The proposed approach relies on a low noise detector sampled at a high rate. The proposed method overcomes limitations of CCD time-and-delay integration slot-scanning systems. The system design and preliminary specifications are described. The results of initial spectral and system simulations in support of system feasibility per the outlined specifications are described. Initial investigations support the potential of the proposed approach to alleviate four shortcomings of the current digital flat-panel approach to chest radiography: (i) by enabling dynamic multi-spectral imaging in a single scan, the approach reduces the time delay between exposures, thus reducing sensitivity to motion; (ii) the approach enables dynamic technique feedback and technique adaptation, eliminating the need for a pre-exposures and reducing the likelihood of poor x-ray techniques in local image areas; (iii) by enabling direct measurement of the scatter field, the proposed method allows further scatter correction resulting in image quality improvements; (iv) finally, full-frame sampling of a digital detector allows imaging of the beam penumbra, thereby reclaiming the detection quantum efficiency loss due to over-collimation in current TDI slot-scanning approach; the resulting DQE potentially exceeds that of flat-panel detectors by a factor up to two.
The FDA has approved the SenoScan slot-scanning Full-field Digital Mammography system. A high power Tungsten-target x- ray tube enables breast imaging with 0.22 s effective exposure time. A 21-cm X 29-cm area is scanned in less than 6 seconds, at a typical clinical technique of 30 kVp, 170 mA. The detector comprises a Thalium-doped Cesium Iodide scintillator coupled to a combination of four CCDs abutted along their narrow dimension to from a 10-mm by 210-mm slot. With the CCDs operated in time-delay-and-integration mode along the narrow dimension, the system functions in a continuous scanning mode. The MTF in the standard and high- resolution modes extend to 10-cycles/mm and beyond 14 cycles/mm respectively. The Detective Quantum Efficiency curve starts at 50 percent at DC and extends to 10 cycles/mm in Standard model. Accordingly the SenoScan system enables screening and diagnostic breast imaging with a limiting resolution approaching that of film-based systems. The overall system design and intrinsic scatter rejection efficiency directly translate in high DQE characteristics that enable screening at a significantly reduced patient dose.
This paper introduces a new multislice helical weighting method. When used in conjunction with rebining from fan-beam to parallel projections, this approach leads to a single- filtering algorithm, that is, each projection needs to be filtered only once independently of the number of image reconstructions performed. This is particularly useful when overlapping images are reconstructed to fully leverage the z- resolution available from the projection data, or when performing CT fluoroscopy, where up to twelve or more images are reconstructed per 360-degree source rotation. A variable helical contribution width method is presented, which allows helical reconstruction for any selected pitch. The algorithm differs significantly from the current published helical weighting methods. In particular it can be applied at any helical pitch in a relevant continuum and to a system with any number of rows (in as much as cone-angles may be ignored). The interpolation method is chosen in such a way that the number of projections contributing to an image is adjustable. This results in user-selectable image thickness as well as user- selectable pitch. Further the minimum number of projections to be used is significantly smaller than that used with the current algorithms.
This paper presents a calibration and correction method for detector cell gain variations. To provide variable slice thickness capability in multislice volumetric scanners, while minimizing costs, it is necessary to combine the signals from several detector cells. The process of combining the output of several detector cells with non-uniform gains can introduce numerical errors when the impinging x-ray signal varies over the range of the combined cells. These scan dependent numerical errors can lead to artifacts in the reconstructed images, particularly when the numerical errors vary from channel-to-channel. A projection data correction algorithm has been developed to subtract the associated numerical errors. For effectiveness and data flow reasons, the algorithm works on a slice-by-slice basis. An initial error vector is calculated by applying a high-pass filter to the projection data. The essence of the algorithm is to correlate that initial error vector, with a calibration vector obtained by applying the same high-pass filter to various z-combinations of the cell gains. The solution to the least-square problem gives the coefficients of a polynomial expansion of the signal z-slope and curvature. From this information, and given the cell gains, the final error vector is calculated and subtracted from the projection data.
This paper presents a modification of the z-slope algorithm introduced in the accompanying paper 'An algorithm to correct for z-detector gain variations in multislice volumetric CT.' The x-expansion coefficients found by SVD are plugged into the equation describing the final estimate of the error vector. It is then found that the algorithm can be expressed in vector form as a matrix equation relating the final error estimate to the initial error estimate, where the final error vector estimate e is subtracted from the projection data for correction. In this form, the matrix coefficients, but not the size, are dependent on the correction order. Also, the final correction matrix MXF (size Nchan by N, where Nchan is the number of channels corrected) is calculated channel-by- channel, by evaluating MX as above for a segment of N channels (with N odd), and extracting the central row to define the corresponding row of MXF. The segment of N channels is then slid by one channel, and the process is iterated till completion of MXF. Although the number of calculations required to obtain the correction matrix is greatly increased, as MXF depends only on the detector cell gains, these calculations can be performed off-line in calibrations. The scan data-dependent calculations are greatly streamlined.
In this paper, a straightforward method of estimating the CT projections is applied to simplified pre-processing, simplified reconstruction filtering, and to low-dose and local CT image reconstruction. The method relies on the projection- to-projection data redundancy that is shown to exist in CT. In the pre-processing application, the output of a few, angularly sparse fully pre-processed projections, is utilized in a linearization model to estimate directly the output of pre- processing for all the other projections. In the reconstruction filtering application, and with projection i and k being fully filtered, intermediate projection j low frequency components are estimated by a linear combination of projections i and k. That estimate is then subtracted from projection j, and the resulting high-frequency components are then filtered without zeropadding. By linearity the same combination of fully filtered projections i and k is added back to projection j. A factor two simplification is obtained, that can be leveraged for reconstruction speed or cost reduction. The local reconstruction application builds on the filtering method, by showing that truncated data is sufficient for calculating a filtered projection high-frequencies, while a very simple projection completion model is shown to be effective in estimating the low frequencies. Image quality comparisons are described.
CT 2(pi) helical weighting algorithms do not lend themselves to fast reconstruction: the weight distributions present a line of discontinuity across the sinogram which defines two separate regions and associated weight expressions. Accordingly, reconstruction of P image planes requires P weightings and filterings of all projections. This paper shows that, by generalizing the concept of the interpolation/extrapolation function to that of distance function, and by selecting particular classes of such functions, the sinogram discontinuity can be eliminated. By imposing specific necessary conditions, single analytical expressions across the entire 2(pi) sinogram are obtained. Decomposition of these particular 'single' functions leads to exact or approximate fast two-filtering algorithms, for which a given projection needs to be filtered only two times for an arbitrary number P of reconstruction planes. Further, another generalization of the concept of helical weighting leads to 'single' weighting functions that depend only on the sum of the projection- and fan-angles. Accordingly, after rebinning the fan-beam projections to parallel projections, weighting commutes with filtering, and reconstruction of an arbitrary number P of image planes requires only one filtering per projection.
A detection-estintation method is introduced for use in digital subtraction angiography (DSA) . The method allows three dimensional (3-D) reconstruction of the vasculature by an image intensifier-based volume irnager using few subtracted projections. In the first step several 2-D projections are processed by vascular segmentation algorithms to define binary vasculature envelopes. In the second step a " logical" backprojection algorithm defines the 3-D binary envelope of the vasculature. The gray levels of the voxels within the 3-D envelope are then estimated by a constrained algebraic reconstruction technique (ART) to refine the vessel boundaries. Results of the reconstruction of an aluminum structure specially designed to test vascular reconstruction algorithms are presented. 1.
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