As stated by Westmore (1997), the variance in the reconstructed value of the differential linear coherent scattering. The beef muscle was positioned so that the axis of the X-ray beam was perpendicular to the muscle grain. A visual representation of the source energy distribution and the surface density of the photons is illustrated in Figure II.2.
A standard scatterplot of the sorted (ordered) laboratory and simulated distributions is shown in part (c) of Figures II.5 to II.7, respectively. An examination of Figure II.4(a) shows that the muscle scattering signal is drawn in the direction of the fat scattering peak for the laboratory samples. Laboratory data plotted as dots with error bars representing one standard deviation of the measured data.
It could also be due to noise in the laboratory data, as the simulated profile appears to fall within two standard deviations of the laboratory measurements. In the laboratory data set, Figure II.4(a), the water profile remains below the fat profile at all scattering angles. The axes of the muscle and air cylinders are parallel to the axis of the fat cylinder.
The experiments in this chapter continued using the CSCT discrete fan beam acquisition geometry developed in the previous chapter.
CTDI 100
Like the contrast measurements, all the CNR measurements (table IV.12) were less than the tungsten and mono-energy spectra CSCT measurements. As can be seen in figure IV.15, the ×10 images contained less noise than the ×1 images, and therefore higher CNR. Due to the low penetration ability of the photons in the Mo spectrum, the CSCT images contained more noise than the previous tungsten and monoenergetic spectra produced CSCT images.
In figure IV.16, a plot of the data in table IV.16 revealed a u-shaped dose deposition profile as beam energy increased. The drop occurs at the effective energy, 60 keV, which appears to coincide with the same energy as the dose trough in figure IV.16. A look back at the keV simulation measurements (table IV.14), which used only 65 sample angles and 1.3×1011 photons, showed that the difference between the maximum and minimum dose was less than a factor of two, but for the Mo 32 kVp doses (table IV.18) the maximum to minimum dose distribution was about four for about the same number of photons used.
The results in Table IV.18 showed a high energy deposition on the phantom surface due to low photon attenuation, as seen in Figure IV.18. The image produced by Mo 32 kVp has a cupping artifact due to the attenuation of lower keV photons at the surface of the phantom. CTDI100 dose measurements were taken on a Philips Brilliance 16 CT scanner using a 10 cm diameter acrylic phantom and are included in Table IV.19 for a 7.5 mm slice thickness.
The associated simulated CTDI100 measurements for the 10 cm acrylic phantom are contained in Table IV.20. In simulations, a certain number of photons can be selected regardless of the energy or thickness of the specified part. The equivalent mAs values for the simulated CSCT geometry and photon source parameters are contained in Table IV.20.
A comparison between the simulated CTDI100, normalized by the equivalent mAs from Table IV.20, and the clinical CTDI100 data from Table IV.19 revealed that the simulated measurements and clinical measurements differed by approximately the same percentage between kVp settings. Any number of images in Figure IV.13 can be processed and combined to create additional images. For example, Figure IV.19 shows a simple manipulation of the fat contrast in the 80 kVp image set: subtracting the 7.4◦ image from the 3.0◦ image times the ratio of the water areas of the two images gives a fat image with high contrast. has been produced.
CTDI
R is the ratio of the average water pixel values in the 3◦ image to average water pixel values in the 7.4◦ image. As the diameter of the acrylic phantom increases, the difference between the periphery increases and the CTDIperiphery measurement is weighted more heavily. Using simple vector geometry, the trajectories of the simulated scattered photons were reversed and extended from their point of detection through the fan-beam plane.
Any intersection that fell outside the support area where the phantom was located was eventually rejected from the final reconstruction. Then, the contours of the image are used to estimate the attenuation path length from each location in the object to its boundaries. Circular ROIs with a diameter of 14 pixels were used to sample the reconstructed SSCSCT images in each of the material regions to calculate the image contrast, CNR, and SNR—similar to the procedure used in Chapter IV.
In emission computed tomography, the linear attenuation coefficient of water chosen for use in the ACF is typically chosen based on the principal energy of the emission photons. However, it was necessary to use attenuation coefficients that were 2.3 times higher to correct for the lower signal from the central region of the phantom. Another correction must be applied to account for the attenuation of source photons to the central regions, which contributed to lower scattering interactions in the central part of the phantom.
Thus, the reconstruction method appears to require two attenuation correction factors to account for signal loss from the central region of the phantom. The following is a brief summary of the research that has progressed in the field of coherence computed tomography. Their calculations showed that coherently scattered photons dominate the scattering in the vicinity of the emitted radiation beam.
The importance of the shape factor and how it modifies the Thomson differential scattering cross section to determine the Rayleigh (coherent) differential scattering cross section for an atom/molecule is discussed. Since the incoming and outgoing EM waves have the same frequency, no energy is lost and the interaction can be described as an elastic interaction. The shape factor approximation satisfactorily describes the differential scattering cross section for X-ray photons with energies greater than the ionization thresholds of atomic shells [19].
References [21] and [43] contain a good description of Compton scattering and the details of the Klein-Hishina form factor. As the phantom radius decreases (ie, 16 cm or lower), the radial dose distribution tends to become more uniform[20].
