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Low kVp settings provide high conspicuity of contrast materials at CT but may result in high image noise, particularly in larger patients. Material decomposition at dual energy CT can differentiate renal stones by their composition, quantify tissue iron stores, improve the detection of pathologic hyperenhancement, and reduce contrast material and radiation dose compared to conventional CT. Further clinical research and technique refinement will be needed as the usage of these exiciting technologies spreads.
Changes are occurring in medical CT imaging that promise to augment the already well recognized CT strengths of speed, high resolution, excellent patient tolerance, and large territory coverage. These changes include dose reduction, noise reduction, and more speed. But perhaps more fundamentally, the ability to perform dual-energy scanning at CT (DECT) is being developed by every major CT scanner manufacturer and provides the ability of CT to differentiate between different materials at imaging. DECT improves our ability to detect contrast material and distinguishes between the high-density areas created by iodine from those created by calcium or other substances – in essence, DECT brings the insight of tissue and material decomposition to standard gray-scale CT images. The basis of DECT has been explored over the past 30 years [1-5] and now CT scanners and software are being engineered to leverage dual-energy imaging for practical clinical applications. This review article will discuss the influence of tube potential on CT images and explore the potential impact of DECT on imaging the abdomen and pelvis.
In conventional X-ray imaging, including CT, the adjustable setting that affects tissue contrast is the tube peak voltage (kVp), which determines the upper limit of the X-ray energy spectrum produced by the X-ray tube. Other settings, such as tube current (typically specified in milliAmperes, mA) and helical pitch affect image noise, but do not affect the X-ray attenuation (also called “CT number” and measured in units of Hounsfield (HU)) of the scanned object. In general, the extent of attenuation of an incident X-ray is affected by its energy and the properties of the intervening matter: its atomic composition, density, and thicknesses. At X-ray energies relevant to diagnostic imaging, the two predominant interactions between X-ray photons and matter are the photoelectric effect (complete X-ray aborption) and Compton interactions (X-ray scatter with fractional loss of X-ray energy), with a shift from a predominance of photoelectric effect to Compton interactions from the lower energy range of medical imaging (20-50 keV) to the higher-energy range (50-150 keV). X-ray attenuation is strongly related to the atomic number of the elements that make up the materials in the object being imaged and is typically higher for lower-energy than for higher-energy photons where attenuation is related to the electron density or mass of the imaged material. Any given material will have a set change in expected X-ray attenuation when imaged with one known X-ray energy spectrum compared to another. When an object with unknown material composition is imaged twice, once with one known X-ray energy spectrum and again with another known spectrum, the materials within that object can be differentiated and quantified based on the change in X-ray attenuation observed between voxels obtained with the two spectra, with reference to the expected change in X-ray attenuations of the base materials. Algorithms that quantify of the materials that make up an imaged object or voxel is called base material decomposition [3, 4, 6]. The evaluation of the attenuation differences by different materials imaged at two X-ray spectra has been exploited in planar imaging in many ways, including to quantify bone mineral density at dual-energy X-ray absorptiometry (DXA) , and to subtract bone from soft tissue, such as in chest radiography to improve parenchymal lung lesion detection.
Base material decomposition at DECT enabled improved CT bone mineral density measurements [2, 8] and liver iron quantification  compared to conventional single energy CT. While the benefits of DECT were recognized over 30 years ago, exploitation of this powerful tool was limited by the slowness of older generation CT since two separate scans were required, one at a high and one at a low kVp setting. A limitation of this method, particularly for soft tissues, is poor image registration due to movement of abdominal structures between the scan acquisitions, which usually required separate breathholds or at minimum, 2 seconds between individual slices. While repeat imaging was of minimal concern when evaluating bony structures [2, 10], the low temporal resolution constraints of early CT scanners prevented DECT utilization for imaging dynamic intravenous enhancement of organs due to rapid changes in the concentration of intravascular contrast material changes between scans. Also, the poor X-ray output of early X-ray tubes and inefficiency of early detectors prevented the use of low kVp settings since those settings resulted in low photon flux and excessive image noise.
However, engineering DECT scanners to be fast and efficient_enough to image contrast material could not be ignored because iodinated contrast material is more conspicuous when imaged at lower kVp settings and shows an approximately 80% increase in CT number at 80 kVp than at 140 kVp. Imaging at lower kVp settings has numerous potential advantages, including an substantial reduction in the radiation dose, reduction in the amount of required intravenous contrast material, and reduction in certain artifacts such as pseudoenhancement . Some reports showed that low kVp settings could emphasize contrast enhancement in certain scenarios such as CT angiography of renal donors  or CT perfusion imaging  and CT intravenous pyelograms (Fig 1) . Unfortunately, low kVp settings frequently result in a substantial increase in image noise due to the relatively low X-ray output of the source and also increased relative X-ray attenuation by the patient, such that image noise is a problem in larger patients. Furthermore, while tissue contrast generally improves with lower kVp settings, certain structures such as gallstones  and certain renal stones have been shown to have improved conspicuity at higher kVp settings, even when image noise was taken into consideration [15, 16]. Combined use of high and low kVp settings in DECT, therefore, aims to capitalize on the strengths of each setting within the same examination for medical imaging
Several design modifications of CT scanners have been implemented to make DECT data acquisition more practical, and these approaches include rotate-rotate software, dual-source scanners, sandwich detectors, rapid kVp switching, and photon counting detectors. Briefly, rotate-rotate software drives the CT gantry to fully rotate once to acquire data at the first kVp setting, then almost immediately again to acquire data at the second kVp setting, and is used when the scanner does not have built-in hardware to allow for simultaneous dual-energy acquisition. The downsides of this method are the low speed and non-simultaneous nature of data acquisition which make this solution a minimal improvement over DECT with conventional CT systems.
Dual-source hardware involves the orthogonal placement of two source-detector units into a single gantry to allow simultaneous acquisition of low- and high-kVp images. Due to mechanical constraints, one unit has a smaller field of view limited to the central portion of the gantry. Slight misregistration artifact may occur between the two energy scans as the two x-ray tubes are offset by 90 degree phase and more sophisticated photon scatter correction is required.
Rapid kVp switching involves a single source that is capable of alternating hundreds of times per second between the two kVp settings during the CT gantry rotation. Sandwich detectors involves placement of one or more layers of detectors immediately on top of each other such that the detector layers deeper in the stack receive progressively higher mean energy photons due to the absorption of lower energy photons by the layers that are closer to the source. In these two solutions, excellent image registration can be obtained with a full field of view. Furthermore, since the source/detector configuration is essentially the same for both kVp settings, material decomposition can be obtained using the projection data rather than the image data and allow for more sophisticated analysis, including extrapolation of monochromatic CT images. A potential drawback of rapid kVp switching may be limited control of tube current for the two kVp settings, which may result in mismatched noise levels between the images acquired at each setting.
The general limitation of all of these approaches to DECT is that the energy spectra generated by the source(s) are polychromatic with overlap of the low-energy spectra with the high-energy spectra. The part of the spectra that overlaps adds noise without any material differentiation. Detectors capable of rapid photon counting combined with special k-edge filters, or sources that produce monochromatic X-rays would most likely provide the highest performance but currently not feasible because of technical limitations.
DECT techniques can be adjusted to deliver similar total radiation doses to the patient as conventional CT scans. A challenge, however, is that the processing of DECT works best if the noise is similar between the high and low kVp images, a scenario which is difficult to achieve particularly for large patients where high noise at 80 kVp imaging in the abdomen is inevitable. To match the image noise at 120 to 140 kVp, the tube current of an 80 kVp scan must be substantially increased. This latter consideration results in the need to select patients and clinical scenarios that are least likely to be impacted by the limitations of 80 kVp imaging. To some extent, use of a higher low kVp threshold, such as 100 rather than 80 kVp, can be utilized in problematic scenarios. Furthermore, newer CT reconstruction algorithms which reduce image noise will likely improve the range of applications available to abdominal DECT in the future.
The intraabdominal DECT application that has received the widest initial attention is the differentiation of renal stones by their composition. Evaluation for possible renal stone disease is one of the most common indications for abdominal CT in the emergency setting, and CT is widely considered the standard of reference for the detection of urolithiasis. For patients where stones are identified and deemed unlikely to pass spontaneously, triage by stone composition could alter patient management because uric acid stones can often be treated by aggressive medical expulsive therapy while other types of stones generally require more urgent surgical consultation for possible mechanical extraction. Differentiation of stone composition plays to the strength of DECT because uric acid stones, which are composed of the small atoms carbon, oxygen, nitrogen, and hydrogen, show relatively little change in X-ray attenuation compared to almost all other stones, which generally include larger atoms including calcium, phosphorous, magnesium, or sulphur (Fig. 2). Limited in vitro and clinical studies have shown that DECT material decomposition can distinguish uric acid stones from other types of renal stones with high accuracy [17, 18]. For example, uric acid and calcium can be semi-automatically distinguished at DECT and color-coded for visual distinction (Fig. 3) .
One of the initial applications developed for material decomposition with DECT was the quantification of liver iron stores [9, 19, 20]. Excessive body iron stores may be due to a primary (idiopathic) cause related to excessive iron absorption from intestinal mucosa, or secondary due to chronic large volume blood transfusion, often associated with hematologic disorders. In these patients, excess iron deposition may be seen in the reticuloendothelial system (liver, spleen, and bone marrow) and other organs, and can lead to cirrhosis and multiorgan failure. The diagnosis of hemochromatosis by liver biopsy or serum ferritin is known to be inaccurate for determining liver or body iron stores. DECT is a method of noninvasively and accurately quantifying liver iron stores based on the fact that differences in hepatic X-ray attenuation at different kVp’s vary linearly with changes in hepatic iron concentration [9, 19], while changes in attenuation of water and normal liver parenchyma are negligible . Differences in hepatic X-ray attenuation between images obtained at two kVp settings are correlated with those of standardized cylinders of known iron content to quantify the liver iron concentration (Fig. 4). The liver iron concentration is then multiplied by the 3D volume of the liver to calculate the liver iron stores. In early studies, excellent correlation (r=0.993) was seen between biopsy and DECT determination of liver iron , which was a marked improvement over what is possible with conventional single energy CT due to the variable contribution of liver fat which may occur concurrently with iron deposition. In addition to the liver, DECT allows for quantification of iron deposition in other tissues including the myocardium .
The clinical utilization and interpretation of unenhanced DECT requires further development and validation. For example, the evaluation of adrenal masses should be approached with caution using DECT until new thresholds are set to reliably differentiate adenomas from non-adenomas. At conventional CT, a low attenuation of a homogeneous adrenal mass consisting of a greater than 10% negative voxels suggests that the mass is an adenoma with microscopic fat . However, these criteria were developed using predominantly 120 kVp settings, hence reference standards appropriate for DECT and attention to technique will be required (Fig. 5).
Conventional CT angiography is increasingly utilized for the assessment of vascular disease throughout the body and in many scenarios has become the first-line test of choice. DECT potentially augments conventional CT angiography by providing improved contrast resolution (Fig 6), potentially reducing the required dose of intravenous contrast material per examination. DECT also addresses the unfortunate drawback of conventional CT angiography of difficulty of manual subtraction of bone and mental subtraction of calcified atherosclerotic plaque from the enhanced arterial lumen. Particularly for 3 dimensional displays, calcifications may obscure the true size of the vascular lumen since both contrast and calcium are radiodense. Using DECT, owing to the substantial difference in the relative change in X-ray attenuation of calcium and iodine when imaged at 80 versus 140 kVp, calcified bone and calcified plaque can be differentiated from intravenous contrast material and automatically or semiautomatically subtracted . Nevertheless, artifacts can occur due to image noise or slight misregistration of the 80 and 140 kVp data which may cause some areas of calcium or iodine to be inappropriately categorized. Furthermore, calcium bloom and small vessels with narrowed lumens may result in partial volume averaging effects .
The value of DECT for contrast enhanced abdominal organ imaging is an exciting area of growth. The improved CT contrast resolution obtained at low kVp settings has been shown to provide higher detection rates of hypervascular lesions in phantom and clinical hepatic imaging, provided that the noise in the low kVp image is not excessive . Areas of hyperemia in bowel and pancreas clearly benefit from enhanced detection of iodinated contrast material [26, 27] (Fig (Fig7,7, ,8).8). While the improved detection of contrast is primarily derived from the low energy (80 or 100 kVp) data, DECT can capitalize on the improved anatomic detail that is generally available from the low-noise high-energy data, particularly in larger patients, to augment the high contrast resolution images obtained from the low kVp settings. Blending strategies are an area of active DECT investigation, and for example, rather than simple linear averaging of the datasets, a sigmoidal blending technique that maximizes the benefits of each dataset appears promising in early experience (Fig. 9) . In this strategy, voxels with low attenuation numbers are preferentially extracted from the high kVp dataset so as to minimize noise in fat, muscle, and homogeneous organs, while high attenuation numbers are obtained from the low kVp dataset so as to emphasize contrast enhancement. Manual blending during image generation and interpretation allows the reader to interact with the two datasets as needed for a given clinical problem.
Conversely, rather than merging the data into a single image set, the high and low kVp datasets can be post-processed to isolate the iodinated intravenous contrast signal from the soft tissue to produce two new sets of images: 1) an “iodine only” image set that provides information regarding contrast material distribution in the end-organ, and 2) a “virtual unenhanced” CT scan which could potentially be used to replace the true unenhanced CT scan and thereby reduce radiation dose and improve patient throughput. These image sets could potentially be used, for example, to obtain a single-scan assessment of the renal parenchyma that would differentiate between high-attenuation hemorrhagic cysts and enhancing renal cell carcinomas  and reveal renal stones that may be otherwise obscured by intravenous contrast material [26, 30] (Fig 10). Similarly, the virtual unenhanced and iodine-only images can confirm the presence of active contrast extravasation, such as for the complication of endoleak from aortic stentgrafts, without the need to obtain a separate conventional unenhanced image for comparison, and thereby reducing radiation dose to the patient . The iodine-only images can also emphasize areas of perfusion abnormality and improve both hyper- and hypo-vascular lesion detection in solid organs. In the thorax, iodine-only images have shown value in improving the detection pulmonary emboli by demonstrating segmental perfusion defects  or to show defects in myocardial tissue perfusion . Similarly, the distribution of inhaled xenon gas can be determined by DECT to provide an assessment of pulmonary ventilation .
Further refinement and clinical evaluation of the iodine only and virtual unenhanced images will be needed because image noise from the source DECT images results in substantially higher noise once they are post-processed. Also, virtual unenhanced images differ from conventional unenhanced CT images in that they do not reflect the X-ray spectra from a single kVp setting and so the interpretation of the X-ray attenuation in these images should be approached with caution. Using a similar example as for unenhanced DECT, it is likely that new thresholds will be required to use the virtual unenhanced images for the assessment of adrenal masses since the currently existing criteria were described for images acquired at 120 kVp, rather than at 80, 140, or blended / virtual images generated from disparate energy spectra.
The ability of DECT to extract basic information regarding iodine concentration and material composition has led to practical abdominal DECT applications that include the triage of patients with nephrolithasis based on stone composition, quantification of liver iron stores, evaluation of tissue perfusion, potential reduction in radiation dose by elimination of a separate unenhanced CT acquisition, and improved detection of pathologic hyperenhancement. As clinical utilization expands, radiologists must critically determine appropriate scenarios for using DECT for lesion detection as well as assess whether new diagnostic criteria unique to DECT should be implemented. Furthermore, to take advantage of the range of data available from DECT, improved display platforms to simplify viewing of the source and post-processed images and further understanding of post-processing artifacts will be needed. Continuing rapid technical developments will improve the acquisition and post-processing of DECT image data and clinical evaluation of this exciting new technology across a spectrum of abdominal diseases will be required.