With the advances in CT technology, the imaging techniques to evaluate the kidneys and urinary tract have progressively changed. State-of-the-art X-ray CT tube/generator (Fig. 1a, b) combinations provide a peak power of 60–120 kW, usually at various user-selectable voltages, for example, 80, 100, 120, and 140 kV. The spatial resolution of the CT images may achieve 24 lines per cm with an image matrix size of 512 × 512, 768 × 768, or 1,024 × 1,024 pixels. Modern MDCT systems use solid-state detectors (Fig. 1c, d) consisting of a radiation-sensitive solid-state material (cadmium tungstate, gadolinium oxide, or gadolinium oxysulfide with suitable dopings), which converts the absorbed X-rays into visible light. The light is then detected by a Si photodiode. The resulting electrical current is amplified and converted into a digital signal. CT detectors must provide different slice widths to optimize scan speed, longitudinal resolution, and image noise for each application.

Nowadays, the visualization of the acquired images at the workstation is supported by simple visualization interfaces and the radiologist can review the images on axial, coronal, sagittal, and multiplanar reformation (MPR), which describes the option of arbitrary plane reconstruction from volumetric datasets. These planes can be chosen in an orthogonal plane (sagittal, coronal), in an oblique plane, or even in a curved-planar reconstruction (e.g., for course of vessels). Coronal images are useful in the evaluation of the kidney because they can be depicted in one single image rather than in multiple axial sections. Although normal coronal images cannot depict the entire ureter in a single image, the evaluation of the urinary tract is, anyway, simple. Moreover, the coronal plane is suitable for reconstruction of an “in situ” perspective resembling the view of the abdominal or urologic surgeon. Maximum intensity projection (MIP) images are more indicated in the evaluation of the urinary tract, in relation with its opacification with contrast media. MIP depicts the structures of highest CT density within a volume of interest. Volume rendering technique (VRT) is an image-processing option that emphasizes regions of selected CT density range, thereby accentuating specific tissues or organs.

Image quality of any given CT system is determined by a number of scanning parameters that the user can modify: (a) collimation, (b) pitch (Fig. 2), (c) reconstruction interval, and (d) timing of scanning after contrast injection. Optimizing these parameters to achieve the best possible image quality for a specific MDCT examination requires thorough understanding of their interrelationships. Collimation is related to spatial resolution, image noise, and length of coverage. Spatial resolution may be increased by means of a decrease in the collimator width, but decreased collimator width also results in increased image noise and decreased length of coverage. The reduced coverage encountered with narrow collimation, however, can potentially be overcome with an increase in pitch.

The effective detector row thickness is simply the sum of the widths of the contributing detector rows for each channel. In multidetector-row CT scan acquisitions, the effective detector row thickness of all channels must be identical. Effective detector row thickness is an important parameter because the reconstructed section thickness cannot be smaller than the effective detector row thickness. The term detector configuration describes a given scan acquisition mode in terms of the number of z-axis data channels being used and the effective detector row thickness of each data channel (Saini 2004). The beam collimation is simply the product of the number of data channels being used and the effective detector row thickness (Saini 2004).

In single-slice helical CT, pitch (p) is defined as the table feed during a 360° rotation of the tube-detector apparatus (mm/rotation) divided by the collimator width in millimeter (collimated beam width). It shows whether data acquisition occurs with gaps (p > 1) or with overlap (p < 1) in the longitudinal direction. In MDCT the beam collimation is different from the slice collimation, and the beam collimation depends on the number of detector rows. In particular, in MDCT there are two distinct versions of pitch, depending on whether the entire beam collimation is considered or the “detector” collimation (Fig. 2) (i.e., beam/4 in 4-detector-row MDCT) is used. The selection of pitch for MDCT is affected by several factors such as the required length of coverage, reconstructed section thickness, and image noise. For a given duration of scanning, minimization of collimation (thus, an increase in pitch) will allow a narrow reconstructed section thickness and potentially improve spatial resolution, particularly in nonaxial planes. However, maximization of pitch may also result in a decrease in contrast resolution. Thus, the most appropriate choice of scanning parameters depends on the imaging problem under consideration. For example, since CT angiography (CTA) allows imaging of structures with very high attenuation against a background of much lower attenuation (i.e., high inherent contrast), the loss of contrast resolution caused by maximization of pitch can be disregarded. The smaller the reconstruction interval, the greater the longitudinal (z-axis) resolution, with a resultant loss in z-axis coverage. If multiplanar reconstructions will be required, a small reconstruction interval with overlapping sections is advantageous. With the evolution of CT from single-slice spiral CT to MDCT, the collimation has been reduced from 3–5 mm to 0.5–1 mm of the most recent 64-row MDCT scanner (0.5–0.625 mm). A decrease in pitch from 1–1.5 to 0.75–0.80 has also been achieved. As a consequence, the thickness of the reconstructed images has been decreased from 5 to 0.5 mm. The final result is the “isotropic voxel,” which provides coronal, sagittal, and oblique multiplanar reformats with the same resolution of the original axial CT sections. A further advantage of state-of-the-art MDCT is the short scan time which varies from 0.27 to 0.5 s for a single 360° rotation.

MDCT examination can be performed before and after intravenous contrast administration. Breath-hold images that are essentially free from motion artifact and respiratory misregistration can be quickly obtained during multiple phases of renal enhancement. The preliminary scan without intravenous contrast administration should be performed also without administration of hyperdense oral contrast agents (e.g., Gastrografin), because these might impair visualization of the kidneys or the ureters in MIP images. The timing of peak aortic and hepatic contrast enhancement is primarily dependent on the rate of injection. Rapid or low-volume (shorter-duration) injections produce earlier peak enhancement, whereas slow or high-volume (longer-duration) injections result in later peak enhancement. These factors must be taken into account before a fixed scanning delay is instituted. A preliminary minibolus (5 mL/s for 4 s), with scanning every 2 s beginning 10 s after the injection is started, can be used. The time to peak aortic enhancement is determined from the resultant time attenuation curve and is used to calculate the scanning delay. The scanning delay can also be accurately timed with a bolus tracking software program.

In CT, the term artifact is applied to any systematic discrepancy between the CT numbers in the reconstructed image and the true attenuation coefficients of the object.

CT scanning provides high-quality diagnostic information, but it is also usually described as a high-dose procedure (Cohnen et al. 2003). Factors that may contribute to the delivery of high radiation dose related to CT evaluation of the urinary tract include replacement of conventional radiography by CT, need for multiple follow-up CT studies, and need for multiphase CT protocols (Kalra and Singh 2008).

Absorbed dose is defined as the energy absorbed per unit mass and is measured in gray (Gy) and its subunit, the milligray (mGy). The effective dose is defined, by the International Commission on Radiological Protection (ICRP 2007), as a single dose quantity reflecting the overall risk from any radiation exposure to a reference person, where the risk is averaged over all ages and both sexes. Effective dose is calculated by summing the absorbed dose in individual organs, weighted by the radiation sensitivity. It is useful for comparing the risks to a reference patient from different imaging techniques and procedures, even though it should not be used for individual patients or detailed risk assessment. The equivalent dose, measured in sieverts (Sv), accounts for differences in the sensitivity of target organs to radiation damage and is calculated by multiplying the absorbed dose to a specific tissue with the radiation weighting factor, which is equal to 1 for X-rays.

Noise has an important influence on image quality, it is inversely related to the radiation dose, and it represents a limiting factor in the improvement of spatial resolution. CT image noise generally depends on the number of X-ray photons interacting with the detector array (quantum noise), electronic noise of the detectors, and reconstruction kernel (sharper kernels give noisier images). Noise is also dependent on the patient’s individual attenuation, convolution filters, slice thickness, pixel dimension, and radiation dose. The level of noise can be quantified easily from images by positioning a standard region of interest in an anatomical structure of known density and measuring the standard deviation of the HUs values. Reduced dose should always be weighed against diagnostic image quality.

MDCT radiation dose and image quality are also affected by helical pitch (table feed/X-ray beam collimation), and radiation dose is inversely related to the pitch value. The radiation dose decreases proportionally with increasing pitch in CT systems with a single detector row, as the tube voltage and current are kept constant, while for multislice CT scanners the relationship between pitch and radiation dose is nonlinear in ECG-gated acquisition and linear in noncardiac mode. MDCT increases dose as a result of thinner collimation, overbeaming, and overranging effects. For example, in four-row scanners, effective dose is about 30 % higher with a collimation of 1 mm than with a collimation of 2.5 mm.

In current MDCT scanners with more than four rows, which allow for two different collimations (millimeter and submillimeter), radiation dose increases only by about 10 % when the smaller collimation is chosen. In spiral technique, additional tube rotations have to be performed at the beginning and at the end of the scan range because adjacent data from both sides are necessary for image reconstruction (interpolation). Therefore, the scanned volume exceeds the reconstructed volume. Moreover, the number of additional rotations depends on pitch, cone beam correction, and scanner type. The relative effect is high with short scan lengths or when a large collimation and high pitch are combined. The overranging effect (Fig. 6a) may cause considerable increase of dose resulting from the difference between the total table feed (table feed × (scan time/rotation time)) and the scan length as planned. A further problem of MDCT is the width of the fan beam in the z-axis direction. As four or more detector chambers have to be exposed, a broader X-ray beam is used compared with the fan beam in single-slice CT. In order to avoid penumbral effects in the external portions of the detector array, the primary collimation must be made wider than necessary to expose only the detector array.

Overbeaming or “penumbra” effect (Fig. 6b) is a specific characteristic of MDCT, which reduces the portion of the beam that is captured by the detector array (beam efficiency) with a consequent increase in noise and decrease in the contrast-to-noise ratio. Even when the incident X-ray beam is collimated to the targeted detector rows, the beam is always slightly wider than the rows. For example, in MDCT, the incident X-ray beam is about 2 mm wider than the selected detector configuration (Saini 2004). Hence, the wider the incident beam, the smaller the percentage of “wasted” radiation due to overbeaming. Since unused radiation is delivered to the patient owing to overbeaming, when the beam collimation is consequently reduced, the number of rotations will need to be increased, contributing “wasted” radiation. Thus, for example, a 16 × 0.625-mm acquisition will require twice as many rotations as a 16 × 1.25-mm acquisition, and, therefore, an approximately 3 % higher radiation dose will be delivered (Saini 2004) (Fig. 6b).

Moreover, the bowtie filters are commonly employed in CT scanners to minimize radiation dose by reducing intensity variations across detector elements in the presence of patient anatomy. This filtration modifies a number of X-ray beam properties (effective energy, flux, first- and second-order statistics), making them nonuniform across the fan beam field of view.

Tube load parameters are very scanner specific, and CT dose index (CTDI) and dose–length product (DLP) are better usable entities to indicate the absorbed dose for protocol comparison and optimization. The fundamental parameter of radiation dose is the CTDI measured by a dose phantom in grays (Gy) and represents the integrated dose along the z-axis from a single rotation of the X-ray tube. CTDI can be measured in air or in a perspex phantom of 16 cm (representing head) or 32 cm (representing body) diameter. Measurements are commonly made both at the center and the periphery (1 cm from the surface). The weighted sum of central and peripheral CTDI values is known as the weighted CTDI (CTDI_w) and represents the mean dose in the x–y plane. However, the CTDI does not indicate the precise dose for any individual patient but is rather an index of the dose as measured and calculated in a phantom, and it does not take into account patient-associated parameters such as size, shape, and inhomogeneous composition (Lee et al. 2008). Furthermore, in helical scanning, the pitch must be taken into account to give the mean dose within a scanned volume, and the volume CTDI (CTDI_vol) corresponds to the weighted CTDI corrected by the pitch factor (dose index divided by pitch). The CTDI_vol can be used to express the average dose delivered to the scan volume for a specific examination. This is the mean dose within a scanned medium. Normalized weighted CTDI (_nCTDI_w) in air is 0.070 mGy/mAs. CTDI_w multiplied to a conversion factor (1/0.38) is converted to _nCTDI_w.

Another parameter used is the DLP, which is related to the total radiation exposure (mGy × cm) and gives an approximation of radiation risk, but it cannot give a direct assessment of the patient dose since it does not take specific organs into account. DLP is calculated from CTDI_w as
where CTDI_w is the effective weighted CTDI (mGy); n is tube rotations; N is the number of detectors; and h is the slice thickness (in centimeters). DLP may be calculated also from CTDI_vol multiplied by the length of the volume scanned:

Patient effective dose may be calculated by multiplying DLP for a specific coefficient representing typical adult and pediatric patients. This provides sufficiently precise estimation of effective dose for standard protocols and reference patients (Deak et al. 2008; Huda 2007; Shrimpton 1997).

Recent advances in CT technology such as automated tube current modulation, optimizing X-ray tube voltage, and the use of iterative reconstruction have allowed reduced radiation dose preserving good image quality.

The aim of automated tube current modulation, ATCM (Fig. 7), is to adjust radiation dose according to the patient’s attenuation. Dose modulation is based on the principle that decrease in body diameter (i.e., the anteroposterior diameter when compared to the lateral diameter) translates into decrease in radiation necessary to obtain a certain contrast-to-noise ratio in the resulting CT image data. Dose management technology in CT has led to the optimization of CT protocols using in-plane (xy-axes), longitudinal (z-axis), or combined (xyz-axes) modulation of the dose through dedicated automated dose modulation algorithms. It is important to realize that ATCM reduces dose in thinner patients, even though the dose may actually increase in thicker patients to compensate for the increase in image noise (Kaza et al. 2014).

Tube voltage has an exponential effect on radiation dose. At constant milliAmperage (mA), radiation dose decreases by approximately 65 % if the kV are reduced from 120 to 80 kV. The optimal kilovoltage for a CT study should be chosen based on imaging task and the habitus of the patient. Reducing kilovoltage implies an increase of the noise and this is especially evident in obese patients. To compensate for the reduced energy of the X-ray beam, the milliamperage needs to be increased. Iodinated structures have higher HU values at lower energies because the effective energy of the X-ray beam is closer to the k-edge of the iodine. Therefore, lesion characterization based on HU measured on CT scans performed at 120 kV would not be applicable (Kaza et al. 2014).

Iterative reconstructions are iterative algorithms to reconstruct 2D and 3D images. They are used to recognize the factor which contributes to noise development using a statistical model of noise and to remove the noise from the CT images improving the image quality. They provide improved image quality without increasing radiation dose or maintenance of image quality at a lower radiation dose (i.e., low kilovoltage) compared to conventional filtered back projection reconstruction (Kaza et al. 2014; Juri et al. 2013; Vardhanabhuti et al. 2013).

Sufficient hydration of the patient is essential to guarantee optimal excretion of contrast media during renal CT. This can be achieved by oral or intravenous administration of fluid volume. Nonetheless, concomitant disease, such as cardiac or renal insufficiency, has to be considered. Hydration may also minimize nephrotoxic effects of intravenous contrast media. Simultaneous oral contrast media administration may be useful; however, the application of positive oral contrast could produce interfering effects with CT arteriography or CT urography.

Table 3 shows the value of the technical parameters suggested in contrast-enhanced CT for renal parenchyma examination. After contrast material injection, different phases of contrast uptake can be discriminated. Usually, 80–120 mL (350–370 mg/I) of nonionic contrast medium is administered through an 18–20-G catheter in the antecubital vein at 4 mL/s, followed by 30–40 mL of saline solution at the same injection rate.

Contrast-enhanced CT scans are acquired in different contrast-enhancement phases: arterial phase, corticomedullary phase, nephrographic phase, and urographic phase. Their relative onset times depend on methods of contrast medium administration and patient characteristics. Not all phases have to be acquired and it is important to avoid unnecessary acquisitions to reduce the radiation dose. Imaging of kidneys in the various phases depends on sequential arrival in different parts of the renal parenchyma of blood carrying contrast media (Fig. 13). With normal renal function, iodinated intravascular contrast medium is excreted by passive glomerular filtration. The amount of a substance undergoing glomerular excretion is determined by the product of its “freely filterable” serum concentration and glomerular filtration rate (GFR). GFR is straightly related with weight, age, and gender; it decreases in systemic hypotension, renal vasoconstriction, or glomerulonephritis (decreased number of functional nephrons). Imaging phases have to be selected carefully in accordance with clinical indication. The comprehensive evaluation of renal masses by CT requires a dedicated renal CT protocol including unenhanced, arterial, corticomedullary, nephrographic, and excretory phases. As previously said, unenhanced CT scans are used to detect calcifications and allow quantification of tumoral enhancement on the postcontrast scans (Prasad et al. 2008).

The arterial CT phase is reached about 15–25 s after intravenous injection of contrast medium (for renal arteriography, a contrast bolus of 3.5 mL/s is desirable). When performing a contrast-enhanced CT of the kidney, it is critical, mainly for the arterial phase, to scan at the peak of contrast enhancement in the renal cortex. Nowadays, this is easily achieved with the “bolus tracking” technique, which has replaced the simple “test bolus” technique. The use of empirical fixed delays (i.e., 25–30 s), with small variations depending upon patient age or body habitus, should be avoided. With state-of-the-art MDCT, the temporal window to perform the acquisition of the examination is very narrow (Bae et al. 2008; Erturk et al. 2008; Kock et al. 2007; Schoellnast et al. 2005; Stacul et al. 2008). Alternatively, a test bolus scan or a bolus tracking program can be applied for optimal start of the arterial scan.

In the arterial phase, renal arterial stenosis or evaluation of stent implantation can be judged, and acute bleeding can be detected or ruled out.

The corticomedullary CT phase (about 25–70 s after i.v. contrast) shows high contrast in the renal cortex (Fig. 13a, b). During the corticomedullary phase, the contrast is located in the capillaries of the renal cortex and peritubular cells, proximal convoluted tubules, and columns of Bertin (Schreyer et al. 2002). The renal cortex becomes clearly distinguishable from the medulla because contrast material fills only the more proximal tubular structures due to greater cortical vascularity. The cortical phase gives a perfect visualization of the cortex (hyperdense) and medulla (hypodense) and, therefore, of the morphology of the kidney. Peak enhancement of the renal vessels during the early corticomedullary phase also provides information on vascular anatomy and patency. Moreover, early opacification of renal veins is also detectable. The optimal delay time for the corticomedullary phase depends on the rate of contrast injection, the amount of contrast material administered, and the patient’s cardiac output (Prasad et al. 2008).

This is the best phase to evaluate the vascularity of a mass. As most renal tumors are hypervascular, this phase is useful in the characterization of renal masses. The corticomedullary phase is also fundamental when evaluating renal arteries (e.g., in patients with hypertension of suspected renovascular origin) and suitable for the depiction of vascular renal anatomy (e.g., aneurysm, arteriovenous malformation, fistula) and the assessment of vessel patency.

The nephrographic (parenchymal) CT phase (80–180 s delay) shows the renal cortex and medulla with equal enhancement due to contrast medium equilibrium between vascular and interstitial compartments entering loops of Henle and collecting tubules. This particular phase may be divided into an early phase and a late phase (Fig. 13c, d), with the latter overlapping the excretory phase. During the nephrographic phase, the renal parenchyma enhances, homogeneously allowing the best opportunity for discrimination between the normal renal medulla and masses. This is the best phase to detect renal masses since all lesions appear hypodense compared to the renal parenchyma. It is also useful for the evaluation of the renal veins and inferior vena cava.

CT examination usually a noncontrast phase followed by the arterial and nephrographic phases. This technique is suggested for the diagnosis of renal tumors, complex renal masses, trauma, and infections. An additional delayed scan in the urographic phase or 3 h after the injection of contrast material may be added to evaluate patients with trauma or infections.

The excretory CT phase (Fig. 13e, f

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