This is the second of a series of four Expert Consensus Documents developed by the Society of Computed Body Tomography and Magnetic Resonance through its task force on dual-energy computed tomography (DECT).
Through conference calls, the writing committee agreed on the topics to be covered in the document, and the lead authors conducted literature searches and drafted the article. Over a series of meetings and conference calls, the writing committee finalized the article with the approval of the Society of Computed Body Tomography and Magnetic Resonance Board of Trustees. This white paper discusses radiation dose and iodine sensitivity in DECT.
RADIATION DOSE IN DECT
Over the past 25 years, radiation exposure from medical imaging has increased substantially, mostly from the proliferation of computed tomography (CT).1–3 Among numerous radiologic indications, CT has become the standard-of-care for evaluating trauma, and for emergency, cardiovascular, and oncologic imaging, with demand for CT accelerating from the early 1980s through the mid-2000s. With increased demand for CT, the awareness of radiation dose has also been heightened in the radiology community, medicine, and the public. The American College of Radiology (ACR) and other professional societies have encouraged the judicious use of CT imaging through campaigns such as Image Wisely (for adults) and Image Gently (for children).4,5 The ACR also published diagnostic reference levels for CT radiation dose in 2008 for all examinations performed in the United States. A reference level of 25 mGy CT dose index-volume (CTDIvol) was established for a standard CT of the abdomen in adults.6
Variations in Radiation Dose
One might expect reproducible and predictable doses from consistent protocols, but radiation doses have been shown to vary significantly from scan-to-scan, even with the same protocol.7,8 Radiation doses often vary in the same patient on the same scanner at different times. A large study presented by Larson and his colleagues at the 99th Scientific Assembly and Annual Meeting of the Radiological Society of North America in 2013 demonstrated that radiation dose can average as much as two times higher for a patient receiving a repeat CT under similar conditions. Larson's data showed a minimum-to-maximum variation ratio of radiation dose ranging from 1.56 to 2.02, citing scanner make/model, patient size, and table height contributing to the unpredictability.9 This potential variation in radiation dose must also be considered when comparing radiation doses between DECT and conventional single-energy CT (SECT) acquisitions.
The concept of performing CT scans at different photon energy levels was first introduced in the late 1970s.10–13 However, implementing these principles was not practical until recently, with robust increases in CT performance and post-processing software. The physical principles, current technology and definitions of terms for DECT are described in part 1 of this white paper series (refer to part 1).
Although acquisitions are obtained with 2 different energy spectra in DECT, radiation dose does not double. The relationship between radiation dose and kVp is quadratic, with the dose at 140 kVp more than 3 times that at 80 kVp ((140/80)2). Various techniques mitigate the radiation exposure from the dual-energy acquisition. In rapid-switching DECT (rsDECT), there is rapid alternation between 80 and 140 kVp data from a single x-ray tube, with more time assigned to the low kVp state. In dual-source DECT (dsDECT), 2 x-ray tubes are nearly perpendicular to one another, and tube current modulation is applied to reduce radiation dose, similar to SECT. Also with dsDECT, a tin filter is used in the higher-energy tube, improving spectral separation and lowering dose.
Dose can also be reduced in rsDECT and dsDECT with radiation protection strategies used in SECT, including iterative reconstruction, automatic tube current modulation, organ-based dose reduction, and adaptive dose shielding.
Over the past few years, radiation doses in DECT have decreased to become comparable or even lower than SECT levels. Several studies between 2012 and 2014 demonstrated CTDIvol values between 12.7 and 21.8 mGy for DECT scans of the abdomen,14–19 all below the ACR's 25-mGy reference level for a single-phase abdominal CT. In some studies, doses achieved are less than half of the reference standard (Table 1). In addition, in multiphase liver CT, Purysko et al20 showed that radiation doses were significantly lower with dsDECT than with SECT (26 mGy vs 38.8 mGy) (P < 0.001). The noise levels with DECT were comparable to those with SECT.
Applications currently under scrutiny for radiation exposure are cardiovascular and thoracic CT (Table 2). Research suggests that there is no increase in radiation exposure with dsDECT for suspected pulmonary embolism. Using dsDECT at 80/140 kVp, Yuan et al21 found no significant difference in dose length product (DLP) and effective dose between dsDECT and SECT pulmonary CTA. Bauer et al22 compared doses for 64-slice dsDECT at 80/140 kVp and SECT for pulmonary CTA and found a significantly lower radiation dose for dsDECT. De Zordo et al23 reported significantly lower radiation dose using dsDECT at 100/140 kVp compared with conventional SECT at 120 kVp in pulmonary CTA. Additionally, Chae et al24 reported no difference in radiation doses using dsDECT and SECT for evaluating the solitary pulmonary nodule.
For abdominal aorta endoleaks after endovascular repair, Chandarana et al25 reported lower effective doses for venous phase dsDECT (mean, 11.1 mSv) than for the single-source triphasic dataset (mean, 27.8 mSv). Thus, venous phase DECT as a standalone acquisition results in a radiation exposure reduction of approximately 61% compared with the exposure from SECT triple-phase acquisition. However, a factor limiting dose comparison studies is incomplete control for equivalent image quality.
The postprocessing flexibility of DECT provides additional options for radiation dose savings. An additional dose benefit of DECT is generating water-based virtual noncontrast (VNC) images, possibly eliminating the need for true unenhanced images with some applications. In dual-phase examinations, this may cut radiation dose by up to one-half, depending on the dose administered for the conventional unenhanced series. Eliminating the unenhanced acquisition may also significantly reduce the radiation dose in 3-phase or 4-phase CT protocols, including renal mass and many vascular CTA protocols. Takeuchi et al19 demonstrated a 52% reduction in radiation dose for their DECT split bolus CT urogram protocol by omitting the true non-contrast exam. Graser et al26,27 published 2 studies using first-generation and second-generation dsDECT systems for abdominal mass characterization. These authors showed that using virtual unenhanced CT resulted in 35% and 50% dose reduction on first-generation and second-generation dsDECT, respectively. Subsequently, Zhang et al28 reported that virtual DECT could replace true unenhanced CT in multiphase liver protocols, with a reduction of 33% for CTDIvol and 34% for DLP. Using VNC images to replace true noncontrast series implies that VNC images are of diagnostic quality. However, because of filtration and reconstruction algorithms, image quality may not be entirely comparable, even if diagnostically satisfactory.
As with conventional CT imaging, performing a dual-energy technique requires assessing risks, including radiation, and benefits. Alternatives to CT scanning may be considered when limiting the patient’s radiation exposure is desired. However, when CT is clearly indicated, such as for evaluating the effect of thromboembolism on lung perfusion or for accurately staging lung cancer, DECT should be strongly considered even if a small increase in radiation dose is incurred. Dual-energy computed tomography does not currently have a role in ultralow dose protocols, such as CT colonography examinations, although DECT may eventually allow minimal-prep CT colonography with subtraction of contrast-tagged fecal material. However, we believe that DECT should be considered in most other scenarios, because it can increase both lesion detection and characterization, and may decrease artifacts and possibly contrast media dose, while maintaining or even decreasing radiation dose.
Although there may be differences in radiation exposure between DECT and SECT, comparing absolute radiation doses is insufficient, because doses can vary considerably even when scanning the same patient on the same scanner with identical technique,9 and doses may vary among scanners. Also, there is a range of acceptable doses, with ACR reference standards providing metrics. Reference doses are based on average patients, but actual exposure varies with patient size. Doses above reference levels are in the top quartile of peer institutions. Although the comparability of image quality has not commonly been addressed, multiple studies nevertheless suggest that doses incurred from dsDECT and rsDECT usually fall within accepted reference levels ranges.14–19 Furthermore, for many protocols, DECT can even yield doses lower than SECT.15,16,19,20,22,23,29
Evidence indicates that radiation doses from DECT are reliably comparable to those of SECT, falling well below the ACR standard reference upper limit of 25 mGy for abdominal CT. Perceived differences in radiation dose should not deter a referring physician or radiologist from ordering or performing a DECT examination.
IODINE SENSITIVITY AND DOSE IN DECT
The advantages of dual-energy acquisition for improving iodine sensitivity through selective material decomposition and monoenergetic imaging30–32 are discussed in this section, including implications for CT angiography.
The sensitivity of DECT for iodine permits sometimes reducing iodine dose and also decreasing radiation exposure by replacing true unenhanced with VNC series, as discussed above. On VNC images, soft tissue attenuation structures, including visceral organs and muscle, are displayed with relatively comparable attenuation and are distinguishable from tissues with high iron content, including hematomas.33
The enhanced iodine images may be displayed separately or may be combined with gray-scale images to create iodine overlay maps displaying perfusion in brain, heart, lung, and solid abdominal organs.34–39 As described in part 1 of this DECT series (refer to part 1), virtual monoenergetic images display the relative attenuation of body tissues that would be generated by selective radiation at a wide range of keV energies, although varying with scanner types.40 There is high attenuation of iodine at lower keV levels and a decrease in blooming artifact at higher keV levels.
Lower keV ranges of 50 to 70, substantially increase contrast in iodine-enhanced structures, including blood vessels and enhancing tissues.41 In SECT, the commonly used 120 kVp corresponds approximately to 77-keV mean effective beam energy. However, with little change in image quality, DECT 70 keV monoenergetic images show a 25% greater intrinsic iodine attenuation than a 120-kVp polychromatic beam with a 77-keV effective beam energy. Therefore, a 25% reduction in iodine load with DECT using a 70-keV effective beam energy results in the same vessel and tissue contrast as a single-energy technique at 120 kVp with a full-contrast load.21,42,43
DECT Angiography Applications
Using a reduced iodine dose for non-CT angiographic applications has not been thoroughly evaluated for relatively low-contrast structures, such as liver metastases, but this approach can be applied across many CT angiographic applications, including cardiac, pulmonary, thoracoabdominal aorta, abdominal viscera, craniocervical, and extremity CT angiography.
In SECT acquisition, a decrease in beam energy from 120 to 100 kVp results in a decrease in the mean effective beam energy from 77 to 65 keV. Dropping the effective beam energy also allows reducting the contrast material load for equivalent intravascular iodine attenuation. However, image quality, in terms of signal-to-noise ratio requires resetting of tube current, pitch, and iterative reconstruction algorithms to maintain the same background noise that can be achieved with 120 kVp.44
With DECT, particularly with a dsDECT approach, the acquisition parameters can be adjusted so that the image signal-to-noise ratio is maintained. Furthermore, the dual-energy approach allows for the flexible choice of keV energy levels to optimize image contrast and contrast-to-noise ratio using keV levels between 50 and 70.41
At higher monoenergetic keV levels, the blooming artifact can be reduced for heavily calcified plaque, resulting in more accurate determination of luminal stenosis for coronary arteries and other heavily calcified vascular tributaries including the aortoiliac vessels and abdominal visceral arterial branches. More accurate estimates of residual lumen diameter are obtained with higher keV levels.45
An additional advantage of DECT monoenergetic imaging is reducing beam-hardening artifact, such as in cardiac CT to provide better evaluation of segmental myocardial perfusion46 and for renal mass evaluation.
Another method for decreasing blooming artifact of calcified plaque is selective material imaging in an iodine-minus-calcium mode.47 Images removing calcified plaque and bone are possible because of the differing K-edges of iodine and calcium of 32 and 4 keV, respectively. The iodine-minus-calcium mode can be used when performing craniocervical CT angiography at the skull base with automated bone subtraction.48 In lower extremity arteriography, calcium subtraction from plaque in the aortoiliac, femoral, popliteal, and tibioperoneal vessels can be implemented in conjunction with automated bone subtraction. For both anatomic areas, DECT for automated bone subtraction is more time efficient and as accurate as manual segmentation.
DECT Angiography Injection Technique
For CT angiography, whether using DECT or SECT, effective contrast agent administration is crucial. Iodine delivery rates sufficient to increase arterial attenuation to 300 to 400 HU optimizes discriminating among vessel lumen, calcified plaque and surrounding tissues.49,50
Key factors for contrast agent administration are iodine concentration, injection rate, and total volume resulting in an adequate arterial circulating contrast agent concentration. With a dual-energy acquisition and monoenergetic display at low keV levels, the contrast agent concentration can be substantially reduced compared with standard techniques, reducing the required iodine load by at least 50%. Therefore, DECT provides a potential benefit with compromised renal function reducing the iodine dose by up to 50%, while maintaining adequate density in the pulmonary arteries on lower keV images. The ability to lower contrast dose while maintaining diagnostic confidence is particularly important in patients with suspected pulmonary embolism because many of them have comorbidities, including renal impairment, and may be at risk for contrast-induced nephropathy. The scope of this problem was reported in the Prospective Investigation of Pulmonary Embolism Diagnosis II trial that established the effectiveness of CTA. In that study, 18.5% of patients were rejected from the trial because of impaired renal function.51
Furthermore, a “dual-display” approach can be implemented, using low keV levels (50 keV) for vascular imaging and higher keV levels (70–77 keV) with lower background noise levels for tissue display.52
Acquisition timing to capture this circulating contrast material bolus within the arterial anatomy-of-interest can be achieved with either bolus-tracking software or preliminary mini-bolus contrast agent injection. The contrast agent becomes dispersed in the central blood volume with a plateau of circulating contrast agent occurring after an initial rise to peak. The slope of this rise depends upon cardiac output, contrast agent concentration, injection rate and injection site. Rapid acquisition can be optimized either with DECT or SECT and tailored contrast material injection and acquisition timing protocols, using relatively short contrast agent boluses.
In summary, the major advantages of DECT angiography are improved iodine sensitivity using monoenergetic reconstruction providing, at the lower end of the keV range, either improved image contrast or lower contrast load and at the higher end of the keV range, decreased blooming and beam hardening artifact. Both ends of the contrast scale can be utilized in each examination. An iodine-minus-calcium mode can be used to remove both arterial calcium and adjacent bone, more applicable to carotid cerebral and extremity CTA. The advantages of DECT angiography will be best exploited using contrast injection and timing techniques that optimize arterial enhancement at the appropriately timed acquisition interval.
Geoffrey Rubin, MD, for conceiving of and supporting this DECT white paper series. [email protected].
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