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Dual-energy CT and its potential use for quantitative myocardial CT perfusion
Journal of Cardiovascular Computed Tomography (2012) 6, 308–317
Technical Notes
Dual-energy CT and its potential use for quantitative myocardial CT perfusion Aaron So, PhDa,b,*, Jiang Hsieh, PhDc, Suresh Narayanan, PhDc, Jean-Baptiste Thibault, PhDc, Yasuhiro Imai, PhDc, Sandeep Dutta, PhDc, Jonathon Leipsic, MDd, James Min, MDe,f, Troy LaBounty, MDe,f, Ting-Yim Lee, PhDa,b a
Imaging Research Laboratories, Robarts Research Institute, 100 Perth Drive London, ON, Canada N6A 5K8; bImaging Program, Lawson Health Research Institute, London, ON, Canada; cCT Engineering, GE Healthcare, Waukesha, WI, USA; dDepartment of Radiology and Medicine, University of British Columbia, Vancouver, BC, Canada; eDepartment of Medicine, Heart Institute, Cedars-Sinai Heart Institute, Los Angeles, CA, USA and fDepartment of Imaging, Heart Institute, Cedars-Sinai Heart Institute, Los Angeles, CA, USA KEYWORDS: Dual-energy CT; Beam hardening correction; Myocardial CT perfusion; Quantitative imaging; Rapid tube potential switching
Abstract. Application of quantitative myocardial CT perfusion (CTP) for the assessment of coronary artery disease may have a significant effect on patient care as the functional significance of a coronary stenosis can be evaluated through absolute measurement of the downstream myocardial perfusion (MP) both at rest and under exercise or pharmacologic stress. A main challenge of myocardial CTP is beam hardening (BH), arising from the polychromatic nature of x-rays used in CT scanning and the presence of highly attenuating contrast agent in the heart chambers during the CT acquisition. The BH effect induces significant nonuniform shifts in CT numbers which, if uncorrected, can lead to inaccurate assessment of MP. With the recent developments of dual-energy CT (DECT) scanning on clinical scanners, the BH effect on MP measurement could be reduced with the generation of monochromatic images relatively free of BH artifacts from the acquired dual-energy data. Here, we review the different techniques of acquiring dual-energy scans and generating monochromatic images, followed by discussion on the progress of developing a DECT technique with reduced radiation dose for quantitative myocardial CTP. Ó 2012 Society of Cardiovascular Computed Tomography. All rights reserved.
Conflict of interest: A. So has no disclosure; J. Hsieh, S. Narayanan, J.-P. Thibault, Y. Imai, and S. Dutta are employees of GE Healthcare; J. Leipsic is on the speakers, bureau and advisory board of GE Healthcare; J. Min receives research grant from and is on the medical advisory board and speakers, bureau of GE Healthcare; T. LaBounty receives research grant from GE Healthcare; T.-Y. Lee has a licensing agreement with GE Healthcare on the CT Perfusion software and received a grant from and is a consultant to GE Healthcare. * Corresponding author. E-mail address: [email protected] Submitted February 10, 2012. Accepted for publication July 30, 2012.
Introduction Since its introduction, dual-energy CT (DECT) has been a source of a great deal of research and interest in clinical cardiac CT.1–4 In conventional CT, tomographic images of the scanned subject depict differences in x-ray attenuation between materials or tissues within the subject. Although the more attenuating materials (eg, bone) can be differentiated from the less attenuating
1934-5925/$ - see front matter Ó 2012 Society of Cardiovascular Computed Tomography. All rights reserved. http://dx.doi.org/10.1016/j.jcct.2012.07.002
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Dual-energy CT for myocardial CT perfusion
ones (eg, water and soft tissues), spectral information of the polychromatic x-ray photons is not used in the image reconstruction process. X-ray attenuation by a material in the diagnostic energy range of CT is mainly contributed by photoelectric absorption and Compton scattering.5,6 These interactions depend not only on the x-ray photon energies but also on the atomic number (absorption) and electron density (scattering) of the material. As such, a CT system capable of resolving energies of individual x-ray photons would allow a better tissue characterization than the conventional CT system that only takes into account the integrated x-ray signals over the entire energy spectrum. Another merit of such an energy resolving CT (eg, photon counting CT [PCCT]) system is that images correspond to a specific x-ray energy can be generated.5 The beam hardening (BH) effect arising from the polyenergetic nature of the x-ray photons7 is therefore reduced, allowing a more accurate measurement of contrast concentration and hence myocardial perfusion (MP) when using dynamic contrast-enhanced (DCE) CT scanning.8–11 Although energy-resolving CT is promising, PCCT scanners remain experimental and not yet commercially available because there are technical difficulties that require resolution. One challenge is the high photon flux required in clinical scanning that exceeds the count rate capability of the current photon counting detectors.12–14 It is widely expected that PCCT remains some years away from being clinically available. By contrast, the recently developed clinical DECT scanners are good alternatives to the ideal PCCT system.15,16 These scanners scan objects with the use of x-rays with 2 different energy spectra generated by 2 distinct tube potentials (expressed in kilovolts or kV). Any material in a voxel can then be represented by a mixture of a pair of basis materials whose attenuation properties are known at the different energy levels.5 By convention, water and iodine are selected as the basis materials because of their vast differences in attenuation properties and because of their prevalence in contrast-enhanced studies. Monochromatic images can also be generated by a linear combination of the water and iodine equivalent density images.17,18 In contrast, monochromatic images with energy-resolved CT may be generated directly and do not require an additional step of estimation of material density images. As such, monochromatic images from energy-resolved CT can be more accurate, at least theoretically, than monochromatic images from DECT. Here, we first review different techniques available for dual-energy (two tube potentials) acquisitions and generating monochromatic images from dual-energy projections. We then review current progress to date of 1 DECT technique for quantitative myocardial CT perfusion (CTP) and efforts of reducing radiation dose associated with that dual-energy CTP protocol.
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Technical methods DECT scanners DECT acquisition can be classified into 2 categories: source-oriented and detector-oriented. The source-oriented approach relies on the x-ray source to produce x-rays with 2 different energy spectra for scanning. This can be achieved by using either 2 independent x-ray tubes or a single tube capable of switching between the low and high tube potential rapidly. The detector-oriented approach relies on the ability of the detector to differentiate low-energy from high-energy x-ray photons in a single x-ray beam. Each acquisition approach is discussed in more detail. Dual-source One source-oriented approach for DECT acquisition is to implement a dual-source system, for example the Siemens’ Somatom Definition Flash (Siemens Healthcare, Henkestrasse, Erlangen, Germany) in which 2 x-ray tubedetector pairs are mounted onto the same rotation gantry at an angular offset of 90 degrees (94 degrees for the secondgeneration scanner), with 1 tube operating at 80 or 100 kV and the other operating at 140 kV (Fig. 1A),15,19 Because of the limited space of the gantry, only 1 detector can cover the full acquisition field of view available (AFOV; 50 cm), and the other detector is restricted to a smaller AFOV (26 cm and 33 cm for first- and second-generation scanners, respectively). Data truncation in projections as a result may lead to image artifacts.20 Another potential problem is the increased scattered radiation from 2 instead of 1 x-ray tube turning on at the same time, which could lead to degraded contrast-to-noise ratio and shading artifacts in reconstructed images.21 These errors can affect the measurements of time-density curves (TDCs) and MP, but they can be corrected with data extrapolation20 and cross-scatter correction22 algorithms. Furthermore, material decomposition cannot be performed in the projection domain because of the mismatch in projection views between the high and low tube potential projection sets when the scanned object is moving (eg, a beating heart) or in helical (spiral) scanning as in coronary CT angiography (CTA).23 This could potentially lead to suboptimal correction of BH artifact in the monochromatic images.16,23 Single-source with rapid tube potential switching between projection views The other source-oriented approach is to use a single x-ray tube capable of rapidly switch between the low (80) and high (140 kV) tube potentials,24,25 as in the case of the GE Healthcare Discovery CT750 HD scanner (GE Healthcare, Waukesha, WI, USA; Fig. 1B).16 Because the tube potential switching occurs as rapidly as every 0.2 milliseconds
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Figure 1 Four technologic solutions for dual-energy CT myocardial CT perfusion. (A) Two source-detector pairs with each source operating at a different tube potential and covering a different scan field of view (SFOV). (B) Single source-detector pair with the source capable of rapidly alternate between 2 different tube potentials (in kV) in a single gantry rotation. (C) Single source-detector pair with the detector made of 2 scintillating materials with different x-ray photon-stopping powers. (D) Single source-detector pair with the source capable of rapid tube potential switching between sequential gantry rotations.
(0.0002 second), each pair of 80 and 140 kV projections are essentially acquired from the same view angle. To avoid spectral contamination between consecutive 80 and 140 kV projections, the scanner uses a scintillating material (gemstone) that has an ultrafast primary decay time (0.03 msecond) and low afterglow (delayed fluorescence). One merit of this CT system is that, because there is minimal view angle mismatch between the successive high and low tube potential projections, material decomposition can be performed in the projection space which would lead to a more exact BH correction, as found in previous phantom experiments.16 A caveat with this CT system is that the tube current cannot be modulated at the same speed as the tube potential. However, an equivalent way to independently optimizing tube current for the high and low tube potential projections is to modulate the duration of each tube potential. With the current implementation the fluence for the low and high tube potential projections are selected to optimize image quality with no dose penalty in coronary CTA studies. Single-source with dual-layer detector Compared with the CT systems discussed earlier, a single-source dual-layer detector system (Philips’ prototype) more closely resembles the ideal energy-resolving
photon counting system. In this system, the x-ray detector consists of 2 different scintillating materials bonded together (Fig. 1C).26 In the direction of transmitted x-rays, the low stopping power material is placed above the high stopping power material. This design would allow higher energy xray photons to pass through the top layer without suffering significant interaction, whereas the lower energy photons are mostly attenuated in the top layer. The 2 signals, one from the top and one from the bottom of this sandwich detector, would correspond to x-rays in 2 different energy ranges. The main advantage of this design is that at each view the high- and low-energy projections are exactly registered with respect to each other. However, one potential problem of this system is that the high- and low-energy projections are spectrally less separated from each other compared with the rapid tube potential switching or dual-source system,27 which could lead to suboptimal material decomposition and BH correction for quantitative MP imaging. Single-source with rapid tube potential switching between gantry rotations Another approach of single-source with fast tube potential switching is to alter the tube potential level between gantry rotations instead of projection views in a single gantry
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Figure 2 Average maps acquired from the dual energy CT Snapshot Pulse 640 mA protocol at (A) 140 kV and (B) 70 keV. Beam hardening induced hypoenhancement in the apical myocardium (arrow) in the 140 kV average map was minimized in the corresponding 70 keV average map.
rotation (Fig. 1D). In this setup, the x-ray tube potential is first set to the high level (or low level) to complete the first gantry rotation (scan). The tube potential is then quickly switched to the low kilovolt setting (or vice versa) for the subsequent scan. The Toshiba’s Aquilion ONE (Toshiba America Medical Systems, Tustin, CA, USA) uses this approach for DECT acquisition, with 80 and 135 kV alternates every 1000 milliseconds (1 second) or less. With this approach, material decomposition can be performed in the projection domain because the 2 projection sets are aligned perfectly. However, a potential challenge to this approach is the patient motion occurring between the 2 scans.
Projection-based versus image-based DECT imaging Two methods are available from which monochromatic images can be synthesized from projection sets acquired with 2 different energy spectra: projection-based and image-based. In the projection-based approach,5 the alternating 80- and 140 kV projections acquired from rapidly switching tube potential scanning first undergo data conditioning and calibration. The 2 corrected projection sets are then mathematically transformed into the density projections of 2 selected basis materials (ie, water and iodine), which are then reconstructed into the corresponding water and iodine equivalent density images free of BH effects. Monochromatic images are then generated from a linear combination of the water and iodine images also free of BH effects. In the image-based approach,6 the conventional 80 and 140 kV single-energy CT (SECT) images with residual BH effects are reconstructed first similar to the case of SECT; monochromatic images and water/iodine equivalent density images are then generated by linear combinations of the 80 and 140 kV images with proper weightings of each image. Compared with the projection-based
technique, the image-based approach is easier to implement because it does not require access to the x-ray projections (raw data). However, the image-based method could be less accurate than the projection-based method in measuring the attenuation (density) of materials because BH is already present in the low and high tube potential images which would affect the accuracy of the derived equivalent density images and monochromatic images, as previously found in independent phantom experiments.16,23 It should be noted that the projection-based technique requires the low and high tube potential projection sets to be obtained from the same view angle to effect BH correction and material decomposition of the projections. This is not possible for a dual-source system with angular offset between the 2 sources, because some of the low and high tube potential projections sets are acquired at different views in partial scans when each x-ray tube does not complete a full rotation; hence, the image-based method is exclusively being applied.23
Projection-based DECT imaging for myocardial CTP We have developed a DECT protocol for myocardial CTP with the use of a single-source CT system capable of rapid tube potential switching and projection-based monochromatic reconstruction. In a previous phantom study,16 we have shown that, with the projection-based DECT technique, monochromatic images at 70 keV exhibited the highest contrast-to-noise ratio (iodine sensitivity) for CT MP imaging. Furthermore, we found in porcine studies that projection-based DECT at 70 keV can minimize the overestimation and heterogeneity of absolute MP measurement in normal (nonischemic) myocardium compared with a conventional single CT protocol without BM correction.16 In our initial porcine experiment, DECT MP imaging could only be performed with the cine (continuous) with
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Figure 3 (A) 140 kV and 70 keV arterial and apical TDCs acquired from the SSP 640 mA. The BH effect significantly distorted the 140 kV apical TDC leading to a steeper upslope and a higher mean MP in the apical ROI (B) compared with those of 70 keV [(A) and (C)].
retrospective electrocardiogram (ECG)–gating technique.16 This resulted in an effective dose of 54 mSv for imaging a 4-cm section of the heart, which is too high for routine clinical use. To reduce radiation dose, we have recently implemented the prospective ECG triggering capability (Snapshot Pulse [SSP]; GE Healthcare) on a dedicated research console connected to our GE Healthcare Discovery CT750 HD scanner (Fig. 2). In the following section, we present our initial findings from a porcine experiment that evaluated a new quantitative myocardial CT perfusion protocol using a number of dose-reduction techniques.
Prospectively ECG-triggered projection-based DECT for MP imaging The animal study was approved by the institution research ethics board. A 37-kg female Landrace cross pig was used in the experiments. A scout and a localization
scan of the pig were taken in a supine position with the use of a 64-slice GE Healthcare Discovery CT750HD scanner. Each axial scan covered 40 mm of the heart. The eight 5-mm thick slices depicting the largest cross-section of the left ventricle (LV) were selected for the subsequent MP studies. Slice thickness was set at 5 mm to ensure a good balance between the signal-to-noise ratio (SNR) and z-axis resolution of the perfusion images. To measure MP, the heart was first scanned with the following scan protocol: 20 ECG-triggered dual-energy scan once every other heart beat (to cover roughly 30 seconds) with the use of 140 and 80 kV alternating at 0.2-millisecond intervals, 640 mA and a 0.35-second rotation period starting at 4 seconds into an intravenous injection of contrast (iopamidol 370; Bracco Diagnostics, Plainsboro, NJ, USA) at a rate of 3.5 mL/s and a dosage of 0.7 mL/kg (26 mL). The MP study was repeated, after 15 minutes, to allow washout of contrast in the heart, using the same scan settings except the tube
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Figure 4 Average maps of Snapshot Pulse 640 mA at (A) 140 kV and (B) 70 keV at a slice location closer to the inferior portion of the heart. (C) Beam hardening (BH) induced hypoenhancement in the basal wall of the left ventricle (arrow; A) led to underestimation the basal time-density curve (TDC; dashed line) compared with the adjacent lateral TDC (solid line). (D) The BH effect was significantly reduced in the 70 keV average map and the corresponding basal TDC.
current was decreased to 375 mA. The MP study was repeated again after 15 minutes with the use of the cine (continuous) scanning protocol with the following settings: 80 and 140 kV at 0.2-millisecond intervals, 640 mA, 0.5second rotation period, 30-second duration. The cine images corresponding to the mid-diastolic phases were retrospectively reconstructed. Finally, to investigate the effectiveness of BH correction at different portions of the heart, the MP study was repeated on another set of 8 ! 5 mm slices that covered the inferior portion of the heart with the use of the DECT SSP 640 mA protocol. Thus, a total of 4 MP studies were acquired on the same pig. For the first 3 MP studies, 140 kV images and monochromatic 70 keV DCE CT images at 5-mm slice thickness were reconstructed with full 360 degrees of BH un/ corrected projections to avoid significant ‘‘half-scan’’ shading artifacts in the peripheral slices (the middle slices are less affected by the half-scan artifacts because of the more exact beam geometry).28 For the last MP study covering the lower portion of the heart, 140 kV and 70 keV DCE CT images were generated with half-scan (180 degrees 1 beam fan angle) reconstruction, which offers a superior temporal resolution than full-scan reconstruction, to minimize the larger degree of motion at the apex of the heart. The slice selected for comparison between 140 kV and 70 keV were aligned to the center of the fan beam to minimize
the half-scan effect. TDCs from different parts of the myocardium were generated from both image sets and were compared. The average map of each image set, generated by averaging all the DCE CT images over the 30-second acquisition, as well as the MP map generated with a kinetics model–based deconvolution method29 (CT Perfusion software; GE Healthcare) were compared. The SNR of the MP studies corresponding to DECT SSP 375 mA and 640 mA were also compared. The SNR of a MP study was estimated as the ratio of the peak signal in the measured TDC to the root mean squared deviation of the best-fit curve obtained from deconvolution from the measured TDC.
Technical results Figures 2 and 3 show the 140 kV and 70 keV average maps and TDCs of the DECT SSP 640 mA protocol, respectively. The BH effect in the 140 kV average map induced a severe hypoenhancement in the apical myocardium (Fig. 2A), leading to a significant distortion in the apical TDC (Fig. 3A). The BH effect in the 70 keV average map was minimized, resulting in a more uniform contrast enhancement throughout the left ventricular myocardium (Fig. 2B), and the apical TDC exhibited a more gradual increase in enhancement over time rather than the abrupt
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Figure 5 Dynamic contrast-enhanced CT images acquired from the (A) Snapshot Pulse (SSP) 375 mA and (B) SSP 640 mA protocols. The corresponding apical time-density curve (TDC; C; dashed) and its deconvolution-fitted curve (D; dashed) are shown.
increase in enhancement as seen in the corresponding 140 kV apical TDC (Fig. 3A). As such, the overestimation of MP in the apical region of interest was minimized in the 70 keV MP map (92 vs 110 mL/min per 100 g; Fig. 3B and 3C). Figure 4 also shows the 140 kV and 70 keV average maps and TDCs of the DECT SSP 640 mA protocol but generated with half-scan reconstruction at a slice closer to the inferior portion of the heart. The basal myocardium in the 140-kV average map exhibited a severe hypoenhancement similar to that in the apical myocardium during the peak contrast enhancement in the LV, which in turn led to underestimation of the basal TDC compared with the adjacent lateral TDC (Fig. 4A and 4C). The BH effect in the 70 keV average map was significantly reduced. The basal TDC was more similar to that in the adjacent lateral wall which is typically less affected by the BH effect (Fig. 4B and 4D). Figure 5 shows the 70 keV DCE CT images during the peak contrast enhancement in the LV and apical TDCs corresponding to the SSP 375 mA and SSP 640 mA protocols. Visual comparison suggested the quality of images acquired with the 2 protocols was comparable (Fig. 5A vs 5B). The apical TDC of SSP 375 mA was noisier than that of SSP 640 mA (Fig. 5C vs 5D), and this impression was confirmed by a lower SNR of the 375 mA TDC (10.3 vs 19.2). Figure 6 compares the 70 keV MP maps and MP values obtained with the 2 dose-saving SSP protocols (11 and 23 mSv) and the higher dose cine protocol (54
mSv). The mean MP values in the lateral and apical wall of the LV ranged from 72 to 83 mL/min per 100 g and 79 to 87 mL/min per 100 g, respectively, among the 3 groups.
Discussion In this report we document the implementation of prospective ECG triggering with a rapid tube potential switching DECT system to facilitate quantitative MP imaging with reduced radiation dose. We have shown in a porcine model that BH effects could be significantly minimized with projection-based monochromatic 70 keV images generated from prospectively ECG-triggered dualenergy scans, leading to more accurate measurements of the iodine TDC and MP as found in our previous porcine studies.16 Our initial findings also indicate that the BH reduction capacity of this scan protocol has been shown to be accurate across multiple cardiac territories, allowing for more accurate MP assessment for the whole heart. We have previously published an image-based BH correction method for MP measurement with SECT30 which requires the geometric configuration of the high attenuating regions to be reproduced in phantoms, which are then scanned to derive the optimized settings of the correction algorithm. Because the shape of the heart and surrounding highattenuating materials can be largely different among different sections (eg, superior vs inferior) of the heart, the
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Figure 6 Myocardial perfusion (MP) functional maps acquired from 3 different rapid tube potential switching dual-energy CT imaging protocols: (A) Snapshot Pulse (SSP) 375 mA, (B) SSP 640 mA, and (C) Cine 630 mA. (D) Mean MP values in the apical and lateral regions of interest shown in the MP maps (A–C). LV, left ventricle.
image-based SECT correction method would require a whole-heart phantom to be fashioned. Notwithstanding the complexity in the fabrication of such a phantom, it can only represent a ‘‘population’’ average geometry not the one specific to an individual patient. As such, DECT could provide a more exact BH correction for MP measurement in individual patients than the image-based SECT correction method. MP measurements with the newer prospective ECGtriggering and faster gantry rotation techniques were comparable with those of the cine protocol, whereas the radiation dose was reduced by .200% (23 mSv vs 54 mSv for SSP 640 mA at 0.35 s/rotation vs Cine 630 mA at 0.5 s/rotation). However, this dose remains high relative to other imaging modalities most commonly used for MP assessment in patients with coronary artery disease. For instance, the effective dose of a whole heart MP imaging with SPECT 99mTc-sestamibi is approximately 10–20 mSv.31 As a result, we further investigated the effect of the use of a significantly reduced tube current.
Importantly, our results suggest that image quality is comparable between the 2 protocols despite the reduction in tube current. In addition, the TDCs acquired from the 375 mA protocol exhibited a good SNR (.10) despite being lower than that of the 640 mA TDCs. The resulting MP measurement was comparable with the SSP 640 mA and cine protocols (15% difference in MP is within the normal range of physiologic variation). These results suggested that rapid tube potential switching DECT scanning with prospective ECG-triggering for MP imaging can be performed at one-quarter the radiation dose (11 mSv vs 54 mSv) as the cine protocol with retrospective ECG-gating technique. A study involving a larger number of animals is needed to confirm these initial findings. Studies are also required to investigate the feasibility of further reducing the tube current (dose) for quantitative MP imaging with the use of methods, including iterative reconstruction (eg, adaptive statistical iterative reconstruction32 or modelbased iterative reconstruction33; GE Healthcare) and other image processing methods such as principal component
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analysis34 which minimize noise in either the projections or images to ensure the accuracy and precision of MP measurement from low tube current acquisitions.
10.
Conclusion 11.
We have developed a DECT scanning technique for quantitative myocardial CTP with the use of a single-source rapid tube potential switching approach and projectionbased material decomposition which permit reconstruction of monochromatic cardiac images relatively free of BH artifacts. Several dose-saving approaches, including prospective ECG triggering, faster gantry rotation speed, and reduced x-ray tube current, have also been implemented to facilitate quantitative myocardial CTP with a lower radiation dose.
Acknowledgment This work was supported by the Canadian Institutes of Health Research, Canadian Foundation for Innovation, Ontario Research Fund: Imaging for Cardiovascular Therapeutics, Research and Education Foundation of the Radiological Society of North America, and GE Healthcare.
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317 32. Thibault JB, Sauer KD, Bouman CA, Hsieh J: A three-dimensional statistical approach to improved image quality for multislice helical CT. Med Phys. 2007;34:4526–44. 33. Yu Z, Thibault JB, Bouman CA, Sauer KD, Hsieh J: Fast model-based X-ray CT reconstruction using spatially nonhomogeneous ICD optimization. IEEE Trans Image Process. 2011;20: 161–75. 34. Flury B: Common Principal Components and Related Multivariate Models. New York: John Wiley & Sons; 1988.
Technical Notes
Dual-energy CT and its potential use for quantitative myocardial CT perfusion Aaron So, PhDa,b,*, Jiang Hsieh, PhDc, Suresh Narayanan, PhDc, Jean-Baptiste Thibault, PhDc, Yasuhiro Imai, PhDc, Sandeep Dutta, PhDc, Jonathon Leipsic, MDd, James Min, MDe,f, Troy LaBounty, MDe,f, Ting-Yim Lee, PhDa,b a
Imaging Research Laboratories, Robarts Research Institute, 100 Perth Drive London, ON, Canada N6A 5K8; bImaging Program, Lawson Health Research Institute, London, ON, Canada; cCT Engineering, GE Healthcare, Waukesha, WI, USA; dDepartment of Radiology and Medicine, University of British Columbia, Vancouver, BC, Canada; eDepartment of Medicine, Heart Institute, Cedars-Sinai Heart Institute, Los Angeles, CA, USA and fDepartment of Imaging, Heart Institute, Cedars-Sinai Heart Institute, Los Angeles, CA, USA KEYWORDS: Dual-energy CT; Beam hardening correction; Myocardial CT perfusion; Quantitative imaging; Rapid tube potential switching
Abstract. Application of quantitative myocardial CT perfusion (CTP) for the assessment of coronary artery disease may have a significant effect on patient care as the functional significance of a coronary stenosis can be evaluated through absolute measurement of the downstream myocardial perfusion (MP) both at rest and under exercise or pharmacologic stress. A main challenge of myocardial CTP is beam hardening (BH), arising from the polychromatic nature of x-rays used in CT scanning and the presence of highly attenuating contrast agent in the heart chambers during the CT acquisition. The BH effect induces significant nonuniform shifts in CT numbers which, if uncorrected, can lead to inaccurate assessment of MP. With the recent developments of dual-energy CT (DECT) scanning on clinical scanners, the BH effect on MP measurement could be reduced with the generation of monochromatic images relatively free of BH artifacts from the acquired dual-energy data. Here, we review the different techniques of acquiring dual-energy scans and generating monochromatic images, followed by discussion on the progress of developing a DECT technique with reduced radiation dose for quantitative myocardial CTP. Ó 2012 Society of Cardiovascular Computed Tomography. All rights reserved.
Conflict of interest: A. So has no disclosure; J. Hsieh, S. Narayanan, J.-P. Thibault, Y. Imai, and S. Dutta are employees of GE Healthcare; J. Leipsic is on the speakers, bureau and advisory board of GE Healthcare; J. Min receives research grant from and is on the medical advisory board and speakers, bureau of GE Healthcare; T. LaBounty receives research grant from GE Healthcare; T.-Y. Lee has a licensing agreement with GE Healthcare on the CT Perfusion software and received a grant from and is a consultant to GE Healthcare. * Corresponding author. E-mail address: [email protected] Submitted February 10, 2012. Accepted for publication July 30, 2012.
Introduction Since its introduction, dual-energy CT (DECT) has been a source of a great deal of research and interest in clinical cardiac CT.1–4 In conventional CT, tomographic images of the scanned subject depict differences in x-ray attenuation between materials or tissues within the subject. Although the more attenuating materials (eg, bone) can be differentiated from the less attenuating
1934-5925/$ - see front matter Ó 2012 Society of Cardiovascular Computed Tomography. All rights reserved. http://dx.doi.org/10.1016/j.jcct.2012.07.002
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ones (eg, water and soft tissues), spectral information of the polychromatic x-ray photons is not used in the image reconstruction process. X-ray attenuation by a material in the diagnostic energy range of CT is mainly contributed by photoelectric absorption and Compton scattering.5,6 These interactions depend not only on the x-ray photon energies but also on the atomic number (absorption) and electron density (scattering) of the material. As such, a CT system capable of resolving energies of individual x-ray photons would allow a better tissue characterization than the conventional CT system that only takes into account the integrated x-ray signals over the entire energy spectrum. Another merit of such an energy resolving CT (eg, photon counting CT [PCCT]) system is that images correspond to a specific x-ray energy can be generated.5 The beam hardening (BH) effect arising from the polyenergetic nature of the x-ray photons7 is therefore reduced, allowing a more accurate measurement of contrast concentration and hence myocardial perfusion (MP) when using dynamic contrast-enhanced (DCE) CT scanning.8–11 Although energy-resolving CT is promising, PCCT scanners remain experimental and not yet commercially available because there are technical difficulties that require resolution. One challenge is the high photon flux required in clinical scanning that exceeds the count rate capability of the current photon counting detectors.12–14 It is widely expected that PCCT remains some years away from being clinically available. By contrast, the recently developed clinical DECT scanners are good alternatives to the ideal PCCT system.15,16 These scanners scan objects with the use of x-rays with 2 different energy spectra generated by 2 distinct tube potentials (expressed in kilovolts or kV). Any material in a voxel can then be represented by a mixture of a pair of basis materials whose attenuation properties are known at the different energy levels.5 By convention, water and iodine are selected as the basis materials because of their vast differences in attenuation properties and because of their prevalence in contrast-enhanced studies. Monochromatic images can also be generated by a linear combination of the water and iodine equivalent density images.17,18 In contrast, monochromatic images with energy-resolved CT may be generated directly and do not require an additional step of estimation of material density images. As such, monochromatic images from energy-resolved CT can be more accurate, at least theoretically, than monochromatic images from DECT. Here, we first review different techniques available for dual-energy (two tube potentials) acquisitions and generating monochromatic images from dual-energy projections. We then review current progress to date of 1 DECT technique for quantitative myocardial CT perfusion (CTP) and efforts of reducing radiation dose associated with that dual-energy CTP protocol.
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Technical methods DECT scanners DECT acquisition can be classified into 2 categories: source-oriented and detector-oriented. The source-oriented approach relies on the x-ray source to produce x-rays with 2 different energy spectra for scanning. This can be achieved by using either 2 independent x-ray tubes or a single tube capable of switching between the low and high tube potential rapidly. The detector-oriented approach relies on the ability of the detector to differentiate low-energy from high-energy x-ray photons in a single x-ray beam. Each acquisition approach is discussed in more detail. Dual-source One source-oriented approach for DECT acquisition is to implement a dual-source system, for example the Siemens’ Somatom Definition Flash (Siemens Healthcare, Henkestrasse, Erlangen, Germany) in which 2 x-ray tubedetector pairs are mounted onto the same rotation gantry at an angular offset of 90 degrees (94 degrees for the secondgeneration scanner), with 1 tube operating at 80 or 100 kV and the other operating at 140 kV (Fig. 1A),15,19 Because of the limited space of the gantry, only 1 detector can cover the full acquisition field of view available (AFOV; 50 cm), and the other detector is restricted to a smaller AFOV (26 cm and 33 cm for first- and second-generation scanners, respectively). Data truncation in projections as a result may lead to image artifacts.20 Another potential problem is the increased scattered radiation from 2 instead of 1 x-ray tube turning on at the same time, which could lead to degraded contrast-to-noise ratio and shading artifacts in reconstructed images.21 These errors can affect the measurements of time-density curves (TDCs) and MP, but they can be corrected with data extrapolation20 and cross-scatter correction22 algorithms. Furthermore, material decomposition cannot be performed in the projection domain because of the mismatch in projection views between the high and low tube potential projection sets when the scanned object is moving (eg, a beating heart) or in helical (spiral) scanning as in coronary CT angiography (CTA).23 This could potentially lead to suboptimal correction of BH artifact in the monochromatic images.16,23 Single-source with rapid tube potential switching between projection views The other source-oriented approach is to use a single x-ray tube capable of rapidly switch between the low (80) and high (140 kV) tube potentials,24,25 as in the case of the GE Healthcare Discovery CT750 HD scanner (GE Healthcare, Waukesha, WI, USA; Fig. 1B).16 Because the tube potential switching occurs as rapidly as every 0.2 milliseconds
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Figure 1 Four technologic solutions for dual-energy CT myocardial CT perfusion. (A) Two source-detector pairs with each source operating at a different tube potential and covering a different scan field of view (SFOV). (B) Single source-detector pair with the source capable of rapidly alternate between 2 different tube potentials (in kV) in a single gantry rotation. (C) Single source-detector pair with the detector made of 2 scintillating materials with different x-ray photon-stopping powers. (D) Single source-detector pair with the source capable of rapid tube potential switching between sequential gantry rotations.
(0.0002 second), each pair of 80 and 140 kV projections are essentially acquired from the same view angle. To avoid spectral contamination between consecutive 80 and 140 kV projections, the scanner uses a scintillating material (gemstone) that has an ultrafast primary decay time (0.03 msecond) and low afterglow (delayed fluorescence). One merit of this CT system is that, because there is minimal view angle mismatch between the successive high and low tube potential projections, material decomposition can be performed in the projection space which would lead to a more exact BH correction, as found in previous phantom experiments.16 A caveat with this CT system is that the tube current cannot be modulated at the same speed as the tube potential. However, an equivalent way to independently optimizing tube current for the high and low tube potential projections is to modulate the duration of each tube potential. With the current implementation the fluence for the low and high tube potential projections are selected to optimize image quality with no dose penalty in coronary CTA studies. Single-source with dual-layer detector Compared with the CT systems discussed earlier, a single-source dual-layer detector system (Philips’ prototype) more closely resembles the ideal energy-resolving
photon counting system. In this system, the x-ray detector consists of 2 different scintillating materials bonded together (Fig. 1C).26 In the direction of transmitted x-rays, the low stopping power material is placed above the high stopping power material. This design would allow higher energy xray photons to pass through the top layer without suffering significant interaction, whereas the lower energy photons are mostly attenuated in the top layer. The 2 signals, one from the top and one from the bottom of this sandwich detector, would correspond to x-rays in 2 different energy ranges. The main advantage of this design is that at each view the high- and low-energy projections are exactly registered with respect to each other. However, one potential problem of this system is that the high- and low-energy projections are spectrally less separated from each other compared with the rapid tube potential switching or dual-source system,27 which could lead to suboptimal material decomposition and BH correction for quantitative MP imaging. Single-source with rapid tube potential switching between gantry rotations Another approach of single-source with fast tube potential switching is to alter the tube potential level between gantry rotations instead of projection views in a single gantry
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Figure 2 Average maps acquired from the dual energy CT Snapshot Pulse 640 mA protocol at (A) 140 kV and (B) 70 keV. Beam hardening induced hypoenhancement in the apical myocardium (arrow) in the 140 kV average map was minimized in the corresponding 70 keV average map.
rotation (Fig. 1D). In this setup, the x-ray tube potential is first set to the high level (or low level) to complete the first gantry rotation (scan). The tube potential is then quickly switched to the low kilovolt setting (or vice versa) for the subsequent scan. The Toshiba’s Aquilion ONE (Toshiba America Medical Systems, Tustin, CA, USA) uses this approach for DECT acquisition, with 80 and 135 kV alternates every 1000 milliseconds (1 second) or less. With this approach, material decomposition can be performed in the projection domain because the 2 projection sets are aligned perfectly. However, a potential challenge to this approach is the patient motion occurring between the 2 scans.
Projection-based versus image-based DECT imaging Two methods are available from which monochromatic images can be synthesized from projection sets acquired with 2 different energy spectra: projection-based and image-based. In the projection-based approach,5 the alternating 80- and 140 kV projections acquired from rapidly switching tube potential scanning first undergo data conditioning and calibration. The 2 corrected projection sets are then mathematically transformed into the density projections of 2 selected basis materials (ie, water and iodine), which are then reconstructed into the corresponding water and iodine equivalent density images free of BH effects. Monochromatic images are then generated from a linear combination of the water and iodine images also free of BH effects. In the image-based approach,6 the conventional 80 and 140 kV single-energy CT (SECT) images with residual BH effects are reconstructed first similar to the case of SECT; monochromatic images and water/iodine equivalent density images are then generated by linear combinations of the 80 and 140 kV images with proper weightings of each image. Compared with the projection-based
technique, the image-based approach is easier to implement because it does not require access to the x-ray projections (raw data). However, the image-based method could be less accurate than the projection-based method in measuring the attenuation (density) of materials because BH is already present in the low and high tube potential images which would affect the accuracy of the derived equivalent density images and monochromatic images, as previously found in independent phantom experiments.16,23 It should be noted that the projection-based technique requires the low and high tube potential projection sets to be obtained from the same view angle to effect BH correction and material decomposition of the projections. This is not possible for a dual-source system with angular offset between the 2 sources, because some of the low and high tube potential projections sets are acquired at different views in partial scans when each x-ray tube does not complete a full rotation; hence, the image-based method is exclusively being applied.23
Projection-based DECT imaging for myocardial CTP We have developed a DECT protocol for myocardial CTP with the use of a single-source CT system capable of rapid tube potential switching and projection-based monochromatic reconstruction. In a previous phantom study,16 we have shown that, with the projection-based DECT technique, monochromatic images at 70 keV exhibited the highest contrast-to-noise ratio (iodine sensitivity) for CT MP imaging. Furthermore, we found in porcine studies that projection-based DECT at 70 keV can minimize the overestimation and heterogeneity of absolute MP measurement in normal (nonischemic) myocardium compared with a conventional single CT protocol without BM correction.16 In our initial porcine experiment, DECT MP imaging could only be performed with the cine (continuous) with
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Figure 3 (A) 140 kV and 70 keV arterial and apical TDCs acquired from the SSP 640 mA. The BH effect significantly distorted the 140 kV apical TDC leading to a steeper upslope and a higher mean MP in the apical ROI (B) compared with those of 70 keV [(A) and (C)].
retrospective electrocardiogram (ECG)–gating technique.16 This resulted in an effective dose of 54 mSv for imaging a 4-cm section of the heart, which is too high for routine clinical use. To reduce radiation dose, we have recently implemented the prospective ECG triggering capability (Snapshot Pulse [SSP]; GE Healthcare) on a dedicated research console connected to our GE Healthcare Discovery CT750 HD scanner (Fig. 2). In the following section, we present our initial findings from a porcine experiment that evaluated a new quantitative myocardial CT perfusion protocol using a number of dose-reduction techniques.
Prospectively ECG-triggered projection-based DECT for MP imaging The animal study was approved by the institution research ethics board. A 37-kg female Landrace cross pig was used in the experiments. A scout and a localization
scan of the pig were taken in a supine position with the use of a 64-slice GE Healthcare Discovery CT750HD scanner. Each axial scan covered 40 mm of the heart. The eight 5-mm thick slices depicting the largest cross-section of the left ventricle (LV) were selected for the subsequent MP studies. Slice thickness was set at 5 mm to ensure a good balance between the signal-to-noise ratio (SNR) and z-axis resolution of the perfusion images. To measure MP, the heart was first scanned with the following scan protocol: 20 ECG-triggered dual-energy scan once every other heart beat (to cover roughly 30 seconds) with the use of 140 and 80 kV alternating at 0.2-millisecond intervals, 640 mA and a 0.35-second rotation period starting at 4 seconds into an intravenous injection of contrast (iopamidol 370; Bracco Diagnostics, Plainsboro, NJ, USA) at a rate of 3.5 mL/s and a dosage of 0.7 mL/kg (26 mL). The MP study was repeated, after 15 minutes, to allow washout of contrast in the heart, using the same scan settings except the tube
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Figure 4 Average maps of Snapshot Pulse 640 mA at (A) 140 kV and (B) 70 keV at a slice location closer to the inferior portion of the heart. (C) Beam hardening (BH) induced hypoenhancement in the basal wall of the left ventricle (arrow; A) led to underestimation the basal time-density curve (TDC; dashed line) compared with the adjacent lateral TDC (solid line). (D) The BH effect was significantly reduced in the 70 keV average map and the corresponding basal TDC.
current was decreased to 375 mA. The MP study was repeated again after 15 minutes with the use of the cine (continuous) scanning protocol with the following settings: 80 and 140 kV at 0.2-millisecond intervals, 640 mA, 0.5second rotation period, 30-second duration. The cine images corresponding to the mid-diastolic phases were retrospectively reconstructed. Finally, to investigate the effectiveness of BH correction at different portions of the heart, the MP study was repeated on another set of 8 ! 5 mm slices that covered the inferior portion of the heart with the use of the DECT SSP 640 mA protocol. Thus, a total of 4 MP studies were acquired on the same pig. For the first 3 MP studies, 140 kV images and monochromatic 70 keV DCE CT images at 5-mm slice thickness were reconstructed with full 360 degrees of BH un/ corrected projections to avoid significant ‘‘half-scan’’ shading artifacts in the peripheral slices (the middle slices are less affected by the half-scan artifacts because of the more exact beam geometry).28 For the last MP study covering the lower portion of the heart, 140 kV and 70 keV DCE CT images were generated with half-scan (180 degrees 1 beam fan angle) reconstruction, which offers a superior temporal resolution than full-scan reconstruction, to minimize the larger degree of motion at the apex of the heart. The slice selected for comparison between 140 kV and 70 keV were aligned to the center of the fan beam to minimize
the half-scan effect. TDCs from different parts of the myocardium were generated from both image sets and were compared. The average map of each image set, generated by averaging all the DCE CT images over the 30-second acquisition, as well as the MP map generated with a kinetics model–based deconvolution method29 (CT Perfusion software; GE Healthcare) were compared. The SNR of the MP studies corresponding to DECT SSP 375 mA and 640 mA were also compared. The SNR of a MP study was estimated as the ratio of the peak signal in the measured TDC to the root mean squared deviation of the best-fit curve obtained from deconvolution from the measured TDC.
Technical results Figures 2 and 3 show the 140 kV and 70 keV average maps and TDCs of the DECT SSP 640 mA protocol, respectively. The BH effect in the 140 kV average map induced a severe hypoenhancement in the apical myocardium (Fig. 2A), leading to a significant distortion in the apical TDC (Fig. 3A). The BH effect in the 70 keV average map was minimized, resulting in a more uniform contrast enhancement throughout the left ventricular myocardium (Fig. 2B), and the apical TDC exhibited a more gradual increase in enhancement over time rather than the abrupt
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Figure 5 Dynamic contrast-enhanced CT images acquired from the (A) Snapshot Pulse (SSP) 375 mA and (B) SSP 640 mA protocols. The corresponding apical time-density curve (TDC; C; dashed) and its deconvolution-fitted curve (D; dashed) are shown.
increase in enhancement as seen in the corresponding 140 kV apical TDC (Fig. 3A). As such, the overestimation of MP in the apical region of interest was minimized in the 70 keV MP map (92 vs 110 mL/min per 100 g; Fig. 3B and 3C). Figure 4 also shows the 140 kV and 70 keV average maps and TDCs of the DECT SSP 640 mA protocol but generated with half-scan reconstruction at a slice closer to the inferior portion of the heart. The basal myocardium in the 140-kV average map exhibited a severe hypoenhancement similar to that in the apical myocardium during the peak contrast enhancement in the LV, which in turn led to underestimation of the basal TDC compared with the adjacent lateral TDC (Fig. 4A and 4C). The BH effect in the 70 keV average map was significantly reduced. The basal TDC was more similar to that in the adjacent lateral wall which is typically less affected by the BH effect (Fig. 4B and 4D). Figure 5 shows the 70 keV DCE CT images during the peak contrast enhancement in the LV and apical TDCs corresponding to the SSP 375 mA and SSP 640 mA protocols. Visual comparison suggested the quality of images acquired with the 2 protocols was comparable (Fig. 5A vs 5B). The apical TDC of SSP 375 mA was noisier than that of SSP 640 mA (Fig. 5C vs 5D), and this impression was confirmed by a lower SNR of the 375 mA TDC (10.3 vs 19.2). Figure 6 compares the 70 keV MP maps and MP values obtained with the 2 dose-saving SSP protocols (11 and 23 mSv) and the higher dose cine protocol (54
mSv). The mean MP values in the lateral and apical wall of the LV ranged from 72 to 83 mL/min per 100 g and 79 to 87 mL/min per 100 g, respectively, among the 3 groups.
Discussion In this report we document the implementation of prospective ECG triggering with a rapid tube potential switching DECT system to facilitate quantitative MP imaging with reduced radiation dose. We have shown in a porcine model that BH effects could be significantly minimized with projection-based monochromatic 70 keV images generated from prospectively ECG-triggered dualenergy scans, leading to more accurate measurements of the iodine TDC and MP as found in our previous porcine studies.16 Our initial findings also indicate that the BH reduction capacity of this scan protocol has been shown to be accurate across multiple cardiac territories, allowing for more accurate MP assessment for the whole heart. We have previously published an image-based BH correction method for MP measurement with SECT30 which requires the geometric configuration of the high attenuating regions to be reproduced in phantoms, which are then scanned to derive the optimized settings of the correction algorithm. Because the shape of the heart and surrounding highattenuating materials can be largely different among different sections (eg, superior vs inferior) of the heart, the
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Figure 6 Myocardial perfusion (MP) functional maps acquired from 3 different rapid tube potential switching dual-energy CT imaging protocols: (A) Snapshot Pulse (SSP) 375 mA, (B) SSP 640 mA, and (C) Cine 630 mA. (D) Mean MP values in the apical and lateral regions of interest shown in the MP maps (A–C). LV, left ventricle.
image-based SECT correction method would require a whole-heart phantom to be fashioned. Notwithstanding the complexity in the fabrication of such a phantom, it can only represent a ‘‘population’’ average geometry not the one specific to an individual patient. As such, DECT could provide a more exact BH correction for MP measurement in individual patients than the image-based SECT correction method. MP measurements with the newer prospective ECGtriggering and faster gantry rotation techniques were comparable with those of the cine protocol, whereas the radiation dose was reduced by .200% (23 mSv vs 54 mSv for SSP 640 mA at 0.35 s/rotation vs Cine 630 mA at 0.5 s/rotation). However, this dose remains high relative to other imaging modalities most commonly used for MP assessment in patients with coronary artery disease. For instance, the effective dose of a whole heart MP imaging with SPECT 99mTc-sestamibi is approximately 10–20 mSv.31 As a result, we further investigated the effect of the use of a significantly reduced tube current.
Importantly, our results suggest that image quality is comparable between the 2 protocols despite the reduction in tube current. In addition, the TDCs acquired from the 375 mA protocol exhibited a good SNR (.10) despite being lower than that of the 640 mA TDCs. The resulting MP measurement was comparable with the SSP 640 mA and cine protocols (15% difference in MP is within the normal range of physiologic variation). These results suggested that rapid tube potential switching DECT scanning with prospective ECG-triggering for MP imaging can be performed at one-quarter the radiation dose (11 mSv vs 54 mSv) as the cine protocol with retrospective ECG-gating technique. A study involving a larger number of animals is needed to confirm these initial findings. Studies are also required to investigate the feasibility of further reducing the tube current (dose) for quantitative MP imaging with the use of methods, including iterative reconstruction (eg, adaptive statistical iterative reconstruction32 or modelbased iterative reconstruction33; GE Healthcare) and other image processing methods such as principal component
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analysis34 which minimize noise in either the projections or images to ensure the accuracy and precision of MP measurement from low tube current acquisitions.
10.
Conclusion 11.
We have developed a DECT scanning technique for quantitative myocardial CTP with the use of a single-source rapid tube potential switching approach and projectionbased material decomposition which permit reconstruction of monochromatic cardiac images relatively free of BH artifacts. Several dose-saving approaches, including prospective ECG triggering, faster gantry rotation speed, and reduced x-ray tube current, have also been implemented to facilitate quantitative myocardial CTP with a lower radiation dose.
Acknowledgment This work was supported by the Canadian Institutes of Health Research, Canadian Foundation for Innovation, Ontario Research Fund: Imaging for Cardiovascular Therapeutics, Research and Education Foundation of the Radiological Society of North America, and GE Healthcare.
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