Radiological Physics and Technology

, Volume 3, Issue 2, pp 127–135

Demonstration of iodine K-edge imaging by use of an energy-discrimination X-ray computed tomography system with a cadmium telluride detector

Authors

    • Faculty of Software and Information ScienceIwate Prefectural University
  • Masashi Kameda
    • Faculty of Software and Information ScienceIwate Prefectural University
  • Eiichi Sato
    • Department of PhysicsIwate Medical University
  • Purkhet Abderyim
    • Department of PhysicsIwate Medical University
  • Toshiyuki Enomoto
    • The 3rd Department of SurgeryToho University School of Medicine
  • Manabu Watanabe
    • The 3rd Department of SurgeryToho University School of Medicine
  • Keitaro Hitomi
    • Department of Electronics and Intelligent SystemsTohoku Institute of Technology
  • Etsuro Tanaka
    • Department of Nutritional Science, Faculty of Applied Bio-scienceTokyo University of Agriculture
  • Hidezo Mori
    • Department of PhysiologyTokai University School of Medicine
  • Toshiaki Kawai
    • Organization for Hamamatsu TechnopolisHeadquarters of Hamamatsu Knowledge Cluster
  • Kiyomi Takahashi
    • Department of Microbiology, School of MedicineIwate Medical University
  • Shigehiro Sato
    • Department of Microbiology, School of MedicineIwate Medical University
  • Akira Ogawa
    • Department of Neurosurgery, School of MedicineIwate Medical University
  • Jun Onagawa
    • Department of Electronics, Faculty of EngineeringTohoku Gakuin University
Article

DOI: 10.1007/s12194-010-0088-8

Cite this article as:
Abudurexiti, A., Kameda, M., Sato, E. et al. Radiol Phys Technol (2010) 3: 127. doi:10.1007/s12194-010-0088-8

Abstract

An energy-discrimination K-edge X-ray computed tomography (CT) system is useful for increasing the contrast resolution of a target region by utilizing contrast media. The CT system has a cadmium telluride (CdTe) detector, and a projection curve is obtained by linear scanning with use of the CdTe detector in conjunction with an X-stage. An object is rotated by a rotation step angle with use of a turntable between the linear scans. Thus, CT is carried out by repetition of the linear scanning and the rotation of an object. Penetrating X-ray photons from the object are detected by the CdTe detector, and event signals of X-ray photons are produced with use of charge-sensitive and shaping amplifiers. Both the photon energy and the energy width are selected by use of a multi-channel analyzer, and the number of photons is counted by a counter card. For performing energy discrimination, a low-dose-rate X-ray generator for photon counting was developed; the maximum tube voltage and the minimum tube current were 110 kV and 1.0 μA, respectively. In energy-discrimination CT, the tube voltage and the current were 60 kV and 20.0 μA, respectively, and the X-ray intensity was 0.735 μGy/s at 1.0 m from the source and with a tube voltage of 60 kV. Demonstration of enhanced iodine K-edge X-ray CT was carried out by selection of photons with energies just beyond the iodine K-edge energy of 33.2 keV.

Keywords

X-ray CTCdTe detectorPhoton countingEnergy discriminationIodine K-edge CT

Introduction

Monochromatic X-rays are very useful for carrying out energy-selective imaging, and various monochromatic X-ray generators have been developed corresponding to specific radiographic objectives. Currently, quasi-monochromatic K-series characteristic X-rays are selected by use of a K-edge monochromatic filter, and the average photon energy of the Kα rays is determined by the target element. Without using the K-edge filter, we have developed a K-ray generator [1, 2] utilizing the angular dependence of bremsstrahlung X-rays. The bremsstrahlung intensity decreases with increasing electron-accelerating (tube) voltage.

For performance of high-speed radiography with X-ray durations below 1 μs, several different flash X-ray generators have been developed [35]. In particular, linear plasma X-ray generators [68] produce extremely clean nickel and copper K-rays without using filters because bremsstrahlung rays are absorbed effectively by weakly ionized metal plasma. From a spherical plasma X-ray generator [9, 10], intense tungsten and tantalum K-rays are produced. Because these rays are absorbed effectively by gadolinium-based contrast media, high-speed gadolinium K-edge angiography has been carried out with X-ray durations of approximately 100 ns.

In conjunction with single silicon crystals, synchrotrons produce quite clean monochromatic parallel X-ray beams, and the photon energy is selected by use of Bragg’s angle. By use of X-ray photons with energies just beyond the iodine K-edge energy of 33.2 keV, enhanced iodine K-edge angiography [1113] has been performed for observation of small blood vessels below 100 μm in diameter with high contrast. On the other hand, we have developed a steady-state cerium X-ray generator [14, 15] and have succeeded in performing cone-beam K-edge angiography by using cerium Kα rays with an average energy of 34.6 keV.

Monochromatic radiography is realizable by means of energy-discriminating imaging, and a reflection-type X-ray camera [16] utilizing a cadmium telluride (CdTe) detector has been developed for carrying out two-dimensional X-ray fluorescence analysis (XRF). Next, we developed a penetration-type, energy-discriminating X-ray camera [17], and iodine K-edge radiography has been performed. Recently, a 64-channel CdTe linear detector has been developed and applied to a material-discriminating X-ray computed tomography (CT) system [18]. However, it is not easy to perform iodine K-edge CT because of low photon energy resolution. Therefore, the energy resolution should be improved so that molecular-level X-ray CT can be performed.

In our research, the major objectives were as follows: development of an energy-discriminating X-ray CT system with an energy resolution of 1.2 keV, decreases in the absorbed dose for patients by a decrease in the tube current, and performing enhanced K-edge imaging by using iodine contrast media. Therefore, we developed an energy-discrimination X-ray CT system utilizing a CdTe detector, which can be used for enhanced iodine K-edge CT.

Energy-discrimination X-ray CT system

X-ray CT system

Figure 1 shows a block diagram of an energy-discrimination X-ray CT system with a CdTe detector, and an experimental setup for counting X-ray photons is shown in Fig. 2. The CT system consists of an X-ray generator (RXG-1050-SP, R-tec, Yokohama), a turntable, a linear-scanning stage (X-stage), a two-stage controller, a CdTe detector system with charge-sensitive and shaping amplifiers (XR-100T, Amptek, Belford, USA), a multi-channel analyzer (MCA) (MCA4000, 4,096 channels, Princeton Gamma-Tec, Rocky Hill, USA), a counter card (CC) (CNT32-4MT, Contec, Osaka), and a personal computer (PC). CT is carried out by repeating the linear scanning and the rotation of an object. Both the stage and the turntable are driven by the stage controller. The scanning and rotation steps are selected corresponding to the spatial resolution, and the resolution improves with decreasing the two steps value. The distance between the X-ray source and the center of the turntable (object) was 900 mm, and the distance between the table and the CdTe detector was set to 41 mm to prevent magnification of the object. Penetrating X-ray photons from an object are detected by the CdTe detector system, and both the photon energy and the energy width are selected by use of the MCA. A 1.0-mm-thick 0.7-mm-diameter lead pinhole is set in front of the CdTe detector, and the event signals from the MCA are counted with the CC and the PC. For performing enhanced iodine K-edge X-ray CT, optimum photon energies just beyond the K-edge energy of 33.2 keV are selected.
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Fig. 1

Block diagram of an energy-discrimination X-ray computed tomography (CT) system with a cadmium telluride (CdTe) detector. MCA and PC are a multi-channel analyzer and a personal computer, respectively

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Fig. 2

Experimental setup of the energy-discrimination X-ray CT system with the CdTe detector

During the scanning for obtaining X-ray projection curves, the counting time per point tc has a constant value of 0.20 s, and the scanning velocity vt is given by
$$ v_{t} = s_{s} /t_{c} , $$
(1)
where ss is the scan step. Next, the X-ray exposure time T for obtaining a tomogram is approximated by
$$ T \cong \pi \left( {L_{ 1} /v_{t} + s_{r} /\omega } \right)/s_{r} , $$
(2)
where L1 is the scanning length, sr (rad) is the rotation step, and ω is the angular velocity (rad/s). In this experiment, if ss, L1, sr, and ω are 1.0, 60 mm, π/60 (rad), and π/180 (rad/s), respectively, vt and T are calculated as 5.0 mm/s and 900 s, respectively.
Figure 3 shows the principle of the X-ray CT system with a single detector and a cone beam. The CT system employs the simplest convolution back-projection method utilizing the Shepp–Logan function for image reconstruction. The scanning length of the detector and the diameter of the turn-table are 60 and 60 mm, respectively. Thus, the allowable diameter Dal of the object is approximated by
$$ D_{al} \cong 2L_{ 2} { \tan }\left( {\theta / 2} \right), $$
(3)
where θ (=64 mrad) is the cone-beam angle, and L2 is the distance between the X-ray source and the table center. Thus, Dal is calculated as 58 mm. Next, the magnification ratio M of the object on the detector-scanning line varies with the distance d on the table, and M is written as
$$ M = L_{ 3} /\left( {L_{ 2} + D_{t} / 2- d} \right), $$
(4)
where L3 is the distance between the X-ray source and the detector, and Dt is the table diameter. Because d ranges from 0 to 60.0 mm, M varies from 1.01 to 1.08
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Fig. 3

Principle of the X-ray CT system with a detector

Novel X-ray generator for photon counting

To perform energy-discrimination X-ray imaging, we developed an extremely low-dose-rate X-ray generator for photon counting. This generator consists of a main controller, an X-ray tube with focus dimensions of 0.5 mm × 0.5 mm, negative and positive Cockcroft–Walton circuits, an isolation transformer, and a double ammeter (Figs. 4, 5). The tube voltage, current, and exposure time can be controlled by a main controller. The high-voltage line employs the Cockcroft–Walton circuits, and positive and negative high voltages are applied to the anode and cathode electrodes, respectively. The filament heating current is supplied by an AC power supply with an isolation transformer which is used for isolation from the high voltage from the Cockcroft–Walton circuit. The double ammeter employs 1.0 and 100 kΩ resistors. The 1.0 kΩ resistor is used for measuring currents within a range from 0.03 to 3.00 mA, and the 100 kΩ resistor is used in the photon counting range from 1.0 to 30.0 μA. The tube voltage ranges from 40 to 110 kV, and the tube current can be regulated within a range from 1.0 μA to 3.0 mA. In this research, the tube voltage applied ranged from 45 to 75 kV, and the tube current was set at 20.0 μA.
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Fig. 4

Block diagram of the X-ray generator for energy-discrimination X-ray CT. Microampere-range tube current can be regulated accurately

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Fig. 5

Main controller of the X-ray generator

X-ray intensity and spectra

The X-ray intensity was measured with an ionization chamber (Victoreen 660, Freedom Imaging, Anaheim, USA) with a 400-ml-volume probe (660-5) at 1.0 m from the X-ray source and at a tube current of 20.0 μA (Fig. 6). When the tube voltage was increased, the X-ray intensity increased. At a tube voltage of 60 kV, the X-ray intensity was 0.735 μGy/s.
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Fig. 6

X-ray intensity measured at 1.0 m from the source and at a tube current of 20.0 μA

In order to measure X-ray spectra, we employed the CdTe detector at a photon energy range from 15 and 70 keV. It is difficult to measure real X-ray spectra because the CdTe detector has two K edges, of cadmium (Cd) and tellurium (Te). At a tube voltage of 60 kV, the bremsstrahlung peak energy was approximately 30 keV when the K-edge energies of Cd and Te were considered (see Figs. 7, 8).
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Fig. 7

Relationship between mass attenuation coefficients of iodine and effective X-ray spectra for iodine K-edge CT

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Fig. 8

Discriminated X-ray spectra for X-ray CT at a tube voltage of 60 kV and a tube current of 5.0 μA. Iodine K-edge CT can be carried out by use of X-ray photons with energies just beyond 33.2 keV

Tomography

Figure 7 shows relation between the mass attenuation coefficients of iodine and X-ray spectra at selected energies; the coefficient curve is discontinuous at the iodine K-edge of 33.2 keV. Figure 8 shows X-ray spectra used for tomography; low-density iodine molecules can be detected for X-ray spectra with energies just beyond the K-edge. In energy-discrimination CT, X-ray photons with an energy range from 33.3 to 43.3 keV were selected by use of the MCA. These photons are absorbed effectively by iodine molecules, and iodine-based contrast media are observed with high contrast. X-ray photons with a range from 43.4 to 60.0 keV reduce the image contrast of iodine molecules. In the X-ray spectra, because two K-edges of cadmium and tellurium are observed, the sensitivity of the CdTe detector varies around the two edges. However, the energies of the two edges are below the iodine K-edge energy, and the decreases in the sensitivity beyond tellurium K-edge energy are negligible for carrying out iodine K-edge imaging.

For performing energy-discrimination CT, the absorbed dose for patients should be minimized by decreasing the tube voltage. Next, the effective photon count for K-edge imaging should be maximized, and X-ray spectra with a peak energy of approximately 30 keV will be useful for performing dual-energy subtraction with use of photons having energies around the K-edge energy of 33.2 keV. Therefore, the tube voltage was determined as 60 kV for carrying out enhanced K-edge CT by use of the CdTe detector.

Energy-discrimination CT was performed at a tube voltage of 60 kV and a current of 20.0 μA. The maximum and minimum densities of tomogram correspond to 255 (black) and 0 (white), respectively, in the JPEG file system. Figure 9 shows three tomograms of glass vials filled with iodine media (Omunipaque, Daiichisankyo, Tokyo) of densities of 15 and 30 mg/ml. With the iodine K-edge CT method, the image density of the iodine media decreased substantially with increasing iodine density. However, the density seldom varied when we used spectra with an energy range from 22.2 to 32.2 keV (below the K-edge energy). The spatial resolution improved with decreases in the scan and rotation steps.
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Fig. 9

Tomograms of two vials filled with iodine media with densities of 15 and 30 mg/ml

Tomograms of a polymethyl methacrylate (PMMA) phantom with two 8.0-mm-diameter holes are shown in Fig. 10. The image density of the PMMA portion decreased for X-rays below the K-edge, and the image density difference between the two holes was large for X-ray spectra with energies beyond the K-edge. Of course, the spatial resolution improved when we decreased the two steps value. Figure 11 shows four tomograms of a PMMA phantom with one 8.0-mm-diameter hole. When K-edge tomography was used, the image contrast of the hole increased substantially. At the scan and the rotation steps of 0.5 mm and π/180 (rad) (=1°), the image contrast decreased with decreasing iodine density.
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Fig. 10

Tomograms of PMMA phantom with two holes filled with two-different-density iodine media

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Fig. 11

Tomograms of PMMA phantom with one hole. Image contrast of hole decreased with decreasing iodine density

K-edge tomography of a dog heart is shown in Fig. 12. The radiogram was obtained with a flat panel detector (1024 eV, Rad-icon Imaging, Santa Clara, USA). The coronary arteries were filled with iodine-based microspheres 15 μm in diameter for making the phantom, and the thick coronary arteries were observed with high contrast. Next, a tomogram of rabbit dactyls (phantom) is shown in Fig. 13. Blood vessels were filled with iodine microspheres. The bones and the muscles are seen, and blood vessels around the dactyls are visible.
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Fig. 12

Tomography of a dog heart. Blood vessels are filled with iodine microspheres, and coronary arteries are observed with high contrast

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Fig. 13

Tomography of rabbit dactyl. Blood vessels around dactyl are visible

A tomogram of a rabbit ear with a cancerous portion is shown in Fig. 14. In this cancer, iodine contrast medium remained after vein injection, and blood vessels for the growing cancer and the cancerous portion could be seen with high contrast. Although the average thickness of the ear was low, the iodine density distribution in the rabbit-ear section could be seen.
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Fig. 14

Tomography of a rabbit ear with cancer. Remaining iodine molecules are detected in the cancer

Discussion

We developed an energy-discriminating X-ray CT system that utilize a CdTe detector with an energy resolution of 1.2 keV to perform enhanced iodine K-edge CT by using photons with energies just beyond the K-edge energy of 33.2 keV. In this CT, the X-ray exposure time for obtaining one tomogram was 15 min for scan and for rotation steps of 1.0 mm and π/60 (rad), respectively, and the time increased with decreases in the two steps. However, the exposure time decreased with increases in both the scanning and angular velocities.

Energy-discrimination CT is realizable with use of the photon-counting CT in conjunction with an MCA for determining both the photon energy and the energy width. In the photon-counting X-ray CT, the image quality improves with increasing photon count per scan step. In our research, the photon counting time and the maximum count were constant, and their values were 0.2 s and 1.0 kilo-count (kc) per second, respectively. Thus, the photon count per measuring point is 0.2 kc; this value seems to be a lower limit for photon-counting imaging. On the other hand, the maximum count rate for energy discrimination is limited by the maximum rate of the MCA. A high-count-rate MCA for reducing exposure time is desirable for the future. In this CT system, the X-ray projection curve is obtained by use of one detector, and multi-slice CT is realizable with multiple detectors.

The spatial resolution improves with decreases in the scan step, the rotation step, and the diameter of the lead pinhole which is used for preventing pileups of the event signal by decreasing the photon count rate. In this CT, because a 0.7-mm-diameter pinhole was used, the diameter should be decreased corresponding to the scan step. In addition, the photon count per scan step should be increased to beyond 0.2 kc by increases in the tube current when a small-diameter pinhole is used.

To prevent signal pileups, the photon-counting X-ray generator is driven at a low tube-current range from 1.0 to 100 μA, and the X-ray flux for performing CT can be decreased easily at the photon-counting. Although iodine K-edge CT was carried out, gadolinium-based contrast media for magnetic resonance angiography [19] can easily be used for gadolinium K-edge CT with X-ray photons just beyond the gadolinium K-edge energy of 50.3 keV.

Acknowledgments

This work was supported by Grants-in-Aid for Scientific Research and Advanced Medical Scientific Research from MECSST, Health and Labor Sciences research grants, grants from the Keiryo Research Foundation, The Promotion and Mutual Aid Corporation for Private Schools of Japan, the Japan Science and Technology Agency, and the New Energy and Industrial Technology Development Organization.

Copyright information

© Japanese Society of Radiological Technology and Japan Society of Medical Physics 2010