To image blood vessels, we performed fundamental experiments of a red-ray computed tomography (RRCT) scanner using 650-nm-laser and high-sensitivity-photodiode (PD) modules. The line laser beam is irradiated to an object, and the photons penetrating through the object are detected using the PD module through a 1.0-mm-diameter graphite pinhole and a 0.7-mm-diameter 5-mm-length graphite collimator for the PD. The spatial resolutions were primarily determined by the collimator diameter for the PD and were approximately 0.7×0.7 mm2. RRCT was performed by repeating the reciprocating translations and rotations of the object, and the ray-sampling-translation and rotation steps were 0.1 mm and 0.5°, respectively. The image contrast was regulated using the digital amplifier, and the visible diameter of object was 0.5mm.
Photon-counting x-ray computed tomography (PCCT) is useful for selecting optimal energy photons to image various portions of the target object, and we performed fundamental experiments of PCCT to carry out gadolinium (Gd) K-edge CT using Gd-based contrast media. The scanner mainly consists of an x-ray generator with a 0.1-mm-focus tube, a turntable, a cadmium-telluride (CdTe) flat panel detector (FPD) with pixel dimensions of 100 Pm, and a personal computer. An object on the turntable is irradiated by the x-ray generator, 720 radiograms are taken using the FPD, and tomograms are reconstructed. We used 1.3-time magnification tomography, the effective pixel dimensions were approximately 80 Pm, and Gd-K-edge CT was carried out using Gd-based contrast media at a tube voltage of 100 kV, a tube current of 0.40 mA, and a threshold energy of 50 keV.
To perform energy-dispersive x-ray computed tomography (EDCT), we constructed a computer program to amplify the digital values of raw radiograms. The CT scanner consists of an x-ray generator with a 0.1-mm-focus tube, a turntable, a flat panel detector (FPD), and a personal computer (PC). An object on the turntable is irradiated by the x-ray generator, 1.3-magnified 720 radiograms are taken by the FPD, and tomograms are reconstructed using the PC. Utilizing the digital amplifier, the object projections obtained using low-energy photons disappeared with increasing amplification factor at a constant maximum value, and the effective energy increased according to increases in the amplification factor by beam hardening. Using the beam-hardening CT (BHCT) scanner, high-contrast tomography for various objects was performed by controlling effective energy. In particular, fine blood vessels were observed by K-edge CT using iodine media.
To realize novel energy-dispersive X-ray computed tomography (CT) and to reduce the incident dose for the object, we have developed a low-dose low-scattering CT scanner with high spatial resolutions using a room-temperature cadmium telluride (CdTe) detector. X-ray photons are absorbed by the CdTe crystal, and the electric charges flowing through the CdTe crystal are amplified using a current-to-voltage and voltage-to-voltage amplifiers. The first-generation CT is accomplished by repeated translations by the detector and rotations of the object, and the effective photon energy increases with increasing amplification factor of the digital amplifier at a constant maximum output voltage of 5.0 V. The tripleenergy computed tomography (TE-CT) is performed utilizing the beam hardening by the object. In the TE-CT, the scattering photon count is reduced using a 0.5-mm-diam pinhole behind the object, and the spatial resolution is improved by a 0.25-mm-diam pinhole. The exposure time for TE-CT was 9.8 min at a total rotation angle of 180°.
To realize novel photon-counting energy-dispersive X-ray computed tomography (CT), we have developed a low-dose CT scanner using a detector consisting of a cerium-doped yttrium aluminum perovskite [YAP(Ce)] crystal and a small photomultiplier tube (PMT). X-ray photons are absorbed by the YAP(Ce) crystal, and negative outputs are produced from the PMT. The PMT outputs are amplified by an inverse voltage-to-voltage amplifier, and the event pulses are counted by the counter with a low threshold energy of 20 keV. First, almost all the photons are counted without the photon-energy dependence, since the scintillation-photon number produced by one X-ray photon is proportional to the photon energy. Second, the energy-dispersive imaging is performed using the self-beam hardening by the object. The maximum photon count of the projection data is determined after the air absorption, and the effective photon energy increases with increasing digital-amplification factor at the constant maximum count. In the triple-energy CT, the X-ray beam diameter was 0.5 mm, and the spatial resolutions were approximately 0.3×0.3 mm2 . The exposure time for DE-CT was 9.8 min at a total rotation angle of 180°.
Human body mainly consists of muscle, bone and air, and we name penetrating photons the human-body-window (HBW) rays. The HBW spectra were measured using a white power light-emitting diode (LED) and a spectrometer. The photons from the LED penetrate the human body, and reflect, refract and scatter. Therefore, we measured only penetrating HBW spectra from the human body. In the computed tomography (CT), we used a 1.0-mm-diam graphite collimator, two 1.0- mm-diam copper pinholes, and a 1.0-mm-diam aluminum pinhole. The white beam diameter is reduced using the collimator and the first copper pinhole, and the 1.0-mm-diam beam is irradiated to the object. The penetrating HBW photons are selected out using the second copper pinhole behind the object and detected by the photodiode through the aluminum pinhole. HBW-CT is accomplished by repeated translations and rotations of the object. The peak wavelength of the HBW spectra was 610 nm. The translation and rotation steps were 0.25 mm and 1.0º, respectively, and the spatial resolutions were determined as 1.0×1.0 mm2. The scanning time and a total rotation angle for CT were 9.8 min and 180º, respectively.
We have constructed a triple-energy (TE) X-ray photon counter with a room-temperature cadmium telluride (CdTe) detector and three sets of comparators and microcomputers to obtain three kinds of tomograms at three different X-ray energy ranges simultaneously. X-ray photons are detected using the CdTe detector, and the event pulses produced using amplifier module are sent to three comparators simultaneously to regulate three threshold energies of 15, 33 and 50 keV. Using this counter, the energy ranges are 15-33, 33-50 and 50-100 keV; the maximum energy corresponds to the tube voltage. The photon-energy resolution was 3.5% at 59.5 keV. We performed TE computed tomography (TE-CT) at a tube voltage of 100 kV. Using four lead pinholes, three tomograms were obtained simultaneously. Gadolinium-K-edge CT was carried out utilizing an energy range of 50-100 keV. At a tube voltage of 100 kV and a current of 1.60 mA, the count rate was 59 kilocounts per second (kcps).
The linear-plasma flash X-ray generator consists of a high-voltage power supply, a 200-nF high-voltage condenser, a turbomolecular pump, a trigger-pulse generator, and a demountable flash X-ray tube. In the flash X-ray generator, the condenser is charged up to 50 kV by the power supply, and flash X-rays are then produced by the vacuum discharging. The X-ray tube is a demountable triode with a rod-shaped nickel (Ni) target, a zinc (Zn) reflector and a trigger electrode, and the turbomolecular pump evacuates air from the tube at a pressure of approximately 1 mPa. The Ni-target evaporation leads to the formation of weakly ionized linear plasma, consisting of Ni ions and electrons, around the target. In the plasma, K-series characteristic X-ray photons (K photons) are produced, and bremsstrahlung photons with energies beyond the Ni-K-edge energy are absorbed by the plasma and converted into Ni-K photons. Subsequently, Zn-K photons from the Zn reflector are absorbed by the linear Ni plasma and converted into Ni-K photons. Thus, intense Ni-K photons (rays) are irradiated from the plasma axial direction by K-ray amplification by spontaneous emission of radiation (KASER).
We have constructed a dual-energy (DE) high-speed X-ray photon counter with a high-count-rate detector system and energy-range and -region selectors. The detector system consists of a cerium-doped yttrium aluminum perovskite [YAP(Ce)] crystal, a small photomultiplier tube (PMT), and an inverse amplifier for the PMT with a pulse-width extender. X-ray photons are detected using a YAP(Ce)-PMT detector, and the negative output pulses from the PMT are input to the inverse amplifier. The 400-ns-width amplifier-output pulses are sent to the pulse-width extender to measure the pulse height correctly. The event pulses from the extender are sent to the DE counter. In DE-CT, both the X-ray source and the detector module are fixed, and the object on the turntable oscillates on the translation stage. A line beam for DE-CT is formed using two lead (Pb) pinholes in front of the object. The scattering-photon count from the object is reduced using a Pb pinhole behind the object. To improve the spatial resolution, a 0.5-mm-diam Pb pinhole is attached to the YAP(Ce)- PMT detector. The tube voltage and the maximum current were 100 kV and 0.60 mA, respectively. The energy range and region for iodine- and gadolinium-K-edge CT are 35-60 and beyond 50 keV (50-100 keV), respectively. The maximum count rate of DE-CT was 80 kilocounts per second, and the exposure time for tomography was 19.6 min at a total rotation angle of 360°.
To obtain three kinds of tomograms at three different X-ray energy ranges simultaneously, we have constructed a triple-energy (TE) X-ray photon counter with a cooled cadmium telluride (CdTe) detector and three sets of comparators and microcomputers. X-ray photons are detected using the CdTe detector, and the event pulses produced using amplifiers are sent to three comparators simultaneously to regulate three threshold energies of 15, 33 and 50 keV. Using this counter, the energy ranges are 15-33, 33-50 and 50-100 keV; the maximum energy corresponds to the tube voltage. We performed TE computed tomography (TE-CT) at a tube voltage of 100 kV. Using four lead pinholes, three tomograms were obtained simultaneously. Iodine-K-edge CT was carried out utilizing an energy range of 33-50 keV. At a tube voltage of 100 kV and a current of 0.11 mA, the count rate was 21 kilocounts per second (kcps).
To obtain two kinds of tomograms at two different X-ray energy ranges simultaneously, we have constructed a dualenergy (DE) X-ray photon counter with a room-temperature cadmium telluride (CdTe) detector. X-ray photons are detected using the CdTe detector system, and event pulses from an amplifier module are sent to three comparators simultaneously to determine three threshold energies of 33, 48 and 50 keV. The DE counter has energy-range and - region selectors, and the energy range and region are 33-48 and beyond 50 keV (50-100 keV); the maximum energy corresponds to the tube voltage. We performed DE computed tomography (DE-CT) using four lead pinholes at a tube voltage of 100 kV. In Gd-K-edge CT at a range of 50-100 keV, Gd media were observed at high contrasts. The spatial resolutions were 0.5×0.5 mm2, and the exposure time for DE-CT was 19.6 min at a total rotation angle of 360°. At a tube voltage of 100 kV and a current of 0.22 mA, the count rate was 36 kilocounts per second.
To perform low-dose low-scattering X-ray computed tomography (CT), we have constructed a dual-energy (DE) X-ray photon counter with a high-count-rate detector system and energy-range and -region selectors. The detector system consists of a cerium-doped yttrium aluminum perovskite [YAP(Ce)] crystal, a small photomultiplier tube (PMT), and an inverse amplifier for the PMT with a pulse-width extender. In DE-CT, both the X-ray source and the detector module are fixed, and the object on the turntable oscillates on the translation stage. A line beam for DE-CT is formed using a two lead (Pb) pinholes in front of the object. The scattering-photon count from the object is reduced using a Pb pinhole behind the object. To improve the spatial resolution, a 0.5-mm-diam Pb pinhole is attached to the YAP(Ce)-PMT detector. X-ray photons are detected using the detector system, and the event pulses are input to the two energy selectors. In DE-CT, the tube voltage and the maximum current were 100 kV and 0.60 mA, respectively. The energy range and region for soft and gadolinium-K-edge CT are 20-40 and beyond 50 keV (50-100 keV), respectively. The maximum count rate of DE-CT was 84 kilocounts per second, and the exposure time for tomography was 19.6 min at a total rotation angle of 360°.
In the near-infrared-ray computed tomography (NIR-CT) scanner, NIR rays are produced from a light-emitting diode (LED) and detected using a phototransistor (PT) and an infrared filter. The LED-peak wavelength is 850 nm, and 850- nm-peak NIRs are detected using the filtrated PD. The photocurrents flowing through the PT are converted into voltages using an emitter-follower circuit, and the output voltages are sent to a personal computer through an analog-digital converter. The NIR projection curves for tomography are obtained by repeated translations and rotations of the object, and the translating is conducted in both directions of its movement. The 850-nm NIRs easily penetrated living bodies, and the NIR-CT was performed with changes in the sensitivity at relative sensitivities of 1 and 21.
To measure X-ray spectra with high count rates, we developed a detector consisting of a cerium-doped yttrium aluminum perovskite [YAP(Ce)] crystal and a recent multipixel photon counter (MPPC). Scintillation photons are detected using the MPPC, and the photocurrents flowing through the MPPC are converted into voltages and amplified using a high-speed current-voltage (I-V) amplifier. The MPPC bias voltage was set to a value at the pre-Geiger mode to perform zero-dark counting. The event-pulse widths were approximately 200 ns, and the widths were extend to approximately 1 μs. X-ray spectra were measured using a multichannel analyzer (MCA) for pulse-height analysis. The photon energy was roughly determined by the two-point calibration using tungsten K photons and iodine K fluorescence. Using the YAP(Ce)-MPPC detector, first-generation dual-energy computed tomography was accomplished using iodine and gadolinium contrast media.
To develop a dual-energy X-ray CT (DE-CT) system, we have performed investigation of high-speed dual-energy photon counting using two comparators and a low-dark-counting LSO-MPPC (multipixel photon counter) detector. To measure X-ray spectra, electric charges produced in the MPPC are converted into voltages and amplified by a highspeed current-voltage amplifier, and the event pulses are sent to a multichannel analyzer. The MPPC was driven under pre-Geiger mode at an MPPC bias voltage of 70.7 V. The event pulses are sent to two high-speed comparators for selecting two threshold energies to perform DE-CT. The ED-CT is accomplished by repeated linear scans and rotations of the object, and two sets of projection curves of the object are obtained simultaneously by the linear scan. In the DECT, two different-energy tomograms are obtained simultaneously, and photon-count energy subtraction imaging was carried out.
X-ray photons are detected using an Lu2(SiO4)O [LSO] single-crystal scintillator with a decay time of 40 ns and a multipixel photon counter (MPPC). The photocurrent from the MPPC is amplified by a high-speed current-voltage
amplifier with an 80 MHz-gain-band operational amplifier, and the 200-ns-width event pulses are sent to a multichannel
analyzer to measure X-ray spectra. The MPPC is driven in the pre-Geiger mode at a bias voltage of 70.7 V and a
temperature of 23°C. Photon-counting computed tomography (PC-CT) is accomplished by repeated linear scans and
rotations of an object, and projection curves of the object are obtained by linear scanning. The exposure time for
obtaining a tomogram was 10 min with scan steps of 0.5 mm and rotation steps of 1.0°. At a tube voltage of 100 kV, the
maximum count rate was 350 kcps/pixel. We carried out PC-CT using gadolinium media and confirmed the energydispersive
effect with changes in the lower level voltage of event pulses using a comparator.
X-ray photon counting was performed using a silicon X-ray diode (Si-XD) at a tube current of 2.0 mA and tube voltages
ranging from 50 to 70 kV. The Si-XD is a high-sensitivity Si photodiode selected for detecting X-ray photons, and Xray
photons are directly detected using the Si-XD without a scintillator. Photocurrent from the diode is amplified using
charge-sensitive and shaping amplifiers. To investigate the X-ray-electric conversion, we performed the event-pulseheight
(EPH) analysis using a multichannel analyzer. Photon-counting computed tomography (PC-CT) is accomplished
by repeated linear scans and rotations of an object, and projection curves of the object are obtained by the linear scan.
The exposure time for obtaining a tomogram was 10 min at a scan step of 0.5 mm and a rotation step of 1.0°. In PC-CT
at a tube voltage of 70 kV, the image contrast of iodine media fell with increasing lower-level voltage of the event pulse
using a comparator.
X-ray photon counting was performed using a readymade silicon-PIN photodiode (Si-PIN-PD) at tube voltages ranging
from 42 to 60 kV, and X-ray photons are directly detected using the 100 MHz Si-PIN-PD without a scintillator.
Photocurrent from the diode is amplified using charge-sensitive and shaping amplifiers. Using a multichannel analyzer,
X-ray spectra at a tube voltage of 60 kV could easily be measured. The photon-counting computed tomography (PCCT)
is accomplished by repeated linear scans and rotations of an object, and projection curves of the object are obtained
by the linear scan. In the PC-CT, we confirmed the energy-dispersive effect with changes in lower-level voltage of the
event pulse using a comparator.
A low-dose-rate X-ray computed tomography (CT) system is useful for reducing absorbed dose for patients. The CT
system with a tube current of 1.91 mA was developed using a silicon-PIN X-ray diode (Si-PIN-XD). The Si-PIN-XD is
a selected high-sensitive Si-PIN photodiode (PD) for detecting X-ray photons. X-ray photons are detected directly using
the Si-PIN-XD without a scintillator, and the photocurrent from the diode is amplified using current-voltage and
voltage-voltage amplifiers. The output voltage is converted into logical pulses using a voltage-frequency converter with maximum frequency of 500 kHz, and the frequency is proportional to the voltage. The pulses from the converter are sent to differentiator with a time constant of 1 μs to generate short positive pulses for counting, and the pulses are counted using a counter card. Tomography is accomplished by repeated linear scans and rotations of an object, and projection curves of the object are obtained by the linear scan. The exposure time for obtaining a tomogram was 5 min at a scan step of 0.5 mm and a rotation step of 3.0°. The tube current and voltage were 1.91 mA and 100 kV, respectively, and gadolinium K-edge CT was carried out using filtered X-ray spectra with a peak energy of 52 keV.
A high-sensitive X-ray computed tomography (CT) system is useful for decreasing absorbed dose for patients, and a
dark-count-less photon-counting CT system was developed. X-ray photons are detected using a YAP(Ce) [cerium-doped
yttrium aluminum perovskite] single crystal scintillator and an MPPC (multipixel photon counter). Photocurrents are
amplified by a high-speed current-voltage amplifier, and smooth event pulses from an integrator are sent to a high-speed comparator. Then, logical pulses are produced from the comparator and are counted by a counter card. Tomography is accomplished by repeated linear scans and rotations of an object, and projection curves of the object are obtained by the linear scan. The image contrast of gadolinium medium slightly fell with increase in lower-level voltage (Vl) of the comparator. The dark count rate was 0 cps, and the count rate for the CT was approximately 250 kcps.
X-ray fluorescence (XRF) analysis is useful for mapping various atoms in objects. Bremsstrahlung X-rays are selected
using a 3.0 mm-thick aluminum filter, and these rays are absorbed by indium, cerium and gadolinium atoms in objects.
Then XRF is produced from the objects, and photons are detected by a cadmium-telluride detector. The Kα photons are
discriminated using a multichannel analyzer, and the number of photons is counted by a counter card. The objects are
moved and scanned by an x-y stage in conjunction with a two-stage controller, and X-ray images obtained by atomic
mapping are shown on a personal computer monitor. The scan steps of the x and y axes were both 2.5 mm, and the
photon-counting time per mapping point was 0.5 s. We carried out atomic mapping using the X-ray camera, and Kα photons from cerium and gadolinium atoms were produced from cancerous regions in nude mice.
10 Mcps photon counting was carried out using a detector consisting of a 2.0 mm-thick ZnO (zinc oxide) single-crystal
scintillator and an MPPC (multipixel photon counter) module in an X-ray computed tomography (CT) system. The
maximum count rate was 10 Mcps (mega counts per second) at a tube voltage of 70 kV and a tube current of 2.0 mA.
Next, a photon-counting X-ray CT system consists of an X-ray generator, a turntable, a scan stage, a two-stage
controller, the ZnO-MPPC detector, a counter card (CC), and a personal computer (PC). Tomography is accomplished
by repeated linear scans and rotations of an object, and projection curves of the object are obtained by the linear scan
with a scan velocity of 25 mm/s. The pulses of the event signal from the module are counted by the CC in conjunction
with the PC. The exposure time for obtaining a tomogram was 600 s at a scan step of 0.5 mm and a rotation step of 1.0°,
and photon-counting CT was accomplished using iodine-based contrast media.
An energy-discrimination K-edge x-ray computed tomography (CT) system is useful for controlling the image contrast of a target region by selecting both the photon energy and the energy width. The CT system has an oscillation-type linear cadmium telluride (CdTe) detectror. CT is performed by repeated linear scans and rotations of an object. Penetrating x-ray photons from the object are detected by a CdTe detector, and event signals of x-ray photons are produced using charge-sensitive and shaping amplifiers. Both photon energy and energy width are selected out using a multichannel analyzer, and the number of photons is counted by a counter card. In energy-discrimination CT, the tube voltage and tube current were 80 kV and 20 μA, respectively, and the x-ray intensity was 1.92 μGy/s at a distance of 1.0 m from the source and a tube voltage of 80 kV. The energy-discrimination CT was carried out by selecting x-ray photon energies.
We developed an embossed radiography system utilizing single- and dual-energy subtractions for decreasing the
absorption contrast of unnecessary regions, and contrast resolution of a target region was increased using image-shifting
subtraction and a linear-contrast system in a flat panel detector (FPD). To carry out embossed radiography, we
developed a computer program for two-dimensional subtraction, and a conventional x-ray generator with a 0.5-mm-focus tube was used. Energy subtraction was performed at tube voltages of 42.5 and 70.0 kV, a tube current of 1.0 mA, and an x-ray exposure time of 5.0 s. Embossed radiography was achieved with cohesion imaging by use of the
FPD with pixel sizes of 48 ×48 μm, and the shifting dimension of an object in the horizontal and vertical directions
ranged from 48 to 144 μm. We obtained high-contrast embossed images of fine bones and coronary arteries approximately 100 μm in diameter.
An energy-discriminating x-ray camera is useful for performing monochromatic radiography using polychromatic x rays. This x-ray camera was developed to carry out K-edge radiography using iodine-based contrast media. In this camera, objects are exposed by a cone beam from a cerium x-ray generator, and penetrating x-ray photons are detected by a cadmium telluride detector with an amplifier unit. The optimal x-ray photon energy and the energy width are selected out using a multichannel analyzer, and the photon number is counted by a counter card. Radiography was performed by the detector scanning using an x-y stage driven by a two-stage controller, and radiograms obtained by energy discriminating are shown on a personal computer monitor. In radiography, the tube voltage and current were 60 kV and 36 µA, respectively, and the x-ray intensity was 4.7 µGy/s. Cerium K-series characteristic x rays are absorbed effectively by iodine-based contrast media, and iodine K-edge radiography was performed using x rays with energies just beyond iodine K-edge energy 33.2 keV.
A high-speed x-ray tomography system is useful for observing high-speed phenomena. The experimental setup for
tomography consists of a tungsten-target x-ray generator, a tungsten collimator, and a computed radiography system.
An object was exposed by a 2-mm-thick fun beam from the x-ray generator, and scattering x-rays from the slice plane
were detected using an imaging plate through a tungsten collimator with hole diameters of 0.8 mm. Because the
exposed dose for tomography was almost equal to those obtained using two intense flash x-ray generators,
ultra-high-speed tomography could be performed.
Embossed radiography is an important technique for imaging target region by decreasing absorption contrast of objects. The ultra-high-speed embossed radiography system consists of a computed radiography system, an intense flash x-ray generator, and a computer program for shifting the image pixel. In the flash x-ray generator, a high-voltage condenser of 200 nF was charged to 50 kV, and the electric charges in the condenser were discharged to the flash x-ray tube after triggering the cathode electrode. The molybdenum-target evaporation lead to the formation of weakly ionized linear plasma, and intense molybdenum K-series x-rays were produced. High-speed radiography was performed using molybdenum K-rays, and the embossed radiography was carried out utilizing single-energy subtraction after the image shifting. The minimum spatial resolution was equal to the sampling pitch of the CR system of 87.5 μm, and concavoconvex radiography such as phase-differential imaging was performed with an x-ray duration of approximately 0.5 Μs.
In the plasma flash x-ray generator, a 200 nF condenser is charged up to 50 kV by a power supply, and flash x-rays are
produced by the discharging. The x-ray tube is a demountable triode with a trigger electrode, and the turbomolecular
pump evacuates air from the tube with a pressure of approximately 1 mPa. Target evaporation leads to the formation of
weakly ionized linear plasma, consisting of ferrum ions and electrons, around the fine target, and intense K-series
characteristic x-rays are produced from the plasma axial direction. At a charging voltage of 50 kV, the maximum tube
voltage was almost equal to the charging voltage of the main condenser, and the peak current was about 15 kA. In the
spectral measurement, Kβ rays were intense, and higher harmonic x-rays were observed. The pulse widths were 0.5 μs,
and the maximum x-ray intensity was approximately 300 μGy.
Characteristic x-ray generator consists of a constant high-voltage power supply, a filament power supply, a
turbomolecular pump, and an x-ray tube. The x-ray tube is a demountable diode which is connected to the
turbomolecular pump and consists of the following major devices: a pipe-shaped molybdenum hole target, a tungsten
hairpin cathode (filament), a focusing (Wehnelt) electrode, a polyethylene terephthalate x-ray window 0.25 mm in
thickness, and a stainless-steel tube body. In the x-ray tube, the positive high voltage is applied to the anode (target)
electrode, and the cathode is connected to the tube body (ground potential). In this experiment, the tube voltage applied
was from 25 to 35 kV, and the tube current was regulated to within 10 μA by the filament temperature. The exposure
time is controlled in order to obtain optimum x-ray intensity. The electron beams from the cathode are converged to the
target by the focusing electrode, and sharp K-series characteristic x-rays are produced through the focusing electrode at
a tube voltage of 35 kV. Using this generator, we performed monochromatic radiography, monochromatic x-ray
computed tomography, and x-ray fluorescence analysis.
A photon-counting K-edge x-ray Computed Tomography (CT) system is useful for discriminating photon energy and for
decreasing absorbed dose for patients. The CT system is of the first generation type and consists of an x-ray generator, a
turn table, a translation stage, a two-stage controller, a multipixel photon counter (MPPC) module, a 0.5-mm-thick
zinc oxide (ZnO) scintillator, a counter board (CB), and a personal computer (PC). Tomography is accomplished by
repeating the translation and rotation of an object. Penetrating x-ray photons from the object are detected by the
scintillator in conjunction with the MPPC module, and the event signals are counted by the CB. Without using energy
discriminating, photon counting CT was carried out by controlling x-ray spectra.
Energy-discriminating x-ray camera is useful for performing monochromatic radiography using polychromatic x-rays.
The x-ray camera was developed to carry out K-edge radiography using iodine-based contrast media. In this camera,
objects are exposed by a cerium x-ray generator, and penetrating x-rays are detected by a cadmium telluride (CdTe)
detector with an amplifier unit. The optimal x-ray photon energy and energy width are selected out using a multichannel
analyzer (MCA), and the photon number is counted by a counter board (CB). Radiography was performed by the
detector scanning using an x-y stage driven by a two-stage controller, and x-ray images obtained by energy
discriminating are shown in a personal-computer (PC) monitor. Cerium K-series characteristic x-rays are absorbed
effectively by iodine based contrast media, and iodine K-edge radiography was performed using x-rays with photon
energies just beyond K-edge energy 33.2 keV.
An energy-discriminating K-edge x-ray Computed Tomography (CT) system is useful for increasing contrast resolution
of a target region and for diagnosing cancers utilizing a drug delivery system. The CT system is of the first generation
type and consists of an x-ray generator, a turn table, a translation stage, a two-stage controller, a cadmium telluride
(CdTe) detector, a charge amplifier, a shaping amplifier, a multi-channel analyzer (MCA), a counter board (CB), and a
personal computer (PC). The K-edge CT is accomplished by repeating translation and rotation of an object. Penetrating
x-ray spectra from the object are measured by a spectrometer utilizing the CdTe detector, amplifiers, and MCA. Both
the photon energy and the energy width are selected by the MCA for discriminating photon energy. Enhanced iodine
K-edge x-ray CT was performed by selecting photons with energies just beyond iodine K-edge energy of 33.2 keV.
Digital subtraction is useful for carrying out embossed radiography by shifting an x-ray source, and energy subtraction
is an important technique for imaging target region by deleting unnecessary region in vivo. X-ray generator had a
100-μm-focus tube, energy subtraction was performed at tube voltages of 40 and 60 kV, and a 3.0-mm-thick aluminum
filter was used to absorb low-photon-energy bremsstrahlung x-rays. Embossed radiography was achieved with cohesion
imaging using a flat panel detector (FPD) with pixel sizes of 48×48 μm, and the shifting distance of the x-ray source in
horizontal direction and the distance between the x-ray source and the FPD face were 5.0 mm and 1.0 m, respectively.
At a tube voltage of 60 kV and a tube current of 0.50 mA, x-ray intensities without filtering and with filtering were 307
and 28.4 μGy/s, respectively, at 1.0 m from the source. In embossed radiography of non-living animals, the spatial
resolution measured using a lead test chart was approximately 70 μm, and we observed embossed images of fine bones,
soft tissues, and coronary arteries of approximately 100 μm.
X-Ray Fluorescence (XRF) analysis is useful for measuring density distributions of contrast media in vivo. An XRF
camera was developed to carry out mapping for iodine-based contrast media used in medical angiography. In this
camera, objects are exposed by an x-ray beam formed using a 3.0-mm-diameter lead hole. Next, cerium K-series
characteristic x-rays are absorbed effectively by iodine media in objects, and iodine fluorescences are produced from
the objects. Iodine Kα fluorescences are selected out using a 58-μm-thick stannum filter and are detected by a cadmium
telluride (CdTe) detector. Kα rays are discriminated out by a multichannel analyzer (MCA), and photon number is
counted by a counter board (CB). The objects are moved and scanned using an x-y stage driven by a two-stage
controller, and x-ray images obtained by iodine mapping are shown in a personal computer (PC) monitor. In particular,
iodine fluorescences were produced from remanent iodine elements in a cancer region of a rabbit ear.
An x-ray fluorescence (XRF) computed tomography (CT) system utilizing a cadmium telluride (CdTe) detector is
described. The CT system is of the first generation type and consists of a cerium x-ray generator, a turn table, a
translation stage, a two-stage controller, a CdTe spectrometer, a multichannel analyzer (MCA), a counter board (CB),
and a personal computer (PC). When an object is exposed by the x-ray generator, iodine K-series fluorescences are
produced and are detected from vertical direction to x-ray axis using the spectrometer. Fluorescent photons are selected
out using the MCA and are counted by the PC via CB, and XRF CT is performed by repeating translation and rotation
of an object.
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