- Email: [email protected]
Gaseous detectors for synchrotron radiation applications in medical radiology
Nuclear Instruments
and Methods in Physics Research A 360 (1995) 277-282
NUCLEAR INSTRUMEN’W 8 METHODS IN PHYSICS RESEARCH SectIonA
ELSEVIlER
Gaseous detectors for synchrotro~ radiation applications in medical radiology H.-J. Besch Unir~ersituet-Gesamthochschule Siegen, Siegen, Germany
Abstract Gaseous detectors for medical radiology in combination with synchrotron radiation are used for coronary angiography with intravenous application of the contrast agent. The method is explained, the detectors used are presented and the parameters determining their design are discussed. The detective quantum efficiency, the balance between spatial and contrast resolution and the choice between counting and integrating detectors are considered. Finally cardiac images of diagnostic quality are shown.
1. Introduction Gaseous detectors, especially xenon filled ionization chambers, are widely used in modern tomographic X-ray scanners for medical applications. Since even fast tomographs are relatively slow devices, the detecting systems can afford to read the slow ionic signal and they achieve good efficiency and moderate spatial resolution. Synchrotron radiation is used very little for medical applications, since it is not generafly available. It has, however, a spectral power density per solid angle many orders of magnitude larger than X-ray tubes can provide. So it is used only in those applications where focused or collimated narrow band beams of very high intensity are required. This is the case for transvenous digital subtraction angiography of coronary arteries. Presently the standard method to image a coronary artery is to insert a catheter via the right femoral artery and the aorta into the coronary to be examined and inject an iodine-based contrast agent through the catheter while a series of X-ray images is taken at a speed of up to 40 frames per second. This method gives, due to the high iodine density of 370 mg/ml, excellent images but involves a serious risk to the patient. Consequently the indication threshold for this operation is set to a high level. A less invasive technique, where the contrast agent is injected into a peripheral vein, would be a substantial progress. On its long way from the peripheral vein to a coronary vessel the contrast agent is heavily diluted to a level of 10-20 mg/ml. This leads to such a reduction of contrast that in ordinary X-ray images only large structures like ventricles or the aorta can be seen. In addition the complex structure of the chest in the region of the heart is Oi6S-9~~2/95/$09.50 0 199.5 Elsevier Science R.V. All rights reserved SSDI 0168-9002195~00098-4
superimposed to the faint image of the coronaries. So it has been proposed to use dichromography with synchrotron radiation at the iodine k-edge to recover partially the lost contrast [ 11.
2.
Transvenous coronary chrotron radiation
angiography
with
syn-
In Fig. 1 photon absorption coefficients of bone, fat, muscle and iodine are shown as a function of energy. If one takes two simultaneous images at slightly different energies E, and E, just below and above the iodine k-edge at 33.17 keV and (logarithmically) subtracts these images, one strongly suppresses all contrast other than the contrast due to the iodine k-edge jump. Fig. 2 shows an imaging system (NIKOS III at HASYLAB at DESY, Hamburg, Germany) making use of this method. In the storage ring DORIS a IO-pole wiggler is inserted and the beams at energies E, and E, are filtered out from the wiggler beam with monochromators of the “bent Laue” type. The beams are flat, about 0.5 mm high, 120 mm wide, have a bandwidth of about 70 eV each and are 300 eV apart in energy. The intensity is about 10” photons/mm’s Due to the beam geometry the detector is a two line detector. Two-dimensional images are taken by moving the patient through the beam with a hydraulic chair. At NSLS, Brookhaven, USA, there is a similar system, with the exception of the detector type used [2]. There are no other systems sufficiently advanced to permit examinations of patients. It has been worked out in a number of publications, e.g. Refs. [3,4], that from Poisson statistics the minimum num-
VII. MEDICAL APPLICATIONS
3. Characterization
of noise limited imaging systems
Noise limited imaging systems are usually discussed in terms of detective quantum efficiency (DQE). modulation transfer function (MTF) and Wieners noise power spectrum. Due to space limitations the reader is referred to Refs. [4,5]. In addition linearity and dynamic range are used to describe the performance of a detector.
4. Detector design considerations
10
-’
EIErb~[MeV] Fig. 1. Photon absorption
coefficients.
ber of counts to view a 1 mm artery with a signal to noise ratio of 3 is 20000 counts per pixel, with a pixel size of less than 0.5 X 0.5 mm’. A field of about 12 X 12 cm’ is required. To avoid motion blurring the total exposure time should not exceed 250 ms, resulting in an exposure time of 1 ms per line and a counting rate of 2 X 107/pixel s in the dark parts of the image and 2 X lO’/pixel s in the bright parts of the image, where the photons pass a long distance through the lung. From these numbers one can easily deduce that the dynamic range of the imaging system, defined as maximum linear signal divided by the system noise, has to be at least 1.5 bit to exclude adverse effects of the system noise on the quality of the images. The resulting skin dose to the patient is about 50 mSv (5 rem) per frame.
Following the title of this summary only gaseous detectors are considered. To achieve a high DQE and the required spatial resolution only krypton or xenon can be used at moderate pressure of around 12 bar. This gives a photon absorption length of about 2 cm and a photoelectron range of about 200 km at 33 keV. So the depth of the active volume has to be 5-6 cm for full efficiency. Because the detector has to be housed in a pressure vessel, the dead space between entrance window and active volume has to be short in order to avoid a loss in DQE. Now it has to be decided if a counting or integrating detector is more appropriate, having in mind the counting rates already quoted. At low rates an integrating detector has a low DQE because of the noise of the integrator and at high rate a counting detector has a low DQE because of deadtime losses. This behaviour is shown quantitatively in Fig. 3. Curves a to c show the DQE of an integrating system as a function of the number of photons integrated. with the integrator noise given in equivalent photons as a parameter. For a minimum number of photons of 20000 per pixel (see Section 2) an integrator noise of somewhat less than 100 equivalent photons is required, electronically not a difficult task. For a counting detector curves e to i show the DQE with the ratio of exposure time to dead time as a parameter. At the upper end of the dynamic range
Two line ian chamber E’ I--? Image processing loi .~
from monochromator
;”
Shutter and safety system intensity monitors Fig. 2. Angiographic
Hydraulic scanner imaging system for synchrotron
radiation.
H-J. Besch / Nucl. Ins@. and Meth. in Phys. Res. A 360 (1995) 277-282
beam at the velocity of the drifting ions such that contributions from the same part of the object are correctly superimposed. Such a detector has been tested at the NSLS angiographic facility at Brookhaven [7]. In this case the principle is reversed and the ion drift velocity is synchronized to the scan speed of the patient’s chair. The detector has afterwards not been used for human studies. It is likely that the very low drift field - the drift velocity is less than 50 cm/s - leads to excessive buildup of space charge and subsequent recombination.
008 04
2
0
279
0,4 02 n
“0
2
4
6
8
10
6. Multichannel
ionization chamber
Log Nin (counts) Fig. 3. DQE as function of integrated count rate for integrating detectors with noise (a to d) and counting detectors with deadtime (e to il. Noise in equivalent photons: (a) 1; (b) 10’; (c) 104; (d) 106. In e to i the parameter is the ratio of exposure time to deadtime: (e) 103; (f) 10”; (g) 10’; (h) 106: (i) 10’.
of less than 100 ps per pixel would be required, well below typical signal development times of gaseous detectors and therefore not possible. So after some initial effort to explore the maximum rates of counting gaseous detectors [3] only integrating systems have been developed. deadtimes
5. Kinestatic charge detector This detector has been built for digital radiography with X-ray tubes [6]. It is a kind of moving time projection chamber, consisting of a two-dimensional ionization chamber with the electric field arranged normally to the beam and a one-dimensional strip readout for the positive ions. During the exposure the detector is moved through the
This detector has been developed at the University of Siegen, Germany [8] and is used for patient examinations within the angiographic system NIKOS III at HASYLAB in Hamburg [9], as already shown in Fig. 2. The detector itself is presented in Fig. 4, without pressure vessel and vertically pulled apart for clarification. The beams enter between the Frisch grids and the drift cathode. The gas used is a mixture of xenon or krypton with CO, at a pressure of 12 bar. Electrons are collected at the anode strips. The strip pitch is 400 pm, the distance between drift cathode and grid is 2 mm, and, between grid and anode it is 0.8 mm, and the active volume is 5.5 cm deep. The width is 10 cm. Resuming the ideas outlined in Sections 1 to 4 it has the following properties: _ operation as ionization chamber; - high pressure for efficiency and spatial resolution; - strip anode for pixel definition; _ electron collection and Frisch grid for speed; _ thin active layer for fluorescence suppression (see below); - very large dynamic range; - deliberately poor spatial resolution for good contrast resolution and low dose. The associated electronic system consists - per strip - of a
./ 71
Plane 1
Frisch grid Gi-) /’ Drift cathodeQ?= Plane 2
Fig. 4. Multichannel
ionization chamber with Frisch grid.
VII. MEDICAL APPLICATIONS
280
H.-J. Besch/Nucl. Imtr. undMeth.inPlays. Rex A 260 (19951 277-.X2
resetable integrator and a sample and hold stage. followed by a multiplexer system and a cable driver feeding a 16 bit ADC. The electronics are built in ceramic hybrid technique. The dynamic range is 15.5 bit, the noise is 43 equivalent photons and the system can be used for exposure times between 100 ~LSand 1 s with little change in noise (demonstrating, incidentally, that the system is not noise optimized). All control signals and addresses are transferred via fiber optic links to avoid pickup and ground loops. The DQE of the detector has also been measured and follows the general pattern of Fig. 3b. There is very little difference between the DQE as calculated from electronics noise and the DQE as measured from the image noise, indicating the absence of other noise sources. The DQE is 0.65 at 20000 photons per pixel and achieves a maximum of 0.8, due to losses in the window and the dead space between window and active volume. This is difficult to improve. A linearity measurement has been performed by moving a water filled wedge through the beam using the patient’s chair. There is a nonlinearity at low intensities. This is due to a third harmonic (99 keV) contribution to the beam, originating in the monochromator. The fraction of this contribution is about 1% at the exit of the monochromator and is increased as the beam is selectively attenuated in the water absorber. If this is corrected for, detector and electronics are linear over more than four orders of magnitude. At very high intensities, at least two orders of magnitude higher than those used in coronary angiography, a different type of nonlinearity occurs. Then the density of positive ions between beam and drift cathode becomes so large that the electric field in the beam region is considerably decreased leading to an even higher ion density, increased recombination and, therefore, nonlinearity. This effect has been carefully studied both experimentally and theoretically [lO,ll]. It is most easily observed at the moment when the beam is turned on, e.g. with a mechanical chopper. Because of the Frisch grid the induced signal of the ion cloud is shielded from the anode and the fast moving electrons arriving at the readout strips report the instantaneous situation in the beam region. Fig. 5 shows the result of such a measurement. One can see the full electron signal at first, and its decrease during the buildup of the ion cloud. The full line is the result of a calculation of this complex feedback mechanism. In Fig. 6 the results of two measurements of the MTF of the detector are presented. The sine-function also shown is the MTF of an ideal strip detector of 400 pm pitch. To understand the curves of Fig. 6 one has to remember that the iodine k-edge is just below the xenon k-edge. Consequently 33 keV photons are not absorbed by the photoeffeet from k-shell electrons of xenon. On the contrary, 99 keV photons are mainly absorbed from the k-shell and produce, due to the high k-fluorescence yield, a fluorescence background. This background can be minimized by
I.20 --------
1.08r 096 2 0.84 .i! 2 0.72 ' 2 .g : 0.60 : g 0.48 2 0.36 ,
~~+~~ 0 0
400
800
1200 1600 2000 t[usl
Fig. 5. Increased recombination beam is turned on at 200 ks.
due to space charge buildup. The
the use of a thin active layer. So the difference between the ideal curve and the 2% curve is mainly due to the photoelectron range, and the difference between the 2% curve and the 19% curve is mainly due to fluorescence background, manifesting itself at low spatial frequencies. Around 20% 99 keV fraction will be found behind a moderately fat patient. Because the third harmonic contribution adds to the noise but not to the signal, it also leads to a deterioration of the DQE. It is a regular point of dispute if the radiation dose to the patient. for a given contrast, depends only on the size
0.6 -
k I
0.5 0.4
0.3 .
02. 0.1 ._~~~
1.25
0.625 he
Fig. 6. MTF for different fractions of third harmonic
pairsimm
contribution.
281
H.-J. Besch / Nucl. Instr. and Meth. in Phys. Res. A 360 (1995) 277-282
. 7. Energy images and subtraction
image of a patient. Visible are the aorta (A), pulmonary
veins (PV) and the right coronary
artet ‘Y
(RCA).
of the object or also on the size of the pixels. Since the unnecessary fraction of the radiation dose is 1 - DQE, the answer can be taken from Fig. 3. At low intensity the radiation dose does depend on the pixel size for integrating systems, and it does not for counting systems. At high intensity, vice versa, the radiation dose does not depend on the pixel size for integrating systems, but it does for counting systems. Consequently. for this integrating detector the pixel size has been made as large as possible to maintain a good DQE and contrast resolution at low intensities.
7. Human studies Fig. 7 shows an intravenous angiogram of a patient, taken with the detector described in section 6 and the NIKOS III system 191. Since it is difficult to determine the necessary delay between injection of the contrast medium and the time when the image is taken and because higher iodine densities can be achieved, presently the contrast agent is injected via a catheter in the vena cava superior (this is a low risk procedure). The figure shows the two images at energies E, and E,, respectively, and the resulting subtraction image. The right coronary artery (RCA) is clearly visible down to diameters of about 1 mm. The image was taken 14.6 s after injection, and a second image
was taken 1.7 s later. Another two seconds the iodine contrast would have disappeared.
later most of
8. Outlook Since the detector works close to the quantum limit with sufficient spatial resolution there is not much room for improvements. A new detector has been built with one single difference: it is 130 mm wide instead of 100 mm. This detector is now being used for further human studies on a more regular basis with the main aim to solve the circulation time problem and to synchronize the exposure with the heart cycle. So it can be expected that after many years of hard effort transvenous coronary angiography becomes a valuable medical tool.
Acknowledgements I would like to thank A. DelGuerra and A. Scribano for the invitation to this lively and interesting conference. My special and cordial thanks go to A. H. Walenta, who introduced me to the subject of medical imaging and who spent much time with stimulating and clarifying discus-
VII. MEDICAL APPLICATIONS
282
H.-J. Besch / Nucl. Instr. and Meth. in Phys. Res. A 360 (19951 277-282
sions. I also thank W. R. Dix and W. Kupper, who kept the project going also in difficult periods. Last but not least I would like to thank the students of the detector group at our university for their hard work, their good ideas and their many unbiased questions.
References [1] E. Rubenstein et al., SPIE Proc. Conf. Digital Radiogr. 314 (1981) 42.
[2] W. Thomlinson, Nucl. Instr. and Meth. A 319 (1992) 295. [3] L. Brabetz et al., Atti di Conferenze, Societa Italiana di Fisica 10 (1987) 107. [4] W. Schenk, Thesis, Siegen (1991) unpublished. [S] K. Doi et al. (eds.1 Recent Developments in Digital Imaging New York, (1985). [6] F.A. DiBianca and M.D. Barker, Med. Phys. 12 (1985) 339. [7] H.D. Zeman et al., Conf. Rec. of IEEE NSS (1990) 1070. [8] H.-J. Besch et al., Phys. Med. 9 (2-3) (1993) 171. [9] W.-R. Dix et al., DESY SR 94-01, Hamburg (1994). [lo] M. Wagener, Diplomarbeit, Siegen (1992) unpublished. [ll] H.-J. Besch et al., submitted to Nucl. Instr. and Meth.
and Methods in Physics Research A 360 (1995) 277-282
NUCLEAR INSTRUMEN’W 8 METHODS IN PHYSICS RESEARCH SectIonA
ELSEVIlER
Gaseous detectors for synchrotro~ radiation applications in medical radiology H.-J. Besch Unir~ersituet-Gesamthochschule Siegen, Siegen, Germany
Abstract Gaseous detectors for medical radiology in combination with synchrotron radiation are used for coronary angiography with intravenous application of the contrast agent. The method is explained, the detectors used are presented and the parameters determining their design are discussed. The detective quantum efficiency, the balance between spatial and contrast resolution and the choice between counting and integrating detectors are considered. Finally cardiac images of diagnostic quality are shown.
1. Introduction Gaseous detectors, especially xenon filled ionization chambers, are widely used in modern tomographic X-ray scanners for medical applications. Since even fast tomographs are relatively slow devices, the detecting systems can afford to read the slow ionic signal and they achieve good efficiency and moderate spatial resolution. Synchrotron radiation is used very little for medical applications, since it is not generafly available. It has, however, a spectral power density per solid angle many orders of magnitude larger than X-ray tubes can provide. So it is used only in those applications where focused or collimated narrow band beams of very high intensity are required. This is the case for transvenous digital subtraction angiography of coronary arteries. Presently the standard method to image a coronary artery is to insert a catheter via the right femoral artery and the aorta into the coronary to be examined and inject an iodine-based contrast agent through the catheter while a series of X-ray images is taken at a speed of up to 40 frames per second. This method gives, due to the high iodine density of 370 mg/ml, excellent images but involves a serious risk to the patient. Consequently the indication threshold for this operation is set to a high level. A less invasive technique, where the contrast agent is injected into a peripheral vein, would be a substantial progress. On its long way from the peripheral vein to a coronary vessel the contrast agent is heavily diluted to a level of 10-20 mg/ml. This leads to such a reduction of contrast that in ordinary X-ray images only large structures like ventricles or the aorta can be seen. In addition the complex structure of the chest in the region of the heart is Oi6S-9~~2/95/$09.50 0 199.5 Elsevier Science R.V. All rights reserved SSDI 0168-9002195~00098-4
superimposed to the faint image of the coronaries. So it has been proposed to use dichromography with synchrotron radiation at the iodine k-edge to recover partially the lost contrast [ 11.
2.
Transvenous coronary chrotron radiation
angiography
with
syn-
In Fig. 1 photon absorption coefficients of bone, fat, muscle and iodine are shown as a function of energy. If one takes two simultaneous images at slightly different energies E, and E, just below and above the iodine k-edge at 33.17 keV and (logarithmically) subtracts these images, one strongly suppresses all contrast other than the contrast due to the iodine k-edge jump. Fig. 2 shows an imaging system (NIKOS III at HASYLAB at DESY, Hamburg, Germany) making use of this method. In the storage ring DORIS a IO-pole wiggler is inserted and the beams at energies E, and E, are filtered out from the wiggler beam with monochromators of the “bent Laue” type. The beams are flat, about 0.5 mm high, 120 mm wide, have a bandwidth of about 70 eV each and are 300 eV apart in energy. The intensity is about 10” photons/mm’s Due to the beam geometry the detector is a two line detector. Two-dimensional images are taken by moving the patient through the beam with a hydraulic chair. At NSLS, Brookhaven, USA, there is a similar system, with the exception of the detector type used [2]. There are no other systems sufficiently advanced to permit examinations of patients. It has been worked out in a number of publications, e.g. Refs. [3,4], that from Poisson statistics the minimum num-
VII. MEDICAL APPLICATIONS
3. Characterization
of noise limited imaging systems
Noise limited imaging systems are usually discussed in terms of detective quantum efficiency (DQE). modulation transfer function (MTF) and Wieners noise power spectrum. Due to space limitations the reader is referred to Refs. [4,5]. In addition linearity and dynamic range are used to describe the performance of a detector.
4. Detector design considerations
10
-’
EIErb~[MeV] Fig. 1. Photon absorption
coefficients.
ber of counts to view a 1 mm artery with a signal to noise ratio of 3 is 20000 counts per pixel, with a pixel size of less than 0.5 X 0.5 mm’. A field of about 12 X 12 cm’ is required. To avoid motion blurring the total exposure time should not exceed 250 ms, resulting in an exposure time of 1 ms per line and a counting rate of 2 X 107/pixel s in the dark parts of the image and 2 X lO’/pixel s in the bright parts of the image, where the photons pass a long distance through the lung. From these numbers one can easily deduce that the dynamic range of the imaging system, defined as maximum linear signal divided by the system noise, has to be at least 1.5 bit to exclude adverse effects of the system noise on the quality of the images. The resulting skin dose to the patient is about 50 mSv (5 rem) per frame.
Following the title of this summary only gaseous detectors are considered. To achieve a high DQE and the required spatial resolution only krypton or xenon can be used at moderate pressure of around 12 bar. This gives a photon absorption length of about 2 cm and a photoelectron range of about 200 km at 33 keV. So the depth of the active volume has to be 5-6 cm for full efficiency. Because the detector has to be housed in a pressure vessel, the dead space between entrance window and active volume has to be short in order to avoid a loss in DQE. Now it has to be decided if a counting or integrating detector is more appropriate, having in mind the counting rates already quoted. At low rates an integrating detector has a low DQE because of the noise of the integrator and at high rate a counting detector has a low DQE because of deadtime losses. This behaviour is shown quantitatively in Fig. 3. Curves a to c show the DQE of an integrating system as a function of the number of photons integrated. with the integrator noise given in equivalent photons as a parameter. For a minimum number of photons of 20000 per pixel (see Section 2) an integrator noise of somewhat less than 100 equivalent photons is required, electronically not a difficult task. For a counting detector curves e to i show the DQE with the ratio of exposure time to dead time as a parameter. At the upper end of the dynamic range
Two line ian chamber E’ I--? Image processing loi .~
from monochromator
;”
Shutter and safety system intensity monitors Fig. 2. Angiographic
Hydraulic scanner imaging system for synchrotron
radiation.
H-J. Besch / Nucl. Ins@. and Meth. in Phys. Res. A 360 (1995) 277-282
beam at the velocity of the drifting ions such that contributions from the same part of the object are correctly superimposed. Such a detector has been tested at the NSLS angiographic facility at Brookhaven [7]. In this case the principle is reversed and the ion drift velocity is synchronized to the scan speed of the patient’s chair. The detector has afterwards not been used for human studies. It is likely that the very low drift field - the drift velocity is less than 50 cm/s - leads to excessive buildup of space charge and subsequent recombination.
008 04
2
0
279
0,4 02 n
“0
2
4
6
8
10
6. Multichannel
ionization chamber
Log Nin (counts) Fig. 3. DQE as function of integrated count rate for integrating detectors with noise (a to d) and counting detectors with deadtime (e to il. Noise in equivalent photons: (a) 1; (b) 10’; (c) 104; (d) 106. In e to i the parameter is the ratio of exposure time to deadtime: (e) 103; (f) 10”; (g) 10’; (h) 106: (i) 10’.
of less than 100 ps per pixel would be required, well below typical signal development times of gaseous detectors and therefore not possible. So after some initial effort to explore the maximum rates of counting gaseous detectors [3] only integrating systems have been developed. deadtimes
5. Kinestatic charge detector This detector has been built for digital radiography with X-ray tubes [6]. It is a kind of moving time projection chamber, consisting of a two-dimensional ionization chamber with the electric field arranged normally to the beam and a one-dimensional strip readout for the positive ions. During the exposure the detector is moved through the
This detector has been developed at the University of Siegen, Germany [8] and is used for patient examinations within the angiographic system NIKOS III at HASYLAB in Hamburg [9], as already shown in Fig. 2. The detector itself is presented in Fig. 4, without pressure vessel and vertically pulled apart for clarification. The beams enter between the Frisch grids and the drift cathode. The gas used is a mixture of xenon or krypton with CO, at a pressure of 12 bar. Electrons are collected at the anode strips. The strip pitch is 400 pm, the distance between drift cathode and grid is 2 mm, and, between grid and anode it is 0.8 mm, and the active volume is 5.5 cm deep. The width is 10 cm. Resuming the ideas outlined in Sections 1 to 4 it has the following properties: _ operation as ionization chamber; - high pressure for efficiency and spatial resolution; - strip anode for pixel definition; _ electron collection and Frisch grid for speed; _ thin active layer for fluorescence suppression (see below); - very large dynamic range; - deliberately poor spatial resolution for good contrast resolution and low dose. The associated electronic system consists - per strip - of a
./ 71
Plane 1
Frisch grid Gi-) /’ Drift cathodeQ?= Plane 2
Fig. 4. Multichannel
ionization chamber with Frisch grid.
VII. MEDICAL APPLICATIONS
280
H.-J. Besch/Nucl. Imtr. undMeth.inPlays. Rex A 260 (19951 277-.X2
resetable integrator and a sample and hold stage. followed by a multiplexer system and a cable driver feeding a 16 bit ADC. The electronics are built in ceramic hybrid technique. The dynamic range is 15.5 bit, the noise is 43 equivalent photons and the system can be used for exposure times between 100 ~LSand 1 s with little change in noise (demonstrating, incidentally, that the system is not noise optimized). All control signals and addresses are transferred via fiber optic links to avoid pickup and ground loops. The DQE of the detector has also been measured and follows the general pattern of Fig. 3b. There is very little difference between the DQE as calculated from electronics noise and the DQE as measured from the image noise, indicating the absence of other noise sources. The DQE is 0.65 at 20000 photons per pixel and achieves a maximum of 0.8, due to losses in the window and the dead space between window and active volume. This is difficult to improve. A linearity measurement has been performed by moving a water filled wedge through the beam using the patient’s chair. There is a nonlinearity at low intensities. This is due to a third harmonic (99 keV) contribution to the beam, originating in the monochromator. The fraction of this contribution is about 1% at the exit of the monochromator and is increased as the beam is selectively attenuated in the water absorber. If this is corrected for, detector and electronics are linear over more than four orders of magnitude. At very high intensities, at least two orders of magnitude higher than those used in coronary angiography, a different type of nonlinearity occurs. Then the density of positive ions between beam and drift cathode becomes so large that the electric field in the beam region is considerably decreased leading to an even higher ion density, increased recombination and, therefore, nonlinearity. This effect has been carefully studied both experimentally and theoretically [lO,ll]. It is most easily observed at the moment when the beam is turned on, e.g. with a mechanical chopper. Because of the Frisch grid the induced signal of the ion cloud is shielded from the anode and the fast moving electrons arriving at the readout strips report the instantaneous situation in the beam region. Fig. 5 shows the result of such a measurement. One can see the full electron signal at first, and its decrease during the buildup of the ion cloud. The full line is the result of a calculation of this complex feedback mechanism. In Fig. 6 the results of two measurements of the MTF of the detector are presented. The sine-function also shown is the MTF of an ideal strip detector of 400 pm pitch. To understand the curves of Fig. 6 one has to remember that the iodine k-edge is just below the xenon k-edge. Consequently 33 keV photons are not absorbed by the photoeffeet from k-shell electrons of xenon. On the contrary, 99 keV photons are mainly absorbed from the k-shell and produce, due to the high k-fluorescence yield, a fluorescence background. This background can be minimized by
I.20 --------
1.08r 096 2 0.84 .i! 2 0.72 ' 2 .g : 0.60 : g 0.48 2 0.36 ,
~~+~~ 0 0
400
800
1200 1600 2000 t[usl
Fig. 5. Increased recombination beam is turned on at 200 ks.
due to space charge buildup. The
the use of a thin active layer. So the difference between the ideal curve and the 2% curve is mainly due to the photoelectron range, and the difference between the 2% curve and the 19% curve is mainly due to fluorescence background, manifesting itself at low spatial frequencies. Around 20% 99 keV fraction will be found behind a moderately fat patient. Because the third harmonic contribution adds to the noise but not to the signal, it also leads to a deterioration of the DQE. It is a regular point of dispute if the radiation dose to the patient. for a given contrast, depends only on the size
0.6 -
k I
0.5 0.4
0.3 .
02. 0.1 ._~~~
1.25
0.625 he
Fig. 6. MTF for different fractions of third harmonic
pairsimm
contribution.
281
H.-J. Besch / Nucl. Instr. and Meth. in Phys. Res. A 360 (1995) 277-282
. 7. Energy images and subtraction
image of a patient. Visible are the aorta (A), pulmonary
veins (PV) and the right coronary
artet ‘Y
(RCA).
of the object or also on the size of the pixels. Since the unnecessary fraction of the radiation dose is 1 - DQE, the answer can be taken from Fig. 3. At low intensity the radiation dose does depend on the pixel size for integrating systems, and it does not for counting systems. At high intensity, vice versa, the radiation dose does not depend on the pixel size for integrating systems, but it does for counting systems. Consequently. for this integrating detector the pixel size has been made as large as possible to maintain a good DQE and contrast resolution at low intensities.
7. Human studies Fig. 7 shows an intravenous angiogram of a patient, taken with the detector described in section 6 and the NIKOS III system 191. Since it is difficult to determine the necessary delay between injection of the contrast medium and the time when the image is taken and because higher iodine densities can be achieved, presently the contrast agent is injected via a catheter in the vena cava superior (this is a low risk procedure). The figure shows the two images at energies E, and E,, respectively, and the resulting subtraction image. The right coronary artery (RCA) is clearly visible down to diameters of about 1 mm. The image was taken 14.6 s after injection, and a second image
was taken 1.7 s later. Another two seconds the iodine contrast would have disappeared.
later most of
8. Outlook Since the detector works close to the quantum limit with sufficient spatial resolution there is not much room for improvements. A new detector has been built with one single difference: it is 130 mm wide instead of 100 mm. This detector is now being used for further human studies on a more regular basis with the main aim to solve the circulation time problem and to synchronize the exposure with the heart cycle. So it can be expected that after many years of hard effort transvenous coronary angiography becomes a valuable medical tool.
Acknowledgements I would like to thank A. DelGuerra and A. Scribano for the invitation to this lively and interesting conference. My special and cordial thanks go to A. H. Walenta, who introduced me to the subject of medical imaging and who spent much time with stimulating and clarifying discus-
VII. MEDICAL APPLICATIONS
282
H.-J. Besch / Nucl. Instr. and Meth. in Phys. Res. A 360 (19951 277-282
sions. I also thank W. R. Dix and W. Kupper, who kept the project going also in difficult periods. Last but not least I would like to thank the students of the detector group at our university for their hard work, their good ideas and their many unbiased questions.
References [1] E. Rubenstein et al., SPIE Proc. Conf. Digital Radiogr. 314 (1981) 42.
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