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Instrumentation of the ESRF medical imaging facility
Nuclear Instruments and Methods in Physics Research A 428 (1999) 513}527
Instrumentation of the ESRF medical imaging facility H. Elleaume *, A.M. Charvet , P. Berkvens, G. Berruyer, T. Brochard, Y. Dabin, M.C. Dominguez, A. Draperi, S. Fiedler, G. Goujon, G. Le Duc, M. Mattenet, C. Nemoz, M. Perez, M. Renier, C. Schulze, P. Spanne , P. Suortti, W. Thomlinson, F. Esteve , B. Bertrand , J.F. Le Bas Universite& J. Fourier, Unite& IRM, CHU, B.P. 217, 38043 Grenoble cedex, France European Synchrotron Radiation Facility, B.P. 220, 38043 Grenoble cedex, France Swiss Light Source, PSI, CH 5232 Villigen PSI, Switzerland NSLS, Brookhaven National Laboratory, Upton, NY 11973, USA Received 28 December 1998
Abstract At the European Synchrotron Radiation Facility (ESRF) a beamport has been instrumented for medical research programs. Two facilities have been constructed for alternative operation. The "rst one is devoted to medical imaging and is focused on intravenous coronary angiography and computed tomography (CT). The second facility is dedicated to pre-clinical microbeam radiotherapy (MRT). This paper describes the instrumentation for the imaging facility. Two monochromators have been designed, both are based on bent silicon crystals in the Laue geometry. A versatile scanning device has been built for pre-alignment and scanning of the patient through the X-ray beam in radiography or CT modes. An intrinsic germanium detector is used together with large dynamic range electronics (16 bits) to acquire the data. The beamline is now at the end of its commissioning phase; intravenous coronary angiography is intended to start in 1999 with patients and the CT pre-clinical program is underway on small animals. The "rst in vivo images obtained on animals in angiography and CT modes are presented to illustrate the performances of these devices. 1999 Elsevier Science B.V. All rights reserved. PACS: 07.85.qe Keywords: Synchrotron; Medical imaging; Monochromatic beams; Angiography; Computed tomography; Germanium detector
1. Introduction Synchrotron radiation provides a new source of monochromatic X-ray beams which are tunable * Corresponding author. Tel: #33-04-76-88-23-43; fax: #33-04-76-88-27-85. E-mail address: [email protected] (H. Elleaume) Deceased.
over a broad energy range and well suited to in vivo medical imaging. At the European Synchrotron Radiation Facility (ESRF), one beamline has been dedicated for medical applications of synchrotron radiation. The beamline has been designed for three main programs: coronary angiography, computed tomography applied to the brain, and a pre-clinical radiotherapy program limited to animal studies in the "rst phase; other imaging techniques, such as
0168-9002/99/$ - see front matter 1999 Elsevier Science B.V. All rights reserved. PII: S 0 1 6 8 - 9 0 0 2 ( 9 9 ) 0 0 1 6 7 - 9
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X-ray di!raction imaging [1], are also envisaged. An ambulatory health care unit has been built to house the imaging programs, and the pre-clinical radiotherapy experiments take place in a separate facility at the beamline. In this paper we "rst present the medical imaging programs and give a brief overview of the beamline design, we then describe the beamline instrumentation, in particular the monochromators, the patient positioning system and the data acquisition system, we "nally show the "rst in vivo results obtained in coronary angiography and computed tomography, and conclude by the perspectives foreseen for the two imaging projects.
2. Medical motivations 2.1. Coronary angiography Coronary artery disease motivates a large number of research programs since it is a major health care problem in the industrialized countries. The underlying disease, arteriosclerosis, is characterized by plaque accumulations leading to gradual narrowing of the arterial lumen. This narrowing will induce myocardial infarction or ischemic necrosis of the heart muscle if complete obstruction is reached. Arterial angiography, as an invasive method for the diagnosis of vascular alterations, is a standard procedure established many years ago. The technique provides excellent images of the coronary arteries and their anatomic con"guration, but is too dangerous for general screening or repetitive controls in clinical research [2]. The main risks of the conventional method are related to the use of intra-arterial catheters. As alternatives, non-invasive diagnostic methods such as magnetic resonance angiography, ultrasound techniques and ultrafast CT have been continuously improving during the recent years [3] but still do not provide clinically relevant images. Another approach has been suggested by Rubenstein et al. [4]. They proposed using monochromatic X-rays from synchrotron sources to enhance the vessel contrast by image subtraction procedures after intravenous injection of a contrast medium. Introducing the contrast agent transvenously carries far less risk than arterial injections, but
results in dilution of the contrast agent while it is swept through the central veins, the right heart and the pulmonary circulation. When conventional X-ray sources were used the intravenous method proved to be a complete failure due to motion artifacts and insu$cient #ux. At present, these di$culties can only be overcome by using the high #ux of monochromatic radiation from synchrotron sources. The method is usually known as K-Edge Digital Subtraction Angiography (KEDSA). After intravenous injection of the contrast agent (usually iodine), two images are produced with monochromatic beams, above and below the contrast agent K-edge. The logarithmic subtraction of the two measurements results in an iodine enhanced image, which can be precisely quanti"ed. Iodine will be used as contrast agent (K-edge energy 33.17 keV), in the initial phase of the medical program at the ESRF, but we envisage also to evaluate the potential of gadolinium. Theoretically, gadolinium is more suitable for X-ray images on heavy patients [5,6], since its K-edge is situated at 50.24 keV. At the moment the concentration of gadolinium in the available contrast agents is insu$cient but new products are under development. Worldwide, several systems using synchrotron radiation have been developed for coronary angiography. In 1979 the "rst two systems were installed at the Stanford Synchrotron Radiation Laboratory, USA [4], and at the Institute of Nuclear Physics in Novosibirsk, Russia [7]. In 1981, the NIKOS system was started at Hasylab, Germany [6] followed in 1983 by another project in Tsukuba, Japan [8]. In 1990, the Stanford program was moved to a new medical facility at the NSLS [9]. At the ESRF, the decision was taken in 1990 to build a beamline dedicated to medical applications of synchrotron radiation. Speci"cations of the ESRF angiography system are summarized in Table 1. 2.2. Computed tomography The CT research program is focused on assessing blood volume and tissular perfusion by concentration measurements of contrast agents in the brain. These parameters are the keys to understanding, treating and predicting the outcome of cerebral
H. Elleaume et al. / Nuclear Instruments and Methods in Physics Research A 428 (1999) 513}527 Table 1 ESRF angiography system speci"cations
Table 2 ESRF CT scanner speci"cations
Wiggler gap Machine current Beam height at the patient position Image "eld size Number of lines Width of detector pixel Patient scanning speed Skin-dose (image pair) Photon #ux (image pair)
Slice thickness Field width Number of projections Width of detector pixel Rotation speed Skin-dose (image pair) Photon #ux Data acquisition speed
Data acquisition speed
55 mm 200 mA 500 lm 150 mm;150 mm 300 350 lm 500 mm/s 32 mGy 6;10 photons/ s/mm 1 ms/line
ischemia and brain tumors, which are major public health care problems [10]. Other techniques are used for this purpose, such as magnetic resonance imaging (MRI) and ultrafast CT, but they have limited ability to precisely quantify the measurements. The limits of conventional scanners in measuring attenuation coe$cients in tissues are mainly due to characteristics of the X-ray tube itself: source size, intensity variations, limited #ux and broad energy spectrum. A scanner based on a synchrotron source, which provides monochromatic and parallel X-ray beams, leads to performance close to the theoretical limits and thereby opens new directions for medical imaging. After intravenous injection of a contrast agent, attenuation pro"les above and below the K-absorption edge are recorded over 3603 to reconstruct energy subtraction computed tomography images (KES-CT). This technique allows a precise and quantitative study of contrast agent distribution. Quantities such as blood volume and cerebral perfusion become accessible which in turn could lead to in vivo studies of angiogenesis development of tumors. Another technique called dual-photon absorptiometry (DPA-CT) [11,12] is under development to measure concentration changes of endogene components, such as potassium, which is a neurological signi"cant element. Two images taken at widely di!erent energies are subtracted, and the di!erence images map the low and intermediate-Z elements. The technical speci"cations of the ESRF scanner are given in Table 2.
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1 mm 150 mm 1440 350 lm 723/s 50 mGy 4;10 photons/s/mm 1 ms/projection
3. Beamline design The medical beamline is one of the so-called `long beamlinesa of the ESRF. The radiotherapy program is performed in a dedicated hutch inside the experimental hall, close to the source in order to obtain a maximum X-ray dose rate (see Fig. 1). The imaging programs take place in a satellite building (150 m from the wiggler source). 3.1. Source parameters The ESRF storage ring is operated at 6 GeV and 200 mA with a lifetime more than 50 hs in the 2/3 "lling mode (Table 3). At present, the ESRF has 30 public beamlines operational, and ten collaborating research group (CRG) lines are also available. The wiggler for the medical beamline has been optimized to ful"ll the requirements of the three main experimental programs. The wiggler "eld can reach 1.5 T to produce hard X-rays especially necessary for the radiation therapy program. Parameters of the multi-pole wiggler are given in Table 4. The wiggler will be operated at a "eld of 0.7 T (gap: 55 mm } critical energy: 16.5 keV) for the imaging program, which provides a good compromise between the available #ux and the contributions of the higher energy harmonics to the beam spectrum. 3.2. Design considerations The radiotherapy program and the imaging programs run only alternately. When the beamline is
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Fig. 1. ESRF medical beamline general view: MRT hutch and satellite building.
Table 3 European synchrotron radiation source Ring energy Circumference Horizontal emittance Vertical emittance Operating current Beam lifetime
Table 4 ID17 wiggler characteristics 6 GeV 844 m 4;10\ mrad 30;10\ mrad 200 mA 50 h
Field Period Number of poles Minimum gap Critical energy E K Total power emitted at 200 mA
1.5 T 150 mm 21 24 mm 33.5 keV 19.6 14.3 kW
Note: Numbers are given at maximum wiggler "eld.
set in the imaging mode, a connecting tube replaces the MRT equipment. The "rst instrumentation hutch in the experimental hall (MRT hutch) is connected to the imaging facility by a 100 m long tube in which the white beam radiation travels under vacuum (10\ Pa). The tube is shielded with 20 mm of lead along its path in the experimental hall, and is enclosed in a concrete tunnel between the experimental hall and the satellite building to avoid any radiation hazard. The tunnel walls are 300 mm thick and are made of special concrete
reinforced with hematite. The imaging facility consists of a monochromator hutch and an experimental hutch located in the satellite building. The monochromators have been installed in the satellite building rather than in the experimental hall to minimize the power density on the crystals and the e!ects of ground vibrations. The white beam reaching the imaging facility is a fan with maximum dimensions of 300 mm width
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Fig. 2. Instrumentation layout (angiography con"guration): (a) MRT hutch, (b) Monochromator hutch in the satellite building.
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Fig. 3. Patient positioning system and data acquisition system in the imaging room.
and 20 mm height. Two di!erent monochromators are used alternatively for coronary angiography and computed tomography. The single bounce angiography monochromator de#ects the monochromatic beam upwards (Fig. 2). The computed tomography monochromator is a double-crystal
"xed-exit system. In the imaging room, a patient positioning device is used to pre-align the patient relative to the beam and to perform scans during the imaging procedures (Fig. 3). In the KEDSA mode, the patient is scanned vertically through the beams, and in CT a rotation scan is performed to
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acquire the projections at all angles. A two-dimensional map of the attenuation coe$cient is then reconstructed as in a classical medical X-ray scanner. The data acquisition system is based on a intrinsic germanium detector followed by read-out electronics of a large dynamic range.
4. Beamline components The "rst hutch located inside the experimental hall serves as the microbeam radiation therapy (MRT) facility and houses several beamline components for the imaging programs. The primary slits are used to de"ne the dimensions of the beam vertically and horizontally. Two sets of "lters are available to limit the heat load on the monochromator crystal by "ltering the low energies of the X-ray spectrum; various thickness of carbon, aluminum or copper can be selected. Two "xed apertures have been designed to ensure that the X-ray beam will never hit the vacuum tube which connects the "rst instrumentation hutch in the experimental hall to the satellite building. They provide 2 mrad aperture in the horizontal plane and 0.14 mrad in the vertical plane. Preclinical MRT equipment, at 30 m from the source, includes a multislit collimator, fast shutters and an animal positioning system. In addition, a photon beam shutter located at the exit of the hutch isolates the rest of the beamline when the beamline is set in the MRT mode. When the beamline is set for the imaging programs, the MRT equipment is removed and a connecting tube is inserted to create a vacuum path up to the satellite building. The 250 m two-level satellite building o!ers an ambulatory patient care unit, which includes separate rooms for beamline optics, patient reception, #uoroscopy, imaging and beamline control. At the entrance of the monochromator hutch, a 500 lm thick beryllium window separates the vacuum tube and the secondary slits from the monochromator vessel which is under oxygen free helium #ow at atmospheric pressure. The radiation shielding of this hutch is realized with local shielding of the instrumentation (20 mm of lead) in addition to the 20 cm thick concrete walls.
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Table 5 Angiography monochromator Con"guration Energy range Crystal thickness Crystal height Crystal width Splitter height Asymmetry angle Typical bending radius
Single bent Laue crystal#splitter 17}51 keV 1.0 mm 20 mm 150 mm 3 mm 153 14 m
4.1. Angiography monochromator The monochromator is a single cylindrically bent silicon (1 1 1) crystal in Laue geometry, which focuses the beam vertically [13}15]. The central thin part of the crystal is 1.0 mm thick, 150 mm wide and 20 mm high, and the asymmetry angle is 153 (see Table 5). Two monochromatic beams are produced, with energies above (E ) and below (E ) the contrast material K-absorption edge, by blocking the central part of the exiting beam by a splitter. The total energy bandwidth re#ected by the bent Laue crystal is about 450 eV, overlapping the Kedge of the contrast material. The bandwidth can be adjusted by the vertical height of the beam. The splitter de"nes the energy separation of the two monochromatic beam. Typically, the two beams have an energy bandwidths of 150 eV, and their mean energies are separated by 300 eV. The assembly is cooled by a gravity-fed water system. The silicon crystal and the splitter can be driven out of the beam to allow unhindered passage of the beam through the monochromator vessel for other experiments. The monochromator can be tuned from 17 to 51 keV, which will allow the use of iodine in the initial phase or gadolinium in the future as contrast agents. At the iodine K edge (33.17 keV) the beams are de#ected upwards by 6.843 (Si (1 1 1) Bragg angle "3.423). The radius of curvature (typically 14 m) is adjusted to give a good cross-over and focus of the beams at the patient position 7 m from the monochromator. The monochromatic beam at this position is 2.25 m above the #oor. The two beams produced by the monochromator propagate at slightly di!erent angles. After crossing at the
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Table 6 Angiography monochromator performances Gap (mm) Intensity of one beam [photons/s/mm/100 mA] Harmonics/fundamental
40 5.9;10
50 4.6;10
60 3.1;10
70 1.8;10
3.6%
1.6%
0.33%
0.13%
Note: Vertical opening of the slits: 10 mm } Beam height at patient position: 700 lm.
Table 7 Computed tomography monochromator Con"guration Energy range Crystal thickness Crystal height Crystal width Asymmetry angle Typical bending radius
Two bent Laue crystals 17}80 keV 1.2 mm 20 mm 150 mm 153 30 m
can be adjusted to re#ect energies which are wide apart or bracket the K-edge of the contrast agent. One energy is selected, and the other one blocked by a mechanical chopper [17]. The imaging scan is performed within 5 seconds, and then fast energy switching from the low energy beam to the high energy beam is done by rotation of the chopper. The acquisition time for the complete procedure will thus amount to little more than 15 s. 4.3. Beam monitors and dose monitors
patient position, these beams diverge and are recorded separately by a dual-line detector. The available space between the monochromator and the detector is 13 m at maximum. Measure of the #ux and harmonic content are given in Table 6. The harmonics can impair the image quality by decreasing the contrast in the subtracted image [16]. The amount of the second harmonic at 99 keV relative to the fundamental energy at 33 keV is less than 1 percent when the wiggler gap is larger than 55 mm. 4.2. Tomography monochromator The monochromator is a tunable "xed-exit device based on two bent Si crystals in the Laue geometry, producing a monochromatic beam parallel to the incident white beam with a vertical o!set of 40 mm. The two bent Si (1 1 1) crystals are mounted on independent goniometers. The energy range is 30 to 80 keV. The bending and geometry of the Si crystals are identical to the angiography monochromator design (see Table 7). Provisions have been made to install another similar "xed exit monochromator in the same vessel. The energy range of the second monochromator will be from 50 to 100 keV. The two "xed-exit monochromators
4.3.1. Beam monitors The two X-ray beams, simultaneously produced by the angiography monochromator, are monitored by a dual beam ion chamber. This detector is made of a single chamber "lled with argon at atmospheric pressure. Two detection zones are de"ned by the electrode con"guration; the upper and lower electrodes are the collection plates and the middle electrode is the high voltage anode. The electrodes are 50 mm long in the beam direction. The ion chamber currents are ampli"ed (Novelec) and read out with VME cards (ICV150-ADAS). The acquisition system records simultaneously the monitor signals and the data from the germanium detector. The monitor values are used to normalize for beam intensities #uctuations. 4.3.2. Dose monitors The dose monitor is a single wide ion chamber, which detects both monochromatic fan beams at the same time, and is used to measure the X-ray dose imparted to the patient. Two similar dose monitors have been installed to provide the redundancy necessary for safety reasons. They are located one after the other inside the imaging room, and are the last beamline components before the
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patient. The ion chambers currents are converted to frequencies (Novelec), which are recorded by VME based counters (VCT6-ESRF). If the pre-set dose limit is reached, a signal is sent to the patient safety system which actuates the fast acting shutters to avoid any over exposure of the patient. In the CT mode, the same dual-beam and dose monitors are used. 4.4. Plexiglas xlters A set of Plexiglas "lters is available after the monochromators to adjust the beam intensity without changing the monochromators settings or the wiggler parameters. It is possible to vary the Plexiglas thickness from 2 up to 32 cm by steps of 2 cm, which provides attenuation factors from 5;10\ to 3;10\ at 33 keV. This device will be used for the pre-positioning scans at small X-ray doses, and to adjust the dose for the imaging scans, depending on the patient size.
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4.7. Medical equipment A conventional X-ray unit (Philips BVR 212) is available in a room close to the imaging area. It will be used if venous catheterization is required and to measure the transit time between the injection of the contrast agent in the arm and the arrival of the bolus in the patient's heart. Medical preparation of the patient will also be performed in this room. The intravenous infusion is performed with an injector (Medrad mark V) under remote control. The injection parameters (volume, rate, speed) are preset and the perfusion is triggered at the beginning of the imaging sequence. The data acquisition and patient scan occur after a delay calculated taking into account the transit time of the bolus. The X-ray exposure is triggered by the ECG in order to record the image during the slow motion phase of the heart cycle (diastole). During the imaging sequence, the physician has the possibility of staying close to the patient in a shielded area inside the imaging room.
4.5. Fast safety shutters Two independent fast-acting shutters are used in case of emergency to stop the beam. Each one is capable of stopping the full monochromatic beam; two are provided to satisfy the requirement for redundancy in the radiation safety system. The shutters are actuated simultaneously. The 6 mm thick shutter blades, made of tungsten, are kept out of the beam with springs compressed by electromagnets. If a problem is detected by the patient safety system (PASS), the electro-magnet power supply is switched o!, which releases the tungsten blades and stops the X-ray beam in about 10 ms. 4.6. Imaging shutters A fast shutter is used to limit the X-ray exposure of the patient to the image acquisition time. Two tungsten blades 1 mm thick, 160 mm wide and 30 mm high, separated by 12 mm, are mounted on each side of a rotation axis. The opening/closing rotation of 303 is driven separately by two motors, and the status of the shutter is detected by a pair of switches. The shutter opens or closes in less than 12 ms.
5. Patient positioning system Since the X-ray beam is a fan of about 0.5 mm height and 150 mm width, it is necessary to move the patient through the beam in order to obtain two-dimensional angiogram. For CT scans, attenuation pro"les are acquired over a complete rotation of the patient. The patient positioning system, which allows both for positioning of the patient and the scan motions during the data acquisition, is a high precision stage with seven degrees of freedom (Spretec). It has a very high rigidity during rotation or translation. This is especially important for tomography, since any eccentricity of the axis of rotation during the data acquisition will result in artifacts in the reconstructed image. 3D computed tomography is made possible by acquisition of a large number of images during helical motion of the patient. The design of the patient positioning system is shown Fig. 3. The U-shaped main frame is mounted on four screw jacks (Z ) which allow precise adjustment of the vertical axis (Z) perpendicular to the plane de"ned by the X-ray beam. The
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maximum vertical translation speed is 500 mm/s with a precision of $0.5%. The speed is constant over a length 200 mm (at maximum), for a total displacement of 600 mm including the acceleration and deceleration phases. The rigidity of the structure has been optimized to obtain the above speci"cations and at the same time to guarantee that the circle of confusion of the rotation axis is less than 400 lm in the CT mode. These requirements led to a structure of 4.5 t (total weight) and 2.6 t for the vertically moving part. The vertical movement is assisted by a pneumatic system (7 bar) which compensates for 95% of the total mass. A single brushless motor (5 kW) drives the stage by four screws linked by belts. In the CT mode, the maximum rotation speed is of 1803/s with a precision of $5%. In the angiography mode, after positioning at the desired incidence angle using two angular movements (h and u), and two translation movements (X and >), several images are acquired while the patient is scanned vertically (Z) up and down. For tomography the same system is used, but the acquisition is done during a rotation of the chair (u).
Ge crystal thickness is 2 mm, giving an e$ciency of nearly 100% at 33 keV and 45% at 99 keV (energy of the second harmonic). The major advantages of HPGe detectors over other types of detectors are their very good e$ciency together with a wide dynamic range. In germanium, the photoelectric e!ect is the dominant attenuation process, and the mean-free path of the photo-electrons is small (Table 9). These properties reduce the cross-talk e!ect, which was observed for example with silicon detectors [18]. The direct conversion from photons to electrons is another important asset. The current produced by absorbed X-rays is measured directly with no ampli"cation process. At low counting rate, su$cient output current is generated by two 33 keV photons (see Table 10), and the linearity is still excellent even at high counting rates. The 16-bit dynamic range of the electronics allows the measurement of signals over more than four decades. The spatial resolution of the system is determined by the vertical height of the beam and the 350 lm horizontal spacing of the elements on the detector.
Table 8 Medical beamline germanium detector
6. Data acquisition system
Total active area
6.1. Detector design In coronary angiography, the two X-ray fan beams are recorded by a dual-line detector synchronously with the vertical motion of the patient. A high-purity germanium detector (HPGe) operating at liquid nitrogen temperature is used (Table 8). It is made from a monolithic P-type Ge crystal 160 mm long, electrically segmented into two rows of 432 parallel strips each (Eurisys Mesures). The
Active thickness Number of elements Pixel width Pixel height Distance between 2 pixels Distance between the 2 rows Leakage current Beryllium window: thickness Minimum current Maximum current
20 mm (height), 150 mm (width) 2.0 mm 2 rows of 432 elements 300 lm 10 mm 50 lm 500 lm (5 pA/pixel 500 lm 10 pA/pixel 0.2 mA/pixel
Table 9 Linear attenuation coe$cients for total attenuation (l ), photoabsorption (l ), coherent scattering (l ), and incohrent scattering (l ) at 33.17 keV.
Germanium Silicon
l (cm\)
l
55.376 2.5546
44.539 1.9821
(cm\)
l (cm\)
l (cm\)
1.7241 0.2344
0.6356 0.3553
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Table 10 Equivalent full-scale X-ray counting rates and number of X-rays per digital output for 33 keV X-rays at di!erent gains of the readout electronics Gain
Capacity (F)
Full-scale current (A)
Number of X-rays in 1 ms at full scale
Number of X-rays per output count
1 2 4 8 16 32 64 128
1.60;10\ 1.60;10\ 1.60;10\ 1.60;10\ 1.00;10\ 1.00;10\ 1.00;10\ 1.00;10\
1.68;10\ 8.39;10\ 4.19;10\ 2.10;10\ 1.05;10\ 5.24;10\ 2.62;10\ 1.31;10\
9.47;10 4.73;10 2.37;10 1.18;10 5.92;10 2.96;10 1.48;10 7.40;10
144 72 36 18 9 4.5 2.25 1.13
In order to minimize the electronic pick-up noise, the "rst ampli"cation stage of the electronics is mounted as close as possible to the Ge crystal [19]. The signals are ampli"ed by ADC integrator boards fanned around the detector. A cryostat houses the Ge detector which works at liquidnitrogen temperature. The 30 l dewar above the detector is "lled automatically from a storage 120 l dewar. The detector is mounted on a support for vertical (Z) and horizontal alignment (>). The vertical position of the beam at the detector position varies from 1.44 up to 4.5 m, depending on the monochromator settings. In addition, the detector can be rotated to be perpendicular to the beam (see Fig. 3). 6.2. Electronics The detector is used in current integration mode where the current in each detector element is integrated and digitized every ms. The general architecture of the data acquisition electronics is based on the following scheme: E Preamplixer: The 864 channels are grouped by 16 on 54 integrator boards constituting the preampli"er stage. These boards are directly plugged into the same printed circuit on which the detector is bonded. The key component is the Burr-Brown low noise ACF2101 dual-switched integrator. Two integrating capacitors (one internal, one external) provide the main gain ranges, controlled by a single digital command.
A second digital command controls the integration and reset status. Both commands are common to all channels. Integration times can be programmed from 0.8 to 130 ms. The pre-ampli"ed voltages travel on twisted-pair #at cables (6 m long) to the digital stage. E Amplixer, AD conversion: The rapidly decreasing cost of ADC components allowed the replacement of the classical multiplexed ADC architecture with a `one ADC per channela solution of much easier implementation. Channels are still grouped by 16 and the resulting 54 ADC boards are distributed in 7 ADC crates, stacked in two racks on both sides of the detector inside the experimental hutch. Each channel has one programmable gain ampli"er (Burr-Brown PGA205) for further gain precision, and one 16 bit ADC (Burr-Brown ADS7807). According to the expected signal levels, the full-scale current of the ampli"er can be changed from 0.13 to 16.8 lA in 8 steps (see Table 10). The 128 signals from each ADC crate are read out from the back-plane by a crate controller and stored in a local bu!er. A 35 m serial line connects this bu!er with the VME data processors located in the control room. E Data processing: The ESRF data acquisition and control bus are based on VME technology. Two Motorola MVME162 processors download data through the serial lines from the ADC crate bu!ers and build up raw data lines in a large VME memory. Each MVME162 has four Industry Pack mezzanine locations available. Seven of
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those are "lled with an ESRF designed IP module for automatic read-out the ADC crate bu!er. The last free location is used for the sequencer which collects commands from the VME and other information from the experiment, such as the patient chair movements. The sequencer provides several programmable schemes depending on the application (tomography or angiography). This module has also been developed at the ESRF. The "nal link with the workstation is made through a BIT3 controller. The workstation performs the logarithmic subtraction from the two detector images, together with pedestal subtraction and gain normalization.
7. Radiation dose The injection of the contrast agent in the venous system rather than in the arterial system leads to a large dilution of the contrast agent in the blood circulation. The dilution factor lies typically between 30 and 40 [6]. After injection of a bolus of 45 ml of iodine (Hexabrix 320 mg/ml) in the superior vena cava, the concentration measured in the coronary arteries is about 10 mg/ml. The signal-to-noise ratio (SNR) in the subtracted images is in "rst approximation proportional to the contrast material concentration (c ) and varies as the square H root of the number of photons (N) (see Eq. (1)). It is therefore mandatory to increase the number of photons, and therefore the radiation dose, to compensate for the contrast agent dilution.
SNR"
*k c H o H
N 2
The absorbed dose (Gy) can be expressed as:
L k D"10*t E U (2) G o G G G Here *t is the X-ray exposure time (s), E the photon energy (J), the X-ray intensity [photons cm\ s\], k/o the mass absorption coe$cient for the tissue [cm g\], and i indicates the photon energy (fundamental and harmonics). KEDSA experiments on patients at Brookhaven and Hamburg [9,6] suggest that 3;10 photons/s/mm in front of the patient, in each beam, and a scan speed of 50 cm/s are needed for clinical relevant images of the distal parts of the coronary arteries without movement artifacts. These intensities correspond to a dose rate of 32 Gy/s. For a scan speed of 50 cm/s, a beam height of 0.5 mm and a line acquisition time of 1 ms, the resulting absorbed dose per monochromatic beam is 16 mGy or an absorbed dose of 32 mGy per subtracted image (Eq. (2)). At present, the #ux available at the medical imaging facility is a factor of 3 smaller than the requirements stated above, and the beam height at the patient position is 700 lm. However, the better experimental conditions at the ESRF compensate for the lower intensity. The beams are more parallel, and the larger depth-of-"eld reduces bone artifacts. The contrast is improved by the good e$ciency and small cross-talk of the detector. These allow for a lower scan speed (350 mm/s), longer sampling time and probably for higher harmonic content. 8. Results
(1)
where (*k/o) is the di!erence of the energy-depenH dent mass absorption coe$cients above and below the K-edge of the contrast agent. The limiting factor in intravenous coronary angiography is the X-ray dose delivered to the patient. The maximum X-ray dose will be 200 mGy for the complete procedure at the ESRF, which is comparable with the average dose received during a conventional coronary examination at the hospital.
8.1. Angiography The angiography system has been tested with several samples. A three-dimensional anthropomorphic arterial phantom [20] has been used to characterize the ability of the system to detect stenoses in the coronary arteries through various thickness of tissue simulating a patient examination. In vivo experiments have been performed on animal models. Four pigs weighing from 40 to
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45 kg were imaged. During this experiment, the wiggler gap was set at 60 mm and the ring current was about 170 mA. The monochromator produced two beams at 33.3 and 33.0 keV, above and below the iodine K-edge, with an energy bandwidth of 150 eV. The beam height at the pig position was 0.7 mm. The vertical speed was set at 250 mm/s with an integration time of 1.4 ms corresponding to a vertical displacement of 350 lm per line. The X-ray dose received by the animal was monitored all along the experiment, it amounted 28 mGy/ image in average. Each pig was anesthetized and ventilated with a respirator. The transit time between the injection of the contrast agent (Levovist, Schering) and the arrival of the bolus in the heart was measured with an ultrasound echograph. The catheter was inserted into the jugular vein and advanced to the superior vena cava under X-ray control. The pig was then placed horizontally on its belly on a support mounted on the patient positioning system.
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The heart rate and the blood oxygen saturation were monitored during the procedure. After the intravenous administration of a vaso-dilatator (Corvasal 1 mg/50ml) images were acquired in left anterior and oblique orientations. During imaging, the physician stood next to the animal in a shielded cabin. Thirty milliliters of iodine (Hexabrix威 320 mg/ml) were injected at a rate of 10 ml/s into the superior vena cava using an auto-injector (Mark V, Medrad). The image sequence was started few seconds after the injection of the contrast agent and ten images were acquired to follow the iodine bolus through the pig circulation, while the scanning stage was moving up and down. The time delay between two images was 1.3 s. The right coronary artery and the left anterior descending artery show up clearly on the images (Fig. 4). We also tested the possibility to inject the contrast agent without the use of a catheter in the peripheral circulation (in a vein of the ear). The quality of the resulting images is slightly inferior to those with injection into the superior vena cava, but the coronary arteries are still clearly visible. 8.2. Computed tomography
Fig. 4. K-edge intravenous coronary angiography of a 42 kg pig in LAO orientation. Thirty ml of iodine (Hexabrix 320) were injected within 2.6 s into the superior vena cava. The scan was performed 13 s after injection within 0.6 s with a chair speed of 250 mm/s and 1.4 ms integration time corresponding to a skin dose of 26.8 mGy/frame. The right coronary artery (RCA) and the left anterior descending artery (LAD) are clearly visible.
The goal of this experiment was to demonstrate the possibility of using monochromatic CT to image rat brain tumors in vivo and of precisely quantifying the amount of contrast agent in the tumor. Imaging was performed four weeks after implantation of the tumor and NMR images were acquired the day before the experiment at the synchrotron. The rat was anesthetized by means of an intraperitoneal injection (400 mg/kg of chloralhydrate) before being placed in the vertical stereotactic frame for KES-CT. The contrast agent was injected in the saphena vein. A solution of iodine (Hexabrix 320 mg/cm) was injected at a concentration of 2 ml/kg. Five tubes "lled with known iodine concentrations (5.00, 2.50, 1.25, 1.00 and 0.50 mg/cm) were taped very close to the rat brain, as a reference to cross-check the quanti"cation. For this experiment, a small motorized animal stage was used instead of the patient positioning system. The slice thickness, given by the beam height, was 800 lm and
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method to follow up an angiogenesis inhibitor by comparing the quantitative results obtained on untreated and treated rats bearing gliomas. Acknowledgements This work is the result of a large collective e!orts and we gratefully acknowledge the contribution of the ESRF sta!, in particular G. Blattman, G. Retout, E. Courraud, J.C. Laidet and the safety group. We thank A. Dilmanian, A Thompson, W. Grae! and R. Dix for many fruitful discussions during the design phase. We further acknowledge the contributions of J.L. Lefaix, J.J. Leplat (INRA/CEA) for their participation in the angiography experiment on the pigs. Fig. 5. Reconstructed SRCT image of a rat brain bearing a glioma obtained by logarithmic subtraction between images acquired above and below the iodine K-edge. Note the ringshaped contrast enhancement at the tumor location.
contiguous multi-slices of the brain were performed over 12 mm. The amount of iodine measured in the tumor itself ranges between 600 lg/cm and 750 lg/cm (see Fig. 5).
9. Perspectives The angiography medical program will start in Spring 1999. We intend to perform about ten patient examinations at the iodine K-edge. The patients selected for the KEDSA protocol will have previously undergone a conventional coronary angiography for comparison, to validate the synchrotron based technique. In parallel, we envisage to evaluate the possibility of using gadolinium as contrast agent for heavy patients. The CT program will focused on the development of the pre-clinical K-edge method in comparison with NMR. Quantitative kinetic studies of the contrast agent clearance in the tumor are expected to provide important information for understanding the blood brain barrier rupture and leakage phenomena. Fundamental studies will be "rst performed on rats bearing glioma. We also envisage using this
References [1] D. Chapman, W. Thomlinson, R.E. Johnston, D. Wahburn, E. Pisano, Z. Zhong, R. Menk, F. Arfelli, D. Sayers, Phys. Med. Biol. 42 (11) (1997) 2015. [2] J. Kennedy, Cathet. Cardiovasc. Diagn. 8 (1982) 5. [3] L.F. Rogers, Amer. J. Roentgenology 170 (1998) 841. [4] E. Rubenstein, J.C. Giacomini, H. Gordon, A. Thompson, G. Brown, R. Hofstadter, W. Thomlinson, H. Zeman, Nucl. Instr. and Meth. A 291 (1990) 80. [5] H.D. Zeman, D.P. Siddons, Nucl. Instr. Meth. A 291 (1990) 67. [6] W.R. Dix, Prog. Biophys. Molec. Biol. 63 (1995) 159. [7] I. Dolbnya, G. Kulipanow, S. Kurylo, N. Pindyurin, M. Sheromov, Phys. Med. VI N3-4 (1990) 313. [8] K. Umetani, K. Ueda, T. Takeda, I. Anno, Y. Itai, M. Akisida, T. Nakajima, Rev. Sci. Instr. 63 (1992) 629. [9] W. Thomlinson, N. GmuK r, D. Chapman, R. Garrett, N. Lazarz, H. Moulin, A.C. Thompson, H.D. Zeman, G.S. Brown, J. Morrison, P. Reiser, V. Padmanabahn, L. Ong, S. Green, J. Giacomini, H. Gordon, E. Rubenstein, Rev. Sci. Instr. 63 (1) (1992) 625. [10] H.J. Feldmann, M. Molls, P. Vaupel, in: M. Molls and P. Vaupel (Eds.) Blood Perfusion and Microenvironment of Human Tumors. Implications for clinical radiooncology, Springer, Berlin, 1998, p. 47. [11] E. Nachaliel, F.A. Dilmanian, R.F. Garrett, W.C. Thomlinson, L.D. Chapman, N.F. GmuK r, N.M. Lazarz, H.R. Moulin, M.L. Rivers, H. Rarback, P.M. Stefan, P. Spanne, Nucl. Instr. and Meth. A 319 (1992) 305. [12] A.M. Charvet, C. Lartizien, F. Esteve, G. Le Duc, A. Collomb, H. Elleaume, S. Fiedler, A. Thompson, T. Brochard, U. Kleuker, H. Stetner, P. Spanne, P. Suortti, J.F. Le Bas, in: M. Ando, C. Uyama (Eds.), Medical Applications of Synchrotron Radiation, Springer, Berlin, 1997.
H. Elleaume et al. / Nuclear Instruments and Methods in Physics Research A 428 (1999) 513}527 [13] P. Suortti, W. Thomlinson, Nucl. Instr. and Meth. A 269 (1988) 639. [14] P. Suortti, W. Thomlinson, D. Chapman, N. GmuK r, R. Greene, N. Lazarz, Nucl. Instr. and Meth. A 297 (1990) 268. [15] P. Suortti, W. Thomlinson, D. Chapman, N. GmuK r, D.P. Siddons, C. Schulze, Nucl. Instr. and Meth. A 336 (1993) 304. [16] H.D. Zeman, H. Moulin, IEEE Trans. Nucl. Sci. 39 (5) (1992).
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[17] C. Schulze, P. Suortti, D. Chapman, Synchrotron Radiation News 7 (13) (1994) 8. [18] H.D. Zeman, E.B. Hughes, J.N. Otis, A.C. Thompson, J.T. Walton, Opt. Eng. 25 (4) (1996) 532. [19] J.C. labiche, French patent C 9413614, November 14, 1994. [20] C.P. Renaudin, B. Barbier, R. Roriz, D. Revel, M. Amiel, Radiology 190 (1994) 579.
Instrumentation of the ESRF medical imaging facility H. Elleaume *, A.M. Charvet , P. Berkvens, G. Berruyer, T. Brochard, Y. Dabin, M.C. Dominguez, A. Draperi, S. Fiedler, G. Goujon, G. Le Duc, M. Mattenet, C. Nemoz, M. Perez, M. Renier, C. Schulze, P. Spanne , P. Suortti, W. Thomlinson, F. Esteve , B. Bertrand , J.F. Le Bas Universite& J. Fourier, Unite& IRM, CHU, B.P. 217, 38043 Grenoble cedex, France European Synchrotron Radiation Facility, B.P. 220, 38043 Grenoble cedex, France Swiss Light Source, PSI, CH 5232 Villigen PSI, Switzerland NSLS, Brookhaven National Laboratory, Upton, NY 11973, USA Received 28 December 1998
Abstract At the European Synchrotron Radiation Facility (ESRF) a beamport has been instrumented for medical research programs. Two facilities have been constructed for alternative operation. The "rst one is devoted to medical imaging and is focused on intravenous coronary angiography and computed tomography (CT). The second facility is dedicated to pre-clinical microbeam radiotherapy (MRT). This paper describes the instrumentation for the imaging facility. Two monochromators have been designed, both are based on bent silicon crystals in the Laue geometry. A versatile scanning device has been built for pre-alignment and scanning of the patient through the X-ray beam in radiography or CT modes. An intrinsic germanium detector is used together with large dynamic range electronics (16 bits) to acquire the data. The beamline is now at the end of its commissioning phase; intravenous coronary angiography is intended to start in 1999 with patients and the CT pre-clinical program is underway on small animals. The "rst in vivo images obtained on animals in angiography and CT modes are presented to illustrate the performances of these devices. 1999 Elsevier Science B.V. All rights reserved. PACS: 07.85.qe Keywords: Synchrotron; Medical imaging; Monochromatic beams; Angiography; Computed tomography; Germanium detector
1. Introduction Synchrotron radiation provides a new source of monochromatic X-ray beams which are tunable * Corresponding author. Tel: #33-04-76-88-23-43; fax: #33-04-76-88-27-85. E-mail address: [email protected] (H. Elleaume) Deceased.
over a broad energy range and well suited to in vivo medical imaging. At the European Synchrotron Radiation Facility (ESRF), one beamline has been dedicated for medical applications of synchrotron radiation. The beamline has been designed for three main programs: coronary angiography, computed tomography applied to the brain, and a pre-clinical radiotherapy program limited to animal studies in the "rst phase; other imaging techniques, such as
0168-9002/99/$ - see front matter 1999 Elsevier Science B.V. All rights reserved. PII: S 0 1 6 8 - 9 0 0 2 ( 9 9 ) 0 0 1 6 7 - 9
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X-ray di!raction imaging [1], are also envisaged. An ambulatory health care unit has been built to house the imaging programs, and the pre-clinical radiotherapy experiments take place in a separate facility at the beamline. In this paper we "rst present the medical imaging programs and give a brief overview of the beamline design, we then describe the beamline instrumentation, in particular the monochromators, the patient positioning system and the data acquisition system, we "nally show the "rst in vivo results obtained in coronary angiography and computed tomography, and conclude by the perspectives foreseen for the two imaging projects.
2. Medical motivations 2.1. Coronary angiography Coronary artery disease motivates a large number of research programs since it is a major health care problem in the industrialized countries. The underlying disease, arteriosclerosis, is characterized by plaque accumulations leading to gradual narrowing of the arterial lumen. This narrowing will induce myocardial infarction or ischemic necrosis of the heart muscle if complete obstruction is reached. Arterial angiography, as an invasive method for the diagnosis of vascular alterations, is a standard procedure established many years ago. The technique provides excellent images of the coronary arteries and their anatomic con"guration, but is too dangerous for general screening or repetitive controls in clinical research [2]. The main risks of the conventional method are related to the use of intra-arterial catheters. As alternatives, non-invasive diagnostic methods such as magnetic resonance angiography, ultrasound techniques and ultrafast CT have been continuously improving during the recent years [3] but still do not provide clinically relevant images. Another approach has been suggested by Rubenstein et al. [4]. They proposed using monochromatic X-rays from synchrotron sources to enhance the vessel contrast by image subtraction procedures after intravenous injection of a contrast medium. Introducing the contrast agent transvenously carries far less risk than arterial injections, but
results in dilution of the contrast agent while it is swept through the central veins, the right heart and the pulmonary circulation. When conventional X-ray sources were used the intravenous method proved to be a complete failure due to motion artifacts and insu$cient #ux. At present, these di$culties can only be overcome by using the high #ux of monochromatic radiation from synchrotron sources. The method is usually known as K-Edge Digital Subtraction Angiography (KEDSA). After intravenous injection of the contrast agent (usually iodine), two images are produced with monochromatic beams, above and below the contrast agent K-edge. The logarithmic subtraction of the two measurements results in an iodine enhanced image, which can be precisely quanti"ed. Iodine will be used as contrast agent (K-edge energy 33.17 keV), in the initial phase of the medical program at the ESRF, but we envisage also to evaluate the potential of gadolinium. Theoretically, gadolinium is more suitable for X-ray images on heavy patients [5,6], since its K-edge is situated at 50.24 keV. At the moment the concentration of gadolinium in the available contrast agents is insu$cient but new products are under development. Worldwide, several systems using synchrotron radiation have been developed for coronary angiography. In 1979 the "rst two systems were installed at the Stanford Synchrotron Radiation Laboratory, USA [4], and at the Institute of Nuclear Physics in Novosibirsk, Russia [7]. In 1981, the NIKOS system was started at Hasylab, Germany [6] followed in 1983 by another project in Tsukuba, Japan [8]. In 1990, the Stanford program was moved to a new medical facility at the NSLS [9]. At the ESRF, the decision was taken in 1990 to build a beamline dedicated to medical applications of synchrotron radiation. Speci"cations of the ESRF angiography system are summarized in Table 1. 2.2. Computed tomography The CT research program is focused on assessing blood volume and tissular perfusion by concentration measurements of contrast agents in the brain. These parameters are the keys to understanding, treating and predicting the outcome of cerebral
H. Elleaume et al. / Nuclear Instruments and Methods in Physics Research A 428 (1999) 513}527 Table 1 ESRF angiography system speci"cations
Table 2 ESRF CT scanner speci"cations
Wiggler gap Machine current Beam height at the patient position Image "eld size Number of lines Width of detector pixel Patient scanning speed Skin-dose (image pair) Photon #ux (image pair)
Slice thickness Field width Number of projections Width of detector pixel Rotation speed Skin-dose (image pair) Photon #ux Data acquisition speed
Data acquisition speed
55 mm 200 mA 500 lm 150 mm;150 mm 300 350 lm 500 mm/s 32 mGy 6;10 photons/ s/mm 1 ms/line
ischemia and brain tumors, which are major public health care problems [10]. Other techniques are used for this purpose, such as magnetic resonance imaging (MRI) and ultrafast CT, but they have limited ability to precisely quantify the measurements. The limits of conventional scanners in measuring attenuation coe$cients in tissues are mainly due to characteristics of the X-ray tube itself: source size, intensity variations, limited #ux and broad energy spectrum. A scanner based on a synchrotron source, which provides monochromatic and parallel X-ray beams, leads to performance close to the theoretical limits and thereby opens new directions for medical imaging. After intravenous injection of a contrast agent, attenuation pro"les above and below the K-absorption edge are recorded over 3603 to reconstruct energy subtraction computed tomography images (KES-CT). This technique allows a precise and quantitative study of contrast agent distribution. Quantities such as blood volume and cerebral perfusion become accessible which in turn could lead to in vivo studies of angiogenesis development of tumors. Another technique called dual-photon absorptiometry (DPA-CT) [11,12] is under development to measure concentration changes of endogene components, such as potassium, which is a neurological signi"cant element. Two images taken at widely di!erent energies are subtracted, and the di!erence images map the low and intermediate-Z elements. The technical speci"cations of the ESRF scanner are given in Table 2.
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1 mm 150 mm 1440 350 lm 723/s 50 mGy 4;10 photons/s/mm 1 ms/projection
3. Beamline design The medical beamline is one of the so-called `long beamlinesa of the ESRF. The radiotherapy program is performed in a dedicated hutch inside the experimental hall, close to the source in order to obtain a maximum X-ray dose rate (see Fig. 1). The imaging programs take place in a satellite building (150 m from the wiggler source). 3.1. Source parameters The ESRF storage ring is operated at 6 GeV and 200 mA with a lifetime more than 50 hs in the 2/3 "lling mode (Table 3). At present, the ESRF has 30 public beamlines operational, and ten collaborating research group (CRG) lines are also available. The wiggler for the medical beamline has been optimized to ful"ll the requirements of the three main experimental programs. The wiggler "eld can reach 1.5 T to produce hard X-rays especially necessary for the radiation therapy program. Parameters of the multi-pole wiggler are given in Table 4. The wiggler will be operated at a "eld of 0.7 T (gap: 55 mm } critical energy: 16.5 keV) for the imaging program, which provides a good compromise between the available #ux and the contributions of the higher energy harmonics to the beam spectrum. 3.2. Design considerations The radiotherapy program and the imaging programs run only alternately. When the beamline is
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Fig. 1. ESRF medical beamline general view: MRT hutch and satellite building.
Table 3 European synchrotron radiation source Ring energy Circumference Horizontal emittance Vertical emittance Operating current Beam lifetime
Table 4 ID17 wiggler characteristics 6 GeV 844 m 4;10\ mrad 30;10\ mrad 200 mA 50 h
Field Period Number of poles Minimum gap Critical energy E K Total power emitted at 200 mA
1.5 T 150 mm 21 24 mm 33.5 keV 19.6 14.3 kW
Note: Numbers are given at maximum wiggler "eld.
set in the imaging mode, a connecting tube replaces the MRT equipment. The "rst instrumentation hutch in the experimental hall (MRT hutch) is connected to the imaging facility by a 100 m long tube in which the white beam radiation travels under vacuum (10\ Pa). The tube is shielded with 20 mm of lead along its path in the experimental hall, and is enclosed in a concrete tunnel between the experimental hall and the satellite building to avoid any radiation hazard. The tunnel walls are 300 mm thick and are made of special concrete
reinforced with hematite. The imaging facility consists of a monochromator hutch and an experimental hutch located in the satellite building. The monochromators have been installed in the satellite building rather than in the experimental hall to minimize the power density on the crystals and the e!ects of ground vibrations. The white beam reaching the imaging facility is a fan with maximum dimensions of 300 mm width
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Fig. 2. Instrumentation layout (angiography con"guration): (a) MRT hutch, (b) Monochromator hutch in the satellite building.
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Fig. 3. Patient positioning system and data acquisition system in the imaging room.
and 20 mm height. Two di!erent monochromators are used alternatively for coronary angiography and computed tomography. The single bounce angiography monochromator de#ects the monochromatic beam upwards (Fig. 2). The computed tomography monochromator is a double-crystal
"xed-exit system. In the imaging room, a patient positioning device is used to pre-align the patient relative to the beam and to perform scans during the imaging procedures (Fig. 3). In the KEDSA mode, the patient is scanned vertically through the beams, and in CT a rotation scan is performed to
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acquire the projections at all angles. A two-dimensional map of the attenuation coe$cient is then reconstructed as in a classical medical X-ray scanner. The data acquisition system is based on a intrinsic germanium detector followed by read-out electronics of a large dynamic range.
4. Beamline components The "rst hutch located inside the experimental hall serves as the microbeam radiation therapy (MRT) facility and houses several beamline components for the imaging programs. The primary slits are used to de"ne the dimensions of the beam vertically and horizontally. Two sets of "lters are available to limit the heat load on the monochromator crystal by "ltering the low energies of the X-ray spectrum; various thickness of carbon, aluminum or copper can be selected. Two "xed apertures have been designed to ensure that the X-ray beam will never hit the vacuum tube which connects the "rst instrumentation hutch in the experimental hall to the satellite building. They provide 2 mrad aperture in the horizontal plane and 0.14 mrad in the vertical plane. Preclinical MRT equipment, at 30 m from the source, includes a multislit collimator, fast shutters and an animal positioning system. In addition, a photon beam shutter located at the exit of the hutch isolates the rest of the beamline when the beamline is set in the MRT mode. When the beamline is set for the imaging programs, the MRT equipment is removed and a connecting tube is inserted to create a vacuum path up to the satellite building. The 250 m two-level satellite building o!ers an ambulatory patient care unit, which includes separate rooms for beamline optics, patient reception, #uoroscopy, imaging and beamline control. At the entrance of the monochromator hutch, a 500 lm thick beryllium window separates the vacuum tube and the secondary slits from the monochromator vessel which is under oxygen free helium #ow at atmospheric pressure. The radiation shielding of this hutch is realized with local shielding of the instrumentation (20 mm of lead) in addition to the 20 cm thick concrete walls.
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Table 5 Angiography monochromator Con"guration Energy range Crystal thickness Crystal height Crystal width Splitter height Asymmetry angle Typical bending radius
Single bent Laue crystal#splitter 17}51 keV 1.0 mm 20 mm 150 mm 3 mm 153 14 m
4.1. Angiography monochromator The monochromator is a single cylindrically bent silicon (1 1 1) crystal in Laue geometry, which focuses the beam vertically [13}15]. The central thin part of the crystal is 1.0 mm thick, 150 mm wide and 20 mm high, and the asymmetry angle is 153 (see Table 5). Two monochromatic beams are produced, with energies above (E ) and below (E ) the contrast material K-absorption edge, by blocking the central part of the exiting beam by a splitter. The total energy bandwidth re#ected by the bent Laue crystal is about 450 eV, overlapping the Kedge of the contrast material. The bandwidth can be adjusted by the vertical height of the beam. The splitter de"nes the energy separation of the two monochromatic beam. Typically, the two beams have an energy bandwidths of 150 eV, and their mean energies are separated by 300 eV. The assembly is cooled by a gravity-fed water system. The silicon crystal and the splitter can be driven out of the beam to allow unhindered passage of the beam through the monochromator vessel for other experiments. The monochromator can be tuned from 17 to 51 keV, which will allow the use of iodine in the initial phase or gadolinium in the future as contrast agents. At the iodine K edge (33.17 keV) the beams are de#ected upwards by 6.843 (Si (1 1 1) Bragg angle "3.423). The radius of curvature (typically 14 m) is adjusted to give a good cross-over and focus of the beams at the patient position 7 m from the monochromator. The monochromatic beam at this position is 2.25 m above the #oor. The two beams produced by the monochromator propagate at slightly di!erent angles. After crossing at the
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Table 6 Angiography monochromator performances Gap (mm) Intensity of one beam [photons/s/mm/100 mA] Harmonics/fundamental
40 5.9;10
50 4.6;10
60 3.1;10
70 1.8;10
3.6%
1.6%
0.33%
0.13%
Note: Vertical opening of the slits: 10 mm } Beam height at patient position: 700 lm.
Table 7 Computed tomography monochromator Con"guration Energy range Crystal thickness Crystal height Crystal width Asymmetry angle Typical bending radius
Two bent Laue crystals 17}80 keV 1.2 mm 20 mm 150 mm 153 30 m
can be adjusted to re#ect energies which are wide apart or bracket the K-edge of the contrast agent. One energy is selected, and the other one blocked by a mechanical chopper [17]. The imaging scan is performed within 5 seconds, and then fast energy switching from the low energy beam to the high energy beam is done by rotation of the chopper. The acquisition time for the complete procedure will thus amount to little more than 15 s. 4.3. Beam monitors and dose monitors
patient position, these beams diverge and are recorded separately by a dual-line detector. The available space between the monochromator and the detector is 13 m at maximum. Measure of the #ux and harmonic content are given in Table 6. The harmonics can impair the image quality by decreasing the contrast in the subtracted image [16]. The amount of the second harmonic at 99 keV relative to the fundamental energy at 33 keV is less than 1 percent when the wiggler gap is larger than 55 mm. 4.2. Tomography monochromator The monochromator is a tunable "xed-exit device based on two bent Si crystals in the Laue geometry, producing a monochromatic beam parallel to the incident white beam with a vertical o!set of 40 mm. The two bent Si (1 1 1) crystals are mounted on independent goniometers. The energy range is 30 to 80 keV. The bending and geometry of the Si crystals are identical to the angiography monochromator design (see Table 7). Provisions have been made to install another similar "xed exit monochromator in the same vessel. The energy range of the second monochromator will be from 50 to 100 keV. The two "xed-exit monochromators
4.3.1. Beam monitors The two X-ray beams, simultaneously produced by the angiography monochromator, are monitored by a dual beam ion chamber. This detector is made of a single chamber "lled with argon at atmospheric pressure. Two detection zones are de"ned by the electrode con"guration; the upper and lower electrodes are the collection plates and the middle electrode is the high voltage anode. The electrodes are 50 mm long in the beam direction. The ion chamber currents are ampli"ed (Novelec) and read out with VME cards (ICV150-ADAS). The acquisition system records simultaneously the monitor signals and the data from the germanium detector. The monitor values are used to normalize for beam intensities #uctuations. 4.3.2. Dose monitors The dose monitor is a single wide ion chamber, which detects both monochromatic fan beams at the same time, and is used to measure the X-ray dose imparted to the patient. Two similar dose monitors have been installed to provide the redundancy necessary for safety reasons. They are located one after the other inside the imaging room, and are the last beamline components before the
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patient. The ion chambers currents are converted to frequencies (Novelec), which are recorded by VME based counters (VCT6-ESRF). If the pre-set dose limit is reached, a signal is sent to the patient safety system which actuates the fast acting shutters to avoid any over exposure of the patient. In the CT mode, the same dual-beam and dose monitors are used. 4.4. Plexiglas xlters A set of Plexiglas "lters is available after the monochromators to adjust the beam intensity without changing the monochromators settings or the wiggler parameters. It is possible to vary the Plexiglas thickness from 2 up to 32 cm by steps of 2 cm, which provides attenuation factors from 5;10\ to 3;10\ at 33 keV. This device will be used for the pre-positioning scans at small X-ray doses, and to adjust the dose for the imaging scans, depending on the patient size.
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4.7. Medical equipment A conventional X-ray unit (Philips BVR 212) is available in a room close to the imaging area. It will be used if venous catheterization is required and to measure the transit time between the injection of the contrast agent in the arm and the arrival of the bolus in the patient's heart. Medical preparation of the patient will also be performed in this room. The intravenous infusion is performed with an injector (Medrad mark V) under remote control. The injection parameters (volume, rate, speed) are preset and the perfusion is triggered at the beginning of the imaging sequence. The data acquisition and patient scan occur after a delay calculated taking into account the transit time of the bolus. The X-ray exposure is triggered by the ECG in order to record the image during the slow motion phase of the heart cycle (diastole). During the imaging sequence, the physician has the possibility of staying close to the patient in a shielded area inside the imaging room.
4.5. Fast safety shutters Two independent fast-acting shutters are used in case of emergency to stop the beam. Each one is capable of stopping the full monochromatic beam; two are provided to satisfy the requirement for redundancy in the radiation safety system. The shutters are actuated simultaneously. The 6 mm thick shutter blades, made of tungsten, are kept out of the beam with springs compressed by electromagnets. If a problem is detected by the patient safety system (PASS), the electro-magnet power supply is switched o!, which releases the tungsten blades and stops the X-ray beam in about 10 ms. 4.6. Imaging shutters A fast shutter is used to limit the X-ray exposure of the patient to the image acquisition time. Two tungsten blades 1 mm thick, 160 mm wide and 30 mm high, separated by 12 mm, are mounted on each side of a rotation axis. The opening/closing rotation of 303 is driven separately by two motors, and the status of the shutter is detected by a pair of switches. The shutter opens or closes in less than 12 ms.
5. Patient positioning system Since the X-ray beam is a fan of about 0.5 mm height and 150 mm width, it is necessary to move the patient through the beam in order to obtain two-dimensional angiogram. For CT scans, attenuation pro"les are acquired over a complete rotation of the patient. The patient positioning system, which allows both for positioning of the patient and the scan motions during the data acquisition, is a high precision stage with seven degrees of freedom (Spretec). It has a very high rigidity during rotation or translation. This is especially important for tomography, since any eccentricity of the axis of rotation during the data acquisition will result in artifacts in the reconstructed image. 3D computed tomography is made possible by acquisition of a large number of images during helical motion of the patient. The design of the patient positioning system is shown Fig. 3. The U-shaped main frame is mounted on four screw jacks (Z ) which allow precise adjustment of the vertical axis (Z) perpendicular to the plane de"ned by the X-ray beam. The
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maximum vertical translation speed is 500 mm/s with a precision of $0.5%. The speed is constant over a length 200 mm (at maximum), for a total displacement of 600 mm including the acceleration and deceleration phases. The rigidity of the structure has been optimized to obtain the above speci"cations and at the same time to guarantee that the circle of confusion of the rotation axis is less than 400 lm in the CT mode. These requirements led to a structure of 4.5 t (total weight) and 2.6 t for the vertically moving part. The vertical movement is assisted by a pneumatic system (7 bar) which compensates for 95% of the total mass. A single brushless motor (5 kW) drives the stage by four screws linked by belts. In the CT mode, the maximum rotation speed is of 1803/s with a precision of $5%. In the angiography mode, after positioning at the desired incidence angle using two angular movements (h and u), and two translation movements (X and >), several images are acquired while the patient is scanned vertically (Z) up and down. For tomography the same system is used, but the acquisition is done during a rotation of the chair (u).
Ge crystal thickness is 2 mm, giving an e$ciency of nearly 100% at 33 keV and 45% at 99 keV (energy of the second harmonic). The major advantages of HPGe detectors over other types of detectors are their very good e$ciency together with a wide dynamic range. In germanium, the photoelectric e!ect is the dominant attenuation process, and the mean-free path of the photo-electrons is small (Table 9). These properties reduce the cross-talk e!ect, which was observed for example with silicon detectors [18]. The direct conversion from photons to electrons is another important asset. The current produced by absorbed X-rays is measured directly with no ampli"cation process. At low counting rate, su$cient output current is generated by two 33 keV photons (see Table 10), and the linearity is still excellent even at high counting rates. The 16-bit dynamic range of the electronics allows the measurement of signals over more than four decades. The spatial resolution of the system is determined by the vertical height of the beam and the 350 lm horizontal spacing of the elements on the detector.
Table 8 Medical beamline germanium detector
6. Data acquisition system
Total active area
6.1. Detector design In coronary angiography, the two X-ray fan beams are recorded by a dual-line detector synchronously with the vertical motion of the patient. A high-purity germanium detector (HPGe) operating at liquid nitrogen temperature is used (Table 8). It is made from a monolithic P-type Ge crystal 160 mm long, electrically segmented into two rows of 432 parallel strips each (Eurisys Mesures). The
Active thickness Number of elements Pixel width Pixel height Distance between 2 pixels Distance between the 2 rows Leakage current Beryllium window: thickness Minimum current Maximum current
20 mm (height), 150 mm (width) 2.0 mm 2 rows of 432 elements 300 lm 10 mm 50 lm 500 lm (5 pA/pixel 500 lm 10 pA/pixel 0.2 mA/pixel
Table 9 Linear attenuation coe$cients for total attenuation (l ), photoabsorption (l ), coherent scattering (l ), and incohrent scattering (l ) at 33.17 keV.
Germanium Silicon
l (cm\)
l
55.376 2.5546
44.539 1.9821
(cm\)
l (cm\)
l (cm\)
1.7241 0.2344
0.6356 0.3553
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Table 10 Equivalent full-scale X-ray counting rates and number of X-rays per digital output for 33 keV X-rays at di!erent gains of the readout electronics Gain
Capacity (F)
Full-scale current (A)
Number of X-rays in 1 ms at full scale
Number of X-rays per output count
1 2 4 8 16 32 64 128
1.60;10\ 1.60;10\ 1.60;10\ 1.60;10\ 1.00;10\ 1.00;10\ 1.00;10\ 1.00;10\
1.68;10\ 8.39;10\ 4.19;10\ 2.10;10\ 1.05;10\ 5.24;10\ 2.62;10\ 1.31;10\
9.47;10 4.73;10 2.37;10 1.18;10 5.92;10 2.96;10 1.48;10 7.40;10
144 72 36 18 9 4.5 2.25 1.13
In order to minimize the electronic pick-up noise, the "rst ampli"cation stage of the electronics is mounted as close as possible to the Ge crystal [19]. The signals are ampli"ed by ADC integrator boards fanned around the detector. A cryostat houses the Ge detector which works at liquidnitrogen temperature. The 30 l dewar above the detector is "lled automatically from a storage 120 l dewar. The detector is mounted on a support for vertical (Z) and horizontal alignment (>). The vertical position of the beam at the detector position varies from 1.44 up to 4.5 m, depending on the monochromator settings. In addition, the detector can be rotated to be perpendicular to the beam (see Fig. 3). 6.2. Electronics The detector is used in current integration mode where the current in each detector element is integrated and digitized every ms. The general architecture of the data acquisition electronics is based on the following scheme: E Preamplixer: The 864 channels are grouped by 16 on 54 integrator boards constituting the preampli"er stage. These boards are directly plugged into the same printed circuit on which the detector is bonded. The key component is the Burr-Brown low noise ACF2101 dual-switched integrator. Two integrating capacitors (one internal, one external) provide the main gain ranges, controlled by a single digital command.
A second digital command controls the integration and reset status. Both commands are common to all channels. Integration times can be programmed from 0.8 to 130 ms. The pre-ampli"ed voltages travel on twisted-pair #at cables (6 m long) to the digital stage. E Amplixer, AD conversion: The rapidly decreasing cost of ADC components allowed the replacement of the classical multiplexed ADC architecture with a `one ADC per channela solution of much easier implementation. Channels are still grouped by 16 and the resulting 54 ADC boards are distributed in 7 ADC crates, stacked in two racks on both sides of the detector inside the experimental hutch. Each channel has one programmable gain ampli"er (Burr-Brown PGA205) for further gain precision, and one 16 bit ADC (Burr-Brown ADS7807). According to the expected signal levels, the full-scale current of the ampli"er can be changed from 0.13 to 16.8 lA in 8 steps (see Table 10). The 128 signals from each ADC crate are read out from the back-plane by a crate controller and stored in a local bu!er. A 35 m serial line connects this bu!er with the VME data processors located in the control room. E Data processing: The ESRF data acquisition and control bus are based on VME technology. Two Motorola MVME162 processors download data through the serial lines from the ADC crate bu!ers and build up raw data lines in a large VME memory. Each MVME162 has four Industry Pack mezzanine locations available. Seven of
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those are "lled with an ESRF designed IP module for automatic read-out the ADC crate bu!er. The last free location is used for the sequencer which collects commands from the VME and other information from the experiment, such as the patient chair movements. The sequencer provides several programmable schemes depending on the application (tomography or angiography). This module has also been developed at the ESRF. The "nal link with the workstation is made through a BIT3 controller. The workstation performs the logarithmic subtraction from the two detector images, together with pedestal subtraction and gain normalization.
7. Radiation dose The injection of the contrast agent in the venous system rather than in the arterial system leads to a large dilution of the contrast agent in the blood circulation. The dilution factor lies typically between 30 and 40 [6]. After injection of a bolus of 45 ml of iodine (Hexabrix 320 mg/ml) in the superior vena cava, the concentration measured in the coronary arteries is about 10 mg/ml. The signal-to-noise ratio (SNR) in the subtracted images is in "rst approximation proportional to the contrast material concentration (c ) and varies as the square H root of the number of photons (N) (see Eq. (1)). It is therefore mandatory to increase the number of photons, and therefore the radiation dose, to compensate for the contrast agent dilution.
SNR"
*k c H o H
N 2
The absorbed dose (Gy) can be expressed as:
L k D"10*t E U (2) G o G G G Here *t is the X-ray exposure time (s), E the photon energy (J), the X-ray intensity [photons cm\ s\], k/o the mass absorption coe$cient for the tissue [cm g\], and i indicates the photon energy (fundamental and harmonics). KEDSA experiments on patients at Brookhaven and Hamburg [9,6] suggest that 3;10 photons/s/mm in front of the patient, in each beam, and a scan speed of 50 cm/s are needed for clinical relevant images of the distal parts of the coronary arteries without movement artifacts. These intensities correspond to a dose rate of 32 Gy/s. For a scan speed of 50 cm/s, a beam height of 0.5 mm and a line acquisition time of 1 ms, the resulting absorbed dose per monochromatic beam is 16 mGy or an absorbed dose of 32 mGy per subtracted image (Eq. (2)). At present, the #ux available at the medical imaging facility is a factor of 3 smaller than the requirements stated above, and the beam height at the patient position is 700 lm. However, the better experimental conditions at the ESRF compensate for the lower intensity. The beams are more parallel, and the larger depth-of-"eld reduces bone artifacts. The contrast is improved by the good e$ciency and small cross-talk of the detector. These allow for a lower scan speed (350 mm/s), longer sampling time and probably for higher harmonic content. 8. Results
(1)
where (*k/o) is the di!erence of the energy-depenH dent mass absorption coe$cients above and below the K-edge of the contrast agent. The limiting factor in intravenous coronary angiography is the X-ray dose delivered to the patient. The maximum X-ray dose will be 200 mGy for the complete procedure at the ESRF, which is comparable with the average dose received during a conventional coronary examination at the hospital.
8.1. Angiography The angiography system has been tested with several samples. A three-dimensional anthropomorphic arterial phantom [20] has been used to characterize the ability of the system to detect stenoses in the coronary arteries through various thickness of tissue simulating a patient examination. In vivo experiments have been performed on animal models. Four pigs weighing from 40 to
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45 kg were imaged. During this experiment, the wiggler gap was set at 60 mm and the ring current was about 170 mA. The monochromator produced two beams at 33.3 and 33.0 keV, above and below the iodine K-edge, with an energy bandwidth of 150 eV. The beam height at the pig position was 0.7 mm. The vertical speed was set at 250 mm/s with an integration time of 1.4 ms corresponding to a vertical displacement of 350 lm per line. The X-ray dose received by the animal was monitored all along the experiment, it amounted 28 mGy/ image in average. Each pig was anesthetized and ventilated with a respirator. The transit time between the injection of the contrast agent (Levovist, Schering) and the arrival of the bolus in the heart was measured with an ultrasound echograph. The catheter was inserted into the jugular vein and advanced to the superior vena cava under X-ray control. The pig was then placed horizontally on its belly on a support mounted on the patient positioning system.
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The heart rate and the blood oxygen saturation were monitored during the procedure. After the intravenous administration of a vaso-dilatator (Corvasal 1 mg/50ml) images were acquired in left anterior and oblique orientations. During imaging, the physician stood next to the animal in a shielded cabin. Thirty milliliters of iodine (Hexabrix威 320 mg/ml) were injected at a rate of 10 ml/s into the superior vena cava using an auto-injector (Mark V, Medrad). The image sequence was started few seconds after the injection of the contrast agent and ten images were acquired to follow the iodine bolus through the pig circulation, while the scanning stage was moving up and down. The time delay between two images was 1.3 s. The right coronary artery and the left anterior descending artery show up clearly on the images (Fig. 4). We also tested the possibility to inject the contrast agent without the use of a catheter in the peripheral circulation (in a vein of the ear). The quality of the resulting images is slightly inferior to those with injection into the superior vena cava, but the coronary arteries are still clearly visible. 8.2. Computed tomography
Fig. 4. K-edge intravenous coronary angiography of a 42 kg pig in LAO orientation. Thirty ml of iodine (Hexabrix 320) were injected within 2.6 s into the superior vena cava. The scan was performed 13 s after injection within 0.6 s with a chair speed of 250 mm/s and 1.4 ms integration time corresponding to a skin dose of 26.8 mGy/frame. The right coronary artery (RCA) and the left anterior descending artery (LAD) are clearly visible.
The goal of this experiment was to demonstrate the possibility of using monochromatic CT to image rat brain tumors in vivo and of precisely quantifying the amount of contrast agent in the tumor. Imaging was performed four weeks after implantation of the tumor and NMR images were acquired the day before the experiment at the synchrotron. The rat was anesthetized by means of an intraperitoneal injection (400 mg/kg of chloralhydrate) before being placed in the vertical stereotactic frame for KES-CT. The contrast agent was injected in the saphena vein. A solution of iodine (Hexabrix 320 mg/cm) was injected at a concentration of 2 ml/kg. Five tubes "lled with known iodine concentrations (5.00, 2.50, 1.25, 1.00 and 0.50 mg/cm) were taped very close to the rat brain, as a reference to cross-check the quanti"cation. For this experiment, a small motorized animal stage was used instead of the patient positioning system. The slice thickness, given by the beam height, was 800 lm and
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method to follow up an angiogenesis inhibitor by comparing the quantitative results obtained on untreated and treated rats bearing gliomas. Acknowledgements This work is the result of a large collective e!orts and we gratefully acknowledge the contribution of the ESRF sta!, in particular G. Blattman, G. Retout, E. Courraud, J.C. Laidet and the safety group. We thank A. Dilmanian, A Thompson, W. Grae! and R. Dix for many fruitful discussions during the design phase. We further acknowledge the contributions of J.L. Lefaix, J.J. Leplat (INRA/CEA) for their participation in the angiography experiment on the pigs. Fig. 5. Reconstructed SRCT image of a rat brain bearing a glioma obtained by logarithmic subtraction between images acquired above and below the iodine K-edge. Note the ringshaped contrast enhancement at the tumor location.
contiguous multi-slices of the brain were performed over 12 mm. The amount of iodine measured in the tumor itself ranges between 600 lg/cm and 750 lg/cm (see Fig. 5).
9. Perspectives The angiography medical program will start in Spring 1999. We intend to perform about ten patient examinations at the iodine K-edge. The patients selected for the KEDSA protocol will have previously undergone a conventional coronary angiography for comparison, to validate the synchrotron based technique. In parallel, we envisage to evaluate the possibility of using gadolinium as contrast agent for heavy patients. The CT program will focused on the development of the pre-clinical K-edge method in comparison with NMR. Quantitative kinetic studies of the contrast agent clearance in the tumor are expected to provide important information for understanding the blood brain barrier rupture and leakage phenomena. Fundamental studies will be "rst performed on rats bearing glioma. We also envisage using this
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