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New developments in clinical applications of CdTe and CdZnTe detectors
Nuclear Instruments
and Methods
in Physics
Research
A 380
( 1996) 385-391
NUCLEAN INSTNUMENTS & METNODS IN PNVSICS RESEARCH SectIonA
> __
New developments in clinical applications and CdZnTe detectors
of CdTe
C. Scheiber lnstitut
de Physique
Biologique
FucultP
de Midecine
4, rue Kirschlegrr
67085,
Strashourg
wdex.
Frumpy
Abstract This review about the medical applications of CdTe and CdZnTe is an update on the 1992 paper [I]. This new paper is legitimized by the recent progress which has been made in this field. First of all, the usefulness of a new material, i.e. CdZnTe, has been demonstrated. While the two materials are still being improved, it seems as yet too early to debate which of CdTe:Cl or CdZnTe will be the best choice. Historical applications span over the past I8 years, involving devices like miniature probes for per-operative scintigraphy or the monitoring of physiological functions and, closer to us, appliances dedicated to bone densitometry, and have been expanding as such devices have become commercially available, for many years now. Newly available microelectronic circuitry allows 2D-arrays to be built for digital quantitative X-ray (chest. dental . ) and for high-resolution gamma cameras. The clinical demand is very high, especially in the held of nuclear medicine. Although there already exist clinical demonstrators, the future of such CdTe applications depends on further reduction in material and device mounting costs. New perspectives concern XCT applications, but the data resulting from research work are kept for restricted use within industrial R&D laboratories.
1. Introduction The 90s were marked by the rapid development of non-ionizing radiation imaging modalities (MRI, US). Their medical applications have spread not only to the study of the morphology of soft tissues (formerly the exclusive domain of conventional or computed X-ray tomography) but also to vascular and functional imaging, which heretofore had been the specific domain of nuclear medicine (NM). In fact only 60% of medical investigations are now carried out using ionizing methods (IM). The future of ionising methods and further developments rest on their ability to adapt to this new situation, by improving their specificity (e.g. metabolism and quantitation for NM, full digitisation for radiological systems). This has to be achieved, with at least the same resolution, in a context of gradual yet growing pressure to lower patient dose. Sensors have been an essential element in this challenge. Among the range of solid detectors available for X- and gamma ray detection CdTe [2] and/or CdZnTe [3] have a privileged position, as far as quantitative studies are concerned. Such detectors have a high absorption coefficient (Cd: 48, Te: 52) and their wide band gap (E, = I..?1.7 eV) allows stable counting and good spectrometry at body temperature. However it is important to notice that, in this field of application, the sensor is not always the limiting factor. Each modality, or at least each medical application, has specific characteristics, which could require a specific detector. In some applications improving 0168.9002/96/$15.00 PI/
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the sensor could have a limited impact on the final image quality, in acceptable medical conditions, Both CdTe:CI and CdZnTe materials are still in rapid evolution and it is not the purpose of this review to debate which will be the material of choice. Today, CdZnTe seems to have the following advantages over CdTe:Cl - larger areas are attainable, as is a higher resistivity. by one order of magnitude. On the negative side, poorer hole collection as well as production yields are observed. CdTe:Cl is now much better known and solutions have been found to polarisation problems.
2. Radiology Sensors in conventional radiology (CR) are still made either by a film behind an intensifier screen or a light intensifier coupled to a CCD. CR is characterised by high spatial resolution (better than 20 lp/cm (more than 4 elements/mm’), and high photon fluxes. Large sensitive arrays (up to 50 X 50 cm’) of solid detectors are now commercially available based on amorphous silicon, or selenium [4] materials. The sensor is of limited bulk, flat, and consists of square arrays of pixels. Each pixel contains a light-sensitive photodiode and a TFT switch. The image acquisition process is similar to that of crystalline silicon photodiode arrays. Although this material fits the requirements for digital radiography it only operates in the current mode and quantitative radiography is not possible. The
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latter currently uses a copper filter that divides the film into two regions, or, as for digital chest arrays, two detector elements consisting of low and high atomic number phosphor screens [5]. Up until 1991, CdTe were only proposed to determine film exposition when conventional ionisation chambers were too cumbersome, as in mammography [6]. Since then, many other applications have been looked into but only the results of a few of these investigations have been published. In fact they are still studied by research laboratories and/or at the clinical experimental stage. In using CdTe in CR the following targets are aimed at: to obtain quantitative information (energy) from the X-ray beam transmitted, to reduce radiation exposure for patients while improving the resolution at low contrast by decreasing the scatter fraction, or to improve sensor reliability, and obtain a fully digital set-up. A CdTe imaging sensor is most suitable for digital radiography. The main reasons for this are as follows: first the simple structure of CdTe facilitates down-sizing and the creation of arrays, then high sensitivity and high S/N ratios are obtained by photon counting; in addition to that, the scattered radiation fraction can be reduced by using fan-beams and/or by being subtracted; finally, X-ray photons are directly converted into electrical signals and energy information relative to the incident X-ray photons can be obtained. Nevertheless, CdTe devices still have several drawbacks: each element requires an amplifier and counter, or sophisticated multiplexing circuits have to be used. The mounting of detectors is delicate; for high spatial resolution arrays, elaborate electronic circuits and highdensity mounting technology have to be used with subsequent important repercussions on the costs, reliability etc.
Radiographic images have been obtained by scanning with the imaging sensor around the focus of the X-ray tube (100 kV, 4 mA). The scanning pitch is 0.25 mm, using a 256 matrix (64 mm) at 5 mm/s. 1024 scanned lines produce simultaneously low- and high-energy radiographs of 256 X 1024 pixels. At 30% the modulation transfer function (MTF) is 2.5 Ip/mm along the scanning direction. The multi-channel imaging sensor is able to detect photon counts over an area of 0.5 X 0.2 mm. For these experiments Cl-doped CdTe:CI detectors were used. In a second prototype [I I], consisting of 90 elements ( I .8 X 2.0 X I .2 mm) a resolution (FWHM) of 4 keV has been achieved for 60 keV gamma rays; the count rate was found to vary by 4%. The charge collection efficiency was measured at 60%, at a 24 V bias (90% at 200 V). Recently, another research laboratory has developed an “X-ray quantum Radiography” system. The device consists of a fan-beam X-ray tube and a position sensitive CdTe array. The 512 X 8 CdTe elements were arranged at a pitch of 250 pm with a gap of 70 urn. The noise level of this detector was measured to be around 40 keV. It has been demonstrated that the detector has sufficient spatial resolution for a Compton scattering measurement system [ 121 (Fig. I ). The experimental device was tested on patients in clinical conditions; chest X-ray images have been published [ 131 (Fig. 2). In comparison with conventional film-based methods [ 141, their system offers clinically useful imaging data, especially of the lungs, in areas overlapping the heart or diaphragm. Chest X-ray is one of the most appealing fields of interest for quantitative imaging, but it remains difficult to achieve considering the large region of interest and important physiological motion. 2.2. High resolution dental digital radiography systems
2. I. Chest X-ray systems Further to a feasibility study in which an array [7] of 32 CdTe detectors was used, one research laboratory is developing a quantitative X-ray multi-detector system operating in the pulse counting mode [8]. The linear array consists of 256 CdTe with Pt electrodes (bias 100 V). Each CdTe has an effective detection volume of 1.0 X 0.25 X 0.3 mm’ (width X length X thickness). The average energy resolution is 5.1 keV for 60 keV gamma-rays and the count rate varies by less than 11% [9]. A spectral distortion has been observed and has been taken into account. The most significant causes for spectral distortion are that the energy depends, for photoelectric absorption, on the thickness of the detector, and that K-shell X-rays escape from the element, whereas the contribution of those generated by adjacent channels could be decreased by using tungsten shields [IO]. X-rays were separated into a low- and highenergy fraction using two pairs of discriminators and counters for each detector. The negative feedback circuit amplifier allows a counting circuit with a IO MHz range.
As far as small fields of view are concerned, high-spatial resolution matrices have been described for their use in dental digital radiography [ 151. The advantage of digital radiography is that the treat-
Fig. 1. Schematic diagram of the chest photon counting X-ray radiography system. The position sensitive detectors (32 CdTe detectors, 512 X 8 pixels, pitch = 250 pm, gap= 70 pm) was designed for Compton scattering measurements (reprint from Ref.
[ITI).
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sensor consists of a 400 X 600 array, 20 X 30 mm, with a 5i) pm pitch. Its spatial resolution has been described as being higher than that of existing systems using a scintillator screen coupled to a miniature CCD imaging system by optical fibers (17X26 mm’. 768X512 pixels. pitch: 40 pm. 12 Ip/mm) [17]. 2.3. Portal radiation therapy (PRTI
Fig. 2. Chest radiograph using the photon counting radiography system. Focal length=200 cm. tube voltage= 100 kV, tube current = 64 mA: sampling time = I. 18 ms/pixel.
ment is not delayed by film processing and, possibly, depending on the local context, it might reduce costs by suppressing the use of films; in addition Lo that, it is possible to decrease the patient dose which. with current silver halide-based film emulsions, is fairly high (l-3 mGy) due to low quantum efficiency (DQE=4%). Today, numerous market
digital
radiography
systems
are available
on the
PRT is another domain where CdTe detectors have been evaluated. Currently, new radiotherapy appliances can be equipped with real-time systems to produce portal images. Radiation therapy is an extremely important therapeutic modality in the cure of malignant tumours. One of the most important considerations is the accuracy with which the radiation dose can be delivered to the tumour. The most commonly used positioning system is the “simulator” equipped with high-resolution CR, and software to overlay a patient’s MRI or CT images. The target is delimited by skin markers. Portal imaging allows direct control of patient positioning under the radiotherapy beam in real-time. A linear array of 256 CdTe detectors was capable of generating images in a IO MV radiation therapy accelerator over an area of 40X40 cm’ in 5 s [ 181.Here, CdTe operates in the photovoltaic mode. In another series of experiments [ 191 the S/N ratio was measured at 143 with a resolution of 2 mm FWHM in less than 3 s. Among the advantages of the CdTe is its lower cost, compared to existing set-ups, and the compactness of the sensor and scanner mechanism, making it easy to retrofit to most existing linacs.
( IO 000 sold in Europe). The aim of that work was
to optimise the detection quantum efficiency without losing on spatial resolution. The experimental device is based on a 64X64 pixel CdTe:CI array with 100 km pitch [16] (Fig. 3). Each detector cell is connected to the input pad of a silicon chip via an indium-bump. The charges are integrated in a capacitance of each silicon pixel and sequentially read out after the 2 ms X-ray pulse. The medical
24.
Bone densitometry
Bone Densitometry is an important tool for the diagnosis and follow-up of osteoporosis. CdTe detector arrays are used in commercially available devices which operate in the planar [20] or tomographic mode [21] (see Ref. [I ] for details about methods and devices). The planar dual-energy system allows a CV better than 1% for bone mineral density measurements, in less than 15 s, using a 64 CdTe detector linear array. First-generation XCT devices equipped with 6 CdTe:CI have now been in clinical use for more than 5 years, and their reliability and measurement reproducibility have been demonstrated. 2.5. Cornputerised tomography
Fig. 3. Dry mandibule detector (exposure 0.007 nected to a 2D-electronic technique developed for module has been tested (reprint from Ref. [ 161).
image realized with CdTe prototype Gy). The CdTe:Cl (900 Fm) is conreadout circuit using the indium bumps infrared technology. The experimental with 64X64 pixels (pitch= 100 em)
(XCT)
The evolution of CT concerns the spiral mode, using multi-detector rings for vascular imaging at high temporal sampling. In this mode the detector ( 1024 Xe, CdW04 or ceramic detectors) and the X-ray tube operate continuously while a patient is being driven in the longitudinal direction (spiral acquisition). The use CdTe or CdZnTe in XCT applications is still at the experimental stage in most companies’ R&D laboratories, but few data have been
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published since 1992 [22]. One of the limitations to the use of CdTe appeared to be the rather long afterglow. Recent progress has been made in shortening the photodecay in CdTe X-ray detectors by modifying their structure and contacts [23]. A hybrid XCT and SPECT system has been developed using either Ge or CdTe detectors [24]. The advantages of this concept lie in the inherent coregistration of X-ray and NM images, i.e. accurate anatomic location of tracer distribution, and improved quantitative measurement by accurate correction of photon attenuation. The XCT operates in the third-generation fan beam geometry with an array of photon-counting detectors. The X-ray CT image was used to obtain an anatomically-accurate attenuation map from the heart of living pigs. This map was later used to reconstruct the SPECT image acquired with the same device, after injection of yymT~ Sesta-MIBI [25]. This correction improved the quantitative analysis of the regional myocardial blood flow.
3. Nuclear medicine Nuclear medicine instrumentation consists of conventional NM sensors: probes, gamma cameras and positron emission tomographs. Although miniature nuclear probes are still the only systems available on the market, new perspectives have been emerging. Single photon emission computed tomography (SPECT) gamma cameras are the most frequently used setup in NM and are now equipped with two or three heads. Among the new trends that have been observed in the field of NM clinical applications the following must be mentioned here: a growing interest for positron-emitter tracers, especially “F-fluorodeoxyglucose ( ‘8F-FDG), an increase in the number of brain studies owing to newlydeveloped blood flow radiopharmaceuticals (99mT~HMPAO or y9mT~ ECD), a growing demand for quantitative measurements involving attenuation correction and/or Compton contribution correction. Nuclear medicine physicians demand higher spatial resolution whilst keeping the counting statistics at an acceptable level, for a given irradiation dose. These factors make the new CdTe detectors more and more attractive for future use in gamma-cameras, owing to the recent availability of high performance infrared microelectronic circuitry. 3. I. Nuclear probes Though limited, the domain of their medical applications consists of unique NM procedures and, thus, represents an important potential field of development for this method. Miniaturised CdTe probes are used in two domains of application: the peroperative localisation of small (tumour) masses and the continuous monitoring of physiological functions.
3.2. Intra-operative
probes
There is a somewhat slow yet steady increase in the medical acceptance and use of such appliances as an aid to the surgeon in the detection of small amounts of malignant tissues. The device consists of a hand-held probe which is used as a homing device. The patient is injected prior to surgery with a targeting agent, typically a monoclonal antibody labelled with a low energy gamma- or beta-ray emitter. The ideal targeting agent binds specifically to malignant tissue while the unbound fraction is gradually cleared from the bloodstream and from normal tissues. The characteristics for an ideal probe and vector have been recently updated [26]. Colorectal cancer used to be the main target, but now, radioimmunoguided surgery challenges traditional decision making [27], but throughout the past 10 years of practice, medical applications have been extended to other pathologies (for a review see Refs. [28,29]). Today new applications are considered and new probes developed [30]. For example, recent works concern the detection of tumour lymphnodes (“‘1 MoAb CC49) [3 I], prostate cancer ( “‘1 B72.3) [32] or gastrinomas (“‘I somatostatin) [33]. The CdTe probes currently available (for these applications) are single-detector systems, but miniature multi-detector devices are being tested. A 25 CdTe detector imaging device (IO X 10X 13 cm) [34] suitable for surgical environment (FoV: 15 X I5 mm) has been developed. Each 2X2X 1 mm detector is connected to a specially designed Application-Specific Integrated Circuit, allowing the conception of miniature gamma cameras without multiplexer (MUX) read out. 3.3. Continuous
monitoring
of physiological
functions
CdTe detectors have been used for over 15 years for blood flow measurements, or in the assessment of the renal or cardiac function (for a review see Ref. [I]). The latest developments concern several new CdTe-based multiprobe and/or multidetector systems. The Ambulatory Renal Monitor (ARM) is one of the renal monitoring devices of the latest generation [35] (Fig. 4). The CdTe (16 mm in diameter, 2 mm thick), preamplifier and miniature data logger are embedded in a structure similar to a blood pressure cuff which is wrapped around the patient’s arm. The system is used to detect rapid changes in the glomerular renal function (99mTc DTPA) in patients at risk for acute renal failure during angiography, or in intensive care units, under non-invasive and near real-time conditions. A portable gamma spectrometry system based on CdTe or CdZnTe detectors ( 10 X 2 mm and 10 X 3 mm, respectively) has been used for simultaneous measurements of 9’mTc-platelets and ‘?-fibrin accumulation after arterial injury in rabbits [36] (Fig. 5). Continuous monitoring of the cardiac ventricular function has been successful in patients with coronary artery bypass grafting [37] using a single CdTe-detector probe. The same system has
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system allowing the processing of cardiac function data. A transcutaneous magnetic link allows power transfer in order to energise the electronic implant, and permits the telemetry of electrocardiogram and radionuclide data between the implant and the external receiver. 3.1.
m
Cd-Te DETECTOR
Fig. 4. Schematic representation showing the essential components of ARM. The diagram on the left shows the way in which the different components are arranged in the wall of instrument. The diagram on the right shows a PC displaying the rate constant versus time obtained after the on-line analysis of the data collected by the instrument (reprint from Ref. [35]).
been applied to the measurement of the cardiac functional response to exercise in patients with non obstructive cardiomyopathy [38]. Ambulatory monitoring can be contemplated, using a dedicated system based on a 16 CdTe detector array ( IO X 10X 2 mm) [39] with associated portable autonomous electronics and a self-contained data recording device [40]. As far as animal studies are concerned, CdTe detectors have been used to monitor the early myocardial clearance kinetics of yymT~ teboroxime in a canine model [41]. Recently, a cardiac implantable CdTe detector has been studied [42]. A system was suggested, composed of a cardiac implant, a miniature transmitter positioned externally on the chest and a remote receiver
Fig. 5. The portable gamma taneously monitor radiotracer used to improve the energy measuring the isotope lzzI simultaneously (reprint from
spectrometry system used to simulin vivo. The pulse processors were resolution of CdTe probes when (159 keV) and 99mTc (140 keV) Ref. [36]).
Imuging
devic,es
Despite commercial considerations and industrial issues which remain to be solved, important and rapid developments could be expected in this field. A large-field (60 X 40 cm) gamma camera with high energy (3% ) and spatial resolution (2 mm) remains a NM physician’s dream but one can expect small-field (20X 20 cm) CdTe/CdZnTe gamma cameras or dedicated imaging devices in the very near future, at least in experimental settings. Several groups are working in that perspective. Initially working with CdTe arrays for non-destructive control purposes [43]. Eisen et al. have recently been using such arrays for medical applications. towards the development of a smallfield ( I6 X 16 cm) CdTe gamma camera, devoted to heart and thyroid investigations (this conference). The first prototype was equipped with 40X 32 CdTe:CI detectors with Pt contacts. Each single detector was attached to a low noise preamplifier and an amplifier/shaper. The pixel size was 4X4 mm. Unlike in the Anger camera, the intrinsic spatial resolution of a semiconductor array imager is determined by a well-defined element size: lower limits on this size are. generally, established by the readout circuitry and can be significantly less than I mm. Another approach consists in assembling several subunits of large arrays of CdTe or CdZnTe detectors. coupled to MUX circuitry [44]. The first image was presented in 1993 using a 32X 32 array of CdZnTe detectors 14.51 (Fig. 6). The sub-arrays used nine 5X5X I.5 mm detectors in a 3X3 configuration with approximately 0.1 mm spacing [46]. A 48 X 48 MUX coupled to a Ge [47]. and more recently to a CdZnTe [48], array has been tested. The CdZnTe slab (2.9 cmX2.9 cmX 1.3 mm) was partitioned into 2304 cells (125 km pitch) by photolithography and connected to the MUX for readout using the so-called indium bump technique (Fig. 7). Surface resistivity was computed at IO”
Fig. 6. The first image of a thyroid phantom generated by the CdZnTe prototype imager (top). The image of the same phantom generated by a commercial nuclear medicine imaging (Anger) (bottom) (reprint from Ref. [45]).
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cameras could be hampered by the industrial costs of large field systems, compared to those of competitive sensors, taking into account the actual economic status of health systems. 3.5. Positron emission tomu~rup~y (PET)
Fig. 7. Conceptual schematic of a gamma ray detector which is indium bump bonded to a silicon multiplexer.
uhmecm, and for the 500 pm test pixel from the monolithic array, the energy resolution, for 24’Am, was 3.2 keV and 2.3 keV (FWHM) for the photopeak and pulser peak, respectively, and 3.8 keV and 2.7 keV for “Co (122 keV). Such high-resolution CdZnTe arrays could be assembled to build a small field (20X20 cm) gamma camera. Alternatively, these modules could be used to build dedicated NM devices such as the ultra-high-resolution SPECT system for brain imaging, developed by Barber et al. [49]. The system uses a multiple pinhole aperture and 2.56 modular detectors in a full 3D-SPECT approach (Fig. 8). Single photon detection in NM requires collimation which controls the spatial resolution and collection efficiency. As for any other collimators, reducing the pinhole diameter improves the spatial resolution but reduces the sensitive area and hence collection efficiency. This group has demonstrated that this trade-off could, in fact, be overcome if detectors with very high spatial resolution were available [SO]. Although the demand from the NM physician community is very high for a high-resolution gamma camera, the promising future of CdTe or CdZNTe-based gamma
Clinical applications of positron emission tracers arc expanding. Affordable clinical PET cameras have been on the market and new gamma cameras can be equipped with coincidence detection circuitry. The characteristics of PET detector modules have been updated recently [5 I], CdTe or CdZnTe detectors are currently being tested to find out if they can be useful in the detection of 5 I 1 keV photons, yet their efficiency is rather limited owing to the useful thickness of the current material.
4. Conclusion This update on medical applications since 1991 has highlighted the fact that both CdTe:Cl and CdZnTe material are now being used. Some of the advantages of CdZnTe over CdTe:Cl are the higher resistivity and larger areas which have been obtained, but with lower yields at the present time. CdTe(C1) has a much longer development history. The material costs and/or usefulness of 2D arrays by photolithography on a single large crystal will be the determining factor. Although for the time being no devices dedicated to medical applications, other than miniature probes and bone densitometer, are found on the market, it is still encouraging to note that they are expanding. Moreover, new clinical applications have emerged, such as digital X-ray (chest, dental . . . ). The most striking developments concern 2D arrays for NM gamma-cameras. Although an experimental device is at the clinical experimental stage, there remain several problems which will have to be solved before industrial production can be contempiated, material and mounting costs being among the most significant. New prospects concern CTX applications which are currently being investigated in research laboratories.
References
Fig. 8. Schematic drawing of a gamma ray imager with multiplexer readout for use in ultra high resolution brain SPECT. A hemispheric multiple pin-hole collimator is surrounded by a tiled layer of 256 CdTe detector modules (reprint from Ref. 1491).
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[30] E. Martin, D.A. Martinez, M.S. O’Dorisio, T.M. O’Dorisio. D.R. King, R.C. Orahood, G.H. Hinkle. J.A. Barnes and M.O. Thurston, in: Radioimmunoguided Surgery in the Detection and Treatment of Colorectal Cancer, Ed. E.W. Martins (Landes Company. 1994) 172. [31] P.L. Triozzi. J.A. Kim, W. Aldrich, D.C. Young, J.W. Sampsel and E.W. Martin. Cancer 73 ( 1993) 580. [32] R.A. Badalament. J.K. Burgers, L.R. Petty, C.M. Mojzisik and A. Berens. Cancer 71 ( 19Y3) 226X. [33] E.A. Wotering, R. Barrie, T. O’Dorisio. MS. O’Dorisio. R. Nance and D.M. Cook, Surgery II6 ( 1994) I l3Y. [34] H. Bedoui. J. Rousseau, P. Masqueher and X. Marchandise, Proc. IEEE-EMSBS Conf. Montreal ( 199s ) 34. [3S] C.A. Rabito. R.H. Moore, C. Bouga\ and C. Dragotakes. JNM 34 (1993) 199. [36] M. Lorentr, J. Van Ryn. H. Merk. M. Market? and W.G. Eisert, Nucl. Instr. and Meth. A 353 ( 1994) 448. [37] J. Taki, A. Muramori. K. Nakajima, H. Bunko, M. Kawasuji. N. Tonami and K. Hisada, JNM 33 ( 1992) 441. [38] J. Taki, K. Nakajima. M. Shimizu. N. Tonami and K. Hisada. JNM 35 (1994) 1937. [39] C. Scheiber, M. Karim, C. Drdman. P. Attali, P. Siffert and J. Chambron. Proc. IEEE-EMBS 14th annual Conf., Paris, 29-31 Oct. 1992. 707. [40] K. Karim. C. Scheiber, C. Draman. P. Attali. P. Siffert and J. Chambron. Proc. IEEE-EMBS 14th annual Conf.. Paris. ‘9-31 Oct. 1992. 711. [4l] G. Johnson. D.K. Glover. C.B. Hebert and R.D. Okada. JNM 34 ( I993 ) 630. [42] Z. Hamici and R. Itti, Phys. Med. Biol. 40 (199.5) 609. 1431 Y. Eisen and E. Polak. Proc. of Mat. Re\. Sot. Symp. 302 (1993) 487. [44] F.L. Augustine, Nucl. Instr. and Meth. A 353 ( 1994) 201. [45] J.F. Butler. S.J. Friesenhahn. C. Lingren. B. Apotovsky. R. Simchon and F.P. Doty, Proc. IEEE Trans. Nucl. Sci.. San Francisco. Cal., Nov. 2-6 (1993) 345. [46] J.F. Butler, S.J. Friesenhahn, C. Lingren, B. Apotovaky and F.P. Doty, SPIE 1896 (1993) 30. [47] H.B. Barber, F.L. Augustine, H.H. Barett. E.L. Dereniak. K.L. Matherson. T.J. Meyers, D.L. Pery. J.E. Venzon. J.M. Woolfenden and E.T. Young. Nucl. Instr.and Meth. A 353 (1994) 361. [48] F.P. Doty, H.B. Barber, F.L. Augustine. J.F. Butler, B.A. Apotovaky. E.T. Young and W. Hamilton. Nucl. Instr. and Meth. A 353 ( 1994) 356. [49] H.B. Barber. H.H. Barett, E.L. Dereniak, N.E. Hartsough. D.L. Perry, P.C.T. Roberts, M.M. Rogulski, J.M. Woolfenden and E.T. Young, IEEE Trans. Nucl. Sci. NS-40 ( 1993) I 140. [50] M.M. Rogulski. H.B. Barber, H.H. Barett. R.L. Shoemaker and J.M. Woolfenden. IEEE Trans. Nucl. Sci. NS-40 ( 1993) 1123. [51] W.W. Moses, S.E. Derenzo and T.F. Budinger. Nucl. Instr. and Meth. A 3.53 (1994) 189.
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NUCLEAN INSTNUMENTS & METNODS IN PNVSICS RESEARCH SectIonA
> __
New developments in clinical applications and CdZnTe detectors
of CdTe
C. Scheiber lnstitut
de Physique
Biologique
FucultP
de Midecine
4, rue Kirschlegrr
67085,
Strashourg
wdex.
Frumpy
Abstract This review about the medical applications of CdTe and CdZnTe is an update on the 1992 paper [I]. This new paper is legitimized by the recent progress which has been made in this field. First of all, the usefulness of a new material, i.e. CdZnTe, has been demonstrated. While the two materials are still being improved, it seems as yet too early to debate which of CdTe:Cl or CdZnTe will be the best choice. Historical applications span over the past I8 years, involving devices like miniature probes for per-operative scintigraphy or the monitoring of physiological functions and, closer to us, appliances dedicated to bone densitometry, and have been expanding as such devices have become commercially available, for many years now. Newly available microelectronic circuitry allows 2D-arrays to be built for digital quantitative X-ray (chest. dental . ) and for high-resolution gamma cameras. The clinical demand is very high, especially in the held of nuclear medicine. Although there already exist clinical demonstrators, the future of such CdTe applications depends on further reduction in material and device mounting costs. New perspectives concern XCT applications, but the data resulting from research work are kept for restricted use within industrial R&D laboratories.
1. Introduction The 90s were marked by the rapid development of non-ionizing radiation imaging modalities (MRI, US). Their medical applications have spread not only to the study of the morphology of soft tissues (formerly the exclusive domain of conventional or computed X-ray tomography) but also to vascular and functional imaging, which heretofore had been the specific domain of nuclear medicine (NM). In fact only 60% of medical investigations are now carried out using ionizing methods (IM). The future of ionising methods and further developments rest on their ability to adapt to this new situation, by improving their specificity (e.g. metabolism and quantitation for NM, full digitisation for radiological systems). This has to be achieved, with at least the same resolution, in a context of gradual yet growing pressure to lower patient dose. Sensors have been an essential element in this challenge. Among the range of solid detectors available for X- and gamma ray detection CdTe [2] and/or CdZnTe [3] have a privileged position, as far as quantitative studies are concerned. Such detectors have a high absorption coefficient (Cd: 48, Te: 52) and their wide band gap (E, = I..?1.7 eV) allows stable counting and good spectrometry at body temperature. However it is important to notice that, in this field of application, the sensor is not always the limiting factor. Each modality, or at least each medical application, has specific characteristics, which could require a specific detector. In some applications improving 0168.9002/96/$15.00 PI/
SOl68-9002(96)00310-5
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the sensor could have a limited impact on the final image quality, in acceptable medical conditions, Both CdTe:CI and CdZnTe materials are still in rapid evolution and it is not the purpose of this review to debate which will be the material of choice. Today, CdZnTe seems to have the following advantages over CdTe:Cl - larger areas are attainable, as is a higher resistivity. by one order of magnitude. On the negative side, poorer hole collection as well as production yields are observed. CdTe:Cl is now much better known and solutions have been found to polarisation problems.
2. Radiology Sensors in conventional radiology (CR) are still made either by a film behind an intensifier screen or a light intensifier coupled to a CCD. CR is characterised by high spatial resolution (better than 20 lp/cm (more than 4 elements/mm’), and high photon fluxes. Large sensitive arrays (up to 50 X 50 cm’) of solid detectors are now commercially available based on amorphous silicon, or selenium [4] materials. The sensor is of limited bulk, flat, and consists of square arrays of pixels. Each pixel contains a light-sensitive photodiode and a TFT switch. The image acquisition process is similar to that of crystalline silicon photodiode arrays. Although this material fits the requirements for digital radiography it only operates in the current mode and quantitative radiography is not possible. The
Elsevier Science B.V. All rights reserved IV(a).
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latter currently uses a copper filter that divides the film into two regions, or, as for digital chest arrays, two detector elements consisting of low and high atomic number phosphor screens [5]. Up until 1991, CdTe were only proposed to determine film exposition when conventional ionisation chambers were too cumbersome, as in mammography [6]. Since then, many other applications have been looked into but only the results of a few of these investigations have been published. In fact they are still studied by research laboratories and/or at the clinical experimental stage. In using CdTe in CR the following targets are aimed at: to obtain quantitative information (energy) from the X-ray beam transmitted, to reduce radiation exposure for patients while improving the resolution at low contrast by decreasing the scatter fraction, or to improve sensor reliability, and obtain a fully digital set-up. A CdTe imaging sensor is most suitable for digital radiography. The main reasons for this are as follows: first the simple structure of CdTe facilitates down-sizing and the creation of arrays, then high sensitivity and high S/N ratios are obtained by photon counting; in addition to that, the scattered radiation fraction can be reduced by using fan-beams and/or by being subtracted; finally, X-ray photons are directly converted into electrical signals and energy information relative to the incident X-ray photons can be obtained. Nevertheless, CdTe devices still have several drawbacks: each element requires an amplifier and counter, or sophisticated multiplexing circuits have to be used. The mounting of detectors is delicate; for high spatial resolution arrays, elaborate electronic circuits and highdensity mounting technology have to be used with subsequent important repercussions on the costs, reliability etc.
Radiographic images have been obtained by scanning with the imaging sensor around the focus of the X-ray tube (100 kV, 4 mA). The scanning pitch is 0.25 mm, using a 256 matrix (64 mm) at 5 mm/s. 1024 scanned lines produce simultaneously low- and high-energy radiographs of 256 X 1024 pixels. At 30% the modulation transfer function (MTF) is 2.5 Ip/mm along the scanning direction. The multi-channel imaging sensor is able to detect photon counts over an area of 0.5 X 0.2 mm. For these experiments Cl-doped CdTe:CI detectors were used. In a second prototype [I I], consisting of 90 elements ( I .8 X 2.0 X I .2 mm) a resolution (FWHM) of 4 keV has been achieved for 60 keV gamma rays; the count rate was found to vary by 4%. The charge collection efficiency was measured at 60%, at a 24 V bias (90% at 200 V). Recently, another research laboratory has developed an “X-ray quantum Radiography” system. The device consists of a fan-beam X-ray tube and a position sensitive CdTe array. The 512 X 8 CdTe elements were arranged at a pitch of 250 pm with a gap of 70 urn. The noise level of this detector was measured to be around 40 keV. It has been demonstrated that the detector has sufficient spatial resolution for a Compton scattering measurement system [ 121 (Fig. I ). The experimental device was tested on patients in clinical conditions; chest X-ray images have been published [ 131 (Fig. 2). In comparison with conventional film-based methods [ 141, their system offers clinically useful imaging data, especially of the lungs, in areas overlapping the heart or diaphragm. Chest X-ray is one of the most appealing fields of interest for quantitative imaging, but it remains difficult to achieve considering the large region of interest and important physiological motion. 2.2. High resolution dental digital radiography systems
2. I. Chest X-ray systems Further to a feasibility study in which an array [7] of 32 CdTe detectors was used, one research laboratory is developing a quantitative X-ray multi-detector system operating in the pulse counting mode [8]. The linear array consists of 256 CdTe with Pt electrodes (bias 100 V). Each CdTe has an effective detection volume of 1.0 X 0.25 X 0.3 mm’ (width X length X thickness). The average energy resolution is 5.1 keV for 60 keV gamma-rays and the count rate varies by less than 11% [9]. A spectral distortion has been observed and has been taken into account. The most significant causes for spectral distortion are that the energy depends, for photoelectric absorption, on the thickness of the detector, and that K-shell X-rays escape from the element, whereas the contribution of those generated by adjacent channels could be decreased by using tungsten shields [IO]. X-rays were separated into a low- and highenergy fraction using two pairs of discriminators and counters for each detector. The negative feedback circuit amplifier allows a counting circuit with a IO MHz range.
As far as small fields of view are concerned, high-spatial resolution matrices have been described for their use in dental digital radiography [ 151. The advantage of digital radiography is that the treat-
Fig. 1. Schematic diagram of the chest photon counting X-ray radiography system. The position sensitive detectors (32 CdTe detectors, 512 X 8 pixels, pitch = 250 pm, gap= 70 pm) was designed for Compton scattering measurements (reprint from Ref.
[ITI).
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sensor consists of a 400 X 600 array, 20 X 30 mm, with a 5i) pm pitch. Its spatial resolution has been described as being higher than that of existing systems using a scintillator screen coupled to a miniature CCD imaging system by optical fibers (17X26 mm’. 768X512 pixels. pitch: 40 pm. 12 Ip/mm) [17]. 2.3. Portal radiation therapy (PRTI
Fig. 2. Chest radiograph using the photon counting radiography system. Focal length=200 cm. tube voltage= 100 kV, tube current = 64 mA: sampling time = I. 18 ms/pixel.
ment is not delayed by film processing and, possibly, depending on the local context, it might reduce costs by suppressing the use of films; in addition Lo that, it is possible to decrease the patient dose which. with current silver halide-based film emulsions, is fairly high (l-3 mGy) due to low quantum efficiency (DQE=4%). Today, numerous market
digital
radiography
systems
are available
on the
PRT is another domain where CdTe detectors have been evaluated. Currently, new radiotherapy appliances can be equipped with real-time systems to produce portal images. Radiation therapy is an extremely important therapeutic modality in the cure of malignant tumours. One of the most important considerations is the accuracy with which the radiation dose can be delivered to the tumour. The most commonly used positioning system is the “simulator” equipped with high-resolution CR, and software to overlay a patient’s MRI or CT images. The target is delimited by skin markers. Portal imaging allows direct control of patient positioning under the radiotherapy beam in real-time. A linear array of 256 CdTe detectors was capable of generating images in a IO MV radiation therapy accelerator over an area of 40X40 cm’ in 5 s [ 181.Here, CdTe operates in the photovoltaic mode. In another series of experiments [ 191 the S/N ratio was measured at 143 with a resolution of 2 mm FWHM in less than 3 s. Among the advantages of the CdTe is its lower cost, compared to existing set-ups, and the compactness of the sensor and scanner mechanism, making it easy to retrofit to most existing linacs.
( IO 000 sold in Europe). The aim of that work was
to optimise the detection quantum efficiency without losing on spatial resolution. The experimental device is based on a 64X64 pixel CdTe:CI array with 100 km pitch [16] (Fig. 3). Each detector cell is connected to the input pad of a silicon chip via an indium-bump. The charges are integrated in a capacitance of each silicon pixel and sequentially read out after the 2 ms X-ray pulse. The medical
24.
Bone densitometry
Bone Densitometry is an important tool for the diagnosis and follow-up of osteoporosis. CdTe detector arrays are used in commercially available devices which operate in the planar [20] or tomographic mode [21] (see Ref. [I ] for details about methods and devices). The planar dual-energy system allows a CV better than 1% for bone mineral density measurements, in less than 15 s, using a 64 CdTe detector linear array. First-generation XCT devices equipped with 6 CdTe:CI have now been in clinical use for more than 5 years, and their reliability and measurement reproducibility have been demonstrated. 2.5. Cornputerised tomography
Fig. 3. Dry mandibule detector (exposure 0.007 nected to a 2D-electronic technique developed for module has been tested (reprint from Ref. [ 161).
image realized with CdTe prototype Gy). The CdTe:Cl (900 Fm) is conreadout circuit using the indium bumps infrared technology. The experimental with 64X64 pixels (pitch= 100 em)
(XCT)
The evolution of CT concerns the spiral mode, using multi-detector rings for vascular imaging at high temporal sampling. In this mode the detector ( 1024 Xe, CdW04 or ceramic detectors) and the X-ray tube operate continuously while a patient is being driven in the longitudinal direction (spiral acquisition). The use CdTe or CdZnTe in XCT applications is still at the experimental stage in most companies’ R&D laboratories, but few data have been
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published since 1992 [22]. One of the limitations to the use of CdTe appeared to be the rather long afterglow. Recent progress has been made in shortening the photodecay in CdTe X-ray detectors by modifying their structure and contacts [23]. A hybrid XCT and SPECT system has been developed using either Ge or CdTe detectors [24]. The advantages of this concept lie in the inherent coregistration of X-ray and NM images, i.e. accurate anatomic location of tracer distribution, and improved quantitative measurement by accurate correction of photon attenuation. The XCT operates in the third-generation fan beam geometry with an array of photon-counting detectors. The X-ray CT image was used to obtain an anatomically-accurate attenuation map from the heart of living pigs. This map was later used to reconstruct the SPECT image acquired with the same device, after injection of yymT~ Sesta-MIBI [25]. This correction improved the quantitative analysis of the regional myocardial blood flow.
3. Nuclear medicine Nuclear medicine instrumentation consists of conventional NM sensors: probes, gamma cameras and positron emission tomographs. Although miniature nuclear probes are still the only systems available on the market, new perspectives have been emerging. Single photon emission computed tomography (SPECT) gamma cameras are the most frequently used setup in NM and are now equipped with two or three heads. Among the new trends that have been observed in the field of NM clinical applications the following must be mentioned here: a growing interest for positron-emitter tracers, especially “F-fluorodeoxyglucose ( ‘8F-FDG), an increase in the number of brain studies owing to newlydeveloped blood flow radiopharmaceuticals (99mT~HMPAO or y9mT~ ECD), a growing demand for quantitative measurements involving attenuation correction and/or Compton contribution correction. Nuclear medicine physicians demand higher spatial resolution whilst keeping the counting statistics at an acceptable level, for a given irradiation dose. These factors make the new CdTe detectors more and more attractive for future use in gamma-cameras, owing to the recent availability of high performance infrared microelectronic circuitry. 3. I. Nuclear probes Though limited, the domain of their medical applications consists of unique NM procedures and, thus, represents an important potential field of development for this method. Miniaturised CdTe probes are used in two domains of application: the peroperative localisation of small (tumour) masses and the continuous monitoring of physiological functions.
3.2. Intra-operative
probes
There is a somewhat slow yet steady increase in the medical acceptance and use of such appliances as an aid to the surgeon in the detection of small amounts of malignant tissues. The device consists of a hand-held probe which is used as a homing device. The patient is injected prior to surgery with a targeting agent, typically a monoclonal antibody labelled with a low energy gamma- or beta-ray emitter. The ideal targeting agent binds specifically to malignant tissue while the unbound fraction is gradually cleared from the bloodstream and from normal tissues. The characteristics for an ideal probe and vector have been recently updated [26]. Colorectal cancer used to be the main target, but now, radioimmunoguided surgery challenges traditional decision making [27], but throughout the past 10 years of practice, medical applications have been extended to other pathologies (for a review see Refs. [28,29]). Today new applications are considered and new probes developed [30]. For example, recent works concern the detection of tumour lymphnodes (“‘1 MoAb CC49) [3 I], prostate cancer ( “‘1 B72.3) [32] or gastrinomas (“‘I somatostatin) [33]. The CdTe probes currently available (for these applications) are single-detector systems, but miniature multi-detector devices are being tested. A 25 CdTe detector imaging device (IO X 10X 13 cm) [34] suitable for surgical environment (FoV: 15 X I5 mm) has been developed. Each 2X2X 1 mm detector is connected to a specially designed Application-Specific Integrated Circuit, allowing the conception of miniature gamma cameras without multiplexer (MUX) read out. 3.3. Continuous
monitoring
of physiological
functions
CdTe detectors have been used for over 15 years for blood flow measurements, or in the assessment of the renal or cardiac function (for a review see Ref. [I]). The latest developments concern several new CdTe-based multiprobe and/or multidetector systems. The Ambulatory Renal Monitor (ARM) is one of the renal monitoring devices of the latest generation [35] (Fig. 4). The CdTe (16 mm in diameter, 2 mm thick), preamplifier and miniature data logger are embedded in a structure similar to a blood pressure cuff which is wrapped around the patient’s arm. The system is used to detect rapid changes in the glomerular renal function (99mTc DTPA) in patients at risk for acute renal failure during angiography, or in intensive care units, under non-invasive and near real-time conditions. A portable gamma spectrometry system based on CdTe or CdZnTe detectors ( 10 X 2 mm and 10 X 3 mm, respectively) has been used for simultaneous measurements of 9’mTc-platelets and ‘?-fibrin accumulation after arterial injury in rabbits [36] (Fig. 5). Continuous monitoring of the cardiac ventricular function has been successful in patients with coronary artery bypass grafting [37] using a single CdTe-detector probe. The same system has
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system allowing the processing of cardiac function data. A transcutaneous magnetic link allows power transfer in order to energise the electronic implant, and permits the telemetry of electrocardiogram and radionuclide data between the implant and the external receiver. 3.1.
m
Cd-Te DETECTOR
Fig. 4. Schematic representation showing the essential components of ARM. The diagram on the left shows the way in which the different components are arranged in the wall of instrument. The diagram on the right shows a PC displaying the rate constant versus time obtained after the on-line analysis of the data collected by the instrument (reprint from Ref. [35]).
been applied to the measurement of the cardiac functional response to exercise in patients with non obstructive cardiomyopathy [38]. Ambulatory monitoring can be contemplated, using a dedicated system based on a 16 CdTe detector array ( IO X 10X 2 mm) [39] with associated portable autonomous electronics and a self-contained data recording device [40]. As far as animal studies are concerned, CdTe detectors have been used to monitor the early myocardial clearance kinetics of yymT~ teboroxime in a canine model [41]. Recently, a cardiac implantable CdTe detector has been studied [42]. A system was suggested, composed of a cardiac implant, a miniature transmitter positioned externally on the chest and a remote receiver
Fig. 5. The portable gamma taneously monitor radiotracer used to improve the energy measuring the isotope lzzI simultaneously (reprint from
spectrometry system used to simulin vivo. The pulse processors were resolution of CdTe probes when (159 keV) and 99mTc (140 keV) Ref. [36]).
Imuging
devic,es
Despite commercial considerations and industrial issues which remain to be solved, important and rapid developments could be expected in this field. A large-field (60 X 40 cm) gamma camera with high energy (3% ) and spatial resolution (2 mm) remains a NM physician’s dream but one can expect small-field (20X 20 cm) CdTe/CdZnTe gamma cameras or dedicated imaging devices in the very near future, at least in experimental settings. Several groups are working in that perspective. Initially working with CdTe arrays for non-destructive control purposes [43]. Eisen et al. have recently been using such arrays for medical applications. towards the development of a smallfield ( I6 X 16 cm) CdTe gamma camera, devoted to heart and thyroid investigations (this conference). The first prototype was equipped with 40X 32 CdTe:CI detectors with Pt contacts. Each single detector was attached to a low noise preamplifier and an amplifier/shaper. The pixel size was 4X4 mm. Unlike in the Anger camera, the intrinsic spatial resolution of a semiconductor array imager is determined by a well-defined element size: lower limits on this size are. generally, established by the readout circuitry and can be significantly less than I mm. Another approach consists in assembling several subunits of large arrays of CdTe or CdZnTe detectors. coupled to MUX circuitry [44]. The first image was presented in 1993 using a 32X 32 array of CdZnTe detectors 14.51 (Fig. 6). The sub-arrays used nine 5X5X I.5 mm detectors in a 3X3 configuration with approximately 0.1 mm spacing [46]. A 48 X 48 MUX coupled to a Ge [47]. and more recently to a CdZnTe [48], array has been tested. The CdZnTe slab (2.9 cmX2.9 cmX 1.3 mm) was partitioned into 2304 cells (125 km pitch) by photolithography and connected to the MUX for readout using the so-called indium bump technique (Fig. 7). Surface resistivity was computed at IO”
Fig. 6. The first image of a thyroid phantom generated by the CdZnTe prototype imager (top). The image of the same phantom generated by a commercial nuclear medicine imaging (Anger) (bottom) (reprint from Ref. [45]).
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Incident Gamma Radlatlon
cameras could be hampered by the industrial costs of large field systems, compared to those of competitive sensors, taking into account the actual economic status of health systems. 3.5. Positron emission tomu~rup~y (PET)
Fig. 7. Conceptual schematic of a gamma ray detector which is indium bump bonded to a silicon multiplexer.
uhmecm, and for the 500 pm test pixel from the monolithic array, the energy resolution, for 24’Am, was 3.2 keV and 2.3 keV (FWHM) for the photopeak and pulser peak, respectively, and 3.8 keV and 2.7 keV for “Co (122 keV). Such high-resolution CdZnTe arrays could be assembled to build a small field (20X20 cm) gamma camera. Alternatively, these modules could be used to build dedicated NM devices such as the ultra-high-resolution SPECT system for brain imaging, developed by Barber et al. [49]. The system uses a multiple pinhole aperture and 2.56 modular detectors in a full 3D-SPECT approach (Fig. 8). Single photon detection in NM requires collimation which controls the spatial resolution and collection efficiency. As for any other collimators, reducing the pinhole diameter improves the spatial resolution but reduces the sensitive area and hence collection efficiency. This group has demonstrated that this trade-off could, in fact, be overcome if detectors with very high spatial resolution were available [SO]. Although the demand from the NM physician community is very high for a high-resolution gamma camera, the promising future of CdTe or CdZNTe-based gamma
Clinical applications of positron emission tracers arc expanding. Affordable clinical PET cameras have been on the market and new gamma cameras can be equipped with coincidence detection circuitry. The characteristics of PET detector modules have been updated recently [5 I], CdTe or CdZnTe detectors are currently being tested to find out if they can be useful in the detection of 5 I 1 keV photons, yet their efficiency is rather limited owing to the useful thickness of the current material.
4. Conclusion This update on medical applications since 1991 has highlighted the fact that both CdTe:Cl and CdZnTe material are now being used. Some of the advantages of CdZnTe over CdTe:Cl are the higher resistivity and larger areas which have been obtained, but with lower yields at the present time. CdTe(C1) has a much longer development history. The material costs and/or usefulness of 2D arrays by photolithography on a single large crystal will be the determining factor. Although for the time being no devices dedicated to medical applications, other than miniature probes and bone densitometer, are found on the market, it is still encouraging to note that they are expanding. Moreover, new clinical applications have emerged, such as digital X-ray (chest, dental . . . ). The most striking developments concern 2D arrays for NM gamma-cameras. Although an experimental device is at the clinical experimental stage, there remain several problems which will have to be solved before industrial production can be contempiated, material and mounting costs being among the most significant. New prospects concern CTX applications which are currently being investigated in research laboratories.
References
Fig. 8. Schematic drawing of a gamma ray imager with multiplexer readout for use in ultra high resolution brain SPECT. A hemispheric multiple pin-hole collimator is surrounded by a tiled layer of 256 CdTe detector modules (reprint from Ref. 1491).
[l] C. Scheiber and 3. Chambron, Nucl. Ins&. and Meth. A 322 (1992) 604. 1’21P. Siffert, Nucl. Instr. and Meth. A 253 (1978) 1. [3] F.P. Doty, J.F. Butler, J. Schetina and K. Bowers, Proc. 1991 U.S. Workshop on the Physics and Chemistry of HgCdTe and Other II-VI Compounds, ed. D.E. Seiler, J. Vat. Sci. Technol. B 10 (1992) 1418. [4] H.G. Chotas, C.E. Floyd and C.E. Ravin, Radiology 19.5 (1995) 264.
C. Scheiher
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IV(a).
MEDICAL
APPLICATIONS