Download CdTe and CdZnTe detectors in nuclear medicine

Nuclear Instruments and Methods in Physics Research A 448 (2000) 513}524
CdTe and CdZnTe detectors in nuclear medicine
C. Scheiber*
Institut de Physique Biologique Faculte& de Me& decine, 4, rue Kirschleger, 67085 Strasbourg Cedex, France
Abstract
Nuclear medicine diagnostic applications are growing in search for more disease speci"c or more physiologically
relevant imaging. The data are obtained non-invasively from large "eld c cameras or from miniaturised probes. As far as
single photon emitters are concerned, often labelled with Tc (140 keV, c), nuclear instrumentation deals with poor
counting statistics due to the method of spatial localisation and low contrast to noise due to scatter in the body. Since the
1960s attempts have been made to replace the NaI scintillator by semiconductor detectors with better spectrometric
characteristics to improve contrast and quantitative measurements. They allow direct conversion of energy and thus
more compact sensors. Room-temperature semiconductor detectors such as cadmium tellure and cadmium zinc tellure
have favourable physical characteristics for medical applications which have been investigated in the 1980s. During one
decade, they have been used in miniaturised probes such as for inter-operative surgery guidance which is today in a fast
growing phase. This material su!ers from charge transport problems which has slowed down imaging applications.
Owing to a considerable research work on material, contacts and dedicated electronics small "eld of view compact
pixellated c cameras have been prototyped and one already marketed. Although extended clinical evaluation has to be
conducted and long-term reliability assesed, the available data already con"rm the expected gain in image contrast.
Medical interest for dedicated imaging systems is greater than it was in the 1980s when the "rst mobile c cameras were
marketed. The future of CdTe or CdZnTe-based imager for routine use now relies at "rst on industrial costs. 2000
Elsevier Science B.V. All rights reserved.
Keywords: Semiconductor detectors; CdTe; CZT; Nuclear camera; Imaging
1. Introduction
Nuclear medicine (NM) aims at functional imaging based on radiotracers. Although the 1990s were
marked by the rapid development of non-ionising
radiation imaging modalities (MRI, US) with high
soft tissue contrasts, submillimetric resolution and
capable of functional studies as well as morphology, there is still a fast growing demand for more
* Tel.: #33-3-88-14-48-56; fax: #33-3-88-37-14-97.
E-mail address: [email protected] (C. Scheiber).
physiological investigations and/or disease speci"c
investigations: the domain of NM expertise. The
poor (or even the lack of) anatomical representation in NM studies is not a handicap as multimodal imaging registration and fusion are now
clinically available. This is particularly noticeable
in neurology, cardiology and oncology, as new
highly speci"c tracers have been developed. Today,
detection characteristics are still one of the main
limiting factors of an even more rapid growth of
NM. In general, NM investigations used a stationary device, the c camera, but dedicated devices are
actively investigated. This paper will report that
0168-9002/00/$ - see front matter 2000 Elsevier Science B.V. All rights reserved.
PII: S 0 1 6 8 - 9 0 0 2 ( 0 0 ) 0 0 2 8 2 - 5
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C. Scheiber / Nuclear Instruments and Methods in Physics Research A 448 (2000) 513}524
CdTe and Cd
Zn Te (CZT) detectors are good
\V V
candidates for both approaches. The "eld of CZT
NM applications has been regularly reviewed
[1}5], we will report only the last technical developments, and, from the NM physician point of
view, actual trends in NM, which will naturally, or
already have, in#uenced the development of instrumentation. The literature has reported extensively
the physical characteristics of these semiconductor
detectors working at room temperature and their
dedicated electronics which will not be detailed here.
2. Nuclear medicine: general considerations
In NM, a c ray emitting radiotracer is injected
into the body. Then a device, usually a c camera, is
used to image its biodistribution as a function of
time or after a chosen delay after injection according to the metabolism of the drug. The basis of NM
is that the counting statistics is proportional to the
concentration of the radiopharmaceutical. As far as
the human body and disease's diagnostic or biological function measurements are concerned, this
leads to very di!erent situations regarding the
detection problem.
At "rst, single photon (SP) studies and positron
emission tomography (PET) studies have to be distinguished. PET makes use of coincident detection
of 511 keV positron-annihilation c rays. Clinical
applications of positron emission tracers are expanding. A!ordable clinical PET cameras have been
presented on the market and new c cameras can be
equipped with coincidence detection circuitry. The
characteristics of PET detector modules have been
updated recently [6], CdTe or CZT detectors are
currently being tested to "nd out if they can be useful
in the detection of 511 keV photons, yet their e$ciency is rather limited owing to the thickness of the
current material and photo memory. This modality
will not be discussed in detail here.
In SP studies the tracer is typically a biologically
relevant molecule labelled with Tc (140 keV,
c ray). Although labelling molecules with Tc is
di$cult, compared to I (360 keV, c and b) for
example, it has more favourable dosimetric characteristics (energy, physical half-life: 6 h) and thus
much higher activity (up to 1000 MBq) might be
injected intravenously while keeping the dose in an
acceptable limit (comparable to radiology) regarding the expected bene"ts for the patient. The
counterpart is a high attenuation coe$cient in the
body (0.12 cm\) and a high Compton fraction
which hampers data quanti"cation and lowers the
contrast. Spatial localisation of the primary
photon's source in SP studies is achieved by the use
of a collimator, basically a slab of lead absorber
containing long thin, typically parallel, bores. The
bore con"nes the detector's "eld of view to a narrow, forward cone at each position of its surface.
The contrast of the lesion is a decreasing function of
distance both due to geometrical considerations
and to other spatial super-impositions of various
biological sources that have physiologically, including circulating blood, or pathologically uptake
the tracer. Ampli"cation of the contrast is achieved
by the use of the computed tomography principle
(single photon emission computed tomography or
SPECT) but at the cost of an increased acquisition
time (typically 15}30 min). The second cause of
a low contrast is the presence of a large scatter
fraction in the body. Energy resolution is useful in
rejecting those scattered photons which seems to
originate from the location of their last scatter.
In some situations a single probe is su$cient or
a miniature device is requested, during surgery for
example, but usually an NM imaging system is
preferred. In both cases the detection system is
designed for the best compromise between the
spatial resolution, which is always rather poor
(6}8 mm at 10 cm from the detection plane) and the
counting statistics which is typically very low compared to other ionisation methods. In addition, the
acquisition time is limited. The size of the useful
"eld of view is controlled by the need to follow
simultaneously several areas of interest in the body
or to complete a whole body screening in an acceptable acquisition time (10}15 min). In some favourable situations the acquisition time is controlled by
the patient relative immobility and/or the device
occupational time but very often by the dynamics
of the physiological process of interest (cardiac
motion for example, 40 ms sampling). The compromise will be di!erent according to the clinical
problem. For example the localisation of a
sub-centimetric tumour deep in the body during
C. Scheiber / Nuclear Instruments and Methods in Physics Research A 448 (2000) 513}524
surgery requires a high sensitivity nuclear probe,
heart perfusion imaging a tomographic system
which can get close to the chest, while cardiac
ventriculography pass study "rst requires high sensitivity. Imaging a small super"cial organ like the
thyroid gland requires a system with a small "eld of
view system whereas it is mandatory to use a large
"eld double head for whole body scanning seeking
bone metastasis. The NM physician is looking for
the best spatial resolution and energy resolution
(5}7 keV at 140 keV) while keeping the e$ciency
(counting statistics) high in the available acquisition time. For practical reasons most of the NM
specialists still want to achieve these goals in the
NM department with a single device. However, this
is not always feasible nor desirable, then dedicated
systems are requested.
The clinical demand of NM investigations has
changed these last 10 years, according to several
factors which can be related to the disease (frequency, diagnostic problems, treatment possibilities, etc.), and of the e!ectiveness of NM compared
to others competitive medical imaging methods
(performances but also costs, availability, etc.). The
NM instrumentation has to adapt.
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Fig. 1. A two-head commercial c camera equipped with large
INa scintillator detectors. The detector is cumbersome and
required heavy shielding (courtesy of Siemens AG).
3. General purpose imaging system: c the camera
The 1999 c camera is a two-head SPECT system
(Figs. 1 and 2). Each head (Fig. 3) is equipped with
a large "eld NaI scintillator (40;50 cm). A thickness of 8 mm, a compromise between the e$ciency
and the intrinsic spatial resolution, is optimised for
the 140 keV photons. Each c-ray interacting in the
scintillator produces a #ash of light which is viewed
by an array of photo-multiplier tubes (PMTs).
Digital circuitry and other localisation schemes
might be used instead of the original Anger principle (1958) which consists of a network of calibrated resistances to make an estimate of the
interaction position of a c ray in the crystal from
the relative responses of the various PMTs. The
c-camera has an intrinsic resolution of &3 mm
full-width at half-maximum (FWHM) and a 6}8
mm FWHM at 10 cm with the so-called highresolution collimator. In the clinical conditions
the collimator controls both spatial resolution and
Fig. 2. Schematics of an Anger c camera. The photons emitted
by the medical target are spatially localised by the parallel
collimator (1,4). The scintillation light resulting from the interaction of a c ray in the NaI(Tl) crystal is viewed and localised by
a battery of photomultiplier tubes connected through a resistance network. A good energy resolution is required to reject the
scattered photons (3).
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C. Scheiber / Nuclear Instruments and Methods in Physics Research A 448 (2000) 513}524
Fig. 3. CdTe typical raw spectrum Co (122 keV, 136 keV),
although the energy resolution is improved compared to
INa(Tl), there is a signi"cant tail due to poor hole collection and
trapping.
sensitivity. The energy resolution is rather poor
(&11% FWHM), thus a signi"cant fraction of
scattered photons contributes to decrease the image contrast. The c-camera is designed for bone
whole-body scintigraphy (seeking bone metastasis,
40% of the NM studies) and cardiac SPECT perfusion studies (&25%). The equipment is cumbersome, partly due to the shielding of the detectors
and the PMTs. Room-temperature semiconductor
detectors have several advantages compared to
NaI scintillators: a direct conversion of energy with
the corresponding better energy resolution and less
cumbersome compared to the NaI-PMT assembly,
including less shielding.
4. CdTe and CZT as NM detectors
Among the range of solid detectors available for
c-ray detection, CdTe [7] and/or CZT [8,9] have
a privileged position. They have a high absorption
coe$cient (Cd: 48, Te: 52) and their wide band gap
(E "1.5}1.7 eV) allows stable counting at body
temperature, and their energy resolution is better
by a factor two as compared to NaI-based systems
(Fig. 3). Unfortunately, until recently, these characteristics were obtained for 2 mm thickness, which
result in a low e$ciency. Indeed, these materials
su!er from carrier recombination and trapping
which lead to incomplete charge collection. When
used in the pulse counting mode, the resulting
spectra exhibit low-energy `tailinga that reduces
the photopeak e$ciency. CdTe : Cl, grown by
THM, is now much better known and solutions
have been found to polarisation problems. HighPressure Bridgman CdTe or CZT have more favourable characteristics: larger volumes are obtained and higher resistivity (at least one order of
magnitude, 10 ) cm). On the negative side,
poorer hole collection as well low production yields
are observed (industrial costs). Further improvements have been obtained through extensive research on intrinsic properties of the material
[10,11]. Better hole collection has been achieved by
designing detectors and/or electrodes of various
geometry [12}14]. Compensation of charge trapping by signal processing is feasible at the cost of an
additional electronic circuitry [15]. However, it is
not clear that these methods would be applicable to
a large multi-element detector-array such as a routine NM clinical imager, as alternative, ohmic-contact detectors operating in the photo-conductive
mode are investigated [16]. Much of the published
recent research work was done on CZT but it is still
necessary to further improve the material quality to
exploit it fully [17].
5. Small-5eld pixellated CdTe or CZT c camera: an
obliged step toward the standard c-camera, or an
useful clinical tool?
Today, to the best of our knowledge, the
0.1}0.2 m active area per detector head for a
large-"eld c-camera will be still too expensive for
clinical systems. The "rst prototypes are small-"eld
c-cameras. Small "eld of view cameras are attractive: they can get closer to the patient, they are more
patient friendly and can be mounted on a mobile
cart to reach the operation room and/or the intensive care unit for cardiac and/or pulmonary studies.
Marketed in the 1980s, mobile, small-"eld Anger
c-cameras have received limited enthusiasm from
the medical community. This can be explained
partly by "nancial aspects, NM logistics and radioactive sources regulations. Nevertheless, in 1999
C. Scheiber / Nuclear Instruments and Methods in Physics Research A 448 (2000) 513}524
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Fig. 4. (A) The test object was plunged in a Tc solution and (B) imaged by a commercial HELIX camera (Elscint, Haifa), the 3 mm
sector was not resolved. (C) The "rst images obtained with comparable acquisition times by the BIOMEDII module (2.83;2.83 mm
CdTe detectors), the contrast is better for the 4 mm. (D) MoireH e!ect is seen on the 3 mm sector.
`imaginga patient might require a di!erent approach. An a!ordable high-performance small-"eld
dedicated system, could be attractive, for example,
for orthopaedics, endocrinology or paediatric services in numerous hospitals which cannot house an
NM service. Despite these commercial considerations and industrial issues, which remain to be
solved, important and rapid developments are
observed in this "eld. Several groups are working
in that perspective.
The "rst CdTe medical imaging system was presented in 1996 by Eisen et al. [18] who have developed a small-"eld (16;16 cm) camera equipped
with 40;32 CdTe : Cl detectors with Pt contacts.
Each single detector was attached to a low noise
preampli"er and an ampli"er/shaper. The pixel size
was 4;4 mm. Images from the heart and the thyroid have been presented. The performances of the
NUCAM were comparable to that of Anger
camera although with reduced e$ciency [19]. A
European Consortium (see dedicated paper this
conference) have developed a small-"eld cardiac
(BIOMED II project) based on the same approach
with a comparable "eld of view (15;15 cm) and an
&3 mm pitch. Acquisitions using a 256 pixels
module were obtained on test objects [20], they
look sharper when compared to that taken from
an Helix (formerly Elscint, now General Electric)
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C. Scheiber / Nuclear Instruments and Methods in Physics Research A 448 (2000) 513}524
Fig. 5. The Digirad 2020Tc imager (courtesy of Digirad). The
small "eld of view 20 cm detector can be placed close to the
patient for better geometric conditions. Real clinical interest
exists for low cost small-"eld camera: a challenge for CdTe or
CZT-based systems.
c camera (Fig. 4). Unfortunately, the crystal thickness of 2 mm and tailing result in a lower sensitivity
which could hamper cardiac dynamic studies.
St GOBAIN industrial Ceramics, US (Division
BICRON), CRISMATEC (France) and LETI
(CEA, France) are developing an NM device,
PEGASE, starting from a CdTe : Cl module (4;4
pixels) with a dedicated ASIC [21] developed in the
frame of the INTEGRAL program (CEA, SAGEM,
France). The device will be equipped with new
CdTe (high-pressure Bridgman) detectors and dedicated electronic signal processing circuitry [15].
In 1997 DIGIRAD Corporation (San Diego, CA,
US) has marketed the "rst small "eld of view c camera based on CZT detectors: the 2020tc Imager
(Fig. 5). The active area was 216;216 mm made of
64 modules 25;25;5 mm. Each module was
composed of a CZT wafer with 8;8 pixels
(3;3 mm) and an electronic shaping circuitry. The
resolution in energy of these detector equipped
with an original electrode (Spectrum Plus2+) was
8% (at 140 keV, but should attain 4%) [22]. Clinical images are available on their web site
(www.digirad.com). The system could even perform
heart SPECT studies by rotating the patient using
a rotating chair with respect to the camera head
plane.
Another approach consists in assembling several
sub-units of monolithic arrays of CdTe or CZT
detectors, coupled to MUX circuitry, adapted from
infra-red technology [23], through indium bumps.
Unlike the Anger camera, the intrinsic spatial resolution of a semiconductor array imager is determined by a well-de"ned element size; lower limits
on this size are, generally, established by the
readout circuitry and can be signi"cantly (1 mm.
The "rst image was presented in 1993 using
a 32;32 array of CZT detectors [24].
Another reason to decrease the pixel size is that
in multi-elements with pixel dimensions smaller
than the detector thickness, the deleterious e!ect of
hole trapping can be greatly reduced by taking
advantage of the `near "eld e!ecta [25]. The "rst
detector array of this kind was a 48;48 MUX
coupled to a CZT array [26]. The CZT slab
(29;2.9;0.3 mm) was partitioned into 2304 cells
(125 lm pitch) by photo-lithography and connected to the MUX for readout.
Most of the leader manufactures involved in NM
instrumentation have a semiconductor camera project. As an example, the ELGEMS (Haifa, Israel)
prototype, presented at the annual meeting of the
Society of Nuclear Medicine (June, 1999), is based
on IMARAD detectors. IMARAD is a young
Israeli company, which markets since 1999,
40;40;4 mm detector modules of CZT indium
doped detectors [17] with a 16;16 pixels (2.5 mm
pitch) resolution. They claimed a 5% energy resolution (140 keV) and a 70% e$ciency (13% energy
window). Images have been made with the same
resolution test object than for the BIOMED II
module and can be compared (Fig. 6). Clinical data
have been obtained as well. The camera is equipped
with a modi"ed collimator (6.8 cm FWHM at 10
cm). Siemens Med. Sys. Inc. (Ho!man Estates IL,
US), works on an R&D basis with IMARAD, to
develop a large "eld c-camera.
C. Scheiber / Nuclear Instruments and Methods in Physics Research A 448 (2000) 513}524
519
Fig. 6. (A) The same test object as in Fig. 4A and B the thyroid Picker phantom, imaged by the ELGEMS's CZT pixellated small "eld
imager prototype compared to the Anger c camera (C and D) showing an improved spatial resolution in these experimental conditions.
(Presented at the Annual Meeting of the Society of Nuclear Medicine, Los Angeles, June 1999; courtesy of A. Pansky.)
6. Other approaches to NM stationary devices
The spatial resolution/e$ciency compromise is
controlled by the collimator and even if it is feasible
to improve the intrinsic spatial resolution of the
detector it will be necessary to develop another
approach to localise the source to take bene"t from
it. One approach is illustrated by the ultra-highresolution brain SPECT system from the Division
of Nuclear Medicine of the University of Arizona.
A camera dedicated to a single organ is more challenging but justi"ed for brain function. In fact, there
is a rapid growing demand for brain perfusion
studies (psychiatric disease, epileptic foci localisation, etc.) owing to newly-developed blood-#ow
(Tc-HMPAO or Tc ECD) or brain-receptors
speci"c radio-pharmaceuticals.
It has been shown and demonstrated in simulations [27] that a combination of multiple pinhole
imaging system and a high-resolution detector can
result in substantial improvement in counting e$ciency over conventional systems. The original system was equipped with NaI detectors; a prototype
utilising CZT focal-plane arrays is under development [28] (Fig. 7). The system will use 256 modular
detectors in a full 3D-SPECT approach. This group
has developed an NM speci"c hybrid semiconductor detector array consisting of a 64;64 pixels
CZT detector array (1.5 mm thick, 380 lm square
pixel electrodes on one side produced by lithography) mated to a multiplexer readout circuit that
was custom designed for this NM application [29].
The cost of semiconductor detectors is still the
main limitant factor for the achievement of the
project.
Another completely di!erent approach is the
electronically collimated SPECT [30], also known
as the Compton camera. Counts in electronic
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C. Scheiber / Nuclear Instruments and Methods in Physics Research A 448 (2000) 513}524
Fig. 7. Artist view of the ultra-high-resolution brain imager
from the Tucson University (Arizona, US) (courtesy of HB
Barber).
collimation are acquired in a coincidence counting
mode between two positions and energy sensitive
detectors from those c rays which scatter from the
"rst detector onto the second after depositing
a measurable energy in the "rst detector. These
informations su$ce to con"ne the c rays origin to
a cone in space about the initial interaction position. The uncertainty of location perpendicular to
the cone's surface is related to the spatial resolution
and energy resolution of the detectors, the "rst
system used an array of Ge detectors as a "rst
detector and a conventional camera. Recently,
feasibility of using CdTe [31] and CZT [32] detectors has been demonstrated. Although the
CZT-based detector system e$ciency was a factor
of 2.5 lower than the Ge at 140 keV corresponding
to a system spatial resolution which is also a factor
of two lower, the di!erence will decrease at higher
energies. From the physics of Compton scattering it
is expected that this Compton camera will work
better at higher energies such as 511 keV (positron-annihilation c rays). The department of medical physics and Bioengineering at the UCL
(London, UK) has regarded semiconductor detectors to be used in their small prototype for the
Compton imaging of annihilation photons produced from positron emitters [33]. Although CdTe has
shown a better ratio of geometric e$ciencies compared to Si, the latter was chosen owing to a higher
Compton/photoelectric-peak ratio. Some possible
clinical implications of the markedly improved sen-
sitivity are: shorter imaging times (resulting in faster patient throughput, greater patient comfort, reduced motion blur); reduced patient/worker
radiation exposure; and `non-tumour-stunninga
diagnostic imaging work-ups for subsequent radiopharmaceutical cancer therapy. Additionally, the
proposed system will be completely stationary,
more compact than the bulky and heavy rotating
cameras in use.
Simultaneous measurement of X-ray computed
tomography (CT) and SPECT can be used for
structural and functional correlation and for improved quantitation of radionuclide uptake as
compared to SPECT. This is usually obtained by
acquiring transmission imaging using external
radioisotope sources. More detailed anatomical information can be obtained by acquiring a real CT
image. This can be achieved by coregistration of the
two data sets acquired on CT and SPECT systems,
respectively, or by the use of hybrid CT}SPECT
system, as the new ELGEMS's system which combines on the same gantry an X-ray digital CT
system and a dual-head c camera. One would prefer a single detector system such as a pixellated
CdZnTe array [34,35] designed to acquire both
transmission and emission data simultaneously
which is currently evaluated.
7. Miniaturisated probes and imaging systems
In at least two circumstances the NM device has
to be further miniaturised: ambulatory monitoring
of physiological functions and intra-operative
guided surgery. Recent clinical validation of the
sentinel-node concept has burst these applications.
7.1. Continuous monitoring of physiological
functions
In several pathological circumstances the symptoms of malfunctioning of a physiological function
are transient and di$cult to quantify or even to
recognise without multiple measurements. The
most well-known parameters are electrocardiogram, blood pressure, and/or blood glucose. NM
investigations allow precise measurements of
renal and cardiac function even in ambulatory
C. Scheiber / Nuclear Instruments and Methods in Physics Research A 448 (2000) 513}524
conditions. Some commercial systems have been
proposed, but till now, these applications have been
restricted to experimental works and clinical
research.
CdTe detectors have been used for over 15 years
for blood #ow measurements, or in the assessment
of the renal or cardiac function (for a review see
Refs. [1,3]). The latest developments concern several new CdTe-based multi-probe and/or multidetector systems. Continuous monitoring of the
cardiac ventricular function has been successful in
patients with coronary artery bypass grafting [36]
using a single CdTe-detector probe. The same system has been applied to the measurement of the
cardiac functional response to exercise in patients
with non-obstructive cardiomyopathy [37]. Ambulatory monitoring can be contemplated, using
a dedicated system based on a 16-CdTe detector
array (10;10;2 mm) [38] with associated portable autonomous electronics and a self-contained
data recording device [39].
7.2. Intra-operative probes
Nuclear probes have been used for a long time as
an aid to surgeon in the detection of small amounts
of (malignant) tissues [40]. Throughout the past 10
years of practice, radioguided surgery (RGS) applications have been extended to other pathologies
(for a review see Refs. [41}43]. Three main factors
are involved in the interest of using a probe during
surgery: the lesion (or all lesions) has not been
identi"ed and/or localised with su$cient precision
before the surgery by imaging (any modality), the
surgeon wants to be sure that all the malignant
tissue has been excised and/or, more recently, to
guide the surgeon to "nd involved lymph nodes for
lymph node dissection.
In the 1980s the main application was radioimmuno-guided-surgery (RIGS) [43] which focused on the high speci"city of the tracer to identify
a tumour of previously known origin. The patient is
injected prior to surgery with a targeting agent,
typically a monoclonal antibody labelled with
a low-energy c- or b-ray emitter. The ideal targeting agent binds speci"cally to malignant tissue
while the unbound fraction is gradually cleared
from the bloodstream and from normal tissues.
521
Colorectal cancer used to be the main target, now,
RIGS challenges traditional decision making
[44,45]. Besides primary tumour and metastasis,
the detection of tumour lymph-nodes (I MoAb
CC49) [46], prostate cancer (I B72.3) [47] or
gastrinomas (I somatostatin) [48] have been
made.
Lymphoscintigraphy (LS) aims at the mapping
of lymphatic routes using Tc- nanocolloids.
There was "rst a clinical boom at the end of the
1970s on the grounds of its use in the evaluation of
the internal mammary chain in breast cancer patients. Again LS becomes important today in oncology because of the validation of the `sentinel
node concepta [49,50]. Here LS is oriented to
identify the "rst appearing lymph node to receive
drainage from a primary tumour and the visualisation of an a!erent lumphatic vessel from the primary tumour to this node as the two major criteria
in identifying the sentinel node. Once identi"ed
a surgical biopsy technique can predict whether
cancer has spread into the surrounding lymph
nodes, this may result in a change in the surgical
plan. This staging procedure is now validated for
frequent cancer diseases such as melanoma and
breast cancer. Imaging with the c camera is performed after intra-dermal injection of the drug at
a location depending on anatomical lymphatic
routes draining the tumour site. The probe is used
prior to surgery to locate the site with the higher
count in the lymphatic basin and is marked as the
spot where the incision will be made.
The device consists of a hand-held probe which is
used as a homing device (Fig. 8). The probe must
have a high absolute sensitivity, a good side shielding, a high-energy resolution, good spatial localisation (angular sensitivity), ergonomic characteristics
and must be adapted to work in the operating
room (Fig. 9). The RGS probes have been compared based on basic physical characteristics [51]
seven commercially available probes (CdTe, CZT,
NaI, CdTe or CsI) showed the same relative performance in a perspex phantom. The CZT and NaI
based probes were preferred owing to their higher
sensitivity in the 300}400 keV range (I, 360 keV).
A more practical approach is to allow comparison
of probe performance in terms related to the task of
sentinel lymph node localisation in melanoma and
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C. Scheiber / Nuclear Instruments and Methods in Physics Research A 448 (2000) 513}524
environment. New probes devoted to laparoscopy
or endoscopy are under investigation.
LS combined with the gamma probe and/or blue
dye for adequate sentinel node biopsy will become
an important part of a clinical work in NM and
surgical oncology in the next few years [54]. As
stated by Ell [55] at the First International Congress on the Sentinel Node in the diagnosis and
treatment of cancer (Amsterdam, April 1999):
`there appear to be real added value in the application of this methodology in the investigation
but also in the management of large numbers of
patients with operable cancera.
8. Conclusion
Fig. 8. An example of a miniaturised probe used for radioguided
surgery.
Fig. 9. The RGS system is handled by the surgeon during the
surgery in the search for the lymph node(s) draining the primary
tumour directly: the sentinel node.
breast cancer [52] which used Tc tracer; the
best probe tested allows SLN localisation between
20}30 mm closer to the injection site than the
poorest performing probe. CdTe and CZT probes
currently available are single-detector systems, but
miniature multi-detector devices are being tested
such as 25-CdTe (2;2 mm) detector imaging
device (10;10;13 mm) [53] suitable for surgical
The clinical demand of improved systems is
growing as NM applications are expanding. Owing
to extensive research and industrial work conducted by teams of solid state physicists, engineers and
medical physicists CdTe/CZT detectors are now
widely accepted and used. There is still some further improvement of the material to be expected
which is mandatory in the perspective of commercial products at reasonable costs. CdTe/CZT based
surgery probes will have a large impact on patient
management in surgical oncology. Although a large
"eld c camera is still a dream, mainly due to the
cost of the multi-element detector array, small-"eld
cameras with improved energy resolution have
been developed. First, clinical images are promising. From the NM physician point of view, the
future of these systems might depend on both their
cost which should clearly be below that of a large"eld system and also on the capability of NM
investigations to spread to other environment
rather than the NM service where they will have
a large clinical use.
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