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Principles of operation Of Scanning Systems, Gamma Camera
single-detector
and
multidetector
stationary and mobile gammacameras (formerly called Anger or scintillationcameras).
Most of the systems listed are capable of single photon emission
computed tomography(SPECT), also called single photon emission
tomography,and some are capable of dual-head coincidence
imaging with F-18 fluorodeoxyglucose (FDG), a radiopharmaceutical used in
positron emission tomography (PET) imaging. For more information on
PET, see the Product Comparison titled SCANNING SYSTEMS, POSITRON
EMISSION TOMOGRAPHY.
These devices are also called: nuclear imaging gantries, SPECT systems, SPET
systems,stationary gamma cameras, whole-body gamma camera systems.
UMDNS information
This Product Comparison covers the following device terms and product codes as listed in ECRI’s UniversalMedical Device Nomenclature
System™
(UMDNS™):
Scanning Systems, Gamma Camera, Mobile
[16-891]
Scanning Systems, Gamma Camera, Planar
Imaging [16-892]
Scanning Systems, Gamma Camera, Single
Photon Emission Tomography [18-444]
Purpose
Gamma cameras are used to produce images of the
radiation generated by radiopharmaceuticals within a
patient’s body in order to examine organ anatomy and
function and to visualize bone abnormalities.
The wide variety of radiopharmaceuticals and procedures used allows evaluation
of almost every organ system. In addition to producing a conventional planar
image (a two-dimensional [2-D] image of the three-dimensional [3-D]
radiopharmaceutical distribution within a patient’s body), most stationary gamma
camera systems can also produce whole-body images (single head-totoe skeletal
profiles) and tomographic images (crosssectional slices of the body acquired at
various angles around the patient and displayed as a computer-reconstructed
image).
SPECT is most commonly used for whole-body bone
imaging, brain-perfusion studies, and cardiac imaging; 30%of SPECT procedures
are cardiac studies. Through sequential image acquisition, the gamma camera can
image blood flow to various organs, including the brain, lungs, liver, kidneys, and
bones. It also helps physicians detect and identify lesions, such as cysts, tumors,
hematomas, and infarcted tissue, as well as areas of altered osteogenesis and
abnormalities of the cortex and white matter. In addition, the gamma camera can
work in tandem with a computer to evaluate cardiac function and perfusion — for
example, SPECT gamma cameras can perform myocardial perfusion imaging
with thallium-201 and technetium-99m.
SPECT is also used to detect femoral head avascular necrosis, knee osteoarthritis,
metastatic liver disease, temporomandibular joint abnormalities, and deep-seated
small hemangiomas, as well as to assess bone metabolism in hyperparathyroidism
and thyrotoxicosis. Such techniques reduce the need for interventional
radiography, thereby circumventing its associated morbidity. Brain SPECT is
being used in the prognosis of strokes, acquired immunodeficiency syndrome
(AIDS) dementia complex, psychiatric diseases, and Parkinson’s disease.
One study indicates that FDG-SPECT is as effective as PET in detecting yocardial
viability and diagnosing certain malignant tumors (Martin et al. 1995).
Coincidence imaging is useful for certain neurologic, oncologic, and cardiac
applications. FDG tomography performed in coincidence mode has been shown to
be successful in detecting occult primary tumors in head and neck carcinoma and
useful in guiding endoscopic biopsies (see Périé et al. 2000).
Mobile gamma camera images facilitate the assessment of cardiac function and
perfusion in patients with impending myocardial infarction (MI), as well as in
those who have suffered acute MI. Bedside evaluation of these and other critically
ill patients greatly reduces the need to transport them by stretcher to a stationary
gamma camera system.
Principles of operation
The gamma camera detects and counts photons emanating from a target organ and
maps individual scintillation events in a spatial configuration that creates an image
of the organ.
Static images display data acquired at a specific point during an exam, and
dynamic images display a change in data measurements over time.
A gamma camera system is composed of a collimator, a thallium-activated
sodium iodide (NaI[Tl]) crystal detector, photomultiplier tubes (PMTs),
electronic circuitry to determine the location and magnitude of scintillation
events, an imaging omputer, and an operator console. An integral computer
and/or a separate image-acquisition, processing, and display workstation is also
used.
Whole-body imaging requires either a track-mounted movable detector
thatpasses over the patient or a patient table that moves beneath a stationary
detector.
SPECT systems require a mechanical gantry to support and rotate the
camera head and collimators in a circular, body-contour, or elliptical orbit.
Noncircular orbits allow the camera head to be closer to the body, thereby
improving spatial resolution.
Two energy-matter interactions are important to conventional
gamma camera imaging:
the photoelectric effect and Compton scattering.
In photoelectric interactions, an incident (incoming) photon with slightly
more energy than the binding energy of a k-shell electron encounters one of these
electrons and ejects it from its orbit; because all its energy is imparted to the orbital
electron, the photon is absorbed.
The ejected photoelectron possesses kinetic energy equal to the energy from the
incident photon minus the energy required to eject the electron from its orbit. The
resultant vacancy in the k-shell is filled by an l- or m-shell electron, which emits
energy in the form of an x-ray photon. The energy of radiation produced by the
movement of electrons within an atom is characteristic of each element and is
therefore called characteristic radiation.
Compton scattering results from a collision between a high-energy incident
photon and a loosely held outershell electron. The incident photon transfers some
of its energy to the electron, which is ejected from its orbit by the collision.
Because incident photons cannot transfer all their energy to the orbiting electron,
Compton scattering always produces an ion pair — a positive ion and the ejected
negative electron (called a recoil electron) — and always results in the formation
of a scatter photon. An incident photon frequently initiates a chain of Compton
reactions and photoelectric
absorption events, which result in the sequential degradation of photon energy.
Because gamma photons cannot be focused using lenses, as light can, a
collimator is used to selectively absorb unwanted radiation; only photons traveling
along the desired path are allowed to pass through to the detector. The collimator is
usually made of a heavy-metal absorber such as lead, with some tungsten or
platinum parts. The basic types used in conventional gamma camera imaging are
pinhole, parallel-hole, diverging, and converging collimators.
The pinhole collimator, which works much like a pinhole camera, is a lead cone
with a small aperture at the tip. Gamma rays passing through the pinhole
produce an inverted image that can be magnified or minified, depending on the
length of the cone and the distance of the organ from the aperture. Pinhole
collimators are best suited for magnification imaging of small, thin structures, such
as the thyroid.
Most have a removable aperture insert that allows changes in aperture size; a
smaller aperture produces sharper images but also reduces sensitivity and increases
imaging time.
The parallel-hole collimator, which is the most widely used, is a piece of lead up
to a few inches thick containing many parallel holes perpendicular to the
collimator surface. The projected image is the same size as the source distribution
onto the detector.
Gamma rays leaving the organ almost perpendicular to the collimator face pass
through to the detector; all other rays are absorbed by the walls (septa) of the
collimator holes. The use of high-energy radionuclides requires thicker septa to
absorb unwanted photons and to keep photons from crossing from one hole to the
next; however, thicker septa are not as efficient because they allow fewer photons
to pass.
Collimators used specifically with low-energy radionuclides have lead foil
septa that are only a few tenths of a millimeter thick and thus are very fragile.
Hole length and diameter also affect performance:
collimators with long, narrow holes provide better resolution but sacrifice
efficiency.
Septal materials with high atomic numbers and high density provide the best
results. Lead is by far the most popular material because of its cost and
availability, although tungsten, tantalum, and gold have some limited research
applications.
For maximum versatility, gamma cameras usually come equipped with several
parallel-hole collimators, including a low-energy all-purpose (LEAP) collimator
for imaging photons of up to 150 keV, as well as low-energy high-resolution
(LEHR) and medium-energy all-purpose (MEAP) collimators for imaging
photons of up to 1 MeV.
The diverging collimator has angled holes that diverge from a point 40 to 50 cm
behind the collimator.
A minified image of source distribution is projected onto the detector. Particularly
useful when imaging large organs with a standard field-of-view (FOV) detector
(e.g., lung scanning with a portable gamma camera),
the diverging collimator effectively increases the
diameter of the detector field of view by approximately
one-third.
The converging collimator has angled holes that converge at a point 40 to 50 cm
in front of the collimator.The image is magnified but not inverted, provided
that the organ is between the collimator face and the convergence point. At the
convergence point, images are reduced; beyond it, they are magnified but inverted.
Some gamma cameras have a single collimator with a removable center insert that
allows both diverging and converging collimation. Specialty collimators, such as
seven-pinhole, rotating slant-hole, fan beam, and coded-aperture collimators, are
also available; most are used primarily for tomographic cardiac
imaging.
The collimator, projects radiation from the organ to be imaged onto the NaI(Tl)
crystal, which converts incoming gamma ray photons into visible light energy.
The scintillation process involves a series of Compton collisions in the NaI(Tl)
crystal, each producing a scattered photon of lesser energy and a Compton recoil
electron that excites the NaI(Tl) electrons in its path and causes them to scintillate
(produce a flash of light)
at an intensity proportional to the energy of the incident photon. The scattered
photon reacts with another crystal atom, produces another scattered photon and
recoil electron, and causes more scintillations until the photons lose enough energy
to be photoelectrically absorbed.
Lower-energy photons undergo fewer interactions before absorption and
produce fewer scintillations.
Because most scintillations occur in the front part of the detector, thin crystals
provide better resolution by bringing the light flashes closer to the PMTs.
However, thin crystals allow more incident photons to pass through without being
absorbed; therefore, the number of scintillations is reduced. The crystals of
most units are 9.5mm(3/8 in) thick; however, cameras equipped for coincidence
imaging have thicker crystals, typically 15.9 mm (5/8 in) thick. Crystal dimensions
range from 25 25 cm (10 10 in) to 52 64 cm (201/2 25 in). Because sodium
iodide (NaI) absorbs water, a hermetically sealed aluminum housing covers
the sides and front of the crystal. The back is sealed by
a clear Lucite light pipe or is optically coupled directly
to the face of the PMTs.
The light pulse created by the incident photon is converted into an electrical signal
of quantifiable magnitude by the PMT array, which can be composed of 37
to 150 PMTs arranged hexagonally (although several manufacturers use
rectangular arrays).
Each PMT has a preamplifier, a simple circuit that allows the PMT to be tuned so
that each yields the same output for a given scintillation intensity, ensuring
uniform detector performance throughout the entire field of view. Several
cameras have an automatic tuning option that electronically balances PMT output
from a single control on the operator console.
The light photons strike the photocathode in the PMT and form photoelectrons
that are then directed through a series of 10 to 12 dynodes, which boost the
signal. The output is sent to a position-encoding circuit,which determines the twodimensional location of the scintillation event and encodes this position as four
signals: x, x-, y, and y-. These signals are combined to form two signals that are
transmitted to a summation amplifier. All the light pulses viewed by the PMTs are
summed into one pulse, which is transmitted to a pulse height analyzer (PHA) that
accepts only those pulses within a predetermined range of energies. Pulses
accepted by the PHA are transmitted to the cathode ray
tube (CRT), and the electron gun, turned on for a few
microseconds, passes a beam through deflector plates
to be guided to coordinates on the CRT screen that
match the actual scintillation coordinates in the crystal.
Mobile gamma cameras
In mobile gamma cameras, the system components are configured in one of two
ways:
In one configuration, the detector and wheeled detector stand are separate
from the data processing console, which is also mounted on wheels; each
component is manually pushed to the patient’s bedside and interconnected by
coaxial or fiberoptic cable. In another configuration,the detector, detector stand,
and data processing console are integrated into a single, motor-driven, wheeled
unit powered by rechargeable batteries.
Either a chain drive or a friction wheel mechanism delivers power to
the system’s wheels. Images stored by these systems
can be transferred to a workstation via floppy disk or
Ethernet connection at a later time.
The principles of operation and image acquisition
for mobile cameras are identical to those for stationary
models.
SPECT
Apart from some basic models and those intended only for whole-body studies,
most stationary and some mobile gamma cameras can perform SPECT, a nuclear
medicine technique used to create a 3-D representation of the distribution of an
administered radiopharmaceutical.
SPECT cameras detect only radionuclides that produce a cascaded emission of
single photons; the technology is thus distinguished from PET, which uses
radionuclides that simultaneously produce two highenergy
photons 180° from each other.
)POSITRON EMISSION TOMOGRAPHY.)
FDG, a radiopharmaceutical used for PET studies, is also used as an imaging agent
for SPECT. FDGSPECT, also called 511 keV or positron-emitting SPECT, has
been used with dual- or triple-head SPECT systems fitted with specially
designed high-energy collimators that optimize relative resolution and sensitivity.
Clinical applications include the detection of cancerous tumors 2 cm in diameter,
studies of the viability of damaged heart tissue, and brain imaging.
Some manufacturers currently offer optional 511 keV collimators.
SPECT systems can be configured with one, two, or
three camera heads. Single-head gamma camera systems
have one detector mounted on a specialized mechanical
gantry that automatically rotates the camera
360° around the patient. SPECT systems acquire data
in a series of multiple projections at increments of 2°
(In limited-angle systems, the camera is moved a limited
number of times, usually six.) From the sequence
of projection, an image is reconstructed by an algorithm
called filtered-back projection: after nontarget
data is mathematically removed or suppressed (filtered)
for each view, the reconstructed, 3-D image is
derived from back projection, which composites the
multiangled, 2-D views and projects them onto a computer
matrix. The projection data is combined to produce
transverse (also called axial or transaxial) slices;
sagittal and coronal image slices can also be produced
through mathematical manipulation of the data.
SPECT systems with multiple camera heads are
also available. In a dual-head system, two 180°opposed camera heads are used, and acquisition time
is reduced by half, with no loss of sensitivity; a triplehead
SPECT system further improves sensitivity (Patton
2000). Some suppliers also offer variable-angle
dual-head systems for improved positioning during
cardiac, brain, and whole-body imaging. One supplier
offers a triple-head system with the detectors electronically
grouped in pairs for coincidence imaging.
Combining this configuration with improved signal
processing improves sensitivity significantly. Imaging
times can be decreased by using another SPECT configuration
— a ring of detectors completely surrounding
the patient. Although multiple camera heads
reduce acquisition time, they do not significantly
shorten procedure/exam time because of factors such
as patient preparation and data processing.
Image processing
System software allows a variety of image-processing
protocols, many of which are user defined.
Some of the more popular general software applications provided
by manufacturers are:
image smoothing,
normalization,
and interpolation;
image addition or
subtraction;
background subtraction;
contrast enhancement;
cyclic display of sequential images (cine);
region-of-interest construction and display;
curve or histogram construction and display;
and creation of alphanumeric overlays.
Cardiac applications include
first-pass acquisition;
multigated acquisition;
automatic edge detection;
determination of end-systolic
and end-diastolic volumes,
stroke volume,
cardiac output,
global ejection fraction,
regional ejection fraction,
and pulmonary transit time;
shunt quantification;
thallium perfusion profiles; and
rest/exercise thallium image comparison.
Electrocardiographic synchronizers are often offered as optional equipment
for gamma cameras. They are used in gated-acquisition studies to
synchronize image collection with the cardiac cycle defined by
electrocardiogram R waves. The beginning of the R wave
triggers the synchronizer to signal the start of data collection. The computer
divides the interval between R waves into equal subdivisions, usually
between 16 and 32. During each cardiac cycle, data is stored in the
corresponding subdivisions so that a composite image
of the cycle can be developed; a number of quantitative and qualitative
assessments are then possible.
Reported problems
 Gamma camera systems have certain limitations in image linearity, image
uniformity, intrinsic and extrinsic spatial resolution, and efficiency.
Because of limitations in detector electronics, straight-line objects may appear
curved: areas directly in front of the PMTs are subject to pincushion distortion
(inward bowing of lines), whereas areas between the tubes undergo barrel
distortion (outward bowing), neither of which is usually clinically significant.
 Image intensity can also vary — for example, pincushion distortion tends to
concentrate signals in the center of the PMT, resulting in areas of increased
intensity at each PMT location.
Improperly balanced PMTs and imperfections inherent in the NaI(Tl) crystal can
also contribute to field nonuniformity. Edge packing occurs when scintillation
photons near the edge of the crystal reflect off the inside of the aluminum housing
into the outer-edge PMTs, resulting in a field of view outlined by a ring of
increased intensity. Some cameras eliminate this ring by electronically creating an
iris that masks edge packing but reduces the field of view by a few centimeters.
 Optical problems can occur if hydrated spots — small white spots caused
by water absorption — develop on the surface of the NaI(Tl) crystal; these
spots scatter or absorb light and cause a loss of light in some scintillation
events. Off-peak testing can reveal these defects in aged crystals.
Variations in spatial resolution are usually caused by statistical fluctuations in the
distribution of light photons between PMTs. These fluctuations can be as great as
one standard deviation from one scintillation to the next. Intrinsic spatial resolution
also depends in part on crystal thickness; thicker crystals allow photons to spread
out before reaching the PMTs. In addition, lower-energy gamma rays produce
fewer photons, causing greater statistical fluctuations and therefore
decreased spatial resolution. Extrinsic spatial resolution is a function of collimator
and detector resolution and, surprisingly, is less than either one alone. Because
collimator resolution decreases with increasing distance from the source, extrinsic
resolution also decreases. Differences in resolution between gamma cameras,
although detectable on bar-phantom performance checks, are seldom clinically
significant.
 A gamma camera cannot efficiently detect high-energy gamma photons
because they pass through the thin crystal before being absorbed and
produce fewer scintillations. Detector efficiency is also limited by dead time
(a period of a few microseconds during which a scintillation is processed
and no other scintillations can be recorded) and pulse pileup, both of which
can be clinically significant in high-count-rate dynamic studies, such as firstpass cardiac function analysis.
SPECT image quality can be limited by Compton scatter and attenuation of the
radiation beam as it travels through the patient. The patient’s body size and
anatomic structure (e.g., amount of soft tissue, chest or breast size) affect the
degree of scatter and attenuation. Compton scatter reduces the contrast in SPECT
images.
Recently, more advanced scatter correction techniques have been introduced to
minimize the effect of Compton scatter on data acquisition. Attenuation
is caused by the weakening of the radiation beam produced by the
radiopharmaceutical as it passes through the patient’s body. Attenuation
correction techniques to reduce or eliminate artifacts have also been introduced by
some manufacturers. These techniques use hardware that transmits a controlled
radiation beam to the detector(s) during data acquisition.
The signals produced from the control beam and the radiation beam produced
by the radiopharmaceutical are integrated, and patient-specific attenuation is
calculated.
These new attenuation correction techniques are primarily used in cardiac
imaging.
Defects in collimators can cause sensitivity loss, longer acquisition times, errors in
image reconstruction, and image artifacts. Collimators should be checked for roper
angulation, sensitivity contrast, and center-of-rotation offset variations.
Quality-control procedures should be established for planar and SPECT imaging
systems to ensure proper operation and creation of the highest-quality images
possible for the equipment used. Daily tests should include energy peaking and
intrinsic uniformity; intrinsic sensitivity and resolution/linearity should be tested
weekly. In addition, center-ofrotation, uniformity correction, and motion correction
testing should be performed for SPECT systems. For further information, see the American Society of
Nuclear Cardiology 2001 guideline article cited below (see Bibliography).
To obtain optimal image quality, hospitals should carefully select the
appropriate imaging protocol or test, patient position, and collimator. The crystal
and the detector assembly of a mobile gamma camera can be damaged during
transport through hospital corridors.
Existing literature shows that when using gamma cameras for coincidence
imaging, camera-based systems miss a nontrivial number of small but potentially
clinically significant malignant lesions compared with full-ring PET scanners. In
addition, available studies do not clearly show that gamma camera imaging can
replace or add to the diagnostic information provided by conventional imaging.
The studies have only evaluated a small number of camera-based systems in a
limited number of oncologic uses. The clinical implications of the potentially
missed lesions have not been systematically evaluated. In order to determine
whether these systems offer net medical benefit or might inadvertently cause harm,
further studies of these systems are necessary.
Stage of development
The Anger scintillation camera was developed in the 1950s and introduced
commercially in the 1960s. In the late 1980s, multihead SPECT cameras were
introduced, and in early 1994, an FDG imaging agent for SPECT was introduced.
Other significant developments include decreased imaging times, faster and
more powerful computers, new radiopharmaceuticals, new collimators for
ultrahigh-resolution imaging, variable- angle capabilities, and digital features.
Many suppliers are now marketing digital gamma cameras that perform analog-todigital conversion, either within eachPMTor immediately after the signal leaves the
PMT. By digitizing the signal at this point, signal averaging, which affects image
resolution, can be computer controlled. Because digital detection provides
more precise event-positioning information, detector performance characteristics,
such as maximum count rate, intrinsic spatial resolution, intrinsic energy
resolution, intrinsic uniformity, and system sensitivity,
are improved.
Software-controlled operation of digital cameras also improves system reliability
and allows use of remote diagnostics for servicing.
One manufacturer has introduced a mobile camera that uses new solid-state
detectors constructed of cadmium zinc telluride (CZT) that replace the crystal/
PMT structure currently used in other cameras. The solid-state CZT detectors
directly convert gamma rays to electrical pulses. The entire system is
approximately the size of an ultrasound scanner. The smaller detector head has a
20 20 cm field of view for organspecific imaging, although whole-body data can
be acquired by scanning sections.
Clinical applications research is focused on breast cancer imaging and
expanded cardiology, oncology, and neurology applications.
Scintimammography, a technique that uses a gamma camera to image the breasts
of a patient injected with technetium-99m-sestamibi (a radioisotope traditionally
used for cardiac imaging), was recently introduced as an adjunct to conventional
mammography. Initial research suggests that scintimammography may be useful
for imaging patients who have dense breasts, who have had breast surgery, or who
have radiotherapy-altered breasts.
Because the radioisotope identifies malignancies, scintimammography may also
prove useful for targeting malignant tumors, thereby reducing the need for biopsy.
(See the Product Comparison titled RADIOGRAPHIC UNITS, AMMOGRAPHIC; STEREOTACTIC SYSTEMS, BIOPSY,
AMMOGRAPHIC for more information on mammography.)
Research into cardiac and brain SPECT is focused on the development of new
imaging agents, including radiopharmaceuticals, monoclonal antibodies, and
peptides, as well as on new applications of dual-isotope imaging with multihead
cameras. Monoclonal antibodies, which may prove useful for early detection and
staging of tumors and ovarian, colorectal, prostate, and lung cancers, have not been
used clinically on a regular basis. Peptide imaging agents are under development
for tumor, thrombus, atherosclerotic plaque, and infection imaging and are more
promising because they are safer and less expensive than monoclonal antibodies.
Additional efforts are focused on evaluating the effectiveness of FDG-SPECT
compared to PET. CMS will determine if the FDG-SPECT technology is
equivalent in performance and diagnostic quality to a full-ring dedicated PET
system.
A camera system designed for standard SPECT imaging can be upgraded
to perform FDG-SPECT, allowing the same equipment, personnel, and
space to be used for all SPECT procedures, includingFDG-SPECT.
However, the clinical sefulness of FDG-SPECT may be limited by the
spatial resolution of SPECT and the need for a gamma camera system that
can support the heavy (>400 lb) high-energy collimators. FDG is now
available from distribution centers, and small cyclotrons targeted at
producing radiopharmaceuticals are also available. Continued developments
in radiopharmaceuticals, as well as expanding applications, should increase
the attractiveness of multihead SPECT.