Single Photon Imaging and Instrumentation

Anna Celler, Department of Radiology, University of British Columbia, Vancouver, Canada

& 2010 Elsevier Ltd. All rights reserved.

Abbreviations2D, 3D 2-dimensional, 3-dimensional

CZT cadmium zinc telluride

EKG electric signal from the heart

FBP filtered backprojection

FOV field-of-view

GE General Electric

Ge(Li) germanium detector doped with

lithium

MLEM maximum likelihood expectation

maximization

NaI(Tl) sodium iodine doped with thallium

NM nuclear medicine

OSEM ordered subset expectation

maximization

PET positron emission tomography

PMT photomultiplier tube

SPECT single-photon emission computed

tomography

Principles of Nuclear Medicine Imaging

The general term nuclear medicine (NM) encompasses anumber of different techniques that all use unsealedradioactive compounds for diagnosis or therapy. NMimaging belongs to the first application. Its objective is tomeasure and visualize the distribution of a radiotracerthat has been introduced into a human body. The basicprinciple is simple: a disease-specific pharmaceutical,molecules of which have been labeled with radioactiveisotope, is administered to a patient, usually by intra-venous injection. The molecules are designed to go toand localize in an organ or area of interest. This tracer’sradioactive emissions, measured outside the body, areused to create two- or three-dimensional (2 D or 3 D)images that, viewed by an NM physician, provide him/her with information about the organ’s function, normalor altered by disease, its physiology, and metabolism.

There are two types of NM imaging. Ultimately bothof them, although based on different principles, use de-tection of gamma rays to create diagnostic images. Posi-tron emission tomography (PET) uses isotopes decayingby positron emission (bþ decay). An emitted positronalmost immediately annihilates with an electron in themedium. Coincident detection of two collinear 511 keVphotons created in this annihilation is used to produce3 D images of the radiotracer distribution in the body.

Single-photon emission imaging (also called scinti-graphic imaging) uses isotopes that decay with photon

emission. Since these photons are recorded individually,

as opposed to measuring coincident pairs of photons in

PET, the modality is referred to as single-photon im-

aging. The emitted photons are measured outside the

patient body by a gamma camera (also called an Anger or

scintillation camera), which may have one, two, or even

three detectors (camera heads). Depending on the diag-

nostic question, the scan is performed following one of

many clinical protocols. In the simplest acquisition, data

are collected in planar mode where the camera remains

stationary for the whole duration of the scan. Alter-

natively, it can move slowly along the length of the pa-

tient acquiring a whole-body image. In both of these

cases, the resulting image represents a 2D projection of

the 3D distribution of radiotracer in the body. When the

camera rotates and multiple projections are acquired at

different locations around the patient, a 3D radiotracer

representation may be obtained using one of the tomo-

graphic reconstruction methods. This technique is

referred to as single-photon emission computed tomo-

graphy (SPECT), and the camera that is able to acquire

and process tomographic studies is called a SPECT

system. Otherwise, dynamic studies can be performed

where a series of 2D planar images is acquired over time,

allowing physicians to visualize temporal changes in the

tracer distribution, thus providing diagnostic information

about body functions. Finally, dynamic 3D images of a

beating heart may be acquired by correlating data ac-

quisition with the electric signal from the heart (EKG).Figure 1 shows three examples of modern clinical

SPECT systems. The Philips SKYlight system presented

in Figure 1(a) does not have a traditional gantry; instead

its two detectors (camera heads) move relatively freely on

special arms. This design allows for high flexibility in

scanning patients of almost any size and in any position.

Figure 1(b) and 1(c) shows the General Electric (GE)

Infinia Hawkeye and Siemens Symbia SPECT/CT

systems, respectively. Both SPECT cameras have two

detectors rotating in open gantry configurations. The

detectors can be positioned at 901 to each other for

cardiac scans or, alternatively, at 1801 configuration,

parallel to each other, for a number of other imaging

applications. Both cameras are additionally equipped

with CT imaging systems.

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Imaging Equipment

Anger Camera

The most important element of an Anger camera is thedetector, which in the majority of systems currently usedis an NaI(Tl) (sodium iodine doped with thallium)scintillation crystal. Each head in a modern SPECTcamera contains one large rectangular detector with anactive surface of approximately 50� 40 cm and a thick-ness of 0.9 cm. Behind the crystal is an array of 50–90photomultiplier tubes (PMTs), the associated electronics,and analog-to-digital converters. The role of the PMT isto measure the energy of each detected photon and theexact location of its interaction with the crystal material.Currently used cameras are optimized for photons inthe 70–400 keV energy range and have relative energyresolution of approximately 9%. Their intrinsic spatialresolution is of the order of 3–3.5 mm.

The NaI(Tl) crystals are highly efficient at convertinggamma photons into visible light photons. Additionally,since the amount of light is proportional to the energy ofincident photons, it is possible to discriminate detectedevents based on their energy. However, relatively poorenergy resolution makes elimination of scattered pho-tons, which arrive at the camera with energy lower thanthe original ones, difficult. A number of other detectormaterials have also been investigated for potential use ina gamma camera. Although semiconductor crystals, suchas Ge(Li) or high-purity germanium, provide sub-stantially better energy resolution than NaI, they havelow detection efficiency, high cost, and complicatedmaintenance (liquid nitrogen cooling). For a number ofyears, the use of cadmium zinc telluride (CdZnTe orCZT) has been considered, but only very recently haveprototype SPECT imaging systems based on CZTcrystals become available. This very promising solid statedetector has much better energy resolution than NaI(Tl)(approximately 5–6%), works at room temperature,and does not require photomultipliers to read the signal.Another attractive characteristic of the CZT detectoris its excellent intrinsic spatial resolution, presently

reaching 0.5 mm, however, requiring thousands of indi-vidual electronic channels. Additionally, although newSPECT cameras with CZT detectors are very promising,the material is still rather costly and its production dif-ficult. Only small crystals can be grown (approximately2� 2� 0.5 cm), and often there are problems with theconsistency of their spectral performances. Therefore,CZT cameras use pixilated arrays of detectors and arecurrently available only in small field-of-view (FOV)configurations for small animal and cardiac applications.

Another vital component of each Anger camera is itscollimator. Figure 2 presents a schematic drawing of atypical camera configuration showing all its main com-ponents: the collimator, the detector, phototubes, associ-ated electronics, and computer. The role of thecollimator, which is placed in front of the detector, is todetermine the direction of incident photons. This isachieved by restricting the angular distribution of pho-tons that may reach the detector surface. The collimatorconsists of a sheet of a highly attenuating material(usually lead) with many thousands of small hexagonalholes. Only photons that arrive within the acceptanceangle of the collimator can reach the detector surface.There is a small probability that some photons with in-cident angle greater than the acceptance angle will passthrough the collimator walls (septa) and be detected; thiseffect is called septal penetration.

Most clinical applications use parallel-hole colli-mators with holes perpendicular to the surface of thecrystal. Theoretically, such collimators accept onlyphotons traveling normal to the detector; thus, there is adirect relationship between the spatial distribution of theradiopharmaceutical and the resulting image. Obviously,the real collimator accepts not only photons parallel tothe septa but also all those that arrive within a small conecorresponding to the acceptance angle, and as a result thearea seen by each hole increases with distance. Thismeans that the resolution of the camera is distancedependent and deteriorates as the distance between thesource and the collimator increases. This effect is calledcollimator blurring and can be partly compensated for

Figure 1 Three examples of modern clinical SPECT systems: (a) Philips – SKYlight, (b) General Electric – Infina Hawkeye, and

(c) Siemens – Symbia. Photos are courtesy of Philips Healthcare (a), GE Healthcare (b), and Siemens AG (c).

2532 Single Photon Imaging and Instrumentation

when iterative image reconstruction is used. It is gener-ally recognized that the typical resolution of SPECT is ofthe order of 1–2 cm. This value results from combinedresolution of the collimator measured at a clinicallyrelevant distance and the intrinsic resolution of adetector.

As the majority of photons that are emitted by a pa-tient body are absorbed in the collimator septa, thetypical parallel-hole SPECT collimator has low sensi-tivity, accepting approximately only one photon for every104 emitted. Collimators with larger holes provide higherphoton detection efficiency, but increase the acceptanceangle and thereby lower the detector spatial resolution.In practice, each clinical camera has two or three sets ofexchangeable collimators with different hole sizes andseptal thicknesses. The collimator selected depends onthe type of study; for example, different collimators areused in scans involving radiotracers emitting low- orhigh-energy photons, requiring high resolution or highsensitivity.

Other collimator types, such as pinhole, cone beam,and fan beam, produce magnified images of an objectand therefore may be used for imaging of small organs(e.g., the thyroid, wrist, or brain). Although cone-beamand fan-beam collimators usually have slightly highersensitivity than parallel-hole collimators, typically thesensitivity of a pinhole collimator is very low. All of thesecollimators, when used in tomographic (SPECT) imaging,require special algorithms for image reconstruction,which take into account the geometry of data acquisition.

To improve detection sensitivity, collimators withmultiple pinholes have been designed. They offer ex-cellent resolution, while maintaining acceptable sensi-tivity. For this reason, they have become very popularin preclinical (small animal) SPECT systems, wherenonoverlapping or overlapping hole configurations are

used. These systems have much better resolution (0.5–2 mm) and sensitivity (B0.3%) than clinical SPECTcameras (10–20 mm and 0.01–0.03%, respectively). Thesecharacteristics, however, apply only to a very small FOV(5–8 cm).

The light created by photon interactions in the crystalis amplified and converted to an electrical signal by anarray of PMTs. The tightly packed hexagonal PMT array(Figure 3) not only measures the total energy of eachphoton by summing the light deposited on the crystal,but also provides the spatial positioning of the detectedphotons. The location of each interaction is determinedby combining the electrical signal from all PMTs in alogic circuit that weighs these signals appropriately andcompares the relative amplitudes of signals resultingfrom each PMT.

Although PMTs have been in use for a number ofyears, their low quantum efficiency significantly limitsboth the energy and the spatial resolution of the camera,their output is sensitive to changes in temperature andmagnetic fields, and they are bulky and expensive. Re-cently, a number of alternative solutions have been pro-posed ranging from position-sensitive tubes and hybridPMTs to semiconductor avalanche photodiodes.

The last component of the camera is a computer or aworkstation. First, it is used to specify acquisition par-ameters of the study such as the number and size ofdetector bins, the number of projection angles, the tem-poral duration of each view, and other features related tothe motion of the gamma camera. During acquisition, theenergy signal from the camera multichannel analyzer isanalyzed in order to restrict the range of photons that areused to create the image. This is done by defining one ormore energy window(s) corresponding to the photopeakenergy of the radiotracer. Next, the computer digitizesand stores the data that passed the energy test, assigning

Patient

Scanning bed

CollimatorDetector

Photomultipliers

Cameraelectronics

Processing and viewingworkstation

Figure 2 Diagram of a typical Anger camera with all its main components.

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