Planar scintigraphy produces two-dimensional images of three dimensional objects. It is handicapped by the superposition of active and nonactive layers which restricts the accurate measurement of organ functions.

Emission computed tomography (ECT) is based on the production of multi cross sectional images of tissue function which are used to produce by overlay three dimensional images.

Two ECT techniques are currently available:

• Single Photon Emission Tomography (SPECT) which involves the imaging of single 7-ray activity (typically from 99 Tc m ).

• Positron Emission Tomography (PET) which involves the imaging of the 511 keV annihilation radiation originated in positron decay (typically from 18 F)

Image planes are derived by using two different techniques:

• longitudinal ECT (limited-angle technique), photons are detected within a limited angular range from several body sections simultaneously. The reconstructed image planes are positioned parallel tothe detector plane.

• transaxial ECT (transverse section technique), the detector moves by 360 ° around the body to sample photons from multiple body sections. The reconstructed image planes are perpendicular to the detector plane.

Longitudinal

radiation source.

ECT

allows to view the radioactivity from different angles (within a limited angle range) to obtain information about the depth of the This is done by using a rectilinear scanner system coupled with a highly focused collimator who creates a sharp image only from a particular plane at a depth defined by the angle range (focal point).

scanner The multiplane tomographic represents an improved version which replaces the single detector by a gamma camera. It also allows to adjust (by electronic repositioning) the focal distance and can therefore select different image planes.

As example are shown twelve images of the skeleton recorded in longitudinal multiplane tomography technique.

Each image represents a different plane along the body.

(SPECT)

Single-Photon Emission Computed Tomography with a rotating gamma camera (transaxial

ECT) allows multiple views of the three dimensional distribution of the radioactivity from different directions.

The gamma camera is coupled to a parallel hole collimator (no focusing) which allows to produce a 2D image (64x64) consisting of multiple profiles (64), each profile represents a ID projection of the radioactivity in the profile (distributed over 64 channels).

Each point in the profile represents the sum of the activity along the line of sight: with € as detector efficiency and  as solid angle The camera rotates either continuously or in fixed angle steps and repeats the monitoring until the completion of the 360 ° turn.

The three dimensional image is constructed by using similar Fourier analysis techniques as designed for X-ray CT scanning.

Positron Emission Tomography (PET)

operates by using at least two opposite to each other positioned rotateable detector.

PET is based on the principle of detecting annihilation radiation with coincidence techniques.

The injected radionuclide must be a positron (  + ) emitter. The positron annihilates after about 1mm path length (depending on density of tissue material and on the energy of the positron) and emits two 511 keV photons in opposite directions.

Detection of both photons in coincidence defines a line along which the annihilation event has taken place. The position of the radionuclide is within  1mm distance. This distance as well as a slight deviation from the 180 0 emission of the two photons limits the spatial resolution to about 1mm – 2mm.

The use of annihilation radiation coincidence technique in PET improves the quality of image formation considerably compared to collimator techniques used in SPECT.

In SPECT the intensity and the resolution of the  signal degrades with increasing depth, due to attenuation through body tissue of increasing thickness

d

, and due to the degradation of collimator resolution increasing source collimator distance

z

:

€ c

with with constant hole diameter system.

d

and hole length

L

for the collimator

In annihilation radiation coincidence measurements the resolution remains essential constant with depth because of the uniformity of the geometric response defined by the straight line between the two detection processes.

The intensity of the coincidence signal is defined by the attenuation in body material from the point of annihilation at depth in both directions,

d

with T being the thickness of the body along the line and (x)dx

.

 Therefore the intensity for the annihilation signal along the line is independent of the depth.

The absolute count rate for coincidence events is determined by the count rate for true coincidences

I true

(real coincidence events originating from one single annihilation process) and for random coincidences

I random

(fake coincidence events which occur when accidentally each detector records an uncorrelated signal within a time window  ).

The count rate for true coincidences from

I 0

annihilation events is determined by the efficiency e and solid angle  of each detector: for present PET machines the total efficiency for coincidence measurement,

The random coincidence count rate is determined by the count rate for single events: in the two detectors and by the coincidence time window r: This yields a ratio of true to random coincidences: which is independent of efficiencies and solid angle but only depends on the intensity of the emitted annihilation radiation, and the coincidence time window: At these conditions the intensity of the radiation source inside the body must be at least

I 0

 10 6 events/s to obtain a true to random ratio of unity.

This would require a source strength of at least 1 MBq inside the body. A 1 MBq source inside the body (neglecting attenuation effects corresponds to a random count rate of:

To improve the detection conditions and to separate the true from the random coincidences the time structure of the signals can be utilized by decreasing  .

Separation is based on the fact that random events come continuously while true events come within a few nano seconds (speed of light c=3*10 10 cm/s).

is used Standard electronic to coincidence separate events in true the sharp time peak from random events using fast timing conditions on the electronic signals.

For  = 10 ns at a count rate of

I 0 =

10 6 s -1 a true coincidence rate of: is obtained at a random coincidence rate of:

Modern PET systems are based on multicrystal designs.

Instead of a rotateable detector pair the patient is surrounded by a ring of individual small Nal scintillator detectors. Each detector is coupled to his own individual phototube and is in electronic coincidence with any of those detectors at the opposite site of the patient.

With such a device a multiple image can be obtained in one shot.

This is particular important for monitoring physiological processes with short time scales ( t  10 -3 s).

Detecting coincidence events between one detector and two neighboring detectors at the opposite site defines the spatial resolution of the device to  5mm.