A gamma camera is a device used to image gamma radiation emitting radioisotopes in a process. It consists of crystal scintillators optically coupled to an array of photomultiplier tubes in an assembly known as a head, mounted on a gantry.
Scintillation
A scintillator is a material that exhibits scintillation, the property of luminescence when excited by ionising radiation.
Components
Collimator
The collimator is a device that is attached to the front of the gamma camera head and its main purpose is to assist in spatially mapping gamma photons exiting the patient’s body, before they reach the detector. As radiation leaving the patient can occur at any angle, the role of the collimator is to reject non-parallel gamma photons as they do not correspond to their original location.
- Used to define the direction of the photon emission point.
- Collimator resolution is the dominant factor in overall system resolution at clinical distances
The collimator consists of a series of holes separated by lead septae. Radiation entering the collimator at non-perpendicular angles will be absorbed by the lead septa and not contribute to image formation. The septa can be arranged in a range of possible configurations (angles of acceptance):
- Parallel – most common. The FOV and sensitivity does not change with distance.
- Diverging
- Minifies the image
- Allows use of a small crystal to image a large field of view e.g. lungs
- Converging (cone beam, fan beam)
- Results in magnified image, but not inverted
- Pinhole
- Single-hole collimator for magnifying images of superficial, small objects e.g. thyroid and parathyroid glands.
- Results in an inverted and magnified image
The choice of a specific collimator is dependent on the amount of radiation absorption that occurs (which influences the sensitivity of the gamma camera), and the clarity of images (that is the spatial resolution) it produces.
The higher the energy of the emitted gamma photons the thicker the septae need to be to ensure maximum absorption of photons that hit them at an angle and, therefore, better rejection of non-perpendicular photons. Parallel hole collimators are classified as low, medium or high energy according to their septal thickness.
| Classification | Photon energy (keV) | Septal thickness (mm) | Radionuclide |
| Low energy* | 150 | 0.3 | 99mTc |
| Medium energy | 300 | 1 | Indium-111 |
| High energy | 400 | 2 | 131I |
* Most commonly used.
Collimator resolution is distance-dependent, hence the collimator is moved close to the patient.
Collimator sensitivity is a measure of the fraction of total gamma rays falling on the collimator that pass through the holes to the crystal. Increasing the number of holes increases the sensitivity and reduces patient dose at the expense of spatial resolution. Similarly, long and narrow holes result in a narrower angle of acceptance, thus reducing sensitivity but increasing spatial resolution.
Patient dose decreases as there is a lower amount of radionuclide needed in order to form the image.
Sensitivity is inversely proportional to the spatial resolution.
Scintillator crystal
- Converts the high energy photon (gamma rays) into optical photons
- Scintillator material: NaI(Tl) – sodium iodide (6 – 13 mm) doped with thallium
A large flat crystal of NaI(Tl) is hermetically sealed in aluminium housing. The crystal scintillates in response to incident gamma radiation, releasing light photons (visible and UV) which then enter the photomultiplier tubes via light pipes for amplification. The scintillator absorbs gamma rays mainly by photoelectric absorption.
The thickness of the scintillator is chosen such that it provides good detection for the 140 keV gamma-rays emitted from 99mTc.
Increasing the crystal thickness improves sensitivity but decreases spatial resolution. Stopping power (part of system sensitivity) is dependent on crystal thickness
Ideal Properties
- High stopping power (more photons interact)
- Want high light yield (better signal)
- Want fast decay time (better timing resolution for PET)
- Crystal thickness
- thicker: better stopping power, worse resolution due to lateral spreading within the crystal
- Thinner – worse stopping power, better resolution
- Hermetically sealed
- Hydroscopic
Photomultiplier tube
The role of the photomultiplier tube is to amplify the signal received. It achieves this by first converting light into photoelectrons.
- Light photons from the scintillator hit a photocathode at the entrance to the PMT.
- The photocathode converts the light photons into ‘free’ electrons in proportion to the amount of incident light.
- The electrons are electrostatically attracted to the electrodes called dynodes which have an increasingly positive charge along the PMT. This accelerates the electrons and, as they accelerate, they gain kinetic energy resulting in multiple electrons being released from the dynode for each electron that hits it. This serves to amplify the original signal with again equivalent to:
Gain ≈ ασVnwhere α = number of photo-electrons emitted, σ = number of electrons emitted per incident photon at each dynode, V = voltage (potential) applied between each dynode stage, n = number of dynode stages
The amount of light detected by the PMT is a function of the distance between the scintillation event & the PMT.
- The total electrons hit the final anode and the current produced forms the signal that the pre-amplifier receives. The pre-ampilifier converts the current produced at the anode of the PMT to a voltage pulse.
- Analogue-digital converter
Image Formation
For each scintillation formed, the calculated absorbed energy (Z value) that caused it depends on the energy of the gamma photon that was emitted from the patient and the proportion of the energy that was absorbed into the crystal.
The gamma photon energy absorbed by the scintillation crystal depends on its interaction with that photon which results in a spectrum of Z values.
- All energy absorbed: gamma photon interacts with crystal via the photoelectric effect
- Part of the energy absorbed: photon undergoes one or more Compton interactions
The spectrum has a peak (photopeak) that corresponds to the maximum gamma photon energy (for 99mTc this is 140 keV). The Compton band corresponds to photons that have undergone Compton interactions and, therefore, have a lower absorbed energy.
The photopeak should be very narrow but a variety of factors means that it often isn’t. The width of the photopeak is measured as the full width at half maximum (FWHM). This value is used to calculate the energy resolution of the crystal, which is given as a percentage:
Energy resolution = FWHM (keV) / photopeak energy (keV) x 100
Pulse Height Analysis
A pulse height analyzer (PHA) is a device used to determine which portion of the detected spectrum is used for image formation. It electronically filters out scatter by only accepting gamma rays that fall within a small energy window (e.g. ± 10%) around the photopeak of the nuclide used, e.g. 140 keV for 99mTc.
- Increasing the window produces an image more quickly but degrades image quality by including more scatter photons.
Spatial Resolution
- Improves by reducing the distance from the patient
- This means the gamma-camera orbit may be ellipitical
- Improves by using collimators with longer and narrower holes
- Tradeoff:
- Improves by decreasing crystal thickness
- Tradeoff: decreased sensitvity
- Larger patients result in reduced spatial resolution due to increased attenuation and scatter.
Sensitivity
Expressed as total counts per second per megabecquerel of activity (cps MBq-1)
- Improves by using a thicker crystal
- Tradeoff: decreased spatial resolution
SPECT
Uses a rotating gamma camera
- Results in improved target-to-background ratios as it separates overlying structures, which results in better
visualization of the target tissues.