NM image quality


As with other modalities the three major factors that determine image quality are:

  • Contrast
  • Noise
  • Spatial resolution

 

Contrast

In radionuclide imaging contrast is created by the differential uptake of a radiopharmaceutical agent. Lesions can have negative (smaller lesion activity than surrounding tissue) or positive (larger lesion activity) contrast.

Subject contrast

The subject contrast is a property of the imaged object i.e. the radioactivity level in a lesion relative to healthy tissue. It is calculated as:


CS = (AL - AT) / AT
 
where:
CS = subject contrast
AL = activity per unit volume of the lesion
AT = activity per unit mass of the healthy tissue

Image contrast

Image contrast is the difference in the display between the lesion and surrounding healthy tissue. It is represented as the counts per unit area and also called the count density or information density. It can be expressed as counts per pixel.


CI = (SL - ST) / ST
 
where:
CI = image contrast
SL = counts per unit area of the lesion
ST = counts per unit area of the health tissue

Factors reducing image contrast

  • Smaller subject contrast
  • Greater background gamma radiation (for positive contrast radiopharmaceuticals)
    • Background radiation gives a count density that is overlayed onto the whole image.
    • Sources include radioactivity in tissue above and below the lesion and other radioactive sources (in vicinity of patient or in environment)
  • Using a collimator: photons penetrate septae and count density  
  • Scattered radiation from patient
  • Attenuation of the gamma radiation in a deep lesion which is much greater than for surrounding healthy tissue
  • Greater patient movement 
  • Low spatial resolution of the gamma camera

 

Noise

Radionuclide imaging is an inherently noisy investigation. Too much noise will impair detectability of an object, especially if it is a low contrast object. 

Structured noise:

  • Non-random count density that interferes with object of interest due to:
    • Uptake in structure that is not of interest e.g. muscle uptake in PET after excercise, bowel uptake in gallium-67 citrate
    • Imaging system artefacts e.g. non-uniformity of the gamma camera

Random noise:

  • Aka statistical noise or quantum mottle
  • Due to random variations in count density as a result of random activity of radioactive decay
  • More significant contributor of noise

Calculating noise

Relative noise (noise contrast) decreases as the count number (signal) increases.


σ = √N

SNR = N/σ = N/√N = √N

CN = σ/N = √N/N = 1/√N

CN = 1/√AS

 
where:
σ = random noise (a standard deviation)
N = counts
SNR = signal to noise
CN = noise contrast
A = area
S = count density

Factors reducing noise

  • Longer acquisition time
  • Increased activity of radiopharmaceutical (for given acquisition time)
  • More sensitive gamma camera (however, increasing the sensitivity to decrease the relative noise also decreases the spatial resolution and contrast)

 

Spatial resolution

Spatial resolution is quantified as the full width at half maximum on a graph of counts or count rates vs distance. This is either measured when a radioactive point source (point source function, PSF) or when a line source (line source function, LSF) is imaged. The LSF is more commonly used. A Fourier transform of the LSF gives the modulation transfer function (MTF) which quantifies how accurately the image represents the object.

Point source function

Intrinsic spatial resolution (RI)

The intrinsic spatial resolution is the maximum resolution achievable by the detector and electronics and depends upon many factors:

  • Energy and linearity correction of scintillation
  • Range of light in scintillation crystal. A thicker crystal means more spread and variation in the depth of the signal which reduces spatial resolution
  • Higher gamma photon energy = more scintillation photons = smaller statistical variation which improves spatial resolution
  • Optimised collection and detection of scintillation photons: good optical coupling and PMT shape (square or hexagonal better than circular), more PMTs
  • Only PMTs above certain voltage contribute to signal (eliminates noise)

Intrinsic spatial resolution at 140 keV is between 2.5 mm FWHM (0.4 lp/mm) and 4 mm FWHM (0.25 lp/mm).

Collimator spatial resolution (RC)

For a parallel hole collimator the collimator spatial resolution is:


RC ≈ d(1 + b/h)
 
where:
RC = collimator spatial resolution
d = hole diameter
b = distance from radiation source to collimator
h = hole length

From this equation you can see that resolution is improved by using a collimator with long holes of small diameter positioned as close to the patient as possible. However, there is still rapid degradation of spatial resolution the deeper the imaged object lies. Taking images from different orientations helps to minimise this.

System spatial resolution (RS)

The RS is takes into account the intrinsic and collimator spatial resolution to give the spatial resolution of the whole system.


RS = √(RI2 + RC2)
 
where:
RS = system spatial resolution
RI = intrinsic spatial resolution
RC = collimator spatial resolution

Factors reducing resolution

  • Low intrinsic spatial resolution of the gamma camera
    • Thick scintillation crystal
    • Small number of PMTs
    • Low threshold for PMT voltage to contribute to signal
  • Low spatial resolution of the collimator
    • Large diameter holes
    • Short holes
    • Far from patient
  • Increased patient motion
  • Imaging deeper structures
  • Large display pixels
  • Increased scattered radiation
    • Improved with narrower energy acceptance window

 

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PET


Contrast

  • Tomographic technique overcomes reduced contrast caused by radiation in front of and behind the lesion
  • Random and scatter coincidences reduce contrast

Noise

Noise reduced by increased sensitivity of the system which is determined by:

  • Intrinsic detector efficiency
    • Scintillation crystal with higher LAC and more depth = better absorption of gamma photons = greater sensitivity
  • Geometric detection efficiency
    • Higher number of gamma photons that reach detector = greater sensitivity
    • Better in 3D than 2D acquisition
  • Width of photopeak acceptance window
    • Wider photopeak acceptance window = greater sensitivity
    • However, also increases scatter coincidence detection rate which reduces contrast

Resolution

  • Positron range
    • Distance from site of disintegration to annihilation
    • Longer range = poorer spatial resolution
    • 15O is 2 mm, 18F is better at 0.6 mm
  • Non-colinearity of the annihilation photons
    • If positron or electron have residual momentum at time of annihilation the angle between the paths of the two gamma photons produced will not be exactly 180°
    • The greater the deviation the poorer the spatial resolution
  • Detector element size
    • Smaller elements = between spatial resolution
  • Thickness of crystal
    • Thicker crystal = poorer resolution
    • Resolution better through centre than periphery of detector ring
  • Reconstruction filter
    • PET has much higher count rate sensitivity than SPECT and so noise is less of a problem
    • PET images can be reconstructed with much higher spatial frequency

Next page: NM Artefacts


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