Clinical Instrumentation

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        Figure 2 shows the detection scheme of the positron camera. (A) After its origin, the positron emitted from the Fluorine-18, is annihilated with a negative electron, giving rise to two photons of 511 keV range each, leaving in opposite directions almost 180 ° which are detected in coincidence by the crystal ring PET camera (B). This allows their exact location and make it possible the reconstruction of tomographic images.

 

 

 

 

 

 

 


 

        PET systems are composed of a set of rings with block detectors with bismuth germanate crystals (BGO) or Lutetium Oxyorthosilicate  crystals (LSO), coupled to photomultipliers. The BGO crystals are denser than the INa (Tl) and therefore more efficient to detect the 511 keV photons. On the other hand they are not hygroscopic and can therefore be cut into parallelepipeds a few millimeters thick (e.g. 4 or 5 mm), the magnitude of spatial resolution.  PET systems are also built with crystals of INa (Tl), which are cheaper, and despite being less efficient detecting the 511 ke photons, are used for clinical PET. In favor of the INa (Tl) is the fact that emits a higher percentage of light and the signal in the crystal decays faster, therefore better resolution and lower dead time, though has less efficiency than the BGO crystal.

        PET coincident detection  between crystals of the same ring allows a two dimension (2D) reconstruction. Lead septa are installed to separate the rings and  avoid random coincidences and scattered among non-contiguous planes. The latest generation of PET systems may do acqusitions without septa employing 3D reconstruction algorithms. With this technique, the coincidence detected events are possible not only with crystals of same  rings but also between rings, increasing sensitivity. One of the big advantages of PET is its ability to correct for attenuation of radiation in the tissues by the method of transmission using 68Ge or 137Cs sources ot the X-rays used in the CT. The ratio of the absorption with and without the patient is used to calculate the absorption coefficient (m) for each projection. Once attenuation and scattered radiation have been taken care of, activity within the tissues of the patient can be quantified in absolute terms. For example, to obtain quantitative images of FDG with plasma data included into kinetic models from which to calculate glucose uptake into tissues.

        In addition, PET images can be quantified to determine the degree of radiotracer uptake, using the SUV (Standardized Uptake Value)  index . This relates tumor metabolic activity compared to the injected dose and patient weight. As a cutoff point of reference  a value = 2.5 may be helpful. Benign lesions tend to be below and malignant abnormalities frequently over this value, but this only is helpful as a general orientation. Even more useful, is for follow-up. In the case of a good therapeutic response, the value should drop significantly. The patient is his own control.
 

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