In both emission and transmission computed tomography, information concerning the internal structure of a subject of interest is obtained without invasive procedures.
In computed tomography a two dimensional image is generated from multiple one dimensional projections. A source of x-radiation transmits X-rays through a subject of interest and the intensitiy of that radiation is monitored on an opposite side by one or more detectors, which are often single scintillation crystals each optically coupled to a differnt photodiode. If the source is moved about the patient and X-ray attenuation data is obtained from a number of directions, a computed tomography reconstruction process can be utilized to generate an image of the subject cross-section.
The x-ray tube radiation source used in computed tomography emits x-ray photons over polychromatic or continuous energy spectrum which extends from near zero in energy to a maximum which is determined by the voltage (kv) applied to the source. The performance and energy spectra of x-ray tubes is well understood and documented by those of ordinary skill in the relevant art.
The radiation detectors utilized in computed tomography each produce a signal proportional to the total x-ray energy absorbed by the detector. The lower energy x-ray photons striking a detector contribute a relatively small amount to the total output signal of the detector, whereas the higher energy photons incident on the detector produce correspondingly larger contributions to the output signal.
The low energy x-ray photons are more readily absorbed in the patient, while the high energy photons have a greater tendency to pass through the patient's body and to undergo less attenuation than do the low energy x-ray photons. Hence, the low energy radiation contains more contrast information defining the patient's internal body structure than does the high energy radiation, which more uniformly passes through the patient's body, as a result of less attenuation.
Therefore, the detector output signal is more a function of the high energy radiation, which carries relatively less contrast information, than it is a function of the lower energy radiation which is relatively rich in image information. Accordingly, the present detectors that employ a single energy sensitive element (e.g., a single scintillation crystal or a gas filled ion chamber) weights its output signal in favor of the less informative high energy radiation than in favor of the more informative low energy radiation.
The generation of x-ray photons by an x-ray tube anode is a random process. Even where the x-ray tube kV and mA (voltage and current) are held constant, the number of photons emitted from the anode fluctuates statistically in time about an average value with a Poisson probability distribution. The absorption, or attenuation, of the x-ray photons in matter, such as in a patient's body, is also a random process following similar laws of probability theory. Therefore, the number of photons detected during a fixed time period, with all other conditions held constant, will vary from one measurement to the next. This statistical fluctuation in the measurement gives rise to an uncertainty or ambiguity in the true value of the attenuation. This ambiguity is sometimes referred to as "quantum noise" or "quantum statistics". While this quantum noise sets an absolute, or fundamental limit on the quality of an image that can ultimately be obtained, unequal response by the detector to photons of different energies magnifies this noise problem.
Noise in the detector output signal which results from the quantum statistics associated with the detection of N x-ray photons is given by the following relation: ##EQU1## where:
.sigma.=the standard deviation (or RMS noise) of the output signal, expressed as a fraction of the signal (i.e., noise to signal ratio);
W.sub.n =Weight (or signal contribution) of the n.sup.th photon, and
N=Total number of photons detected.
In most present detection systems used in computed tomography, W is proportional to the absorbed energy of the photon. However, less noise (or .sigma.) is obtained if W.sub.n is constant (i.e., every photon is weighted equally). In this optimum case, the standard deviation is given by the familiar formula: ##EQU2##
To illustrate with an example, consider two cases:
Case I. (Unequal Weighting)
Let N=100 and assume that half of the detected photons have a relative weighting of W, and the other half are higher in energy and have a relative weighting of 4W, then ##EQU3##
Case II. (Equal Weighting)
Let N=100 and assume that all photons are detected with relative weighting of W; then ##EQU4##
As can be seen in this set of examples, Case I is 17% noisier than Case II. In order for Case I to achieve the reduced level of noise of Case II, the number of detected photons, i.e., N, would have to be increased to 136. This would correspond to an increase in patient dose of 36%.
Thus, the characteristic of present detectors to weight their response to high energy photons more heavily than to low energy photons results in an exaggerated effect of noise in the output signal, and, in order to compensate for this noise, requires a larger x-ray dose to be applied to the subject.
It is an objective of this invention to provide a detector of penetrative radiation having reduced noise caused by the quantum statistics of the polyenergetic spectrum of the x-ray tube output and which exhibits enhanced response to the more informative lower energy portion of the detected x-ray energy spectrum.
USPTO Backgrounds
USPTO Backgrounds