Understanding the technology behind photon-counting CT
Watch the animation to see what's behind this quantum leap in computed tomography
After the introduction of spiral CT in 1990, wide detector CT in 2004, Dual Source CT in 2005 and Dual Layer CT detectors in 2013, computed tomography is a mature modality that has reached a saturation phase. Despite technological progress, limitations remain for current CT technology.
With photon-counting CT, we have developed a radically new technology for clinical routine. At its core is a new kind of detector that is substantially different from a standard energy integrating detector.
Watch the animation for an explanation of how the technology behind these photon-counting detectors work and how they could overcome the limitations of current CT detectors by providing CT data at high spatial resolution, without electronic noise, with improved contrast-to-noise ratio, at lower radiation dose and with intrinsic spectral information.
What makes photon-counting detectors different?
All medical CT systems today are equipped with solid‐state scintillation detectors. In a two‐step conversion process, the absorbed X‐rays are first converted into visible light in the scintillation crystal. The light is then converted into an electrical signal by a photodiode attached to the backside of each detector cell.
The low‐level analog electric signal of the photodiodes is susceptible to electronic noise, which sets an ultimate limit to potential further radiation dose reduction.
At the same time, it is problematic to significantly increase the spatial resolution of solid‐state scintillation detectors beyond today’s performance levels.
As part of this two‐step conversion process, the light created by thousands of xray photons is accumulated over the integration time and measured as a whole, thereby losing the spectral information of the incoming signal.
Photon-counting detectors, by contrast, can directly transform X‐ray photons into electrical signals.
In a direct conversion process, the absorbed X‐rays create electron‐hole pairs in the semiconductor. The charges are separated in a strong electric field between cathode on top and pixelated anode electrodes at the bottom of the detector.
Compared to solid‐state scintillation detectors, photon-counting detectors have several advantages. The individual detector cells are defined by the strong electric field between common cathode and pixelated anodes (Fig. 2), and there is no need for additional septa between the detector pixels to avoid optical cross talk inherent to scintillation detectors. The geometrical dose efficiency is, therefore, better than that of scintillation detectors and only reduced by the anti‐scatter collimator blades or grids that are also present in scintillation detectors. Furthermore, each “macro” detector pixel confined by collimator blades may be divided into smaller detector sub‐pixels which are read‐out separately to significantly increase spatial resolution.
With a photon-counting detector being able to count the charges created by individual x‐ray photons as well as measuring their energy level, we now have a detector that has intrinsic spectral sensitivity in every scan.
The crystal inside that makes it all possible
The purest cadmium telluride crystal in the world lays the basis for next technology leap in computed tomography. Take a sneak peak into the crystal growing lab where the magic happens – and see how it makes photon‐counting CT a reality for clinical routine.
What does this mean for computed tomography – and for you?
The direct signal conversion of photon-counting detectors can have great impact: They are much more dose efficient than current detectors. Also, their pixels are much smaller, which can significantly increase the spatial resolution. Based on this new technology, patients can expect even further reduction of radiation dose and less use of contrast agent. In addition, physicians can work with images that visualize even very fine tissue structures, such as the smaller bronchi of the lungs or metastases in bones.
- No downweighting of lower energy quanta: Improved image contrast
- Smaller detector pixels: High spatial resolution without losing dose efficiency
- Eliminate electronic noise: Lower radiation dose
- Intrinsic spectral sensitivity: Multi‐energy information