Storage Phosphors for Medical Imaging
Abstract
:1. Introduction
1.1. Medical Radiography
1.2. Screen/Film Radiography
1.3. Computed Radiography (CR)
2. CR Phosphor Requirements
- It should have a high X-ray absorption for medical X-ray spectra, i.e., for energies ranging from 20 keV to 140 keV.
- It should have high conversion efficiency. The amount of energy stored per unit X-ray energy should be high. This implies that a large fraction of the mobile charge carriers created by the absorbed X-ray quanta is converted into trapped electrons and holes forming defect aggregates leading to PSL. The energy needed to create a trapped electron and hole in the storage phosphors in commercial use is estimated to be:
- 100 eV in BaFBr:Eu2+
- 67 eV in CsBr:Eu2+.
- The image information stored in the phosphor should have slow fading in dark at room temperature, i.e., electron- and hole traps should be stable. In practice, the so-called dark-decay of the stored image in a CR plate is between 10 and 25% in the first hour after X-ray exposure.
- The stimulation spectrum should be in the range covered by inexpensive solid state lasers. Currently laser diodes are commercially available with wavelengths ranging from 375 nm to 1800 nm, but red and IR lasers exhibit the best price/performance ratio. In addition, the stimulation light should not produce free charges that can give rise to PSL. This excludes lasers with wavelengths in the blue and UV parts of the spectrum. Finally, the stimulation light and the stimulated emission should be spectrally separable. This is necessary since the emitted light is between 105 and 109 times weaker than the stimulation light, depending on X-ray spectrum and dose used to make the image. Color filters are applied to block the stimulation wavelength and transmit the emission light. Since the slope of the filter curves is finite, a stimulated emission with a wavelength close to that of the laser will also be attenuated. Altogether, this means that efficient stimulation of the trapped charges should be possible in the wavelength range between 500 and 1,500 nm. However, it is unlikely that trapped charges that can be stimulated with 1,000 to 1,500 nm light are stable at room temperature in the dark.
- The CR phosphor emission should match the sensitivity spectrum of the light detector (Figure 5). In flying-spot scanners with a PMT light detector, the emission is preferable below 500 nm and most preferably in the range between 350 and 450 nm.
- The laser power required for CR plate stimulation should be low. This implies that both the oscillator strength and the recombination probability of the stimulated trap should be high. Actual CR scanner laser powers are between 5 and 30 mW.
- Since read-out of a complete image plate in less than one minute is desirable, the luminescence decay time should be short. In current flying-spot CR digitizers, the pixel read-out time does not exceed 2 µs. At the time the next pixel is stimulated, the light emission of the previous pixel should have decayed to at least 1/e of its initial value. Consequently, for use in flying-spot scanners the decay time should be maximum 2 µs.For line scanners phosphor decay time is much less critical. In these devices a complete line is stimulated and read-out at once and the line-time is typically in the millisecond range.
- The IP should be re-usable. Only roughly 50% of the stored energy (trapped charges) is released during scanning. Hence, the IP must be erased for re-use. For erasure the phosphor is illuminated with a strong light source. To allow efficient erasure, the amount of “persistent traps” that is formed upon X-ray radiation should be negligible.
- The phosphor should be stable under normal room conditions, i.e., its performance should not degrade when it is exposed to humidity and daylight.
- It should be possible to manufacture the phosphor on an industrial scale in a commercially viable way.
- Spatial and contrast resolution are important clinical requirements and are affected by phosphor quality as well. Smaller particles lead to higher resolution and to lower light output. The best compromise is obtained when the phosphor particles have a median size in the range between 1 and 10 μ.
3. Powder Imaging Plates (PIP’s) for Medical CR
3.1. PIP Manufacturing
3.2. The BaFBr:Eu2+ Storage Phosphor
4. BaFBr:Eu2+ PSL Mechanism
4.1. Electron and Hole Center Production
4.2. Recombination of Trapped Electrons and Holes
4.3. Spatial Correlation of Defects
4.4. The Formation of Bromine Vacancies in the BaFBr Lattice
5. BaFBr:Eu2+ Engineering
6. Alternative Storage Phosphors for Medical CR
- -
- the most promising candidates to replace BaFBr:Eu2+
- -
- storage phosphors having high X-ray absorption or useable for manufacturing thicker and, therefore, more absorbing phosphor screens.
- Conversion efficiency (CE): the total energy of stimulated light per unit area and per unit of X-ray dose produced by the phosphor in pJ/mm2/mR,
- Stimulation energy (SE): the laser energy per unit area required to release 63% of the stored energy in µJ/mm2.
6.1. BaFBr:Eu2+ Alternatives
6.2. Higher Intrinsic Absorption
6.3. Elpasolite Storage Phosphor
6.4. Alkaline Earth Sulfide Storage Phosphors
6.5. Storage Phosphor Glass Ceramics
7. Needle Storage Phosphors for Medical CR
7.1. Needle Plate Manufacturing
7.2. Needle Phosphors for Medical Imaging
7.3. CR Needle Phosphors
Phosphor | Stimulating laser wavelengtd (nm) | Peak PSL wavelengtd (nm) | Peak stimulation wavelengtd (nm) | CE pJ/mm2/mR | SE μJ/mm2 |
BaFBr:Eu2+ | 633 | 390 | 550 | 20 | 16 |
BaFBr:Eu2+ | 680 | 390 | 550 | 14 | 28 |
RbBr:In+ | 680 | 490 | 700 | 2 | 25 |
RbBr:Ga+ | 680 | 550 | 705 | 6 | 4 |
CsBr:In+ | 680 | 504 | 700 | 3 | 23 |
CsBr:Ga+ | 680 | 515 | 685 | 6 | 4 |
8. The CsBr:Eu2+ PSL Mechanism
9. Image Quality of CR Systems
9.1. Definition of Medical Radiography System’s Image Quality
9.1.1. Signal-to-Noise Ratio (SNR)
9.1.2. Detective Quantum Efficiency (DQE)
9.1.3. Noise Sources
9.1.4. Noise as a Function of Spatial Frequency
9.1.5. Resolution and Modulation Transfer Function (MTF)
9.1.6. DQE as a Function of Spatial Frequency
9.2. Measurement of Image Quality
9.2.1. Measurement of MTF
9.2.2. Measurement of W
9.2.3. Measurement of Air Kerma and Translation to Number of Quanta
9.3. Relation between Imaging Plate Characteristics and Image Quality (DQE)
9.3.1. Simple DQE Model
9.3.2. Relation between DQE and IP Characteristics
- The diameter of the laser beam halo increases with the depth in the layer. Hence, spatial resolution and MTF in CR decrease with increasing phosphor layer thickness.
- Another factor of influence is the color of the substrate on which the phosphor layer is coated. The more the substrate absorbs the stimulation light, the higher MTF will be, because reflected stimulation light will diffuse further in the lateral direction, thereby increasing the size of the stimulating halo.
- Resolution can be improved by mixing anti-halo dye or pigment in the phosphor layer or top-coat. This, obviously, limits lateral light diffusion.
- The shorter the distance between two light scattering events, i.e., the shorter the mean free path of the photons in the phosphor layer, the shorter will be the distance over which light can travel before being absorbed. This explains why the resolution offered by PIP improves with decreasing phosphor particle size and in NIP with decreasing needle diameter.
9.4. Technical and Clinical Image Quality of CR Systems with PIP and NIP
- (1) The Institute for Clinical Radiology in Munich evaluated image quality and anatomical detail depiction [112]. Dose-reduced radiographs were made with NIP and compared to full dose PIP images. 24 Supine chest radiographs were made with PIP at standard dose and compared to follow-up studies with NIP in which the dose was reduced by 50%. In a blind study, 6 independent readers rated the PIP and dose-reduced NIP images equally. The same group also evaluated the low-contrast performance of DX-S® compared to that of a PIP CR system [113]. A total of 36 images of a CDRAD phantom were made using nine different exposure conditions. A blind study of the images by five radiologists and five physicists lead to the conclusion that for all but two of the exposure settings NIP allowed visualization of significantly lower contrast levels. The remaining two settings also showed a trend toward better low contrast depiction with NIP.
- (2) A comparison of imaging performance with different doses in skeletal radiography was made for a PIP based CR system, DX-S® and a DR system [114]. The DR system was a Siemens flat-panel detector based on a CsI:Tl+ scintillator layer and an amorphous silicon sensor layer. Five independent blinded radiologists evaluated 72 plain radiographs of the feet of six human cadavers obtained with four surface entrance doses. The conclusion was that the radiation dose can be reduced by 75% in clinical skeletal imaging of peripheral extremities when NIP is used instead of PIP in CR. The DR system allowed a dose reduction of 50%.
- (3) Physical image quality was compared for two CR and two flat-panel DR systems [115]. The CR systems were DX-S® and the PIP based Agfa ADC Compact Plus®. The first DR system, DR1, was a General Electric Revolution XQ/I® with an a-Si panel and a CsI:Tl+ scintillator; the second, DR2, was a Philips Diagnost®, based on the same technology. Image quality was assessed with a contrast-detail object and acrylic material to simulate clinical conditions. Important image quality differences were observed. DX-S® and DR2 showed similar image quality and were superior compared to DR1 and ADC Compact Plus®. It was further concluded that DX-S® provides better low-contrast detectability and a potential for dose reduction and that for doses over 0.2 mGy it provides even better image quality than DR1 and DR2.
- (4) Smans et al. tested CR image quality in neonatal imaging [116]. For typical acquisition parameters of neonatal chest X-ray examinations, the threshold-contrast detectability in simulated and acquired images of a contrast-detail phantom was compared. This was done for the Agfa ADC Compact Plus® with PIP and for DX-S®. Good agreement was found between the threshold-contrast curves of the simulated and experimentally acquired images and the superiority of NIP for neonatal imaging was confirmed.
- (5) Ludewig et al. simulated the conditions in neonatal radiological practice by making thoracic radiographs of cats. They compared DX-S® to ADC Compact Plus®. The conclusion was that a dose reduction of 50% seems possible without relevant deterioration of image quality [117].
10. Outlook
11. Conclutions
Acknowledgements
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Leblans, P.; Vandenbroucke, D.; Willems, P. Storage Phosphors for Medical Imaging. Materials 2011, 4, 1034-1086. https://doi.org/10.3390/ma4061034
Leblans P, Vandenbroucke D, Willems P. Storage Phosphors for Medical Imaging. Materials. 2011; 4(6):1034-1086. https://doi.org/10.3390/ma4061034
Chicago/Turabian StyleLeblans, Paul, Dirk Vandenbroucke, and Peter Willems. 2011. "Storage Phosphors for Medical Imaging" Materials 4, no. 6: 1034-1086. https://doi.org/10.3390/ma4061034
APA StyleLeblans, P., Vandenbroucke, D., & Willems, P. (2011). Storage Phosphors for Medical Imaging. Materials, 4(6), 1034-1086. https://doi.org/10.3390/ma4061034