Flat-Panel Detectors

Ho Kyung Kimhokyung@pusan.ac.kr

Pusan National University

Lectures on Digital RadiographyCh. 4 HMI

References

J. A. Rowlands and J. Yorkston, "Flat Panel Detectors for Digital Radiography," in Handbook of Medical Imaging, J. Beutel, H. L. Kundel, and R. L. Van Metter, Eds., Bellingham, WA, USA: SPIE, 2000, ch. 1, pp. 223-328.

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X-ray detection media Flat-panel array technology Configuration and operation of a flat-panel x-ray imager Methods of evaluating performance Clinical applications of complete systems Future prospects

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Routine• Load a film into a cassette (in the dark room)• Carry it to the examination room• Insert it into the x-ray table• Position the patient• Make the x-ray exposure• Carry the cassette back to the processor to develop the film• Check if the film is suitable for making a medical diagnosis• Hang the film in a view box

Key issue in the development of digital diagnostic radiology• Large-area availability of a digital sensor enough to cover the patient body• Flat-panel active matrix array, originally developed for laptop-computer displays

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General concepts

Digitization• Sampling in space => pixels checkboard artifacts• Quantization in intensity => bits contouring artifacts

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Dynamic range• Ratio between x-ray attenuation of the most radiolucent & the most radio-opaque paths

– 𝐼𝐼0𝐼𝐼

= 𝑒𝑒𝜇𝜇𝑡𝑡

• Precision of the x-ray signal measured in the most radio-opaque anatomy– e.g., 1% precision in the signal attenuated by a factor of 50 results in the dynamic range of 5000

– Chest: 3000 𝜇𝜇R6 𝜇𝜇R

= 500, mammography: 240mR60 𝜇𝜇R

= 4000, & fluoroscopy: 10 𝜇𝜇R0.1 𝜇𝜇R

= 100

– Corresponding to 54 dB, 72 dB, & 40 dB or 10 bits, 12 bits, & 7 bits with the noise level of 1 LSB

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Imaging field (or field of view, FOV)• Scanning

– Scanning with a 1D array detector– Scanning with a multi-line array (slot) detector– Excellent scatter rejection– Heavy load of x-ray tube

• Optical coupling– Lens, tapered fiber-optic bundles– Secondary quantum noise: secondary quantum sink (𝑁𝑁𝑋𝑋 < 𝑁𝑁𝑜𝑜𝑜𝑜𝑡𝑡) due to optical inefficiency

• Mosaic– Fixed pattern noise, high dark current– Stitching problems

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Flat-panel detectors

Traditional x-ray detection materials + large-area active matrix readout structure

• X-ray interactions with phosphors & photoconductors to generate a measurable response

• Storage of the response with a recording device

• The measurement of the stored response

Integrating the incoming signal over a finite period of time

• X-ray fluence detector or charge-integration detector

• Not a photon-counting detector

Pixel = switch + sensor/storage

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Operation

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Energy band structures

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Resolution losses

Main intrinsic causes• Geometrical blurring

– Oblique incidence of x-ray beam

• Electron-range blurring– ~1 − 3 𝜇𝜇m at 10-30 keV– ~10 − 30 𝜇𝜇m at 50-100 keV

• K-fluorescence reabsorption– ~50 − 200 𝜇𝜇m

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Independence of a-Se thickness on MTF (due to high E-field)

Mainly due to electron range

Mainly due to K fluorescence

Photoconductors

Typical properties• 𝐸𝐸𝑔𝑔~ 2 eV; insulator with negligible free carriers at room temp.

• Sufficient lifetime of radiation-induced carriers to reach the surface of the detection vol.• High E-field => fast traverse of charges => less time for lateral diffusion => high resolution• 𝑊𝑊~3𝐸𝐸𝑔𝑔 to release an e-h pair

Amorphous selenium• 𝑊𝑊~ 50 eV at typ. 10 V/µm due to (germinate & columnar) recombination of e-h pairs

– Decrease with E-field

• Blocking contact (limited by the field strength)• Low surface (transverse) conductivity (by introducing a high density of traps)• a-Se:0.5%As to prevent from crystallization; adversely, As causes large density of ℎ traps (short 𝜏𝜏ℎ)• a-Se:0.5%As + 10-20 ppm Cl (stabilized a-Se) to reduce ℎ traps

– 𝜇𝜇ℎ = 0.13 cm2V−1s−1; 𝜇𝜇𝑒𝑒 = 0.003 − 0.006 cm2V−1s−1

– 𝜏𝜏ℎ = 50 − 500 𝜇𝜇s; 𝜏𝜏𝑒𝑒 = 100 − 1000 𝜇𝜇s– 𝑆𝑆ℎ~6.5 − 65 mm; 𝑆𝑆𝑒𝑒~0.3− 3 mm @ 10 V/µm (Schubweg, 𝑆𝑆 = 𝜇𝜇𝜏𝜏𝐸𝐸)

• 𝑍𝑍 = 34

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Other possible photoconductors

PbI2 & PbO have been used for nuclear radiation detectors & vidicon TlBr has a high ionic conductivity, causing large dark current Further candidates include CdZnTe, CdTe, & HgI2

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Phosphors

Conversion gain (or quantum amplification)• Deexcitation of conduction-band electrons through activators, emitting light (~2-3 eV)

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X-ray screens

Highly scattering (or turbid) rather than transparent to light• Phosphor powder (high refractive index) + binder (optically transparent)• Trade-off between the phosphor thickness (x-ray interaction efficiency) & spatial resolution

Factors affecting image quality• Phosphor grain size, size distribution, bulk absorption, & surface reflectivity• Back-screen configuration improves the spatial resolution compared to the front-screen one• FPDs are configured in the front-screen design because of the thickness (~0.7 mm) of glass

substrate

Poor light-collection eff. can limit the overall performance of the complete system

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A. Transparent w/ reflective backingB. Transparent w/ absorptive backingC. Turbid w/ reflective backingD. Turbid w/ 50% abs. in bulkE. Turbid w/ 50% abs. in abs. backingF. Ideal

Swank's model Design C

Structured phosphors

CsI• Needle-like closely packed crystallites => pillar-like (columnar) structure (~10 µm, reduced density

of ~80-90% of single crystal)• Fiber-optic light guide (𝑛𝑛CsI = 1.78 > 𝑛𝑛air = 1): 83% internal reflection• Activator controls the emission spectrum

– Na: blue (~450 nm)– Tl: green (~550 nm) well matched to the response of a-Si:H layers

• Hygroscopic & mechanically weak

Fiber-optic faceplates with scintillation impurities

Micro-channel array filled with scintillators

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Higher light output

Noise in x-ray converters

Mainly related to the x-ray exposure Further degraded by

• Lack of x-ray absorption (quantum-absorption inefficiency)

– 𝛼𝛼 𝐸𝐸 = 1 − 𝑒𝑒−𝜇𝜇 𝐸𝐸 𝑡𝑡

• Fluctuations in the detector response to the x-ray absorption

– Gain-fluctuation noise or the Swank noise factor, 𝐼𝐼• DQE 0 = quantum absorption inefficiency ×

gain fluctuation = 𝛼𝛼 × 𝐼𝐼

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Measurement of the Swank noise

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Depth-dependent signal transfer• Known as the Lubberts effect• Decreasing DQE as 𝑓𝑓 increases

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Fabrication

Amorphous Si by PECVD• 𝜇𝜇𝑒𝑒 = ~1 cm2V−1s−1; 𝜇𝜇ℎ = 0.003 cm2V−1s−1

• Microcrystalline Si with an order of magnitude higher mobilities• Polycrystalline Si with an order of magnitude higher again• Crystalline Si: 𝜇𝜇𝑒𝑒 = ~1300 cm2V−1s−1; 𝜇𝜇ℎ = ~500 cm2V−1s−1

Difference in the pixel design compared to the conventional LCD process• p+ deposition; thick i layer (≥ 1 µ𝑚𝑚)

Photolithography

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Indirect-conversion pixel

Photodiode• Light absorption coeff. of a-Si:H (~104 to ~106 cm-1) > that of c-Si by an order of magnitude

– a-Si:H 𝐸𝐸𝑔𝑔 = ~1.7 eV whereas c-Si 𝐸𝐸𝑔𝑔 = ~1.1 eV– ~0.5 µm-thick i-layer is sufficient to absorb most visible photons– 1~2 µm is typical to reduce pixel capacitance

• Photoconductive gain mode with ohmic contacts– Large dark current & low dynamic range

• Blocking contacts preventing the injection current

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pin photodiode• ITO (~50 nm) + p+ (µc-Si:H alloy) (~10 − 20 nm) + i (~1.5 µm) + n+ (~10 − 50 nm)

– Lower dark current (~10−11 A cm−2)

Schottky photodiode• No p+ layer

– Higher quantum efficiency; more compatible with the AMLCD process– Higher dark current (~10−9 A cm−2)

MIS photodiode• Insulator instead of the p+ layer; hence reverse n-i-p configuration

– Requiring the refresh cycle to remove hole build-up (switching polarity)

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Direct-conversion pixel

Charge storage capacitor (insulator) + collection electrode (metal) Require an electric field strength ~10 V µm−1

• Additional designs protecting from the high voltage damage

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~1 pF (~10 V)

~5 pF cm-2 => 0.002 pF for 200 µm pixel (~4990 V)

Switch

Diodes• Improvement in device yield because of the same as the sensor• Smaller in area (~5%), increasing pixel fill factor• Larger resistance at lower on bias voltage, causing image lag• Large charge transient (i.e., pixel cross-talk) during switching (due to higher capacitance)

TFT• Structure

– Gate dielectric a-Si3N4:H/intrinsic a-Si:H/drain & source n+ a-Si:H/passivation– Additional metallic light shield

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• On (~+10 – 20 V)/off (~-5 – 10 V)– 𝐼𝐼DSoff~0.1 − 1 × 10−15 A per W in µm– 𝑅𝑅on~1 MΩ; independent of 𝑉𝑉DS in contrast to

the diode switch

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• Coupling capacitance due to geometrical overlaps– Depending on the exact details of the array designs– Can be a significant contributor to the total data line capacitance (seen by the external electronics)

• Charge is injected into the sensor & onto the data line– Small changes in the gate voltage along a data line can result in line-correlated noise– Fluctuations in the bias voltage are coupled into the data line– Changes in the threshold voltage, 𝑉𝑉T

Array design

Aperture: the dimension of active portion of each detector element• Determine the spatial frequency response of the detector; MTF 𝑢𝑢 = sinc(𝜋𝜋𝑎𝑎𝑢𝑢)

Pitch: the sampling interval of the detector• Nyquist frequency = (2𝑝𝑝)−1

• Aliasing• Pre-sampling blurring to reduce aliasing

Fill factor• Geometrical fill factor: the fraction of the pixel area sensitive to the incoming signal• Effective fill factor

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Noise source intrinsic to array

kTC switching (thermal noise on capacitors)• 𝜎𝜎kTC = 2𝑘𝑘𝑘𝑘𝐶𝐶PD/𝑒𝑒• 𝜎𝜎kTC = ~560 𝑒𝑒− if 𝐶𝐶PD = 1 pF @ room temp.

TFT channel resistance (related to the bandwidth of the readout circuit)• Thermal (Johnson) noise• 1/f flicker noise

Dark current• Shot noise• 1/f noise

Distributed resistance of the metallic lines

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Peripheral electronics

Circuitry that supplies the required voltages to the array elements, and amplifies & digitizes the signals from the pixels

Potential noise sources tending to reduce the quality of the final image

Preamplifier• Charge-integrating design (CSPA)• Feedback capacitor determines the electronic gain (mV/pC)

• 𝜎𝜎amp = noise floor 𝑒𝑒− + slope 𝑒𝑒−

pF× 𝐶𝐶data = 400 + 4.5𝐶𝐶in

Correlated-double sampling

Bias voltage• Line-correlated noise to which human vision is extremely sensitive

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Image processing

To obtain a diagnostic-quality radiograph, remove artifacts & adjust the appearance of the raw image information

Flat-fielding correction• Correction of variations in pixel sensitivity & offset (due to thickness & quality of layers)• Dark-field images for variations in pixel & electronic offsets• Flood-field images for variations in pixel sensitivity & electronic gain

Defects• Thresholding to identify defects• Difficult to set up the threshold limits for partially functional or nonlinear response pixels• Median filter

– Difficult to correct the clusters of bad pixels/several adjacent bad lines

Image lag & ghosting

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Performance

Dynamic range

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𝑆𝑆 =100

𝑋𝑋 (in mR)

Imaging performance

DQE 𝑢𝑢 = MTF2(𝑢𝑢)𝑊𝑊(𝑢𝑢) �𝑞𝑞

– 𝑊𝑊(𝑢𝑢) = normalized NPS

• A basis to determine the physical principles involved in the detectors

MTFsys 𝑢𝑢 = MTFconv 𝑢𝑢 × MTFaper 𝑢𝑢• MTFdir 𝑢𝑢 ≈ MTFaper 𝑢𝑢

– ~60% @ 𝑢𝑢Ny– Noise aliasing– 𝛾𝛾 = (𝑢𝑢1𝑝𝑝)−2; 𝑢𝑢1 = the first zero freq.

• MTFind 𝑢𝑢 ≈ MTFconv 𝑢𝑢– ~10% @ 𝑢𝑢Ny

NPSdir 𝑢𝑢 = const. NPSind 𝑢𝑢 ~𝑢𝑢−1

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NPS shows a marked exposure dependence

DQE increases with exposure and a plateau is finally reached where further increase in exposure makes no difference to the DQE

DQECsI 0 > DQESe 0 ; DQECsI 𝑢𝑢Ny < DQESe 𝑢𝑢Ny

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System performance modeling

Fill factor• Sub-unity 𝛾𝛾 reduces DQE(0) of the direct

detector by the same factor, whereas no loss in the indirect detector due to sharing of signals from each x-ray photon with many pixels

Aliasing• Always present in Se detectors• Reduction/removal by blurring prior to

pixel sampling– Inevitable in reduction of the high-freq.

components of signal– Much more susceptible to external noise

(e.g., electronic noise or secondary quantum statistics)

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DQE 𝑢𝑢 losses

• Lubberts effect– Due to the depth-dependent MTF characteristics in an x-ray conversion layer (usually phosphors)– Less or none effective in Se & fiber-optic scintillation layer– Vulnerable to aliasing– Requiring anti-aliasing blurring layer

• K-fluorescence reabsorption– Partial energy absorption with K-fluorescence escape or reabsorption at remote site give rise to substantial

blurring & noise

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Clinical applications

Chest radiography• Difficult because of necessity of large dynamic range considering very radio-lucent (lung fields) &

very radio-opaque (mediastinum) regions• Possible solution with very highly penetrating x-ray beams (130–150 kVp)• Dual-energy imaging to isolate the bony details (i.e., the spine & ribs) from the soft tissues

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Fluoroscopy• The most demanding potential application for

FPDs• Must be x-ray quantum-limited even at

extremely low exposure levels • XRII

– Bulky nature, veiling glare, geometric distortion

• FPDs– Reduction of noise to permit quantum-noise

limited operation at the low end of fluoroscopic rates (i.e., 0.1 µR s−1)

– Reduction of the image carry over or lag

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Mammography• X-ray projection to visualize soft-tissue contrast• Breast compression to equalize the x-ray path length to the point that the whole breast can be

visualized• Film followed by screen to ensure the highest possible image resolution• FPDs with high dynamic range is compatible to visualize dense breasts

Tomography• Blurs out the shadows of superimposed structures to allow better isolation of the structures of

interest• Digital tomosynthesis• Volumetric CT

– Improved z-direction resolution– Problem with the increased level of scatter

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Portal imaging• To confirm the correct positioning of the patient in the output portal of the therapy machine• Not possible to see soft-tissue contrast & difficult to visualize bones because of inadequate

contrast/resolution• Video-based systems

– Secondary quantum sink where each x-ray photon is represented by less than a single light photon– Dominance of electronic noise

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Future prospects

Increased numbers of active elements per pixel, allowing an amplifier at every pixel and integrated readout electronics to make an x-ray imager on glass

Increased x-ray to charge conversion gain (lower W or avalanche gain) More sophisticated switching structures with reduced coupling capacitance, lower leakage

current, smaller physical area, & more robust operating characteristics

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