1. Introduction –Why digital? –Why dual energy? Experimental setup Image acquisition Image processing and results A silicon microstrip system with the RX64DTH ASIC for dual energy radiology

2. 1) University of Eastern Piedmont and INFN, Alessandria, Italy L. Ramello; 2) University and INFN, Torino, Italy P. Giubellino, A. Marzari-Chiesa, F. Prino; 3) University and INFN, Ferrara, Italy; M. Gambaccini, A. Taibi, A. Tuffanelli, A. Sarnelli; 4) University and INFN, Bologna, Italy G. Baldazzi, D. Bollini; 5) AGH Univ. of Science and Technology, Cracow, Poland W. Dabrowski, P. Grybos, K. Swientek, P. Wiacek; 6) University of Antwerp, Antwerp, Belgium P. Van Espen; 7) Univ. de los Andes, Colombia C. Avila, J. Lopez Gaitan, J.C. Sanabria; 8) CEADEN, Havana, Cuba A.E. Cabal, C. Ceballos, A. Diaz Garcia, L. Bolaños; 9) CINVESTAV, Mexico City, Mexico L.M. Montano; The Collaboration

3. Introduction: why digital ? Digital radiography has well known advantages over conventional screen-film systems –Enhance detecting efficiency w.r.t. screen-film –Image analysis –Easy data transfer

4. Dual energy techniques GOAL: improve image contrast Based on different energy dependence of different materials Enhance detail visibility (SNR) Decrease dose to the patient Decrease contrast media concentration Introduction: why dual energy ?

5. Example 1: dual energy mammography

6. E  15-20 keV: Signal from cancer tissue deteriorated by the adipose tissue signal E  30-40 keV Cancer tissue not visible, image allows to map glandular and adipose tissues

7. Example 2: angiography Angiography = X-ray examination of blood vessels  determine if the vessels are diseased, narrowed or blocked  Injection of a contrast medium (Iodine) which absorbs X-ray differently from surrounding tissues Coronary angiography  Iodine must be injected into the heart or very close to it  A catheter is inserted into the femoral artery and managed up to the heart → Long fluoroscopy exposure time to guide the catheter → Invasive examination Why not to inject iodine in a peripheral vein?  Because lower iodine concentration would be obtained, requiring longer exposures and larger doses to obtain a good image  But, if the image contrast could be enhanced in some way…

8. Example 2: angiography at the iodine K-edge (II) Iodine injected in patient vessels acts as radio-opaque contrast medium Dramatic change of iodine absorption coeff. at K-edge energy (  33 keV) Subtraction of 2 images taken with photons of 2 energies (below and above the K-edge) → in the resulting image only the iodine signal remains and all other materials are canceled

9. Experimental setup To implement dual energy imaging we need: a dichromatic beam a position- and energy-sensitive detector Quasi-monochromatic beams ordinary X-ray tube + mosaic crystals instead of truly monochromatic synchrotron radiation Advantages: cost, dimensions, availability in hospitals Linear array of silicon microstrips + electronics for single photon counting Binary readout 1 or 2 discriminators (and counters) per channel Integrated counts for each pixel are readout Scanning required to build 2D image

10. Experimental setup: beam (1) Bragg Diffraction on Highly Oriented Pyrolitic Grafite Crystal W anode tube 1st and 2nd Bragg harmonics  E and 2E are obtained in the same beam Collimator

11. Experimental setup: beam (2) Bragg Diffraction on Highly Oriented Pyrolitic Grafite Crystal W anode tube Double slit collimator Two spatially separated beams with different energies  E-  E and E+  E obtained in 2 separate beams

12. More on the dichromatic beam incident spectra at 3 energy settings … … spectra after 3 cm plexiglass (measured with HPGe detector)

13. Fully parallel signal processing for all channels Binary architecture for readout electronics  1 bit information (yes/no) is extracted from each strip  Threshold scans needed to extract analog information Counts integrated over the measurement period transmitted to DAQ Fully parallel signal processing for all channels Binary architecture for readout electronics  1 bit information (yes/no) is extracted from each strip  Threshold scans needed to extract analog information Counts integrated over the measurement period transmitted to DAQ data, control Silicon strip detectorIntegrated circuit 100  m current pulses X-rays PC N. I. I/O cards PCI-DIO- 96 and DAQCard-DIO-24 Experimental setup: Single Photon Counting System

14. Experimental setup: PCB detector pitch adapter ASICs PCB: - One 400 strip detector - Pitch adapter - 6 RX64 chips  384 equipped channels - connector to DAQ card 2 protoype detectors: a)6 x Single threshold RX64 b)6 x Dual threshold RX64 PCB: - One 400 strip detector - Pitch adapter - 6 RX64 chips  384 equipped channels - connector to DAQ card 2 protoype detectors: a)6 x Single threshold RX64 b)6 x Dual threshold RX64

15. Detecting system Chip RX64 → counts incident photons on each strip of the detector 4 cm 6.4 mm 10 strip = 1 mm micro-bondings Silicon microstrip detector each strip is an independent detector which gives an electric signal when an X- ray photon crosses it and interacts with a silicon atom Knowing from which strip the electric signal comes from,the position of the incoming X-ray phonton is reconstructed.

16. Why silicon detectors? Main characteristics of silicon detectors: speed of the order of 10 ns spatial resolution of the order of 10  m small amount of material  0.003 X 0 for a typical 300  m thickness excellent mechanical properties good resolution in the deposited energy  3.6 eV of deposited energy needed to create a pair of charges, vs. 30 eV in a gas detector

17. Silicon sensor diode The impinging ionizing particles generate electron-hole pairs The impinging photons which interact in the detector volume create an electron (via Photoelectric, Compton or Pair Production) The electron ionizes the surrounding atoms generating electron-hole pairs Electron and holes drift to the electrodes under the effect of the electric field present in the detector volume. The electron-hole current in the detector induces a signal at the electrodes on the detector faces. Metal contact n+-type implant n-type bulk Charged particle -V +V electron hole P+-type implant photon photoelectron Reverse bias E

18. Why reverse biased diode? The amount of charge deposited in the silicon detector is very small ≈5500 electrons are produced by a 20 keV photons making photoelectric effect in the silicon  Forward-biased junction: the signal would be masked by the fluctuations of the current which the applied field makes flow even in high resistivity, hyper-pure silicon.  Reverse-biased junction: allows to obtain the necessary electric field and only a very small “dark” current also at room temperature. -V +V depleted region Increasing the polarization voltage, it is possible to extend the depletion layer down to the backplane. To have full efficiency, the polarization voltage must be high enough to deplete the full detector thickness (typically 300  m) junction NOT GOOD

19. Silicon Microstrips detectors A micro-strip detector is a silicon detector segmented in long, narrow elements. Each strip is an independent p-n reverse-biased junction Provides the measurement of one coordinate of the particle’s crossing point with high precision (down to 1  m). N-type substrate P+ n+ Al P+ SiO 2 AC coupling to electronics SiO 2 Al DC coupling to electronics

20. Experimental setup: silicon detector ParameterValue Depth300 μm Strip length10 mm Number of strips400 Strip pitch100 μm Depletion voltage20-23 V Leakeage curr. (22º C)50-60 pA Inactive region thickn.765 μm Designed and fabricated by ITC-IRST, Trento, Italy

21. Experimental setup: RX64 chip Cracow U.M.M. design - ( 2800  6500  m 2 ) - CMOS 0.8 µm process (1) (1) 64 front-end channels a) preamplifier b) shaper c) 1 or 2 discriminators (2) (2) (1 or 2)x64 pseudo-random counters (20-bit) (3) (3) internal DACs: 8-bit threshold setting and 5-bit for bias settings (4) (4) internal calibration circuit (square wave 1mV-30 mV) (5) (5) control logic and I/O circuit (interface to external bus) Cracow U.M.M. design - ( 2800  6500  m 2 ) - CMOS 0.8 µm process (1) (1) 64 front-end channels a) preamplifier b) shaper c) 1 or 2 discriminators (2) (2) (1 or 2)x64 pseudo-random counters (20-bit) (3) (3) internal DACs: 8-bit threshold setting and 5-bit for bias settings (4) (4) internal calibration circuit (square wave 1mV-30 mV) (5) (5) control logic and I/O circuit (interface to external bus) 12 34 5 Detector

22. Detector efficiency Front geometry –Strip orthogonal to the beam –70  m of Al light shield Edge-on geometry –Strip parallel to the beam –765  m of inactive Si –Better efficiency for E > 18 keV Efficiency calculation –X-ray absorbed if interacts in passive regions –X-ray detected if makes photoelectric effect in active regions

23. Imaging test 1-dimensional array of strips → 2D image obtained by scanning Cd-109 source (22.24 keV) Detector Collimator (0.5 mm) Test Object 5 mm

24. Imaging test 1-dimensional array of strips → 2D image obtained by scanning Scanning

25. System calibration setup in Alessandria Detector in Front config. Fluorescence target (Cu, Ge, Mo, Nb, Zr, Ag, Sn) Cu anode X-ray tube → X-ray energies = characteristic lines of target material

26. Cu K  Mo K  Ge K  Rb K  Ag K  Sn K  Ag K  Mo K  Sn K  SystemTpTp GAIN  V/el. ENCEnergy resolution 6 x RX64 0.7  s 64≈170 el.≈0.61 keV 6 x RX64DTH 0.8  s 47≈ 200 el.≈0.72 keV 241 Am source with rotary target holder (targets: Cu, Rb, Mo, Ag, Ba) Cu-anode X-ray tube with fluorescence targets (Cu, Ge, Mo, Ag, Sn) System calibration

27. K-edge subtraction imaging with contrast medium Cancel background structures by subtracting 2 images taken at energies just below and above the K-edge of the contrast medium Suited for angiography at iodine (gadolinium) K-edge -Cancel background structures to enhance vessel visibility Possible application in mammography (study vascularization extent) -Hypervascularity characterizes most malignant formations Dual energy projection algorythm Make the contrast between 2 chosen materials vanish by measuring the logarithmic transmission of the incident beam at two energies and using a projection algorithm [Lehmann et al., Med. Phys. 8 (1981) 659] Suited for dual energy mammography –remove contrast between the two normal tissues (glandular and adipose), enhancing the contrast of the pathology –Single exposure dual-energy mammography reduces radiation dose and motion artifacts Dual energy imaging

28. X-ray tube with dual energy output Phantom Detector box with 2 collimators 1.X-ray tube + mosaic crystal and 2 collimators to provide dual-energy output - E1= 31.5 keV, E2 =35.5 keV (above and below iodine k-edge) 2.Detector box with two detectors aligned with two collimators 3.Step wedge phantom made of PMMA + Al with 4 iodine solution filled cavities of 1 or 2 mm diameter Angiography setup

29. E = 31.5 keV E = 35.5 keV logarithmic subtraction Phantom structure not visible in final image Angiographic test results (I)

30. Conc = 370 mg / ml Conc = 92.5 mg / ml Conc = 23.1 mg / ml Possible decrease of iodine concentration keeping the same rad. dose Angiographic test results (II)

31. Results with a second phantom Phantom Digital Subtraction Angiography Dual Energy Angiography smaller cavity (  =0.4 mm) visible in DEA and not in DSA Iodine conc. = 95 mg/ml

32. Dual energy projection algorithm The mass attenuation coefficient μ of any material  at a given energy E is expressed as a combination of the coefficients of any two suitable materials  and  : The logarithmic attenuation M = μ ξ t ξ of the material of thickness t ξ is measured at two different energies: low (E l ) and high (E h ): A 1 and A 2 represent the thicknesses of the two base materials which would provide the same X-ray attenuation as material ξ.

33. C  C 90°  M1M1 R 1   M2M2 2 If a monochromatic beam of intensity I 0 goes through material ξ which is partly replaced by another material ψ … I0I0 I1I1 I2I2 ξ ψ … then the vertexes of log. attenuation vectors M 2 (material ξ ) and M 1 (mat. ξ + ψ ) lie on a line R which is defined only by the properties of materials α, β, ξ and ψ. Projecting along direction C, orthogonal to R, with the contrast cancellation angle  : … it is possible to cancel the contrast between materials ξ and ψ : both M 1 and M 2 will project to the same vector A2A2 A1A1 Dual energy projection algorithm The logarithmic attenuation M in a given pixel can be represented as a vector having components A1 and A2 in the basis plane, the modulus will then be proportional to the gray level of that pixel

34. Mammographic phantom Three components: polyethylene (PE), PMMA and water to simulate the attenuation coeff.  (cm -1 ) of the adipose, glandular and cancerous tissues in the breast  S. Fabbri et al., Phys. Med. Biol. 47 (2002) 1-13 E  _fat  _gland  _canc 20.456.802.844 40.215.273.281 E μ_PEμ_PMMAμ_water 20.410.680.810 40.225.280.270

35. Image processing (1) Low thr.High thr. Measured (raw) 16 keV32 keV HE and LE imagesCorrect for: 1.pixels with huge n. of counts (bad counter conversion) 2.dead pixels 3.X-ray beam fluctuations 4.subtract high threshold image from low threshold one 5.correct for spatial inhomogeneities of beam and detector extracted from flat-field profiles

36. 16 – 32 keV 18 – 36 keV 1= PMMA 2=water 3=PE 4=(water+PE) Image processing (2) Low statistics due to: 1)2 nd order harmonic 2)dectecting efficiency

37. Simulation with MCNP 1=detector 2=PMMA 3=water 4=PE MCNP-4C simulation with ENDF/B-VI library Photons and electrons tracked through the phantom and the detector (including the inactive region in front of the strips) Energy deposition in each strip recorded histogram of counts vs. strip number filled Top View Side View

38. Experiment vs. Simulation (1) RX64DTH 16 – 32 keV simulation 16 – 32 keV

39. Experiment vs. Simulation (1)

40. Results (1): SNR vs. proj. angle RX64DTH 16 – 32 keV MCNP simulation Cancellation angle for a pair given by SNR=0 Theoretical cancellation angles: PMMA-water 36.5° PE-water 40.5° PMMA-PE 45°

41. Results (2): SNR summary EnergyCanceledContrastSNR (keV)materialsmaterialRX64*RX64DTH PMMA-water PE8.119.63 16-32 PE-water PMMA2.533.19 PE-PMMA water3.964.72 PMMA-water PE7.435.14 18-36 PE-water PMMA2.702.10 PE-PMMA water3.853.13 PMMA-water PE2.553.27 20-40 PE-water PMMA0.671.07 PE-PMMA water0.891.58 * Previous version of ASIC, exposure with about 2x more incident photons

42. Results (3): Projected images RX64DTH 16 – 32 keVsimulation 16 – 32 keV PMMA-water cancellation PMMA-PE cancellation

43. Conclusion and Outlook We have developed a single photon counting silicon detector equipped with the RX64DTH ASIC, with two selectable energy windows The energy resolution of 0.8 keV (rms) is well adapted for dual energy mammography and angiography We have performed mammography imaging tests with a three-material phantom –We have demonstrated the feasibility of contrast cancellation between two materials, enhancing the visibility of small features in the third one We have performed angiography imaging tests with 2 different phantoms and iodine contrast medium –We have demonstrated the feasibility of logarithmic subtraction between two images, enhancing contrast of small vessels also with lower iodate solution concentrations OUTLOOK: –Increase photon statistics at high energy, optimize exposure conditions –New detector materials, CZT? –Tests with a more realistic mammographic phantom