1. 13-15 July 2005 L. Ramello 1 Medical Imaging with Semiconductor Detectors CINVESTAV 2005 Advanced Summer School L. Ramello – Dip. Scienze e Tecnologie Avanzate, Univ. Piemonte Orientale, ALESSANDRIA (Italy)

2. 13-15 July 2005L. Ramello2 Topics Basic properties of semiconductor detectors Image quality: contrast, SNR, MTF, DQE Recent detector developments: – MEDIPIX (2D pixels) – SYRMEP (Synchrotron Light Source) – High Z semiconductors Dual Energy Mammography Dual Energy Angiography

3. 13-15 July 2005 L. Ramello 3 Basic properties of semiconductor detectors

4. 13-15 July 2005 L. Ramello 4 Advantages for medical imaging with x-rays: – High spatial resolution (down to ~50 micron) – High detection efficiency, especially in the low energy range (mammography) – Combine x-ray conversion and electrical signal generation – Decrease radiation dose and/or improve image quality Semiconductor imaging system concepts: –Digital radiography with scintillator + amorphous silicon (commercially available) –Digital radiography with direct conversion in semiconductor material (R & D) –PET and SPECT with high Z semiconductors (R & D) Why semiconductor detectors ?

5. 13-15 July 2005L. Ramello5 Semiconductor materials Atomic number Z, density and thickness  probability of x-ray photon conversion Average energy loss to create electron-hone pair, W (roughly proportional to E g )  energy resolution MaterialZdensity, g/cm 3 E g, eV W, eV Si142.331.123.6 GaAs31,335.311.424.2 Ge325.320.732.9 Se344.3 1.71-1.75 5.6 CdTe48,526.201.525.0 HgI 2 80,536.362.136.7 PbI 2 82,536.22.317.2

6. 13-15 July 2005L. Ramello6 Energy resolution Radiation ionization energy (W): determines the number of primary ionization events Band gap energy (E g ): lower value  easier thermal generation of e-h pairs (kT = 26 meV for T = 300 K)

7. 13-15 July 2005L. Ramello7 Semiconductor detectors To fully exploit these attractive semiconductor detector features: – High electric field is needed to collect signal – Dedicated, low noise electronics is needed (usually the first element is a charge amplifier) – For silicon, a p-n junction is needed to reduce dark current (operation at room temperature is OK) – For germanium, cryogenic operation (liq. N 2 temperature) is needed Multichannel systems require special care for power density, connection technique, cross-talk

8. 13-15 July 2005L. Ramello8 The p-n junction (1) Net charge density vs. distance Electric field vs. distance Electrostatic potential vs. distance Valence and Conduction band energies vs. distance Abrupt junction approximation

9. 13-15 July 2005L. Ramello9 The p-n junction (2) In reverse polarization (positive voltage to n-side): the diode current density saturates at a low value Js the depletion layer thickness (d) increases with increasing voltage, so does the active volume qV/kT = ratio between potential energy and thermal energy d = (2V B /  eN D ) 1/2  =  r  0  12  0 (Si) N D = donors/cm 3 (n-Si)

10. 13-15 July 2005L. Ramello10 The microstrip detector SIGNAL = number of electron- hole pairs: n e-h =  E/W, where W=3.62 eV for silicon REVERSE POLARIZED DIODE Depletion region => free from charge carriers: e-h pairs may be detected Reverse Bias voltage (V B ) => controls diode depletion thickness, i.e. active volume p-n junction capacitance per unit area C: 1/C 2 grows linearly with V B => C-V measurement determines full depletion voltage V FD

11. 13-15 July 2005L. Ramello11 A microstrip detector AC coupling: Bias Line and resistors to bias each strip, without shorting adjacent strips Guard ring(s) are essential to collect surface currents – This introduces a dead layer for edge-on geometry guard ring bias linefirst strip (AC contact) DC contact (to p+ implant)

12. 13-15 July 2005L. Ramello12 A readout chain This is just one possibility, the binary readout scheme – another one is to put an ADC instead of the discriminator, preserving the full analog information charge preamplifier shaperdiscriminator INOUT vthn vthp calib in

13. 13-15 July 2005L. Ramello13 The RX64 ASIC RX64 - Krakow UMM design - ( 2800  6500  m 2 ) consists of: - 64 front-end channels (preamplifier, shaper, discriminator), - 64 pseudo-random counters (20-bit), - internal DACs: one 8-bit threshold setting and and two 5-bit for bias, - internal calibration circuit (square wave 1mV-30 mV), - control logic, - I/O circuit (interface to external bus).

14. 13-15 July 2005L. Ramello14 Conversion efficiency (1) } 300 μm (standard thickness) } 10-20 mm μm (edge-on) Si (300 μm): efficiency drops to 50 % at 15 keV (Al window limits efficiency at low energies)  Recover efficiency with edge-on orientation

15. 13-15 July 2005L. Ramello15 Conversion efficiency (2) } 300 μm GaAs (Z ~32) } 10-20 mm μm Si (Z =14) GaAs (300 μm): efficiency drops to 50 % at 48 keV  Material of choice for mammography, E ~ 22 keV

16. 13-15 July 2005 L. Ramello 16 Image quality: contrast, SNR, MTF, DQE

17. 13-15 July 2005L. Ramello17 X-ray beams (1) X-rays are generated by bremsstrahlung of electrons emitted from cathode, accelerated by an applied voltage and impinging on the anode The energy spectrum of x-rays is determined by: –Peak kilovoltage (kVp) –Anode material (concerning peaks at characteristic energies) –Intrinsec and added filtration Effect on an 80 kVp x-ray beam of added filtration with a light material (Al) and with a rare earth material (La, K-edge @ 39 keV)

18. 13-15 July 2005L. Ramello18 X-ray beams (2) Most common anode materials: – W (Z=74) for general radiographu (chest, whole body, …) – Mo (Z=42) & Rh (Z=45) for mammography – Cu (Z=29) for diffractometry Energy emitted as x-rays is only 0.5-1% of input energy, the remaining part must be dissipated as heat X-ray tubes with moderate power are with fixed anode, high power ones have a rotating anode to avoid melting Typical currents are 1-5 mA for prolonged exposure (fluoroscopy) and 50-1000 mA for short exposures; exposure is measured in mAs

19. 13-15 July 2005L. Ramello19 X-ray imaging techniques Film: sensitivity is very low, it would require too high a dose to the patient Film + screen: conventional radiography Image intensifier (I.I.): fluoroscopy Photosensitive phosphor (computed radiography) Indirect digital radiography (I.I. or photoconductor coupled to a semiconductor) Direct digital radiography (semiconductor)

20. 13-15 July 2005L. Ramello20 Film + screen (1) X-rays transmitted through patient first screen second screen double coated emulsion / AgBr About 50% of the photons convert in the film-screen, mostly (95%) in the two screens The film exposure is mainly due to the blue-green light emitted by the phosphorescent screens (CaWO 4, Gd 2 O 2 S:Tb, etc.) Film-screen systems are classified according to their speed, with faster systems requiring less incident radiation to obtain same optical density The standard speed is = 100, slower (50) and faster (200, 400, 600) speed film-screen systems are commonly used

21. 13-15 July 2005L. Ramello21 Film + screen (2) X-ray absorption vs. energy by different screens Spectrum of primary and scattered x-rays from a tube operated at 80 kVp, with a Perspex (clear acrylic resin) phantom  usefulness of Gd screen to suppress scattered x-rays

22. 13-15 July 2005L. Ramello22 Exposure and optical density (1) Radiographic film blackening radiografico (mostly due to visible light emitted by screens) may be quantified by optical density (D): D = -log(T) where T is the transmission: T = I 1 /I 0 Useful optical density goes from 0.2 to 2.5-3.0 Exposure X quantifies the number of incoming x-rays

23. 13-15 July 2005L. Ramello23 Exposure and optical density (2) Relation between optical density D and exposure X: Transm. TD = -log(T) 1.0000.0100% transm. 0.7410.13base 0.1001.0good exposure 0.0102.0lung 0.0013.0very dark 0.00033.5maximum darkness 1) Film-screen: D = cX  highly non linear, constants depend on film speed 2) Electronic detector (e.g. phosphor + photodiode): D =kX linear (image may be subsequently processed to “emulate” film of any given speed)

24. 13-15 July 2005L. Ramello24 X-ray film: dynamical range overexposedunderexposed 8 mAs0.5 mAs2 mAs4 mAs16 mAs32 mAs63 mAs M. Overdick (PHILIPS), 11/09/2002, IWORID 2002, Amsterdam

25. 13-15 July 2005L. Ramello25 Flat panel detector: dynamical range typical usage Digital Diagnost (PHILIPS) 43 cm x 43 cm, 143μm x 143 μm M. Overdick (PHILIPS), 11/09/2002, IWORID 2002, Amsterdam

26. 13-15 July 2005L. Ramello26 Image quality Image quality has a decisive impact on the radiologist’s ability to detect pathologies (other factors: visualization conditions, radiologist’s experience) Most important aspects of image quality: – Contrast – Noise (hence signal/noise ratio, SNR) – Spatial resolution (sharpness) Then of course the dose to the patient must be minimized

27. 13-15 July 2005L. Ramello27 Contrast (1) The radiographic contrast C between two areas A (signal) and B (background) of an image may be defined in terms of optical densities: C = D A -D B The radiographic contrast depends from both subject contrast C s and detection method (film- screen, digital detector, etc.) The subject contrast C s depends on the radiation-subject interaction, in the case of x- rays it depends on the linear attenuation coefficient μ and on the thickness x of areas A and B In electronic imaging systems the contrast can be manipulated in a second time

28. 13-15 July 2005L. Ramello28 Contrast (2) Transmission of monochromatic photons of several energies vs. soft tissue thickness: T = exp[-μx] Subject contrast C s : C s = (I 1 -I 2 )/I 1 =ΔI/I 1 with I 1, I 2 representing absorbed energy per unit area of photoreceptor: I 0 = NE I 1,2 = N E ε exp[-∫μdz] (1+R) con N = number of primary photons per unit area, ε = detection efficiency, R = ratio secondary/primary photons I0I0 I0I0 I1I1 I2I2 t x μ1μ1 μ2μ2

29. 13-15 July 2005L. Ramello29 Contrast and Signal Subject contrast C s : Cs = ΔI/I 1 = {1-exp[-(μ 2 -μ 1 )x]}/(1+R) depends on the thickness x of the detail under study (but not on the background tissue tickness t) depends on the difference between linear attenuation coefficients μ 1 and μ 2 decreases as diffused radiation (by Compton effect) impinging on the detector increases: this can be countered by antiscatter grids or exploiting the lesser energy of diffused photons The signal relative to a certain area A may be defined as ΔI·A, and must be compared with fluctuations of the background I 1 ·A (same area)

30. 13-15 July 2005L. Ramello30 Noise and signal-to-noise ratio Fluctuations are due both to quantum noise (fluctuation in the number of converted photons) and to properties of the photoreceptor and of the imaging system Quantum noise in our case follows Poisson statistics: noise = E(I 1 A/E) 1/2 = E[NεAexp(-μ 1 t)(1+R)] 1/2 Taking the ratio of signal: ΔI·A = I 1 CA = CANεEexp(-μ 1 t)(1+R) to noise we get the signal-to-noise ratio: SNR = {1-exp[-(μ 2 -μ 1 )x]}[NεAexp(-μ 1 t)/(1+R)] 1/2 Setting a minimum SNR (Rose criterion: SNR > 5) one can compute the number N of incident photons per unit area necessary to detect a detail of thickness x and transverse area A

31. 13-15 July 2005L. Ramello31 Spatial resolution (1) Every imaging system has intrinsic resolution limits which define the smallest detectable detail For example, in the case of film-screen systems, several factors contribute to the spatial resolution: – finite dimensions of the focal spot and magnification value – possible motion of the patient (breathing, hearth beat) during exposure – resolution loss in the photoreceptor, due e.g. to diffusion of light in screens (or in image intensifiers) Many test objects and procedures have been developed to measure spatial resolution of imaging systems

32. 13-15 July 2005L. Ramello32 Spatial resolution (2) An objective measure of spatial resolution is given by the MTF (Modulation Transfer Function), which quantifies the ratio between output and input contrast vs. spatial frequency The MTF may be measured by taking an image of a lead object having a series of slits with given spatial frequency (lp/mm, line pairs per mm), or an image of a sharp edge

33. 13-15 July 2005L. Ramello33 Spatial resolution (3) Radiographic image of a test object with an array of 3 x 7 groups of slits with different spatial frequencies Optical density profiles of the top-left 3 rows by 4 columns of the test object. The resolution limit (*) corresponds to a spatial frequency of 1.5 cycles/mm

34. 13-15 July 2005L. Ramello34 Detective Quantum Efficiency Detective Quantum Efficiency The Detective Quantum Efficiency (DQE) measures the noise added by the imaging system: DQE(f) = SNR 2 out (f) / SNR 2 in (f) Comparison of DQE among four different imaging systems: Film-screen (speed 400) Computed Radiography Indirect digital radiography (CsI + a-Si) Direct digital radiography (a-Se)

35. 13-15 July 2005 L. Ramello 35 Recent detector developments

36. 13-15 July 2005L. Ramello36 Medipix: Hybrid Pixel Detector M. Campbell, V. Rosso, Rome IEEE NSS-MIC 2004 conference

37. 13-15 July 2005L. Ramello37 Medipix detector - cross section

38. 13-15 July 2005L. Ramello38 ITC-Irst Detector – Si – 300-800  m thick – pixel 170 x 170  m 2 – p+ side 150x150  m 2 – 64 x 64 chs – 1.2 cm 2 area MEDIPIX1 ASIC: SACMOS 1  m technology pixel: 170 x 170  m 2 64 x 64 channels area 1.7 cm 2 threshold adjust 3-bit 15-bit counter VTT Bump-bonding http://medipix.web.cern.ch/MEDIPIX/http://medipix.web.cern.ch/MEDIPIX/ Medipix1 ASIC with silicon pixel detector

39. 13-15 July 2005L. Ramello39 Al thickness 75  m Air X-ray (W-anode) settings : 40 kV, 25 mA, 500 ms Si detector 140 cm Collimator X-ray focus 1.5 cm Al Medipix1 + Si: contrast measurement

40. 13-15 July 2005L. Ramello40  Al (E) and  air (E) are the absorption coefficients at the energy E  (E) is the detector efficiency at the energy E S(E) is the incident spectrum Contrast

41. 13-15 July 2005L. Ramello41 Thickness ratioCalculated ratioExperimental SNR ratio 525/3001.241.25 800/3001.411.42 800/5251.14 75  m Al air Medipix1 + Si: SNR

42. 13-15 July 2005L. Ramello42 Nyquist Freq. (2.94 lp/mm) MTF: 64 % Evaluated aperture 168  m Detector pitch 170  m Thr. (keV) 800  m aperture (  m) 11168 15161 19155 23146 Medipix1 + Si: MTF

43. 13-15 July 2005L. Ramello43 Medipix2: 55 μm x 55 μm pixels

44. 13-15 July 2005L. Ramello44 Calculated x-ray spectrum and energy thresholds used Siefert FK-61-04x12 X-ray tube, W-target, 2.5 mm Al, V peak = 25 kV. Thresholds 9.1 keV 11.3 keV 12.8 keV 18.8 keV

45. 13-15 July 2005L. Ramello45 Medipix2: Measured MTF @ various thresholds

46. 13-15 July 2005L. Ramello46 Medipix2: DQE @ various thresholds

47. 13-15 July 2005L. Ramello47 SYnchrotron Radiation for MEdical Physics The main aim of the SYRMEP beamline is the investigation and the development of innovative techniques for medical imaging. The challenge of mammography – High image quality: Both high contrast and spatial resolution – Very low delivered dose: Breast is very radiosensitive – Very high social relevance After successful feasibility studies on in vitro mammography, the project for synchrotron radiation clinical mammography is under development. R. Longo, C. Venanzi, Rome IEEE NSS-MIC 2004 conference

48. 13-15 July 2005L. Ramello48 SYRMEP Beamline Conceptual Design Sample Holder 5 d.o.f. ionization chamber slit systems monochromator filters Detector Holder 2/3 d.o.f. Fast shutter

49. 13-15 July 2005L. Ramello49 SYRMEP silicon microstrip detector Silicon microstrip detector in edge-on geometry Single photon counting read-out electronics Active area matched with beam cross-section Pixel size 100x300 mm2 Very high scattering rejection Maximum SNR

50. 13-15 July 2005L. Ramello50 SYRMEP digital radiography SR digital image Energy 20 keV 100  m scan step MGD 1.4 mGy Conventional image MGD 1.8 mGy object Si detector Laminar beam

51. 13-15 July 2005L. Ramello51 SYRMEP digital radiography 3 cm thick ‘in vitro’ human breast tissue a) SR digital image Energy 17keV Scan step 100  m MGD 1 mGy b) SR digital image Energy 20keV Scan step 100 mm MGD 0.33mGy clinical mammographic unit 26 kVp MGD 1 mGy

52. 13-15 July 2005L. Ramello52 SYRMEP image acquisition tomography mammography

53. 13-15 July 2005L. Ramello53 SYRMEP mammographic unit Patient support Patient movement stage Detector and Exposimeter holder

54. 13-15 July 2005L. Ramello54 High Z semiconductors Positron Emission Tomography (PET) and Single Photon Emission Computed Tomography (SPECT) make use of high energy photons up to ~500 keV High Z semiconductors are being developed as a replacement for scintillators (BGO, LSO, …) currently used in commercially available systems Due to present limits in the volume (and cost) of semiconductors, the targeted applications are those for small animals with a not too large field of view

55. 13-15 July 2005L. Ramello55 PHILIPS prototype CZT for SPECT Aim: improve energy resolution and spatial resolution Overall size: 20 cm x 48 cm, pixel size 2.4 mm Compare CZT (5 mm thick) and NaI(Tl) (9.6 mm thick) with 3.5 mCi of Tc-99m M. Petrillo, Rome IEEE NSS-MIC 2004 conference

56. 13-15 July 2005L. Ramello56 PHILIPS CZT imaging performance Sensitivity of CZT slightly inferior to NaI(Tl) Contrast and spatial resolution of CZT clearly superior to NaI(Tl)

57. 13-15 July 2005L. Ramello57 CdTe, CZT developments Quite some progress in recent years for CdTe, CdZnTe detectors concerning: – Crystal growth – Electrode design – Interconnect technology (bump bonding, …) – Hybrid and ASIC electronics L. Verger, Rome IEEE NSS-MIC 2004 conference

58. 13-15 July 2005L. Ramello58 Single channel CZT 4x4x6 mm 2 After bi-parametric correction (based on pulse rise time – amplitude correlation) the efficiency at 122 keV rises from 30% to 75-80%

59. 13-15 July 2005L. Ramello59 With bi-parametric correction the energy resolution at 662 keV is improved and tailing is drastically reduced Single channel CZT 8x8x15 mm 2

60. 13-15 July 2005L. Ramello60 CZT-based micro-PET Micro-PET system with pixellated CZT replacing LSO scintillator: – improve spatial resolution (2 mm  1 mm) with depth-of-interaction information