13-15 July 2005 L. Ramello 1

Medical Imaging with

Semiconductor Detectors

 

CINVESTAV 2005 Advanced CINVESTAV 2005 Advanced Summer SchoolSummer School

L. Ramello – Dip. Scienze e Tecnologie Avanzate,Univ. Piemonte Orientale, ALESSANDRIA (Italy)

13-15 July 2005

L. Ramello 2

TopicsTopics

Basic properties of semiconductor detectorsImage quality: contrast, SNR, MTF, DQERecent detector developments:

– MEDIPIX (2D pixels)– SYRMEP (Synchrotron Light Source)– High Z semiconductors

Dual Energy MammographyDual Energy Angiography

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Basic properties of Basic properties of semiconductor detectorssemiconductor detectors

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• 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 ?Why semiconductor detectors ?

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Semiconductor materialsSemiconductor 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 Eg) energy resolution

Material Z density, g/cm3

Eg,

eV

W,

eV

Si 14 2.33 1.12 3.6

GaAs 31,33 5.31 1.42 4.2

Ge 32 5.32 0.73 2.9

Se 34 4.3 1.71-1.75 5.6

CdTe 48,52 6.20 1.52 5.0

HgI2 80,53 6.36 2.13 6.7

PbI2 82,53 6.2 2.31 7.2

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Energy resolutionEnergy resolution

Radiation ionization energy (W):determines the number of primaryionization events

Band gap energy (Eg):lower value easier thermal generation of e-h pairs(kT = 26 meV for T = 300 K)

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Semiconductor detectorsSemiconductor detectorsTo 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. N2

temperature) is neededMultichannel systems require special care for

power density, connection technique, cross-talk

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The p-n junction (1)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

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The p-n junction (2)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 potentialenergy and thermal energy

d = (2VB/eND)1/2

= r0 12 0 (Si)ND = donors/cm3 (n-Si)

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The microstrip detectorThe microstrip detector

SIGNAL = number of electron-hole pairs:

ne-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 (VB) => controls diode depletion thickness, i.e. active volume • p-n junction capacitance per unit area C:1/C2 grows linearly with VB => C-V measurement determines full depletion voltage VFD

2/1

2

B

D

V

eNC

d

C

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A microstrip detectorA 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

guardring bias line first strip (AC contact)

DC contact (to p+ implant)

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A readout chainA 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 shaper discriminator

IN OUT

vthn

vthp

calib in

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The RX64 ASICThe RX64 ASIC

RX64 - Krakow UMM design - (28006500 m2) 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).

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Conversion efficiency (1)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

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Conversion efficiency (2)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

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Image quality: Image quality: contrast, SNR, MTF, DQEcontrast, SNR, MTF, DQE

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X-ray beams (1)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 addedfiltration with a light material (Al) and with a rare earth material (La, K-edge @ 39 keV)

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X-ray beams (2)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

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X-ray imaging techniques 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)

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Film + screen (1)Film + screen (1)

X-rays transmitted through patient

first screen

second screendouble 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 (CaWO4, Gd2O2S: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

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Film + screen (2) 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

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Exposure and optical density (1) 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 = I1/I0

Useful optical density goes from 0.2 to 2.5-3.0

Exposure X quantifies the number of incoming x-rays

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Exposure and optical density (2) Exposure and optical density (2)

Relation between optical density D and exposure X:

Transm. T D = -log(T)

1.000 0.0 100% transm.

0.741 0.13 base

0.100 1.0 good exposure

0.010 2.0 lung

0.001 3.0 very dark

0.0003 3.5 maximum 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)

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X-ray film: dynamical rangeX-ray film: dynamical range

overexposedunderexposed

8

mAs

0.5

mAs

2

mAs

4

mAs

16

mAs

32 mAs 63

mAs

M. Overdick (PHILIPS), 11/09/2002, IWORID 2002, Amsterdam

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Flat panel detector: dynamical rangeFlat 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

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Image qualityImage 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

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Contrast (1)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 = DA-DB The radiographic contrast depends from both

subject contrast Cs and detection method (film-screen, digital detector, etc.)

The subject contrast Cs 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

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Contrast (2)Contrast (2)Transmission of monochromaticphotons of several energies vs. soft tissue thickness: T = exp[-μx]

Subject contrast Cs:

Cs = (I1-I2)/I1 =ΔI/I1

with I1, I2 representing absorbed energy per unit area of photoreceptor:

I0 = NE

I1,2 = N E ε exp[-∫μdz] (1+R)

con N = number of primary photons per unit area, ε = detection efficiency,

R = ratio secondary/primary photons

I0 I0

I1 I2

t xμ1

μ2

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Contrast and SignalContrast and SignalSubject contrast Cs:

Cs = ΔI/I1= {1-exp[-(μ2-μ1)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 μ1 and μ2 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 I1·A (same area)

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Noise and signal-to-noise ratioNoise 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(I1A/E)1/2 = E[NεAexp(-μ1t)(1+R)]1/2

Taking the ratio of signal:

ΔI·A = I1CA = CANεEexp(-μ1t)(1+R) to noise we get the signal-to-noise ratio:

SNR = {1-exp[-(μ2-μ1)x]}[NεAexp(-μ1t)/(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

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Spatial resolution (1)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

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Spatial resolution (2)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

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Spatial resolution (3)Spatial resolution (3)Radiographic image of a test objectwith an array of 3 x 7 groups of slitswith different spatial frequencies

Optical density profiles of the top-left 3 rowsby 4 columns of the test object.The resolution limit (*) corresponds to a spatial frequency of 1.5 cycles/mm

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Detective Quantum EfficiencyDetective Quantum Efficiency The Detective Quantum Efficiency (DQE) measures the noise

added by the imaging system: DQE(f) = SNR2out(f) / SNR2

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)

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Recent detector developmentsRecent detector developments

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Medipix: Hybrid Pixel Detector Medipix: Hybrid Pixel Detector

M. Campbell, V. Rosso, Rome IEEE NSS-MIC 2004 conference

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Medipix detector - cross sectionMedipix detector - cross section

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ITC-Irst Detector–Si <111>–300-800 m thick–pixel 170 x 170 m2

–p+ side 150x150 m2

–64 x 64 chs–1.2 cm2 area

MEDIPIX1 ASIC:SACMOS 1 m technology

pixel: 170 x 170 m2

64 x 64 channelsarea 1.7 cm2

threshold adjust 3-bit15-bit counter

VTT Bump-bonding

http://medipix.w

eb.cern.ch/ME

DIP

IX/

http://medipix.w

eb.cern.ch/ME

DIP

IX/

http://medipix.w

eb.cern.ch/ME

DIP

IX/

http://medipix.w

eb.cern.ch/ME

DIP

IX/

Medipix1 ASIC with silicon pixel Medipix1 ASIC with silicon pixel detectordetector

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Al thickness 75 m

Air

X-ray (W-anode) settings : 40 kV, 25 mA, 500 ms

air

Alair

N

NNC

Si detector

140

cm

Collimator

X-ray focus

1.5

cm

Al

8 10 12 14 16 18 20 22 242.4

2.6

2.8

3.0

3.2

3.4

3.6

3.8 800m detector

Co

ntra

st (

%)

Energy threshold (keV)

Medipix1 + Si: contrast measurementMedipix1 + Si: contrast measurement

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8 1 0 1 2 1 4 1 6 1 8 2 0 2 2 2 4

2 , 8

3 , 0

3 , 2

3 , 4

3 , 6

3 , 8

4 , 0

4 , 2

C (

%)

E n e r g y t h r e s h o l d ( k e V )

xdEESE

dEESEEE airAl

eC

)()(

)()())()((

1

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

ContrastContrast

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22Alair

Alair NNSNR

Thickness ratio Calculated ratio Experimental SNR ratio

525/300 1.24 1.25

800/300 1.41 1.42

800/525 1.14 1.14

75 m Al

air

Medipix1 + Si: SNRMedipix1 + Si: SNR

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Nyquist Freq. (2.94 lp/mm)MTF: 64 %

Evaluated aperture 168 mDetector pitch 170 m

Thr. (keV)

800 m aperture

(m)

11 168

15 161

19 155

23 146

Medipix1 + Si: MTFMedipix1 + Si: MTF

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Medipix2: 55 Medipix2: 55 μμm x 55 m x 55 μμm pixelsm pixels

55 m

55 m

55 m

55 m

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Calculated x-ray spectrum and Calculated x-ray spectrum and energy thresholds usedenergy thresholds used

Siefert FK-61-04x12 X-ray tube, W-target, 2.5 mm Al, Vpeak = 25 kV.

Thresholds9.1 keV11.3 keV12.8 keV18.8 keV

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Medipix2: Measured MTF Medipix2: Measured MTF @ various thresholds@ various thresholds

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Medipix2: DQE @ various Medipix2: DQE @ various thresholdsthresholds

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SYnchrotron Radiation for SYnchrotron Radiation for MEdical PhysicsMEdical 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

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SYRMEP Beamline SYRMEP Beamline Conceptual DesignConceptual Design

SampleHolder5 d.o.f.

ionization chamber

slit systems

monochromator

filters

DetectorHolder

2/3 d.o.f.

Fast shutter

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SYRMEP silicon microstrip SYRMEP silicon microstrip detectordetector

dead volume

pixels

X-rays

read-out electronics

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

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SYRMEP digital radiographySYRMEP digital radiography

SR digital image

Energy 20 keV

100 m scan step

MGD 1.4 mGy

Conventional image

MGD 1.8 mGy

objectSi detector

Laminar beam

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SYRMEP digital radiographySYRMEP digital radiography

3 cm thick ‘in vitro’ human breast tissue

a) SR digital imageEnergy 17keVScan step 100 mMGD 1 mGy

b) SR digital imageEnergy 20keVScan step 100 mmMGD 0.33mGy

clinicalmammographicunit 26 kVpMGD 1 mGy

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SYRMEP image acquisitionSYRMEP image acquisition

tomography

mammography

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SYRMEP mammographic unitSYRMEP mammographic unitPatient support

Patient movementstage

Detector andExposimeterholder

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High Z semiconductorsHigh 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

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PHILIPS prototype CZT for SPECTPHILIPS 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

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PHILIPS CZT imaging PHILIPS CZT imaging performanceperformance

Sensitivity of CZT slightly inferior to NaI(Tl) Contrast and spatial resolution of CZT clearly

superior to NaI(Tl)

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CdTe, CZT developmentsCdTe, 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

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Single channel CZT 4x4x6 mmSingle channel CZT 4x4x6 mm22

After bi-parametric correction (based on pulse rise time – amplitude correlation) the efficiency at 122 keV rises from 30% to 75-80%

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With bi-parametric correction the energy resolution at 662 keV is improved and tailing is drastically reduced

Single channel CZT 8x8x15 mmSingle channel CZT 8x8x15 mm22

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CZT-based micro-PET 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