The Role of Silicon Radiation Sensors and Integrated Front ...€¦ · Some of the Shortcomings of Present Day Clinical PET Scanners A reference for the new generation of PET scanners

3. 05. 2007 Seminar SLAC 1

The Role of Silicon Radiation Sensors and Integrated Front-End Electronics In Medical

Imaging Instrumentation.

SLAC Seminar

2. May 2007

P. Weilhammer

INFN Perugia/CERN


1. Overview of present Medical Imaging Modalities

2. Photon Detection with Silicon Radiation Detectors and Implications for Applications of Silicon in Medicine

3. Examples of Silicon Detectors in Medical Imaging Applications

OUTLINE of Talk


The Famous Picture


Medical Imaging in the 21st Century: “Multi Modality”


While photographic emulsion was for a very long time without competition, important innovations in medical imaging were introduced over the last 50 to 60 years:

•Anger Camera for SPECT

•First attempts on PET with Proportional Wire Chambers

•PET Scanners with High Z Scintillators and PM Readout using Anger Logic

•MRI

•Attempts on Electronic Collimation in SPECT using Compton Scattering of the Gamma Ray with Germanium Detectors

•XR-CT with Silicon Photo Diode arrays in Current Mode

•Ultrasound Scanners

•Etc………..


With the exception of X-Ray CT where Silicon Photo Diodes Play a dominant role, Silicon Radiation Sensors are not very strongly represented in this field so far.

The dominant detector technologies are

•Scintillators

•PM Tubes

•Maybe soon High Z semiconductors in Digital X-Ray CT

One exception:

Low Dose digital Mammography Scanner from SECTRA (single sided strip detectors “edge on” and VLSI Front-end) See www.sectra.com

Very promising for Screening in Mammography


Medical Imaging ModalitiesA short incomplete list of imaging modalities which might be

improved by the implementation of silicon detectors:

1. Field of X-ray imaging

• X-ray radiography; projection images to obtain 2 D anatomical information. Classical X-ray images

• X-ray Computed Tomography; 3-D anatomical information through reconstruction of the distribution of attenuation coefficient μ(x,y,z); Tissue specific contrast is obtained by measuring at the detector

I = I0 exp(-μ(x,y,z) d)

• Most often used in patient diagnostics in all hospitals:Translation-Rotation Scanner. Measure many slices in one or several rotations. Modern scanners have up to 64 slices. Radiation detectors are pixelated matrices of suitable (high Z) scintillators like CsI, BGO, LSO, more recently ceramic based scintillators.


• Preclinical X-ray CT: small animal scanners. Similar to Clinical scanners with emphasis on higher spatial resolution for lower density tissue. High intensity X-ray sources are used with micro-focus or very high resolution synchrotron source.

• X-ray energy ranges: ~ 10 keV to 120 keV

2. Nuclear Medicine Imaging

•Gamma Camera

•Single Photon Emission Computed Tomography (SPECT)

•Positron Emission Computed Tomography (PET)

“High” energy γ rays penetrate tissue with little absorption.

Imaging is performed by injection and take-up in the patient of a radio-ligand containing a meta-stable reporter radionuclide which emits γ-rays or positrons and often also electrons which are absorbed in the surrounding tissue. He gamma rays will exit fromthe body with occasional Compton scatter.


The goal is to reconstruct the distribution of radioactivity within the body, either 2-D or 3-D image, using back projection algorithms.

Traditional detectors are highly segmented scintillation crystal arrays (NaI, CsI, BGO, LSO, La-bromides,…).

The scintillation arrays are readout by Photo Multipliers.

Anger Camera: The direction of the photon absorbed in the detector is determined by the x,y-coordinate of the impact and by tight collimation in front of the scintillator (Pb-collimator with many small holes).

Pin-Hole camera: the collimator is replaced by a arrangement of one or several specially shaped pinholes in a collimator structure.

SPECT Camera

PET Scanner; both SPECT and PET cameras allow direct recording of 2-D projections simultaneously (or by rotating one camera head around the patient in case of SPECT) leading to full 3-D image reconstruction (not slice by slice). Closed ring detector (scintillator) detector geometry to measure 180 degree 511 keV photons from positronium annihilation.

Data Recorded are sinograms


Preclinical PET : high resolution small animal PET.

Compton camera and Compton probes

Autoradiography

Bio-molecular Imaging is emerging

………..

In all imaging detectors and systems the important “quality “ factors are:

Detective Quantum Efficiency (DQE) (Sensitivity)

Spatial resolution

Speed, coincidence window


Detection of Photons and Energetic Electrons in Semiconductor Detectors

Medical imaging requires good ability of detection of photons, in reality detection of energetic electrons created inside the material ( an advantage!), over a wide range of energies.

Energy Ranges:

•Computed Tomography (CT) X-rays: 20 to ~>120 keV

•Single Photon Emission Tomography (SPECT): detect γ-rays for a big variety of isotopes used in different tracer molecules

•99mTc 140 keV

•111In 185 and 245 keV

•31I 360 keV

Positron Emission Tomography: 511 keV γ from e+e- annihilation

Autoradiography: β particles emitted from e.g. Tritium, 14C, 33Ph,…( from 10 keVto several 100 keV)


Photon Interactions in Silicon

Only two out of all photon interactions are important for medical imaging:

The “wanted” one: Photoelectric Absorption (total absorption of γ or X-ray)

σ = 4√2 α4 Εγ−7/2 Ζ5σTh

with the Thomson cross-section σTh.= 8π/3 r20 = 6.652 bars per electron.


The “unwanted” one: Compton scattering

)cos1(12

θγ γ

γ

−+

=′

cm

E

EE

e

⎟⎟⎟⎟⎟

⎠

⎞

⎜⎜⎜⎜⎜

⎝

⎛

−+−=−

)cos1(1

11

2 θγγ

cmEEE

e

e

θ

The recoil electron ( from K-shell or L-shell or valence band) creates (eh) pairs in the semiconductor bulk through ionization

Kinetic energy of recoil electron


Attenuation of incoming photons in material

“Good” photon detector are detectors which absorb most of the incoming photons preferably by photoelectric absorption. Quantitatively the attenuation in the material of a sensor is characterized by the mass attenuation coefficient μ(E) :

eetEtE

NoN )(]/)([)( ••−=•= −

ρρμμ

With t the thickness of the sensor in direction of the photon beam and ρ the density of the material.

Materials with low Z (silicon has Z=14)become quickly impractical with increasing photon energy!


In 1mm thick silicon for 20 keV photons

Photoelectric interaction: ~ 97%

Compton interactions: ~3%

Interactions/m for Si versus photon energy

Interesting region for medical imaging


Range of Electrons in MaterialsThe range of electrons in materials expressed as range * density is very similar for many different materials

Typical Range:

50 keV electron in silicon: ~20 μm

200 keV : ~200 μm

500 keV :~ 600 μm

For Compton interaction the “point-like” domain is between 10 keV and 250 to 300 keV!

10-2

Range*density [g/cm2]

100 keV

Si

NaI


Some inherent physical limitations in different imaging modalities are:

Spatial extension of the photon interaction in the detector material due to the nature of photon interactions (in most materials interaction cascades are frequent before final absorption). The typical extension of a photon interaction in many detector materials ( at 500 keV) can be considered to be confined in a sphere of ~1 cm in diameter.

Uncertainty in Depth of Interaction Parallax error

Finite path length of positrons and recoil electrons

Compton scattering in tissue.

In PET: Finite momentum of e+e- compound at the moment of decay Acolinearity

Accidental coincidences.

………

~1cm through multiple

interactions in scintillator

Incoming γ


Types of silicon radiation sensors which could be interesting in medial applications:

Si strip detectors, single sided and double sided.

Si pixel detectors

Si pad detectors and micro pad detectors

CMOS imagers

Flat panel devices based on amorphous silicon

Sub-micron fast CMOS front-end chips for readout of strips, pixels and pads and others

•Silicon Photo Multipliers(SiPM)

•Variety of Front-end deep submicron circuits developed for HEP


Advantages of Silicon Detectors over “classical”Instrumentation in Medical Imaging?

For Si pixel, pad and strip detectors

•Very high segmentation feasible

•Matching of segmentation of front-end readout electronics

•Excellent energy resolution

•Excellent position resolution

•Possibility of using “counting mode” with energy weighting

•Low voltage operation

•…..


In the following I want to discuss possible applications of Silicon Detectors and Front-End Electronics ASICS which are projects within the CIMA Collaboration.

Emphasis will be on

•New Developments for PET

•Compton Camera SPECT and Compton PET

•Micro X-Ray CT


R&D Projects using Silicon Detectors in Medical Imaging within the CIMA

Collaboration

•Novel axial brain PET Scanner using Hybrid Photon Detectors (HPD)

•Readout of z-coordinate of Axial PET with Wave Length Shifters and Silicon Photo-Multipliers

•Compton Imaging and Probes

•High resolution small animal PET scanner based on Compton interactions

•A High Resolution Micro-CT Prototype Module for Small Animal Imaging


The CIMA Collaboration:

InstitutesLisbon INFN Bari

INFN Rome INFN Perugia

HUG Geneva Phys. Dept. Uni Geneva

University of Michigan University of Ljubljana

Ohio State University LANL

University of Valencia Karlsruhe

Kharkhov Space Institute CERN

Cracow

Industrial Partners:IRST Trento

Gamma-Medica_IDEAS

SINTEF


New Development for PET


Some of the Shortcomings of Present Day Clinical PET Scanners

A reference for the new generation of PET scanners could be the High Resolution Research Tomograph (HRRT) developed by CPS innovations*

Efficiency for the detection of photon pairs is given to be 6.9%, which includes a sizeable fraction of unidentified Compton interactions. Energy resolution is 17% at 511 keV, timing resolution is 2 to 4 ns. The volumetric voxel resolution is given as 20 mm3, corresponding to a trans-axial resolution of in average 2.6 mm and an axial resolution of about 3mm. All quantities referenced here are FWHM.

* K. Wienhard et al, IEEE Trans. Nucl. Sci. 49 (2002) 104-110


The main shortcomings are:

• Relatively low efficiency of photon conversion due to the anti-correlation between accurate knowledge of depth of interaction and thickness of scintillation crystals.

• Parallax error due to limited knowledge of the depth of interaction in the radial direction of the 511 keV photon. Several techniques have been developed to reduce the parallax error, e.g. the Phoswich arrangement of scintillation crystals [HRRT].

• Image smearing due to the physics of the photon interaction. The spatial extension of the 511 keV photon interaction cascade even in high Z scintillation material gives an important contribution to deteriorate the image quality. This is due to the fact that even for scintillation material with the highest density and effective Z the fraction of Comptoninteractions is 60% or more.

• Limited capability to identify and reject events with a Compton interaction in the scintillation material.


Some ideas to improve over present day scanners.


A proposal for a parallax-free Compton enhancedPET camera module for high resolution, high

sensitivity functional brain imaging based on a Hybrid Photon Detector (PET-HPD)

Novel Axial Geometry PET Scanner


hν

e-

Bi-alkali photocathode

16 front-endchips

Ceramic PCB

Si Sensor2048 pads(1 x 1 mm2)

200 300 400 500 600

lambda (nm)

0

4

8

12

16

20

24

28

32

Q.E

. (%

)

HPD PC87(produced Easter Sunday 2001)

Perfect single photon detection

1 ph.e.

2 ph e

3 ph e

12 σ

The HPD


This concept is discussed in detail in

J. Seguinot et al., Il Nuovo Cimento C, Vol. 29 Issue 04 pp 429-463


HPD2

HPD1

x

yz

HPD1Principleof a cameramodule

The Concept


γ reconstruction point

“unambiguous”

Select only events in which Compton scattering happens in forward hemisphere Restrict to Compton angle 10° ≤ θ ≤ 60°Ask for energy deposit in first interaction E ≤ 170 keVThrow out of data sample Compton events which cannot be resoved w.r.t. first vertex

Discriminate Compton interactions: Fine 3D segmentation makes itpossible…

γ

“ambiguous”

which was the point of 1st γ interaction ?

γ


x and y resolution :axially arranged, long LYSO scintillation crystal bars allow to choose x and y resolution according to the chosen lateral dimension (s) of the bars. z-resolution : optimize light yield N1 and N2, read on both sides by HPDsand light absorption along the bars. A well optimized bulk absorption in the scintillation bars and highest possible light yield are the most important parameters.


The z coordinate and the spatial resolution in z is

Monte Carlo simulations showed that a good compromise for LYSO crystals is a crystal length of 15 cm with λ around 100 mm. These simulations indicate that σz = 3.5 – 4.5 mm could be obtained

The axial resolution is not as good as one would desire for an ideal instrument

11. May 2005 EUROMEDIM2006 Marseille 34

The Hybrid Photon Detector: PET-HPD

Some relevant properties of HPD for PET application:

Very good spatial resolution can be chosen; size, geometry and granularity of silicon pad sensor can be chosen according to the requirements.

Very good energy resolution

charge gain in a single stage dissipation

Very good linearity over large dynamic range

Not temperature sensitive


A Prototype PET HPD has been successfully built and tested


Basic Elements: Double metal pad detectors and

VATAGP5 Chip


Some sensor properties:

• Full depletion voltage: V ~ 30 V

•Leakage current per pad: ~500 pA

•Pad and routing line cap.: ~ 5 pF

Self-triggering Front-End Electronics: the VATAGP5 chip

Fast Charge Sensitive Preamplifier

Output of preamp fanned out to

•Slow shaper amplifier (t=220nsec) followed by a S/H to record precisely pulse height (energy)

•Fast shaper (t=40 ns) followed by a discriminator with time walkcompensation and a monostable; firing of discriminator initiates a S/H

Repeat pattern 128 times; all 128 channels have common threshold


VATAGP5 continued

3-bit trim DAC for trimming thresholds for each channel individually

The mono-stable pulse initiates readout clock and S/H

Four readout modes

•“Serial” reads sequentially S/H of all channels

• “Sparse” puts address of hit channel(s) into a register and only those channels will be read

• “Sparse with neighbors” reads hit channel and n neighbors.

• “Sparse” with any pre-defined neighbors

• Dynamic range: up to 1.2 pC for positive polarity.


Results from tests with first Proto-type PET-HPD


First Brain PET HPD works: Hit Distribution from Light Spot

This lego plot shows that a threshold of ~ 20 fC eliminates easily any dark current hits

Background free images


Mean Charge μ and σ/μ of charge distributions as Function of Cathode Voltage

5 7 9 11 13 15 17 19 210

100

200

300

400

500

VPC (kV)

μ (A

DC

cou

nts)

chan. #45

4.0

4.5

5.0

5.5

6.0

6.5

7.0

σ/μ

(%)

μchan. #53σ/μ

Mean charge μ (left axis) and ratio of Gaussian width to mean charge σ/μ (right axis) versus cathode voltage UC (kV).

Note:

that mean charge is linear but intercepts at ~6 keV due to energy loss in dead layer. This can be improved in next sensor production run. Expect intercept at 0.5 keV, which will considerably improve the charge gain in the HPD

σ/μ reaches almost a plateau around 17 keV since energy straggling becomes small.

One can estimate an energy loss of ~ 1.6 keV at 20 kV with nearly negligible straggling.

Gain at 20 kV is 5090

From σ/μ = (ENF/N)1/2

N = 507 photo electrons( ~ N0 of LYSO)

The absolute gain of the chain can now be calculated:

0.94 fC/(ADC count) for chip1


Timing is another crucial problem:Time walk as a function of distance from threshold for Cr-RC shaping:

FWHM time resolution versus “Threshold Distance”

Number of times threshold for coincidence window

One can obtain ~ 5 nsec FWHM

with VATAGP-5 ASIC


Status

A first PET HPD tube for readout of a scintillation crystal matrix developed for use in a novel axial PET concept has been designed, fabricated and tested.

All the relevant features of the mechanical and optical properties of the envelope, the front-end electronics chip and the silicon pad sensors, required for this application have been successfully demonstrated.

The system has an appropriate dynamic range which will allow detecting energy deposits from 30 keV to well above 511 keV in a LYSO crystal with very good linearity.

The required time resolution (~ 5ns FWHM) needed for PET can be achieved with the VATAGP-5 ASIC.

The fabrication of a second PET-HPD tube is under way. The next major step in the project is to assemble a complete camera module to characterize its spatial (axial coordinate) and energy resolution


A New Concept To Obtain Optimal Axial Spatial Resolution


The Principle:LYSO Crystal Bar

Thin WLS Strip

Question: is there enough light from WLS strips for 511 keV photon in LYSO or even for 50 to 100 keV Compton recoil electrons?


For a Brain Pet: Crystal size: 3 mm x 3 mm x 150 mm

WLS size: 1 mm x 3 mm x 35 mm

Measurement of z-coordinate:

either take WLS strip with highest hit σz = 3mm/Sqrt(12) = 0.9 mm

Or measure analog values on more than one strip: center of gravity; should in general be better than digital resolution


Experimental Verification: Two different and independent methods to establish validity of new concept

Adjustable pulsed low energy electron beam

22Na source in coincidence


Measure

• Photoelectric yield of WLS strips

• Achievable spatial resolution along the crystal axis

• Timing resolution


Measured charge distribution in each of the 2 WLS strips with the electron beam moved across the WLS strips (~350 keV)


With the existing set-up one can generate a charge in the LYSO equivalent to a photon energy of up to 400 keV.

Extrapolate curve below to 511 keV measured photo electric yield in WLS is 42 p.e.

The WLS were read with Hamamatsu PM’s with 15% Photo efficiency.

In a real device one will use SiPM for WLS readout; Hamamtsu quotes 40% yield for their MPPC; this will give ~ 100 p.e. in two WLS strips

Signals from both strips summed


Experimental results for z spatial resolution

In this setup always 2 WLS strips hit: calculate z with formula:

Beam moved across the 2 WLS strips

Correlation between beam spot and measured z-coordinate


Z-resolution as a function of electron beam energy

The real value of sigma(z) after deconvolution of beam spot size is ~800micron

Z-resolution is dominated by the p.e. statistics in the WLS strips. Expect a 1/SQRT(Econv) dependence.


GM-APD Arrays (just a few words before discussing results with SiPM Set-Up)

Commercial SiPM from Hamamatsu

S. Uozumi, Talk at VCI, 2007


Main blockWafer

A full wafer with Si-PM structures ; produced by IRST, Trento ,Italy


-5.00E-009 -4.00E-009 -3.00E-009 -2.00E-009 -1.00E-009 0.00E+000 1.00E-0090

1000

2000

3000

4000

5000

6000

7000

8000Plot 032

No.

of c

ount

s

Amplitude

Pedestal

1 photon

IRST 1mm x 1mm SiPM read with P/N MSA 0886- BLK HP fast Ampl., Single Photon Response

Vbkd = 35 V, Vbias = 38 V Gain = 1.25 x 106

Our Measurements


-3.00E-009 -2.00E-009 -1.00E-009 0.00E+000 1.00E-0090

50

100

150

200

250

300

350Plot 031

No.

of c

ount

s

Amplitude

Vbkd = 69.7 V, Vbias = 71.2 V, Gain 6.7 x 105

Single Photon Response of Hamamatsu 3mm x 3mm SiPMRad with Fast HP Ampl.


Results from the second method using 22Na source and G-APD (Hamamtsu MPPC) readout on LYSO and on WLS strip

Pulse Height Spectrum with G-APD and LYSO: ΔE/E ~12%


Measured Pulse Height Spectrum fro WLS with G-APD (For photo-electric events)

~ 35 -40 p.e.


Timing of LYSO w.r.t. WLS Strip with G-APD:

FWHM ~700ps


The concept of axial (z) coordinate measurements using a WLS strip matrix looks very promising for PET imaging:

Next step is the construction of two prototype crystal stacks with WLS matrix readout. Test both HPD and SiPM readout of LYSO


This novel concept could solve most of the problems inherent in present day PET systems.

Summary of (expected) Performance:

•Full 3D reconstruction of the 511 keV photons

•No parallax error

•Spatial resolution (x, y, and z) can be chosen according to requirements by selecting Crystal and WLS dimensions .

•The total thickness of the scintillation detector stack can be chosen independently of other device parameters, which allows in principle to choose the efficiency according to the requirements of specific applications.

•Total uniformity of spatial resolution over the complete field of view.

•Capability to distinguish photon interactions with Compton cascades from photo-absorption events with nearly 100% efficiency.

.


•Increase of sensitivity by including in the final event sample events with a primary Compton interaction exploiting the constraints given by energy deposited in the scintillation crystals and the position measurement of both observed interactions. About 25% of the Compton interactions can be kinematically fully resolved. This will increase for LYSO the number of coincidences to be used for chord reconstruction by a factor 1.6 to 1.8, depending on the recoil electron energy cut-off.

•Axial arrangement of the scintillation crystals can reduce the number of electronic readout channels, while maintaining high granularity.

•Very good energy resolution in the order of 8% at 511 keV if the LYSO crystal matrix is readout with a HPD and ~12% with SiPM readout.

•Competitive timing resolution of ~700ps for SiPM readout of LYSO, maybe also useful for TOF PET.

•The spatial resolution which can be obtained with the new concept (for scintillation crystal dimensions proposed in[5]) will result in a voxel precision of 9 mm3 FWHM, close to the limitations imposed by the inherent physical limits from a-co-linearity and range of positron.

•Silicon Photo Multipliers (SiPMs) are an option to readout the LYSO crystals in a strong magnetic field, hence opening the possibility of co-registration with MRI.


Compton Imaging


The main features of Compton Imaging are:The Mechanical collimator in the Anger Camera is replaced by “Electronic

Collimation”. This removes the coupling between sensitivity and spatial resolution.

This is achieved by having two detectors in coincidence:In the first detector the gamma rays are scattered by Compton Scattering

on electrons in the detector materialIn the second detector the scattered gamma ray is absorbed


The measured quantities in Compton imaging are:x, y, z-co-ordinates in the first detector

x, y, z-co-ordinates in the second detectorEnergy of recoil electron in first detector

Energy of scattered photon in second detector

Not measurable with Compton Cameras for medical applications: Direction of recoil electron, which leads to the conical ambiguity. This

leads to more complicated image reconstruction algorithms.

Expected improvements over Anger Camera:

•Factor ~5 in spatial resolution for probes

•Factor 5 to 50 improvement in sensitivity

Due to Doppler effect smearing of the recoil energy resolution:

Silicon is the only realistic semiconductor detector material for first detector


Results from a Demonstrator Test for a Compton Prostate Probe in 2005


Silicon detector and stack of 5 detectors


A Demonstrator set-up with stack of 5 Silicon pad sensor and 3 camera heads


Main ResultsSpatial resolution was measured for 4 energies; 57Co (122 keV) and 133Ba (272,302 and 356 keV). For the highest energy with a source-first detector distance of 11.3 cm: 5mm FWHM

With a source Si distance of 3 cm this gives (simulation) 2 -3 mm FWHM


Status:

Spatial resolution in Silicon Demonstrated

Next Demonstrator test foreseen before summer of 2007 with much improved camera head and improved silicon ( lower thresholds possible)


A High Resolution Small Animal PET Scanner based on Compton Scatter Events in Silicon Pad

Detectors


BGOdetector

Sidetector

Si-Si

BGO-BGO

Si-BGO

Si-Si :Very High Resolution

Si-BGO : High Resolution

BGO-BGO : Conventional PET Resolution

Three Major Coincidence Events

A Very High Resolution PET Scanner for small animals based on Compton Scattering

events is proposed:

The Concept


Detection Efficiency (%)

Radial Posn. (mm)

Single – Single Single – BGO BGO - BGO

0 1.05 8.83 20.84

6 0.96 8.96 20.69

12 1.04 8.94 19.70

18 1.19 9.06 18.17

Calculated for point source in center plane. Only single scattering or absorption interactions in the silicon detector are included.Back scattered photons from BGO and events without full energy deposition are excluded.

BGO ring

Simulation results with this configuration

Efficiency for different event classes

Harris Kagan Imaging 2006, June 26-30, Stockholm

Compton PET Test Bench

Silicon detector BGO detector

4.5 cm × 2.2 cm and 1 mm thick 32×16 (512) pads, 1.4 mm × 1.4 mm pixel sizeEnergy Resolution 1.39 keV FWHM for Tc 99m

5.3 cm × 5 cm and 3 cm thick 8×4 array, 12.5 mm × 5.25 mm crystal sizeEnergy Resolution 22% FWHM for Na-22

HAMAMATSU PMT R2497

VATAGP3


Prototype PET InstrumentSingle-slice instrument using silicon and BGO

Disassembled Assembled

Silicon detector Silicon detector



Resolution Uniformity

0 1 2 3 4 5 cm

5

4

3

2

1

0

Source pairs at 5, 10, 15, & 20mm off-axis

Sinogram

The sources in each pair are clearly separated at appropriate sinogram angles


Compton PET: Intrinsic Resolution

Needle 25G (ID = 0.254 mm, OD = 0.5mm, SS_steel wall = 0.127 mm)

5

4

3

2

1

0

0 1 2 3 4 5 cm

0.254 mm

0.127 mm

Image Resolution= 700 μm FWHM

SS_steel wall

F-18

5

4

3

2

1

00 1 2 3 4 5 cm


A Prototype Module for Small Animal X-Ray CT

•Fast Counting Chip with Energy Window

•“Edgeless” Micro Pad 1mm thick Silicon Detectors; In prototype module pitch 130 micron. Probably needs to be reduced.


The complete module with four ASICs and two detectors, 512 pixels. Fixed on the alu base plate with 3 screws.Note that the 2 detectors are slightly wider than the PCB and the alu support allowing in principle to arrange several modules side by side ( two sided “Butting”) with minimum distance between detectors

A Prototype Module


Photograph of middle of module



ConclusionThe silicon sensor with a dominant role in medical imaging is still the photo diode array for readout of plastic scintillators in X-ray CT. Replacement of this technology might come in the form of Cd(Zn)Te readout with very high speed counting ASICs

There are many attempts and projects to apply HEP developed technology ,based on silicon detectors, in medical imaging and develop instruments for imaging and actually get those into hospitals.

So far only one new device (to my knowledge) , a digital mammography low dose (exposure) X-ray CT scanner (SECTRA [www.sectra.com]) based on silicon microstrip detectors, is on its way to become a standard in hospitals.

Other promising applications are studied with intensive R&D efforts.

Most important impact of silicon radiation sensors and submicron front-end electronics will be in PET and SPECT with impressive performance improvements of SiPM processing (my prediction).There might be a hard time coming up for Photo Multiplier tubes which dominate the readout concepts of present day nuclear imaging devices.

Documents

The Role of Silicon Radiation Sensors and Integrated Front ...€¦ · Some of the Shortcomings of Present Day Clinical PET Scanners A reference for the new generation of PET scanners