Synchrotron light and imaging systems for medical radiology · 2004-01-13 · Synchrotron light and...

*Corresponding author. Tel.: #39-040-375-6233; fax: 39-040-375-6258.

E-mail address: [email protected] (F. Arfelli).

Nuclear Instruments and Methods in Physics Research A 454 (2000) 11}25

Synchrotron light and imaging systems for medical radiology

Fulvia Arfelli*

Dipartimento di Fisica, Universita+ di Trieste, Via A. Valerio 2, I-34127 Trieste, Italy

Abstract

In the past years several successful experimental studies have proven the suitability of synchrotron light in the "eld ofmedical radiology and in diagnostic medical imaging. To be mentioned and reviewed in the following are transvenouscoronary angiography with synchrotron radiation, bronchography, single- or multi-energy tomography and synchrotronradiation-based mammography. Amongst those the most demanding application is X-ray mammography. Here,a dedicated system has to ful"l the demanding requirements on the source and the X-ray detector. Synchrotronradiation-based mammography, which enables improved image quality maintaining a low dose level, takes advantage ofthe tunable monochromatic energy and of the peculiar laminar geometry of the beam. In addition, further improvementscan be achieved with the development of novel detectors for full-"eld digital mammography in order to overcome theintrinsic limitations of the conventional screen}"lm systems. Encouraging results in digital mammography have beenobtained at the medical beamline of the synchrotron radiation facility ELETTRA, Trieste. Besides, the high degree of thespatial coherence of the synchrotron source allows new imaging modalities based on phase contrast e!ects that producesuperior radiographs featuring highly enhanced contrast with a low delivered dose. ( 2000 Elsevier Science B.V. Allrights reserved.

Keywords: Synchrotron radiation; Mammography; Digital imaging detectors; Phase imaging techniques

1. Introduction

Worldwide synchrotron radiation (SR) facilitiesprovide an extraordinary kind of X-ray source,which o!ers a large variety of application in several"elds such as material science, crystallography,micro-spectroscopy, X-ray di!raction and manyothers.

The main feature of these sources is the wide andcontinuous energy spectrum providing a very highphoton #ux over an energy range up to some

50keV or even higher. Moreover, the beam com-prises a high natural collimation at least in thevertical direction and a high degree of coherence inboth, space and time. In the horizontal directionone can produce a large-area, fan-like photon beamshape.

All the aforementioned unique features in combi-nation with the sophisticated optics make synchro-tron sources well-suited instruments, e.g. formedical applications [1]. Thus during the pastyears several synchrotron radiation (SR) facilitieshave developed dedicated medical beamlines. Theynot only make use of the excellent source character-istics but also the fact that the high-intensitySR spectrum allows to select and tune monochro-matic photon beams with a very narrow energy

0168-9002/00/$ - see front matter ( 2000 Elsevier Science B.V. All rights reserved.PII: S 0 1 6 8 - 9 0 0 2 ( 0 0 ) 0 0 8 0 3 - 2 INVITED LECTURE

Fig. 1. The typical experimental set-up for transvenous coron-ary angiography with synchrotron radiation.

bandwidth. As a consequence an enhancement ofthe image quality is observed while the dose isconserved or is even reduced at the same time. Thisis simply due to the fact that for each applicationthe right energy window can be chosen. Besides thelaminar geometry reduces scattered radiation andthus avoids image blurring. Moreover, the mono-chromaticity can be utilized to implement multipleenergy techniques.

So far transvenous coronary angiography, com-puted tomography, bronchography and mammog-raphy have been implemented at several medicalbeamlines, just to mention some imaging applica-tions.

A dedicated imaging system is required for eachspeci"c medical application and various ap-proaches have been pursued with the developmentof di!erent X-ray digital detectors with appropriatefeatures like the geometry, the X-ray detectionmode and the electronics. In general a high spatialdetective quantum e$ciency (DQE) close to 1 isessential for detectors used in medical radiology.Note, that (1!DQE) is the additional dose ap-plied to the patient that does not contribute to anincrease in image quality [2].

In particular in the last 10 years successful e!ortshave boosted the development of systems for clini-cal and synchrotron-based full-"eld digital mam-mography using di!erent technologies.

Besides conventional radiography proceduresnew imaging modalities have been recently de-veloped aiming at further improvement in the im-age quality based on the detection of phase shiftinformation, an important e!ect especially for lowabsorbing biological materials.

2. Coronary angiography

Coronary angiography is a useful examination toreveal coronary artery disease that could lead tomyocard infarct and subsequently even to death ofthe patient. Here the main interest is the visualiz-ation of occlusions (stenoses) in arteries down toa size of 1mm in diameter. The conventional hospi-tal procedure consists of an injection of an iodine-based contrast agent (with 370mg/ml iodine)directly into the coronary arterial system via

catheter while an X-ray transmission image/videois taken. As a screening procedure, however, thismethod is not useful since it bears a high risk (evendeath in up to 0.1% of the examinations dependingon the age and the health condition [3]) due to theinsertion of the catheter.

SR-based transvenous coronary angiography(certainly the most advanced SR-based medicalprogram) is a less invasive method since the con-trast agent is injected in the veins with a sensiblereduction of risk. On its travel through the venoussystem and subsequently through the lungs theiodine bolus is rather diluted (around 10mg/ml).Using conventional equipment this method turnsout to be unsuitable for imaging fast moving ob-jects like the heart and only a SR source can pro-vide su$cient photon intensity in order to obtaina reasonable contrast between small arteries andthe surrounding areas.

Basically the set-up (Fig. 1) consists of twomonoenergetic photon beams bracketing the iodineK-edge and intersecting at the heart position. Twoimages are taken simultaneously by a dual linedigital detector while the patient is translated invertical direction through the cross-over point ofthe fan beams. In order to avoid artifact due to themovement of the heart the scan should be com-pleted within 250 ms. Once the images are takena digital logarithmic subtraction on a pixel base isperformed that allows to extract the iodine signal(Fig. 2). The following algorithm (1) is based on the(simpli"ed) assumption that the patient consistsonly of water and iodine [4]. In this fashion one

12 F. Arfelli / Nuclear Instruments and Methods in Physics Research A 454 (2000) 11}25

Fig. 2. Comparison between the mass attenuation coe$cients ofiodine and of water at the iodine K-edge.

obtains a system of two unknowns, namely thewater image (o

8!5%3x8!5%3

) and the iodine image(o

*0$*/%x*0$*/%

) by the simple matrix inversion of

!AS1

S2B"A

k1*0$*/%

o*0$*/%

k18!5%3

o8!5%3

k2*0$*/%

o*0$*/%

k28!5%3

o8!5%3

BAo*0$*/%x*0$*/%

o8!5%3

x8!5%3

B (1)

where subscripts 1 and 2, respectively, refer to theenergy below and above the iodine K-edge, S is thelogarithmic and normalized images taken at thetwo energies, o, k are the densities and absorptioncoe$cients for water (tissue) and iodine and x de-scribes the thickness of the speci"c material.

Several groups have taken various approaches touse synchrotron for intravenous coronary angiog-raphy with di!erences in X-ray optics and the typesof digital detectors summarized in Table 1.

The concept of coronary angiography with syn-chrotron radiation was developed at Stanford Uni-versity [5] and in 1979 the system was installed atthe Stanford Synchrotron Radiation Laboratory(SSRL) where the "rst human studies were per-formed. In 1990 the equipment was moved to theSynchrotron Medical Research Facility (SMERF)at the National Synchrotron Light Source (NSLS)in Brookhaven where the experiments continueduntil 1996 [6]. The imaging system at NSLS wasa 5mm thick dual-linear lithium-drifted silicon

detector with horizontal resolution of 250lm whilethe vertical resolution was de"ned by the beamheight [7].

Pioneering experiments at the Institute of Nu-clear Physics in Novosibirsk (Russia) started in1979 [8]. They used a double-line scintillationcounter detector based on NaI scintillator crystalscoupled with individual photo-multipliers workingin single-photon counting mode with a countingrate of about 1MHz per channel.

Di!erent from other projects the programs atPhoton Factory, Tsukuba (Japan), started in 1993,used a large beam obtained by an asymmetricallycut crystal and a two-dimensional detector. Oneexample is the system working above the K-edgeusing an X-ray image intensi"er and a CCD videocamera [9].

The "rst experiments with the NIKOS system atHASYLAB at DESY (Germany) started in 1981 andbetween 1997 and 1999 the beamline was dedicatedto extensive patient studies with some 400 patients[10] (Fig. 3). The focus here is to validate the diag-nostic sensitivity and the medical impact with rea-sonable statistics. Until now this system is the mostadvanced. The detector consists of a two-line high-pressure KrCO

2ionization chamber with a horizon-

tal pixel size of 400lm de"ned by the distance of thestrips [11]. A comparison in terms of spatial fre-quency-depending DQE between the NIKOS andNSLS detectors has been performed [2].

At the medical imaging beamline at EuropeanSynchrotron Radiation Facility (ESRF), Grenoble(France) successful human studies started in thebeginning of 2000 after preliminary results havebeen obtained on animals trials [12] (Fig. 4).A dual-line high-purity germanium detector(HPGe) with 350 lm pixel width made by Eurisysmeasures is used as an imaging device.

3. Bronchography

K-edge xenon dichromography for imaging thebronchial tree in humans is based on the sameprinciple used for synchrotron radiation trans-venous coronary angiography. Pilot studies [13]have been carried out on human volunteers at theSMERF facility at the NSLS using the same system

F. Arfelli / Nuclear Instruments and Methods in Physics Research A 454 (2000) 11}25 13

INVITED LECTURE

Table 1

Facility Detector Monochromator Operation and status

SSRL Dual line Si(Li) 2 Si(1 1 1) Started in 1979Stanford Integrating mode Asymmetrical cut Animal and human studies(USA) 5mm thick, 500lm pixel In 1990 moved the equipment

to NSLS4ms integration time/line14 bit dynamic

NSLS Dual line Si(Li) 1 bent LAVE Si(1 1 1) Started in 1990 Human studiesBrookhaven Integrating mode(USA) 5mm thick, 250lm pixel

4ms integration time/line14bit dynamic

VEPP-3 Dual line scintillation counters 2 LAVE Si(1 1 1) Started in 1979Novosibirsk Single photon counting mode In vivo studies(Russia) Spatial resolution 0.1}2mm

20 bits dynamic

Hasylab Dual line ionization chamber 2 bent LAVE Si(1 1 1) Started in 1981Hamburg Integrating mode Dedicated beamline(Germany) 400lm pixel for a large program on human

studies since 1997Integration time 0.8 ms perline19 bit dynamic

ESRF Dual line HPGe 1 bent LAVE Si(1 1 1) Started in 1998 "rstGrenoble Integrating mode animals. Human(France) 2mm thick, 350lm pixel studies in the beginning of 2000

integration time 0.8ms per line16 bit dynamic

Photon X-ray image Intensi"ed#videocamera 1 Si(3 1 1) Started in 1983Factory 2-dimension detector Asymmetrical cut In vivo studiesTsukuba spatial resolution 250lm (large beam) Human studies in 1996(Japan) Integration time 2ms

9 bit dynamic

as that for angiography, which was recon"gured forthe operation at the K-edge of xenon (34.56 keV). Inthis case a medical grade, non-radioactive xenongas mixture (80% xenon, 20% oxygen) is used asa contrast agent. A time series of images above andbelow the xenon K-edge of xenon are taken simul-taneously by the aforementioned Si(Li) dual linedetector after the inhalation of the contrast agent.The bronchi down to the tertiary level were visible.

4. Computed tomography

A SR beam o!ers a suitable planar geometry forcomputer tomography (CT). In this case the sampleis rotating while the laminar beam is crossing a slice

of the object to be imaged. Due to the very smallangular divergence the X-rays can be consideredparallel and the detector has a simple linear ge-ometry. Besides the ease of the geometry also themonochromaticity of the photon beam further sim-pli"es the reconstruction [14] which is based on theassumption that the dependence from the energy ofthe attenuation coe$cient does not change throughthe integration path. Thus artifacts due to beamhardening, which arise in normal tomography espe-cially at the borders of high absorbing materialslike bones and decrease the image quality, can besuppressed signi"cantly. Another advantage is thepossibility to choose the appropriate energy fora given size of the sample, resulting in larger signalsfor the same delivered dose.


Fig. 3. A picture of a human heart obtained with subtractioncoronary angiography at HASYLAB at DESY showing clearlythe right coronary artery (RCA) after an intravenous injection ofthe contrast agent.

Fig. 4. First experiments for transvenous coronary angiographyat the medical beamline at the ESRF: shown is the heart of a pigafter an intravenous injection of the contrast agent.

Moreover multiple-energy imaging can be per-formed since the small-energy bandwidth allowsenergy-selective methods like K-edge subtraction(KES) imaging and dual-photon absorptiometry(DPA). The K-edge subtraction method (describedin Section 3 for coronary angiography), can also beimplemented for CT at the iodine K-edge for imag-ing the human neck, for example. However it issuboptimal for head imaging and impractical forchest or abdomen imaging because of the lowtransmission of the 33 keV X-rays. For whole-bodyCT the use of gadolinium as contrast agent(K-edge"50.23 keV) is more suitable.

DPA is a subtraction method based on the di!er-ent behavior of the photoelectric and Comptone!ects at two widely separated energies: two imageshave to be acquired, for instance at 40 and 100keV.This method allows the di!erentiation of concen-tration of low- and intermediate-Z elements. DPAis also performed with X-ray tubes using "lteredbremsstrahlung radiation for measuring bone min-eral density, but the SR source can improve theimage quality.

Multiple Energy Computer Tomography(MECT) was "rst developed at medical facilitySMERF at NSLS with a program for imaginghuman neck and head with single- and dual-energymethods [15]. Experiments were carried out onphantoms and small mammals using a linearmodular CdWO4-pin photodiode detector witha beam provided by two LAVE #at Si(1 1 1) crys-tals. DPA will be applied in the study of tissuecharacterization as carotid artery atheroscleroticplaque composition.

The CT program at ESRF [16] is focused on theevaluation of blood volume and tissuelar perfusionby concentration measurements of contrast agentin the brain in order to predict the outcome ofcerebral ischemia and brain tumors. Here two bentLAVE Si(1 1 1) crystals are used as monochromatorand the aforementioned high-purity germanium de-tector as imaging system.

In Tsukuba (Japan) a CT system has also beendeveloped [17]. It was designed for performingKES with low-concentration contrast materials(iodine, barium, gadolinium) for research applica-tion like functional analysis imaging and bone min-eral distribution imaging.


INVITED LECTURE

Fig. 5. The theoretical calculation of the Figure of Merit (FOM)for a detail of 5mm diameter adipose tissue embedded ina breast with standard composition (50% adipose#50%glandular) for di!erent breast thicknesses.

5. Mammography

Mammography is a screening technique with thehighest sensitivity for detecting early breast cancer,which statistically leads to a signi"cant improve-ment in breast cancer survival [18]. The demandson the image performance in mammography, interms of contrast and resolution are beyond thoseof any other medical imaging modality. Sub-sequently the requirements on the X-ray source aswell as on the detector are very high especially forthe detection of low-contrast and small-size detailssuch as clusters of micro-calci"cations and low-contrast masses, possible indicators of early breastcancer. At the same time special care has to betaken for an e$cient use of the radiation dose to thepatient: since the breast is one of the most radiosen-sitive organs, the risk of cancer induced by X-rayexposure has to be minimized.

Since the "rst mammographic units becameavailable in the 1970 s several technology advance-ments have changed and have boosted both theequipment and the examination procedure withappropriate X-ray beam quality, adequate breastcompression and automatic exposure control. Un-fortunately, there are still several limitations andapproximately 10% of the obvious clinical breastcancers are not visible with mammography [19].

The image quality is a!ected, besides the breastcomposition and thickness, by the radiation qual-ity, the scattered radiation and the receptor perfor-mance.

The contrast of structures depends essentially onthe di!erence of attenuation coe$cient that, in caseof normal and neoplastic breast tissue, can be inthe order of few percent. In order to reveal these"ne structures a dedicated X-ray source is required,since the optimization of the X-ray source canenhance the visibility of details. In conventionalmammography a molybdenum target X-ray tube(Mo K-edge at 20 keV) in conjunction with a mol-ybdenum "lter is used. The energy spectrum con-sists of the typical continuous bremsstrahlungbackground with the 17.9 and 19.5 keV character-istic #uorescence line of the molybdenum target.The polychromatic spectrum can be modi"edchanging the tube voltage. However, the best en-ergy optimization can be achieved with a mono-

energetic beam. This becomes obvious regarding theFigure of Merit (FOM), a parameter which describesthe image quality versus the energy. According toEq. (2) it is de"ned as the signal-to-noise ratio nor-malized to the mean glandular dose (MGD):

FOM"

SNR

JMGD(2)

In an image where a detail is present in thebackground area the SNR measures the visibility ofthat detail and in the case of Poisson distribution itis de"ned as

SNR"

N1!N

2JN

1

(3)

where N1

is the average transmitted #uence ofphotons on the image in the background regionand N

2is the average transmitted #uence in the

area of the detail. From Eq. (3) it turns out that theSNR is higher at lower energies because the di!er-ence of attenuation coe$cients decreases with theenergy. On the other hand lower energy meanshigher absorbed dose in the organ. Inserting Eq. (3)in Eq. (2) the FOM curve shows (Fig. 5) a maximumcorresponding to a certain energy and conse-quently at that energy the image quality is maxi-mized for the same radiation dose. Since the FOM


depends on the tissue thickness and the tissue com-position, a selectable monochromatic energyshould be chosen according to the breast thicknessand variation in the composition between moretransparent adipose and denser glandular tissue.The example in Fig. 5 shows how the maximummoves to higher energies as the thickness of thesample increases.

Another important aspect is the beam geometry.In principle the spot size of the X-ray source shouldbe pointlike to avoid penumbra e!ects increasingthe image blurring. Moreover the wide-sized beamof an X-ray tube in#uences scattered radiationwhich decreases the contrast adding a uniformnoisy background to the image. In order to reducescattering anti-scatter grids are used. However, theinsertion of the grid results in an increase of theradiation dose applied to the patient.

5.1. Synchrotron radiation mammography

In order to improve the source quality SR-basedmammography experiments have been carried out.Comparing the energy spectrum of a SR source andof a mammographic X-ray tube it turns out that theSR intensity is orders of magnitude higher thanthat of a clinical unit. A monochromator can beutilized in order to obtain a beam with high #uxand with a narrow energy bandwidth. The appro-priate energy can be tuned according to the densityand thickness of the breast. In addition the highlyvertical collimated SR geometry reduce inherentlythe scattering in the image thus avoiding the use ofanti-scatter grids. Approximately a decade agoa "rst pioneering work was done at ADONE (Fras-cati) using a channel-cut Si(1 1 1) crystal as mono-chromator and a conventional screen}"lm system[20]. The conclusion of these experiments was thatthe images of test objects as well as breast tissuespecimens taken with the monochromatic SR havehigher contrast resolution and similar or less radi-ation dose. Over the past years other preliminaryinvestigations have been carried out by the medicalresearch group at NSLS (Brookhaven) withscreen}"lm and imaging plate detectors [21] and atSRS facility (Daresbury) with screen}"lm system[22] using in both cases #at Si(1 1 1) crystal mono-chromators.

Furthermore at NSLS additional studies forproducing scatter-free images with the help of ananalyzer crystal have been done. These studiesopened the way to the DEI technique (see Section7) based on the di!raction e!ects. Also at Dares-bury di!raction imaging studies have shown thepossibility of characterization of selected tissuetypes [23].

In order to fully validate and exploit the methoda dedicated beamline for mammography calledSYRMEP was designed and built at the SR facilityELETTRA, Trieste (Italy) and it is under operationsince 1996 [24,25]. A Si(1 1 1) channel-cut crystalmonochromator provides a 100 mm wide beam atthe experimental hall in the energy range from 10 to35 keV. The project studies the advantages of SR inmammography in conjunction with a novel digitaldetector based on a linear silicon array of pixels (seeSection 6.2). Joining the SR features with the digitalmammography approach promising experimentalresults have been obtained with both, phantomsand in-vitro full breast tissue samples (Fig. 6). Forcomparison images were taken with the digital de-tector as well as a classical clinical screen}"lmcassette [26]. For the images taken with the digitaldetector the applied dose to the sample was smalleror comparable to that of the clinical unit, howeverin terms of contrast resolution the digital ones aresuperior.

According to the speci"c experimental require-ments the fan-like beam is "rst monochromizedand then is shaped by a micro-metric tungsten slitsystem in the experimental hall. When using thelinear digital detector the beam height has to matchthe vertical pixel size to avoid undesirable dosedelivered to the sample. In all the SR based mam-mography experiments the image acquisition is ob-tained scanning the sample through the beam dueto the laminar shape of the beam.

6. Screen}5lm system limitations and digitalmammography

Screen}"lm systems are currently dominatingthe market and o!er excellent detection capabili-ties. However several limitations still a!ect thesedetection systems [27].


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Fig. 6. An in-vitro full breast tissue sample image acquired at20keV with the digital detector at the SYRMEP beamline atELETTRA.

Fig. 7. A characteristic curve of a screen}"lm system with thetypical sigmoidal shape.

One of the main drawbacks of "lm is its non-linear response. The contrast properties of a radio-graphic "lm are described by its characteristiccurve (Fig. 7) which has typically a sigmoidal shape.This means that the range of the X-ray exposurewith linear correspondence with the optical densityis limited to a factor of about 25.

The noise in this system is due to the structure ofthe #uorescent screen and, mainly, to the granular-ity of the "lm emulsion that prevents the system tobe &quantum limited' decreasing the contrast res-olution.

An important challenge in screen-"lm mammo-graphy is the capability to diagnose early stages ofbreast cancer in radiologically dense breast (orbreast types that contain a large amount of "bro-glandular tissue). Subtle attenuation variation in

breast tissue requires a high contrast sensitivitythat means a steep slope in the characteristic curveof the "lm. On the other hand a trade-o! betweenthe contrast sensitivity and the "lm dynamic range(latitude) has to be considered. High contrast sensi-tivity is not suitable for a large range of transmittedexposure due to the presence of dense (radiographi-cally opaque) and adipose (transparent) part inbreast composition. As a consequence dense partscan be underexposed or transparent parts can beoverexposed corresponding to the toe and to theshoulder of the characteristic curve.

Another limitation of screen-"lm mammographyis the compromise between spatial resolution andX-ray interaction e$ciency. To perform a radio-graph at a reasonable dose the "lm is used incontact with a #uorescent screen that converts theX-rays in the visible light. The absorption e$ciencydepends on the phosphor thickness, but as thethickness increases also the screen blur increasesreducing the spatial resolution. This is due to the


di!usion of light from the point of absorption of theX-ray to the "lm. In mammography high spatialresolution (18}20 lp/mm) is essential thus the screenmust be relatively thin with a consequent reductionof the detection e$ciency.

The contrast resolution is also a!ected by theintrinsic noise of this system which is due to thestructure of the #uorescent screen and, mainly, tothe granularity of the "lm emulsion that preventsthe system to be &quantum limited'. The DQE ofscreen}"lm system reaches a maximum of 0.25}0.50at low spatial frequencies and drops to less than1% at higher frequencies [28]. The loss of DQE isattributable primarily to "lm granularity noise.

In the digital approaches several limitations canbe overcome [27}29]. It should be pointed out thatin digital mammography the image acquisition, dis-play and storage are performed independentlyallowing optimization of each procedure. The re-sponse of detector is linear over a much widerdynamic range (12 bits or more), therefore moreideal for radiologically dense breast types with im-proved lesion detectability.

To ful"l the severe requirements of mammogra-phy the digital detectors must feature low intrinsicnoise and a very high detection e$ciency in orderto keep the dose to minimum necessary, hence,a high DQE system is necessary.

Furthermore, in a digital system the exposure isterminated when a su$cient signal-to-noise ratio isobtained and it is not de"ned by the optical densityof the "lm. The spatial resolution should be alsovery high and reasonable pixel size is supposed tobe about 100}50lm2. Moreover, a digital imagepresents all the advantages of the computer manip-ulations to improve the detectability of subtle andsmall suspicious lesions with o!-line contrast en-hancement procedures.

6.1. Full-xeld detectors

The main challenge in digital detector develop-ment is the implementation of a large area (to becomparable to the 18]24 cm2 "eld size of themammographic "lm cassette) together with a highspatial resolution. Di!erent solid-state detectorscould be available for this purpose, but it is nota trivial problem to design one that could produce

a full breast image with resolution matching that ofthe screen}"lm system (18}20 lp/mm). All the de-tectors under development have a spatial resolu-tion lower than that of the "lm. However the bettercontrast resolution achieved in all cases is expectedto compensate for the reduced spatial resolution.Di!erent approaches are pursued in detector de-sign employing either area or linear solid-state de-vices [27}30].

Small charge-coupled devices (CCDs) forstereotactic breast biopsy with a limited "eld ofview have been available for several years and re-cently a development of these systems has lead tonew designs for full "eld of view mammography.They all work in indirect-conversion mode thatmeans they utilize a scintillator, which is coupledwith the CCD by means of a "ber-optic taper toconvert the X-rays into visible light. If an unstruc-tured phosphor scintillator is used the trade-o!between spatial resolution and absorption e$cien-cy, correlated with its thickness, must be kept inmind. Several systems have employed columnar-structured cesium iodide scintillators with a re-duced lateral spread of the light.

One approach is to use a mosaic of 12 (3]4)charge-coupled devices (CCDs) bonded togetherand coupled to an 18]24 cm2 cesium-iodide scin-tillator to create a large imaging area [31]. Anothermethod is a slot}beam system, a good compromisebetween an area and a linear detector, developed bythe University of Toronto [32}34]. In the proto-type unit a gadolinium phosphor screen and a 2:1demagnifying "ber-optic taper and a pixel of50lm2 was used. The scanned-slot mechanism con-sists of a fan-shaped beam of mammographic X-raytube scanned in continuous mode over the breastsimultaneously with a slot-shaped CCD detector.This system can reduce the amount of scatteredradiation in the image without the use of grids. Ofcourse, in order to complete the scan without a lon-ger exposure time, the X-ray #ux must be increasedcompared to that used with an area detector. Forthis systems a time-delay integration (TDI)technique is applied in order to allow a smoothmechanical movement and an averaging of pixelresponse (defects and dishomogeneities) in the rowalong the scan direction [29,34,35]. As the detectorand the fan-shaped beam are scanned across the


INVITED LECTURE

Fig. 8. A schematic layout of a pixel in an amorphous silicon #atpanel detector.

Fig. 9. The scheme of an active matrix of Thin Film Transistor(TFT). For description see text.

breast the charges collected in each pixel are shiftedin the opposite direction and synchronized at thesame speed as that of the scan.

In a more recent development cesium-iodide iscoupled to a slot-scan CCD detector in TDI modewith 1.0]24 cm2 sensitive area [36]. The DQEexceeds 50% and the pixel size is 50]50lm2.

Another method for full-"eld mammography isbased on the use of large-area, #at-panel amorph-ous-silicon image sensors that have recently beendeveloped for X-ray imaging applications (radi-ography and #uoroscopy) [37,38]. The develop-ment in the technology of amorphous silicon isprimarily due to the commercial interest in #at-panel displays. They o!er the advantage (with re-spect to crystalline silicon) to be fabricated atreasonable low cost in large sheet (20]25 cm2),recently covering an area up to 30]40 cm2 [39].They can achieve a high DQE (around 65}75%)and a wide dynamic range. Recently high spatialresolution with a pixel size of about (100lm)2 hasbeen achieved [40].

These active-matrix, #at-panel imagers consist ofan array of pixels, each one formed by a single-amorphous-silicon (a-Si : H) thin-"lm transistor(TFT) which is coupled with a discrete a-Si : Hn}i}p photodiode working in indirect-conversionmode. As in the case of the previously discussedCCD detectors a scintillator such as a phosphor orcesium-iodide is coupled to the #at panel (Fig. 8).

The electronic signal accumulated in each pixelcapacitance is read-out by the TFT switches con-nected to the gate lines and the source lines [29](Fig. 9). When one gate line is activated all theTFTs in that row are switched on and the signalsfrom all the pixels in the row can be read-out by the

source lines. Because of its architecture the activefraction ("ll factor) of the pixel is less then 100% anddecreases with the pixel size resulting in a limit forthe spatial resolution (35% for (127lm)2 pixel [28]).However very recent studies [41] describe a newdesign for an amorphous silicon sensor array witha pixel size of about (70lm)2 with a high "ll factor.

In the last few years growing interest and pro-gress have been reported on the development ofdirect-detection selenium #at-panel digital radio-graphic imaging devices [42,43]. An amorphousselenium photo-conductor layer is coated overa TFT readout active matrix.

The charges, generated by means of photoelectricinteraction of the X-rays with the selenium layer,are integrated on storage capacitors that arelocated at each pixel. A high electric "eld (typically10V/lm) must be applied across selenium to collectcharges and to avoid their recombination. In thisway the lateral spreading of the signal is minimizedresulting in a high spatial resolution.

A 500lm thick selenium device has been de-veloped [44]. It has large active imaging area(35]43 cm2) with a pixel size of (139lm)2 anda high "ll factor of 86%.

Selenium detector prototypes with a design suit-able for digital mammography performing highDQE and high special resolution (optimized


Fig. 10. A picture of the three-layer SYRMEP detector duringthe assembly of the "rst two layers: visible are the silicon stripdetector, the fan-out and the electronics.

conversion layer thickness and pixel size smallerthan (100lm)2) are currently under study [45}48].

6.2. The digital detector at the SYRMEP beamline

The digital detector utilized at the SYRMEPbeamline at ELETTRA for mammography consistsof a micro-strip silicon detector made by Canberrawhich is irradiated from the side in an &edge on'geometry [49]. In this con"guration it acts as a lin-ear pixel detector with a horizontal pixel size of200lm de"ned by the strip pitch and a verticalpixel size of 300lm due to the wafer thickness. Thiscon"guration allows an absorption e$ciency ofabout 100% in the energy range of interest since thebeam impinges parallel to the longitudinal dimen-sion (10mm) of the strips which becomes the e!ec-tive detector thickness. This pixel depth ensuresa high detection e$ciency (about 80% for X-rays at20 keV) which is anyway lower than the absorptione$ciency because of the presence of an insensitivevolume. This is due to the fact that the strips areimplanted 250lm from the edge [50]. A single-layer detector consists of an array of 256 stripscovering an area of 51.2]0.3mm2.

The electronics works in single-photon countingmode; hence the detector is &quantum limited' be-cause the image noise is only due to the photonPoisson statistics resulting in a high dynamic rangeand in a better contrast resolution. The readoutelectronics is realized in the form of a low noiseVLSI chip [51], which is made by LEPSI (Stras-bourg). It is a 32-channel mixed analogue-digitalcell custom designed for applications requiringlow-noise ampli"cation and counting of singlepulses. Each channel features an analogue part,which consists of a charge-sensitive preampli"er,a CR-RC shaping ampli"er, an analogue bu!er,a threshold discriminator, and a 16 bit pulsecounter. The read-out of all counters in the circuitis done in serial up to a rate of 20 MHz. Thecounting rate using periodic signals is close to100kHz. During the data acquisition, however, it iskept lower (about 10 kHz) to avoid partial satura-tion of the read-out chain due to the statistical timedistribution of the photons.

In order to achieve a shorter acquisition timea multilayer detector with a larger sensitive area

has been implemented [52]. To allow the stackingof devices one on the top of each other theSYRMEP detector was produced in a special de-sign. The "rst &stacked' assembly prototype consistsof three layers covering a cross section of about50]1mm2 (Fig. 10).

7. New imaging techniques

In the recent years with the increasing use ofsynchrotron sources several researchers have inves-tigated the potential of di!erent imaging techniquesbased on phase e!ects. The contrast enhancementin comparison with the normal absorption tech-niques can be achieved by exploiting refractiondi!erences in those heterogeneous samples, whichnormally show only little absorption contrast.These di!erences can be related to a phase shift ofthe incident wave "eld propagating inside thesample and can be described with correspondingvariations in the real part of the complex refractiveindex n"1!d#ib. In the energy range suitablefor the X-ray imaging of biological samples theycan be considerably larger (&103) than the imagi-nary ones, which are related to the absorption.


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Fig. 11. Simulated tissue in a mammographic phantom ob-tained with a clinical mammographic unit (upper image) and atthe SYRMEP beamline (ELETTRA) with the phase contrasttechnique at 20keV (lower image). In both cases the sameclinical screen}"lm system was used with a comparable dose.

One modality that exploits these phase e!ects isthe Phase Contrast `in linea imaging method thatrequires high spatial coherence of the X-ray sourcelike micro-focus X-ray tubes or SR sources [53}55].

Unlike in optical in-line holography the phaseinformation is converted here into an intensity in-formation generated by the interference between anunperturbated reference wave "eld and a wave "elddi!racted by small objects inside the sample. Sincea larger gradient of the refractive index is present atthe borders of those details this technique producesa tremendous edge enhancement e!ect in the im-ages. The intensity information that can be resolvedis just a question of the right distance between thesample and the detector and eventually of the spa-tial resolution.

Usually the detectors utilized for this techniquehave a very high resolution; hence a high radiationdose is delivered to the sample. However, studies onmammographic phantoms and breast tissue sam-ples at the SYRMEP beamline at ELETTRA havedemonstrated the possibility of performing phasecontrast imaging using clinical mammographicscreen}"lm system. That means an improved detailvisibility with a dose fully comparable to the clini-cal one [56] (Fig. 11).

A phase contrast X-ray-computed tomographymethod for observing biological soft tissue has beenrecently developed and demonstrated by Momosefollowing an idea of using an X-ray Bonse}Hart-type interferometer [57,58].

Another technique considers that a phase gradi-ent across a wavefront is equivalent to a variationin the direction of propagation of the wave. Thesevariations can be resolved with an analyzer crystalof high angular sensitivity due to its narrow re#ec-tivity curve [59,60]. One of those methods is calledDi!raction Enhanced Imaging (DEI) and was de-veloped at NSLS for mammography with im-proved image contrast [61]. DEI is based on theuse of an analyzer crystal and a post-processingelaboration. It is capable of producing images ofthe refraction and absorption independently ina sample acquiring two images, one on the left andone on the right hand side of the crystal re#ectivitycurve. Images of a standard mammographic phan-tom and breast tissue sample have been obtainedusing a Si(3 3 3) crystal.

Studies of di!raction imaging with phantomsand breast tissues using a Si(1 1 1) analyzer imagingcrystal have been carried out at the SYRMEPbeamline at ELETTRA [62]. A clinical screen}"lmcassette and a Fuji BAS 1800 imaging plate systemfeaturing a pixel size of 50lm have been used asdetectors. Images obtained with the analyzer per-fectly aligned with the primary monochromator arein principle scatter-free while the images obtainedwith di!erent degrees of misalignment showa weighted contribution of the transmitted beamand of the beam refracted at a certain angle. As anexample the radiographs of nylon strings of 100 lmdiameter and quartz spheres of 150 lm diameterare depicted in Fig. 12. The X-ray energy was set to17 keV and in all three cases the applied dose wasalmost the same. Note that in the transmissionimage the strings are not visible while the imagestaken with the analyzer show all the details includ-ing the strings with a high SNR.


Fig. 12. Images of 100lm nylon wires and 150lm quartzspheres obtained at the SYRMEP beamline (ELETTRA) intransmission geometry (upper image), with the analyzer alignedwith the primary monochromator (middle) and with a misalig-ned analyzer (lower).

8. Summary

X-ray synchrotron radiation sources have a wide-spread "eld of interest and in the past years itwas demonstrated that synchrotron radiation isa powerful tool in medical/radiological applications.

Coronary angiography is the most advanced medi-cal application and recently the validity of themethod has been tested on human patients ina large program at HASYLAB at DESY. Clearlymammography can take advantage of the peculiarcharacteristic of the synchrotron radiation sources.In order to ful"l the demanding requirements ofthese experimental researches a big e!ort in a par-allel development of imaging system technologywith digital detector and electronics has to be laun-ched. In particular, crucial will be the full-"elddigital mammography and di!erent solutions havebeen investigated. Some systems produced for stan-dard medical X-ray equipment are already com-mercially available. For SR-based full-"eldmammography a novel &edge on' Si-strip detectorhas been developed.

Recently new promising phase imaging modali-ties have been developed which can highly enhancethe image contrast hence opening interesting per-spective also in medical diagnostic imaging. Heresynchrotron light o!ers a suitable X-ray source toadvance these techniques.

Acknowledgements

I would like to thank the members of theorganizing committee of the SAMBA Confer-ence for their kind invitation and their warmhospitality. It is a pleasure for me to thank Dr.H.J. Besch for many enlightening discussions andfor the good times during the common experi-ments.

All my gratefulness are due to Dr. R.H. Menk forall the interesting discussions and all his teachingand advises from his experience in SR applicationsand detectors. A special thanks for his careful revi-sion of my paper.

I am indebted to Dr. W. Thomlinson and theother researchers of the medical facility at NSLSfrom whom I gained a fundamental experience inSR medical applications and my knowledge of theDEI technique.

I would like to thank Prof. E. Castelli and allmy colleagues and friends for all the importantwork we have done together at the SYRMEPbeamline.


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