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Current Medical Imaging Reviews, 2007,3, 45-59 45

1573-4056/07 $50.00+.00 2007 Bentham Science Publishers Ltd.

Small Animal Computed Tomography Imaging

Soenke H. Bartling1,2,*, Wolfram Stiller2,*, Wolfhard Semmler2and Fabian Kiessling1,2

1Junior Group Molecular Imaging, 2Department of Medical Physics in Radiology, German Cancer Research Center

(DKFZ), Heidelberg, Germany

Abstract: Micro Computed Tomography (micro-CT) was suggested in biomedical research to investigate tissues andsmall animals. Its use to characterize bone structures, vessels (e.g. tumor vascularization), tumors and soft tissues such as

lung parenchyma has been shown. When co-registered, micro-CT can add structural information to other small animal

imaging modalities. However, due to fundamental CT principles, high-resolution imaging with micro-CT demands for

high x-ray doses and long scan times to generate a sufficiently high signal-to-noise ratio. Long scan times in turn make the

use of extravascular contrast agents difficult. Recently introduced flat-panel based mini-CT systems offer a valuable trade-

off between resolution (~200 m), scan time (0.5 s), applied x-ray dose and scan field-of-view. This allows forangiography scans and follow-up examinations using iodinated contrast agents having a similar performance compared to

patient scans. Furthermore, dynamic examinations such as perfusion studies as well as retrospective motion gating are

currently implemented using flat-panel CT.

This review summarizes applications of experimental CT in basic research and provides an overview of current hardware

developments making CT a powerful tool to study tissue morphology and function in small laboratory animals such as

rodents.

Keywords:Small animal imaging, Computed Tomography (CT), micro-CT, mini-CT, flat-panel detector, motion gating.

INTRODUCTION

Micro- and mini-CT are scaled down CT-imaging moda-lities for small animals, which in principle provide the sameinformation about morphology and disease status or disease

progression for animals as clinical-scale CT does forhumans. Nonetheless, several major differences compared toclinical CT scanning exist.

This review provides an overview of basic concepts andtechniques that apply to CT when used for in-vivo smallanimal imaging. Technical and physical limitations of theseconcepts as well as applications of micro- and mini-CT and

an outlook on future developments in the growing field ofresearch in small animal CT imaging are given anddiscussed.

Definition

There is no unique definition of mini-CT and micro-CT.Often, all CT scanners that provide higher spatial resolutionthan current clinical scanners are named micro-CT. Thisdefinition results in a very broad range of different CTscanner design concepts ranging from CT scanners designedand used for non-destructive material-testing with aresolution in the order of a few m up to dedicated smallanimal CT scanners with a resolution of a few hundred m.Taking the different design constrains of these various types

of scanners into account, the term mini-CT should be used todescribe CT systems with a resolution ranging from 100 m

*Address correspondence to either author at the Molecular Imaging Group,Department of Medical Physics in Radiology, German Cancer ResearchCenter (DKFZ), Im Neuenheimerfeld 280, 69120 Heidelberg, Germany;Tel: +49 (0) 6221 422686; Fax: +49 (0) 6221 422557; E-mail:[email protected]; [email protected]

*These authors contributed equally to this article.

to 500 m; micro-CTs are then scanners with a resolution ofbelow 100 m [Kalender W (2006), personal communica-tions] [1] (Fig. (1), Table 1).

Small Animal Imaging

Due to the fact that more and more research is based onanimal models the interest in in-vivo small animal imagingrises. With small animal imaging methods cross-sectionalstudy designs (e.g. where cohorts of animals are killed ateach time point and histology is examined) can be replaced

by an in-vivostudy design permitting repeated measurementsof the same animals. Each animal acts as its own control,

therefore the variability that would normally be presentwithin a cohort of animals is accommodated and changes aremore readily detected by paired comparisons [2].Additionally, CT imaging scaled down to small animals canhelp to define, optimize and test future performance goals ofclinical-scale CT.

Scaling down CT imaging to the size of small animals ischallenging: The volumes of mammals morphologicalstructures and organs are proportional to their weight [3]. So,to acquire CT data, e.g. of mice, that show their internalorgans with detail comparable to clinical CT scans of human

patients, a small animal imager needs a resolution of about100 m [1].

For motion-compensated imaging (i.e. taking intoaccount cardiac and respiratory motion present during imageacquisition) stakes are similar. The heart rate of a mouse isabout 400-600 min-1 and respiration frequency ranges from30-60 per minute. To use the diastole as the phase of theheart cycle that shows a minimum amount of motion, a basaltemporal resolution of the CT scanner in the order of 50 msin mice in comparison to 300 ms in humans is necessary [4].

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46 Current Medical Imaging Reviews, 2007, Vol. 3, No. 1 Bartlinget al.

BASIC DESIGN CONCEPTS

Several design factors of CT scanners such as geometry,the employed x-ray source and the detector technologyinfluence the fundamental characteristics of CT scanners.Inherent relationships between resolution (spatial, temporalas well as soft-tissue contrast), noise, scan time, scan field-of-view (FOV) and applied x-ray dose are imposed by the

fundamental laws of physics. Therefore there is not onescanner design that is able to optimize all of thesefundamental scanner characteristics. Instead, several diffe-rent scanner designs exist which are suited for differentkinds of diagnostic questions (Fig. (1), Table 1). Fundamen-tal CT design concepts will be discussed in the next

paragraphs.

Fig. (1).Examples for three different design concepts of small animal CT scanners. A bench-top micro-CT (A, B) with rotating sample

holder, stationary area detector and micro-focus x-ray source offering variable magnification. Such a setup is mostly used for in-vitro

imaging. An optimized relationship between scan field-of-view and resolution as well as good animal handling due to a non-rotating sample

bed is provided by a rotating gantry based concept (C, D). Imposing fewer demands on spatial resolution, faster scanning and a bigger scanfield-of-view can be achieved with the displayed flat-panel detector, rotating gantry-based design with stationary sample bed (E, F).

Reprinted with permission from Willi Kalender, Computed Tomography, Publicis Corporate Publishing.

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Small Animal Computed Tomography Imaging Current Medical Imaging Reviews,2007,Vol. 3, No. 1 47

Rotating Sample vs. Rotating Gantry Systems

Small animal CT-scanners can be classified into twomajor categories: rotating sample and rotating gantrysystems [5].

So-called rotating sample systems feature a stationary x-ray source and a stationary x-ray detector usually mountedon a mechanical bench. Both detector and x-ray tube arefacing each other and have to be precisely aligned with thecentral axis of the CT system. A rotating sample holder is

placed on the mechanical bench in between source anddetector, also aligned with the central axis of the scannersystem. The sample holder can usually be precisely rotated

perpendicular to the central axis of the scanner system by acomputer-controlled motor-driven rotating stage (Fig. (1), A,B). For in-vivo animal imaging this implies that theanesthetized animal has to be immobilized and fixated in anupright position (e.g. head up) on the rotating sample stagewhich demands careful animal handling [6]. In addition,administration of inhalation anaesthetics and injection ofcontrast agents during a scan is difficult since the neededapparatus may not inhibit or conflict with the samplerotation.

Most of the rotating sample systems have the advantagethat the scanner geometry can easily be modified in betweenscans. The source-to-isocenter (also called source-to-object)distance (SOD) can be varied by shifting the x-ray tubealong the central axis of the system relative to the sampleholder defining the isocenter of the scanner. The source-to-detector distance (SDD) can be varied by either shifting onlythe detector relative to the scanners isocenter along itscentral axis while leaving the source fixed or by moving bothsource and detector relative to the isocenter. This flexible

design enables the user to choose the optimum scannergeometry for the imaging application.

Systems with rotating gantry but stationary sampleaccommodate x-ray source and detector mounted exactlyfacing each other on the inside of a ring-shaped mechanicalsupport (gantry). Here, the gantry containing source anddetector is no longer stationary but rotates around the centralaxis of the scanner. The sample which is to be imaged can be

placed in a prone or supine position on a patient bed orpatient table. (Fig. (1), C-F). The longitudinal axis of this

table is parallel to the central axis of the CT scanner and thetable is usually driven by a computer-controlled micrometerstage, allowing exact and reproducible positioning of thesample within the scanner system. Small animal CT systemsfeaturing a rotating gantry assembly can easily be identifiedas scaled-down clinical CT scanners adapted to the specificrequirements of small animal imaging (Fig. (1), C-F).Compared to scanners having a rotating sample holder,rotating gantry based systems usually do not offer the

possibility to change the scanner geometry easily, normallySDD and SOD are fixed numbers of the particular scannersystem. Animal handling during examination is much easierin rotating gantry systems. In addition, these systems supportfaster rotation times than rotating sample systems. However,they usually are more complex mechanically and more

intricate from the point of view of systems control andtherefore more expensive.

No matter if the small animal CT scanner is rotatinggantry based or not, both system types can be either bench-top systems as shown in Fig. (1) A-D or systems requiringmechanical support structures too large to be scaled to fit a

bench top, e.g. if modified clinical x-ray tubes offeringhigher x-ray photon fluxes than micro-focus tubes or verylarge area detectors e.g. to cover large FOV sizes (Fig. (1),

Table 1. Comparison of Micro-, Mini- and Clinical-Scale CT

Micro-CT Mini-CT Clinical-scale CT scanner

Suited for Tissue samples, insects, mice, rats Mice, rats, rabbits, primates,

mini-pigs

Up to humans

Spatial resolution (isotropic) 5 m (single limbs) 100 m (whole animals) 100-450 m > 450 m (z-axis > 600 m)

Transaxial scan field-of-view

(FOV)

1-5 cm 5-20 cm > 20 cm

Time to acquire a standard

volume (e.g. a whole animal)

Seconds to hours (CT scanners with single

slice acquisition within subsecond times exist)

Subsecond (0.5 s) to a few

seconds

A few seconds (with rotation

times down to 0.33 s)

Radiation dose > 1 Gy can be reached ~ 10-500 mGy < 50 mGy

Design Bench-top, rotating sample (with variable

geometry, resolution, scan field-of-view, etc.)

or rotating gantry

Rotating specimen or rotating

gantry (fixed geometry)

Rotating gantry (fixed geometry)

Cardiac- & respiratory motion

compensation

Prospective triggering Prospective triggering,

retrospective gating

Scan modulation, retrospective

gating

Example figures Fig. (1) A, B, C, D, (3), (4) Fig. (1) E, F, (2), (5), (6)

Overview of different in-vivosmall animal CT scanner designs (values given are approximations). The term micro-CT is often used for all scanners that are smaller than clinical

scale scanners, however, taking the performance differences into account the term mini-CT could be used for scanners with a resolution of around 100 m.

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E, F) are to be used. An overview over commercial andexperimental laboratory micro-CT scanners of both types can

be found in [7], a compilation of the most recent smallanimal CT scanner models, including those featuring flat-

panel detector systems can be found in [8].

Scanner Geometry

Independent of the x-ray source employed or the detectortechnology chosen for a particular small animal CT-scannerthere are two possibilities for CT scanner geometry: the so-called short scanner geometry where the SOD is smallcompared to the object-to-detector distance (ODD) and theso-called long scanner geometry where SOD and ODDhave about the same size [Kachelriess M (2006), personalcommunication]. Both geometries have their advantages anddisadvantages on the imaging properties of the scanner butshare the requirement that the small animals imaged should

be entirely covered by the available FOV.

Short scanner geometries place the animal close to the x-ray source, the ODD being larger than the SOD. The image

projected onto the detector can thereby be magnified by thefactor

M= ODD

SOD

,

improving spatial resolution of the scanner system.However, the resulting FOV is being decreased by the samefactor requiring larger detectors. With the advent of large-area flat-panel detectors higher magnification factors will be

possible, since the active detector area will increasesignificantly. Placing the animal close to the x-ray source,however, requires the use of micro-focus x-ray tubes sincelarger focal spot size will introduce a significant image blur(so-called penumbral blurring), which degrades resolution.The problem of penumbral blurring is aggravated by high

magnification of scanner systems since the penumbra of thesources focal spot is also magnified by the factorM.

Long scanner geometries also have their advantages.Image blur caused by the finite size of the x-ray source focalspot is decreased if the object is placed closer to the detectorinstead of being placed close to the source. However, the x-ray photon flux will decrease exponentially with

1

SOD2

,

which can be compensated by x-ray tubes with higher output(but normally bigger focal spot because of anode heating).Otherwise, the scan time per projection needs to be increasedor contrast resolution will be degraded since (statistical)image noise (being related to the attenuated photon flux

attreaching the detector by1

att) will increase.

The larger the distance between source and object (SOD)is, the smaller the skin entrance dose to the animal imagedwill be. This implies that although scanners with shortgeometry usually are more dose efficient and allow togeometrically magnify the object, thereby improving spatialresolution, long scanner geometries can significantly reducethe skin entrance dose of the animal under investigation.

X-Ray Source

The scanner geometry and the desired spatial, low-contrast spatial and temporal resolution govern the choice ofx-ray tubes for small animal CT imaging. An x-ray sourcefor micro- and mini-CT has to fulfill three demands: thefocal spot size has to be as small as possible, the source hasto emit a high photon flux , and x-ray energies should beselectable over a suitable range.

As already mentioned above the focal spot size should beas small as possible in order not to have a negative influenceon spatial resolution by image blurring. However, the size ofthe focal spot limits the available x-ray photon flux. Forsmall focal spot sizes and stationary targets, heat dissipationfrom the target area is proportional to the diameter of the

focal spot and radial to first order, the maximum poweremission being

Pmax

1.4 x f,FWHM( )0.88

, with x f,FWHM

being the focal spot size in microns [9]. Since the emitted x-ray photon flux is roughly proportional to the product ofthe x-ray anode current Iand the square of the tube voltageUand the tube power is P = UI, the available x-ray flux islimited by the size of the focal spot, which can only absorb acertain amount of heat (anode heating). The maximum powerthat an x-ray tube can emit thus also depends on the heatcapacity of the anode material and the tube technique used:

tubes with rotating anode support much higher output thantubes with stationary anodes since heat is evenly distributedalong the focal spot trajectory.

High x-ray flux is desirable for small animal CT imagingto achieve high temporal resolution and short scan times: ifthe photon flux is high enough, sufficient x-ray photonsreach the detector and can be collected in short times foreach projection. A sufficient amount of x-ray photons isrequired to limit image noise and allow good low-contrastspatial resolution [6]. Through short scan times in-vivoscanning of animal subjects is eased since the animals do nothave to be anesthetized for long times. Also, high photonfluxes are needed to ensure high temporal resolution for

perfusion studies or motion compensated imaging (ECG orrespiratory gated imaging).

In order to achieve the best low-contrast spatialresolution in small animal CT imaging, the x-ray tube shouldallow choosing a range of tube voltages, thereby selectingappropriate x-ray energies E. For higher x-ray energies theenergy-dependent absorption coefficient (E) is small andlow-contrast spatial resolution is limited due to the smallnumber of x-ray photons absorbed in the animal; if x-rayenergy is low and thus (E) large, most photons are absorbed

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in the animal and the contrast resolution is limited by thesmall amount of x-ray photons reaching the detector [9]. Insmall animal CT imaging photon energies in the range of 30-50 kV are commonly applied. Hereby, the advantage is thatx-ray attenuation is up to two orders of magnitude higherthan in clinical CT systems usually operated at 80-140 kVwhich allows better discrimination of soft tissue types [1].These low-energy x-ray photons are still able to penetrate theanimal since its examined volume is much thinner than that

of human patients. Additionally, the highest absorptiondifferences of iodine lie in the low energy range, leading toan improvement in contrast-enhanced scans for the contrast

between iodinated contrast agents to the surrounding tissue.

For small animal CT scanning three different types of x-ray sources have been used so far: micro-focus x-ray tubes,diagnostic x-ray tubes as in clinical-scale human patientradiological equipment and diagnostic x-ray tubes modifiedto fulfill the special requirements of small animal CT.

Micro-focus x-ray tubes with focal spot sizes ranging for5-50 m (~ 5-200 W power emission) have mostly beenemployed in micro- and mini-CT and have the advantage ofallowing high isotropic spatial resolution down to a range of

10-50 m for FOV sizes of 30-50 mm [7]. Others haveemployed diagnostic x-ray tubes with larger focal spot sizes(0.3 mm at 9 kW and 1.0 mm at 11 kW) in order to profitfrom the high x-ray flux at short exposure times [6, 10, 11]to increase temporal resolution for cardiac and respiratorygating in small animal CT while still achieving an isotropicspatial resolution of 100 m. Experimental small animal CTsystems with modified diagnostics x-ray tube exist [12].Here, the tube filament was shortened to decrease the focalspot size of the rotating anode x-ray tube while still offeringconsiderable photon fluxes which make high scan speed

possible and are needed for the flat-panel area detector.

Detector Technology

Detector technology plays a crucial role for the

performance characteristics of all small animal CT systems.Together with the focal spot size of the x-ray source the sizeof the detector elements has the biggest influence on thespatial resolution of the images, their size also influences thelow-contrast resolution together with x-ray flux.

No matter what detector technology is used (see below)in a particular small animal CT system, all of its parts (e.g.image intensifier screen, optical coupling medium, photodetector, etc.) have to be optimized for the following factors:high quantum efficiency and uniform response throughout itswhole surface at the chosen x-ray energies and good pixelresolution through small detector pixelation. The detectorsdynamic range needs to be as large as possible to ensuregood low-contrast resolution and high dose efficiency,

whereas its noise and dark current should be as low aspossible. A high read-out rate is favorable for the systems toachieve a maximum temporal resolution and to minimizescanning times. Of course the area of the detector should beas large as possible so large FOVs can be employed andnone of the detector elements should introduce geometricaldistortions [9].

From the first developments of small animal CT systemsonwards different detector types have been used. Apart from

an overview over the different types that have been used todate we limit our focus to the emerging flat-panel x-raydetectors, which will be elaborated on in more detail. For adetailed look on digital x-ray detector technology pleaserefer to [13].

As in clinical-scale CT systems, detector systems inmicro- and mini-CT can consist of linear arrays of

photodiode detector elements illuminated by a so-called fan

beam. This approach to small animal CT was used in someof the early systems but is still used by some commercialsmall animal CT manufacturers [7]. After the developmentof the so-called cone-beam reconstruction algorithm [14],the first three-dimensional small animal CT systems weredescribed end of the 1980s and beginning of the 1990s untilthe early 2000s. They employed two-dimensional detectorsconsisting of phosphor screens which were optically coupledto cooled charge-coupled detector (CCD) arrays via anoptical lens or a fiber-optic taper (e.g. [15, 16]). Though

being inefficient, the optical coupling via lenses offers thepossibility of variable image magnification, whereas fiber-optic coupling is very efficient but uses fixed magnification[1]. This detector setup is technically demanding but remainsone of the most sensitive x-ray detection methods [7].

Holdsworth and co-workers (1993) used x-ray imageintensifier (XRII) screens coupled to a video readout [17].The rapid illumination response to x-ray radiation is themain advantage of image intensifiers [1]. Since the 1990sCCD-based detector arrays with scintillating plates (e.g.made of CsI(Tl), GdO2SO4, etc.) have become common insmall animal CT systems [6]. Recently, flat-panel areadetectors have been introduced to small animal CT imaging[10, 12, 18-20].

An overview of indirect (x-rays are converted to light inthe scintillating layer first, then the light is detected by

photodiodes and subsequently amplified) and directconversion (x-ray photons are directly detected) flat-paneldetector types and their technical differences can be found in

[21]. The flat-panel detectors used in recent flat-panel smallanimal CT systems are of the indirect conversion type. Theyusually consist of a scintillating crystal layer like thallium-doped caesium iodide (CsI(Tl); or GdO 2SO4, etc.) coupled toa photodiode array in form of an amorphous silicone (a-Si)wafer [10, 22, 23]. Others have employed photodiode arraysfabricated by a complementary metal-oxide semiconductor(CMOS) process [18, 19]. Compared to XRII detectors flat-

panel detectors have some advantages: their structure is thinand they allow for large-area detection (see below) withoutgeometrical distortions [24]. Because of technologicaladvances they can be produced in good quality and with high

precision. The hydrogenated a-Si material used has theadvantage that it can be deposited on very large areas

(making large FOVs possible) at relatively low temperatures(200-250 C, easing fabrication), has the properties of asemiconductor (photoconductivity in the visible spectralrange) and does not age through x-ray exposure [25], whichis a drawback of CCD- and CMOS-based detectors.

However, a-Si flat-panel detectors have several draw-backs: the pixel size of around 100-200 m2 is relativelylarge compared to pixel sizes of ~2.5 m2 achievable withCCD detectors [16], due to low fill-factors of ~ 45-70 %

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([18, 26]) the geometric efficiency is not as high as in CCDdetectors, and the so-called image lag (ghost images of

previously acquired projections due to charge traps in the a-Si layer, e.g. dangling bonds in the amorphous structure) hasa negative influence on spatial, low-contrast and temporalresolution [27]. The effects of image lag can be quite seriouson image quality. They are described for flat-panel computedtomography in general in [28]. Algorithms are applied tocorrect for the effect of image lag; projection frames pre-

viously acquired are weighted and subsequently subtractedfrom the currently acquired frame [23]. Compared to otherdetector types, flat-panel detectors in small animal CT alsosuffer from their smaller dynamic range which candeteriorate image quality: insufficient resolution of pixelgain and offset normalization causes ring artefacts andtruncation of low-density anatomical detail due toinsufficient dynamic range results in shadow artefacts andincorrect Hounsfield numbers (HU) [29]. In addition, flat-

panel detectors need frequent and careful recalibration foroffset and pixel gain factors [23, 29].

Nonetheless, the advantages of flat-panel detectors usedin small animal CT imaging by far outweigh the disadvan-tages stated and cited here. Further developments in flat-

panel detector technology ([21, 29]) will allow even morestudies with and new applications of small animal CT. Theuse of other detector materials like CdTe or CdZnTe for flat-

panel detector fabrication might increase detector efficiencyfor the x-ray energies employed and thus lead to improvedsoft tissue contrast [30].

Spatial Resolution

In small animal CT spatial resolution is one of the keyparameters since the size of morphological structures isapproximately proportional to the weight of the (laboratory)animal examined. This leads to the requirement that smallanimal CT systems should be capable of providing a spatialresolution in the order of 100 m. Spatial resolution of small

animal CT scanner systems is mainly determined by thefollowing factors [9]: the geometry of the scanner system,which can be chosen to allow geometrical magnification ofthe specimen imaged, the size of focal spot having aninfluence on the magnitude of penumbral blurring (see

below) and the chosen detector technology, whichdetermines the minimum image pixel or voxel size throughits intrinsic resolution. Thus, the size of the single detectorelements sets the maximum spatial resolution a particularmicro- or mini-CT can theoretically achieve, even if imagereconstruction features image matrices with smaller pixel orvoxel sizes. Reconstruction voxel sizes chosen too largecompared to the intrinsic detector pixel size reduce thespatial resolution of a CT scanner, while very smallreconstruction voxels size do not increase its intrinsicresolution. Ideally, the reconstruction voxel size should bechosen to be half the size of the intrinsic resolution of a CTscanner complying with Nyquists sampling theorem.

Per definition spatial resolution is the smallest distancebetween two objects still being distinguishable as twoseparate entities. In order to describe the spatial resolution ofCT images obtained from a particular scanner, themodulation transfer function (MTF) is employed. The MTF

plots the percentage of contrast that is transferred from the

imaged object (e.g. a line-pair phantom) to the CT image as afunction of spatial frequency (expressed as the number ofline pairs per unit length). It is common use to state thespatial frequency which corresponds to 10 % contrast in theCT image as the spatial resolution of the scanner. I.e.: thefiner the CT image detail still resolvable the higher thespatial resolution of the scanner and thus the spatialfrequency at 10 % modulation transfer [12].

For small animal CT scanners the so-called slant-slitmethod is often used to calculate the MTF of a scanner [18,19]. Here the phantom consists of an acrylic plate having arectangular slit covered by a very thin metal foil (e.g. 18 mof aluminium [18]). The cross-section of the foil is the slitwhich can be slightly tilted with respect to the horizontal orvertical axis of the image plane so that the line-profile of theslit in the image domain can be sampled at sub-pixel level.Then the MTF can be calculated as the Fouriertransformation of the line profile.

In practice, the spatial resolution of clinical scale CTscanners and small animal CT flat-panel scanners featuring aclinical-scale gantry is usually measured by scanning line-

pair phantoms, which provide sets of metal line-pairs (e.g.

tungsten bars) of various thicknesses at decreasing distancescast into plastic material. Here the smallest feature size (inmillimeters) that is still visible can easily be calculated bydividing 10 mm by the amount of distinguishable line-

pairs/cm times two.

The maximum theoretically achievable spatial resolutionaccording to the employed detectors intrinsic resolution andthe geometrical magnification of a particular small animalCT system can seldom be achieved [5]. Apart from themechanical stability (e.g. vibrations, etc.) of the scanner andthe influence of the reconstruction algorithm, the image blurintroduced by the finite size of the focal spot ( x f,FWHM) ofthe x-ray tube limits the spatial resolution of the real system(penumbral blurring is

b= ODD

SDDODDx f,FWHM=

ODD

SODx f,FWHM

[6]). To reduce the influence of the focal spot shape and theassociated penumbral blurring Popescu and co-workers [23]have modified and measured the focal spot shape and size ofthe customized clinical x-ray source employed in the flat-

panel CT used for small animal imaging [31]. Themeasurements of Popescu and co-workers [23] included aninvestigation of the focal spot shapes throughout the wholedetector plane. These authors have shown that smalldistortions of the focal spot introducing a negligible amount

of image blurring can still be observed even if the tube ismodified to match source and detector setup.

In order to achieve the spatial resolution of interest,fluxes of 104-105 photons have to traverse a particularregion of interest (ROI) per resolution unit at minimum [32]so a detector signal significantly larger than the statisticalnoise due to the quantum nature of the x-ray photons can begained. Higher photon fluxes allowing shorter exposurestimes (and thus better temporal resolution) require larger x-

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ray tube focal spot sizes but dictate a relaxation of therequirements of spatial resolution and geometricalmagnification of a small animal CT scanner system.

Low-Contrast Resolution

The highest spatial resolution achievable in small animalCT is of little use unless the scanner is able to resolve smalldifferences between structures and neighbouring tissueshaving a poor inherent contrast. In any CT system, contrastin the images is a consequence of the difference in theenergy-dependent linear absorption coefficient (E) (alsocalled attenuation coefficient) of different tissues [7].Thus, low-contrast detectability (LCD) can be defined as theability to detect fine variations in the (electron) density of anobject over the background [12].

The low-contrast resolution capabilities of a small animalCT scanner largely depend on the type of detector used: themore quantum efficient the detector is, the lower the amountof incident photons to generate a detector signal can be; andthe larger the detectors dynamic range is, the smaller thedifferences in photon flux still detectable and differentiablecan be. This implies that the low-contrast resolutioncapability of a scanner also depends on the x-ray doseapplied per projection (view) of a scan (see below), sinceresolving low-contrast structures actually means collectingdetector signals whose difference is significantly larger thanthe noise level present. Micro-CTs that are not specificallydesigned for small animal imaging usually have inherentlimitations in signal-to-noise ratio (SNR) performance

because of their small detector voxel size (to achieve highspatial resolution) and their low x-ray exposure level (caused

by the employed micro-focus x-ray tubes). Therefore, theirlow-contrast resolution is traded for high spatial resolution ofhigh-contrast structures such as bones [18]. Otherinvestigators trying to achieve high low-contrast detectabilityhave employed x-ray tubes featuring large foci [6, 33]. Inaddition, a careful selection of the x-ray energy appropriate

to the imaging task is necessary if low-contrast detectabilityshould be optimal.

Experimental evaluations of the low-contrast resolutionof small animal CT systems usually employ custom madecontrast phantoms. These may consist of water-filled acrylichollow cylinders with small cylindrical low-contrast insertswith densities close to the density of water [18]. They allowto either determine the minimum dose level at which thelow-contrast inserts can be discriminated from the surroun-ding water or used to quantify the low-contrast resolution

performance of a small animal CT at a fixed x-ray exposureby measuring its contrast-to-noise ratio (CNR):

CNRi =Si

Sb

i2 + b

2 ,

with S and being the mean CT number and the standarddeviation of the image pixel values in a specific ROI inHU,the subscripts stand for the inserts (i) and the background ( b)[19]. Ideally, the CNR in CT images is proportional to thesquare root of the applied dose [18, 34].

Contrast agents are used in small animal CT like inclinical CT to increase low-contrast resolution. The increasein CT number (CTNumber) caused by the contrast agentapplied can be approximated by

( )( )

CE

ENumberCT

water

mediumcontrast

([9, 35]).

Here C is the concentration of the contrast agent in the

tissue (in mg/ml) and ( = cm2g[ ]) is the

attenuation coefficient normalized for density.

Image Reconstruction

In small animal CT the same image reconstructiontechniques as in clinical-scale CT are applied. CT systemsusing fan-beam geometries apply fan-beam reconstructions

based on filtered back-projection algorithms originating fromthe early days of clinical CT. The recently introduced flat-

panel based micro- and mini-CTs employ variants of theFeldkamp-David-Kress (FDK) algorithm [14] for cone-beamreconstruction. Both reconstruction techniques have to beadapted to the specific small animal scanner geometry. Anoverview over the underlying principles of imagereconstruction can be found in [9, 35].

Contrast Media

Non-ionic extravascular water-soluble contrast agents, astypically used for clinical CT examinations, rapidly clearfrom the blood within minutes after intravenous injection.The first-pass of the contrast media with the highest vascularcontrast is only retained for a few heart beats. This timeframe is too short for image acquisition with most micro-CT

scanners.Contrast media that remain in the vasculature for longer

are called blood-pool contrast agents. In order to achieveblood-pool effects the size of contrast media molecules canbe increased to larger sizes than that of capillary fenestra-tions. Furthermore, phagocytosis by the reticuloendothelialsystem should be avoided as long as possible through thechemical design of the contrast media. Several differentagents have shown potential as blood-pool agent for CTimaging [36].

A water-soluble macromolecular agent (dysprosium-DTPA-dextran) was reported to provide blood-pool contrastenhancement for up to 45 minutes [37]. Liposomal encapsu-lation of iohexol resulted in a blood-pool effect for up to 3hours after injection in rabbits [38]. In rats, iodine-containingmicelles caused enhancement in the blood, liver and spleenfor more than 3 hours [39].

An iodinated triglyceride emulsion (ITG-LE) packed intothe lipophilic core of a synthetic chylomicron remnant has

been described in [40]. It can be used to image liverparenchyma because it is internalized by the hepatocytes,whereas liver tumor cells do have less functional lipoproteinreceptors and therefore only show low enhancement. The

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liver sequestration of ITG-LE from the blood-pool can bedelayed by adding polyethylene glycol-modified phospho-lipids (ITG-PEG) to the polyiodinated triglyceride emulsionresulting in a longer blood-pool time. The combined applica-tion of both formulations of ITG enhances the healthy liver

parenchyma as well as the intrahepatical vessels at the sametime. This reduces the risk to misinterpret vessels as tumorlesions. False-positive results may occur with hepatocyte-specific contrast media alone, because liver vessels as well

as liver tumors are non-enhancing and can therefore beconfounded [41].

Dual-phase contrast enhancement can be achieved withcontrast media that first circulate in the blood and aresubsequently cleared via the hepatobiliary pathway herebyenhancing the liver and the spleen for hours [36].

Also nanoparticles may deal as CT contrast agents. Anexample is nanosized bismuth-sulfide with a polymercoating. These nanoparticles were described to have x-rayabsorption characteristics that are five times better than thatof iodine-containing compounds, which reduces the risk ofviscosity problems [42]. The vascular half-life is longer thanthat of iodine-based contrast media and their efficacy/safety

profile is comparable or better. Imaging of tracheobronchiallymph nodes with nanoparticles in CT has already beenshown in dogs [43].

X-Ray Dose

Dose applied during image acquisition is of specialconcern in small animal CT, especially if follow-upexaminations of the same animal are necessary. When allother hardware and acquisition parameters are kept constanta good low-contrast spatial resolution requires theapplication of high x-ray doses: in order to be able todifferentiate neighboring tissues with similar (electron)densities with a significant SNR, a particular voxel of theimaged animal being the smallest unit of resolution needs tointeract with a certain minimum amount of photons. The

minimum amount of required photons is constant regardlessof the voxel size. However, dose is a function of the energydeposited through photon absorption and energy loss viaCompton scattering interactions per volume, therefore doseneeds to increase with the power of four with the resolutionelement if photon noise is supposed to be constant. In otherwords: An increase in resolution from 1.0 mm to 0.5 mmhas to be paid for by a 16-fold increase in exposure if weexpect to see low-contrast lesions of 0.5 mm diameter withthe same clarity as lesions of 1.0 mm diameter in the

previous case [34].

Small animal CT needs to be performed with higherresolution than human CT scanning because anatomicalstructures within laboratory animals are smaller than in

human bodies. In order to achieve spatial and low-contrastresolution (and thus noise levels) comparable to human CTimaging higher x-ray doses are physically needed. Forexample, mice imaging with an isotropic resolution of 135m requires a dose of 250 mGy supposing an ideal smallanimal CT-scanner is used for the examination [44], whichamounts to a 10-30 fold increase in dose compared to a CTexamination of a human patient [45, 46]. An isotropicresolution of 135 m results in a three-fold increase in

spatial resolution compared to clinical CT, whereas thetheoretically expected x-ray dose would increase by a factorof 34=27 (cf. [34]). This would mean that imaging mice withthe same image quality at an isotropic resolution of 35 mwould require a dose of 2500 Gy. From these studies one canconclude that fundamental laws of physics impose the limitsof dose in small animal imaging and thereby the limits of in-vivosmall animal imaging [44]. Depending on the diagnosticdemand, however, x-ray dose can be decreased while

maintaining high spatial resolution if cut-backs in otherimage quality parameters such as image noise and soft-tissuecontrast can be accepted. Much higher image noise isacceptable e.g. when high-contrast structures such as bones,lung or contrast media filled vessels are the focus of theexamination. So far only whole body doses have beendiscussed, which means that depending on the experimentalfocus higher doses could be applied to organs or parts of theanimal which are less dose-sensitive.

Dose limitations for imaging living animals are usuallystated in terms of the lethal dose (LD): LD50/30 is the whole

body radiation dose which would kill 50% of an exposedpopulation within 30 days of exposure. Previous studiesindicated that the lethal dose of mice ranges between 5-7.6

Gy [44, 47, 48], depending on the strain [49], age and otherfactors. This range certainly sets the upper limit for thewhole body x-ray exposure of a single examination of aliving animal. Fortunately, small animal CT examinationsusually do not even get close to this limit.

Since the main advantage of in-vivosmall animal CT liesin the fact that follow-up studies can be performed theeffects of sublethal dosages become relevant. The sum of the

biological effects of many sublethal doses is a complexfunction that is influenced by many factors. While the lethaldose might give an absolute limit, exposure to much smallerx-ray dosages can already have biological effects, whichmight interfere with experimental results. Several low-doseradiation effects are described including effects on tumor

growth, hematopoiesis [45, 50] and bone growth [2]. Iffollow-up studies are performed over a long timeframe a

possible contribution of radiation effects to the observedresults needs to be assessed: a group of animals that is onlyimaged at the end of the time series could act as a control incomparison to the scanned animals [51] or the contralaterallimb not having been exposed to radiation could be taken asa non-irradiated control [2].

Single organ doses for mice imaging were calculatedusing a Monte Carlo simulation. Here the typical whole

body dose was 80 mGy (at 80 kV tube voltage) to 160 mGy(at 50 kV tube voltage). Furthermore, a strong dependanceon the tube voltage as well as on the position within theanimal was found [50]. To allow the computation of x-ray

exposure doses from air kerma measurements dose coeffi-cients for standardized mice, typical scanner geometries andx-ray spectra were calculated using Monte Carlo simulations[45].

NEW DEVELOPMENTS

Flat-Panel Based Small Animal CT

Flat-panel detector based micro- and mini-CT scannersfor small animal imaging have been described in bench-top

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systems [18] as well as in gantry-based systems [10, 12, 20,52]. As discussed in Detector Technology, flat-panel detec-tors have several benefits over detectors that were used formicro-CT earlier.

Large flat-panel detectors are ideally suited for gantry-based small animal CT imaging systems with a geometrythat allows a reasonably high resolution with a large gantry

bore and a wide z-coverage so that bigger animals such as

rabbits and piglets [10, 12] can be scanned. The data of evenbigger animals can be acquired in one rotation.

A decrease in spatial resolution allows a relaxation indose performance and therefore a very fast acquisition ofsingle projections and therefore the whole volume in shortscan times. Fast, continuous scanning becomes possible.Acquiring the same volume in short succession is a

prerequisite for perfusion imaging of fast contrast mediadynamics. Slow contrast media dynamics such as kidney orliver uptake and bladder or gall bladder filling could already

be assessed with slower micro-CT scanners. However, fastcontrast media dynamics such as brain, tumor and scar

perfusion is of much higher interest. Using flat-paneldetector CT scanners, exemplary perfusion imaging of fast

dynamics for (whole) small animals has been shown [52].This technique could provide valuable physiological andfunctional information of many small animal models anddiseases. Several post-processing methods could be appliedto the contrast media time course data. First experimentswere promising: the time course of contrast media could betraced, and parameters such as Mean Transit Time, BloodVolume and Blood Flow of experimental tumors could becalculated [52].

Fast and continuous scanning could also make retros-pective cardiac and respiratory gating possible. For this goal,the short exposure time of flat-panel CT that virtually freezesthe heart motion on the level of projections is especiallyadvantageous. Earlier micro-CT scanners were limited with

regard to motion gating by the long exposure times perprojection.

Compared to prospective triggering, retrospective gatingcould decrease the scan times significantly. This could makethe use of standard, iodinated contrast media possible andmay open perspectives to lung or cardiac perfusion studies.

Future Perspectives

For a good overview of future developments of micro-CTwith focus on in-vivo small animal imaging please read thereview from Ritman [53]. If big technical and engineeringchallenges could be solved, techniques such as x-ray fluore-scence, x-ray diffraction or x-ray scatter imaging would be

big steps for CT imaging. Their potential for the improve-

ment of soft-tissue contrast and contrast media sensitivity areenormous. However, a long time might still pass until thesetechnologies will reach operational readiness.

Another method possibly available in the near future isK-edge subtraction imaging. It is based on synchrotrongenerated narrow bandwidth rays that allow imaging ofcertain chemical elements by imaging the volume twice atdifferent photon energies (once with a frequency just aboveand once with a frequency just below the characteristic K-

edge of the selected element). When both data sets aresubtracted, density differences only remain for the selectedelements. However, the availability of synchrotron radiationis currently limited to expensive synchrotron generators.Furthermore, multi-source CT imaging could decrease thescan time by radiating the object from several directions atonce.

APPLICATIONS

The following chapter gives an overview of theapplications of micro- and mini-CT in small animals. In thecompiled studies herein various scanner designs with a rangeof scan parameters, scan times and spatial resolution areused. Most studies are proof of principle studies. Never-theless, the conclusions drawn from one study hold true forother studies in case that similar mini-/micro-CT scannerswith comparable characteristics were used.

Lung Imaging

In mice, the total lung volume is 1.3 cm3. The thorax ofmice is about 2 cm in diameter. Proportionally, a mini-CTscanner that provides an insight on the mice lung structuresas in a human lung should offer a resolution in the order of

75 m [54]. In vivo lung imaging is challenged by therespiratory and cardiac motions and its compensated imagingwas described by Badea and co-workers [33].

Beside the lung parenchyma, the larger thoracic struc-tures such as heart, oesophagus, trachea, bronchi and largevessels can be imaged in non-enhanced scans (Fig. (2)).

Lung Tumors

The detection of experimentally induced lung tumors ispossible in mice using micro-CT. This was shown in a study,in which urethane induced lung tumours were imaged inliving mice. The scans were performed without motioncompensation and with an isotropic resolution of 35 m [55].Only few, small tumors that were found in histologic slices

were missed in the micro-CT scans.

Fig. (2).In-vivothoracic imaging of a mouse (A) and a rat (B) fromflat-panel based CT. Data acquisition time was 5 s. However,

motion artefacts close to the ribs and diaphragm are pronounced.

Bronchi (white arrow) and vessels (black arrow) can be

distinguished from the lung parenchyma.

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Follow-up studies of lung tumor size changes over timecould be performed [56-58]. Size distortions of the lungtumors due to motion artefacts should be taken into accountwhen micro-CT without appropriate motion-compensationalgorithms is used for absolute measurement studies.

The minimal sizes of lung tumors that were found inseveral studies were: 6.6 mm3 in volume and 0.85 mm indiameter in the free lung parenchyma and 1.4 mm in

diameter perihilar [58]. In other studies the size of lungnodules regardless of their localization was reported to be0.63 mm3 minimum in volume [57] and 0.5 mm [56],respectively, smaller than 0.2 mm in diameter [55]. In thiscontext, using a ventilator and prospective respiratory gatingthe accuracy of tumor volume determination could beimproved [57].

In the reviewed studies almost all tumors that wereidentified by micro-CT were confirmed histologically, butvice versa not all tumors that were found in histology werereliably identified in micro-CT even if they had the samesize as tumors that were found. Juxtahilar tumors were mosteasily to detect, while the tumors next to the thoracic walland the big vessels were most difficult to visualize, which is

most likely due to motion artefacts that lead to a smearing ofthe thoracic walls. Furthermore, it is difficult to differentiatetumors from hilar structures and from vessels. Similarly,

because the lung tissue in the reviewed studies was normalapart from the induced tumors, it remains to be proven byother studies, how accurate micro-CT can differentiate

between scar tissue, microgranulomas, hyperplastic lympho-cellular or other non-malignant lesions in lung tissue. Theuse of macroscopic differential features to distinguish bet-ween tissues as known from human thoracic imaging iscertainly limited by the resolution. Blood-pool contrastmedia proved to be valuable to improve differentiation oftumors and vessels [58].

Even pleural effusions occurring as the consequence oftumor disease could be successfully imaged with micro-CT[58].

General Lung Parenchyma Changes (Emphysema andFibrosis)

Using micro-CT the assessment of generalized lungparenchyma changes is possible as shown in followingexperiment: For example lung emphysema was induced in

C57BL/6J mice by intratrachael instillation of pancreaticelastase. The mice were scanned using a micro-CT withoutmotion compensation and an isotropic resolution of 35 m.On CT images mice treated with the highest dosage ofelastase showed the highest amount of pixels with low HUvalues within their lungs. Furthermore, the lung volume waslarger in this group. In conclusion, lung emphysema can bedetected by micro-CT indirectly by the lower density of thelung and by an increase in lung volume [59]. Unfortunately,despite the sufficiently up-scaled spatial resolution a directassessment of the altered emphysematous lung structure as inhuman high resolution (HR)CT-imaging is not possible [54].This is mainly due to motion artefacts and resolutionlimitations.

Vice versa, also the assessment of lung fibrosis in mice ispossible by analysis of CT numbers. Here the diseaseprogress was shown to be associated with an increase in CTnumbers on a clinical CT-scanner [60]. Using mini-CT itwas shown that not only the general density could beassessed but also structural changes of bleomycin-inducedfibrosis such as ground glass opacity, septal thickening,fibrotic strands and secondary dilation of the bronchialsystem (bronchiectasis) [52].

Bone Imaging

Similar to its applications in very high resolution ex-vivobone imaging micro-CT can provide information about bone

Fig. (3).3Dreconstruction (A) and 2D slice (B) of a rat knee and its adjacent bones scanned with a bench-top micro-CT shown in Fig. (1) B.The rat was rotated between the detector and the x-ray source while only the limb was radiated. High spatial resolution of 6 m 3could be

achieved because a single limb fits into a very small scan field-of-view (image courtesy Prof. W. Kalender, Institute of Medical Physics,

Erlangen, Germany).

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structure and its changes in vivo in small animals, however,with stricter resolution constraints.

Microstructural Bone Imaging

Using micro-CT a morphological bone analysis of itsmicro-architecture can be performed in vivo providingseveral descriptors such as bone volume ratio, bone surfaceratio, trabecular thickness, trabecula separation, trabecularnumber, connectivity density and structural model index [61-

63].

CT density values as a questionable surrogate marker forbone status could also be acquired using micro-CT. Dual-energy scans as used in clinical CT setups to assess the bonedensity have not been reported using micro-CT yet.

Small animal bone scans can be performed on singlelimbs. Customized jigs can assure immobilization. Due tothe low diameter of those limbs resolution optimized scannergeometries (e.g. 15 m isotropic [2]) with only a small scanFOV can be applied (Fig. (3)).

Macroscopic Skeletal Imaging

Micro-CT can be used to image the whole macroscopic

animal skeleton. Applications are studies on the skeletaldevelopment [64] or studies in which structural changes inthe skeleton anatomy occur as a consequence of disease. Forthese applications, the demand for large scan volumes as

provided by clinical-scale CT or flat-panel based mini-CTscanners is higher than the need for high resolution. Clinicalscale scanners as well as flat-panel based mini-CT systemscan provide an FOV that is big enough to image a wholemouse or rat during one acquisition. Using appropriate post

processing the whole skeleton can be displayed in highspatial resolution three dimensionally [20, 52].

Bone Metastasis Screening

Micro-CT can be used to screen with a high sensitivityand specificity for osteolytic bone metastases in wholemouse skeleton [58]. Visualization of the soft tissue of themetastases that extend beyond the bone surfaces, however, isa challenge for the soft-tissue contrast resolution capabilityof the micro-CT. Its detection has not yet been describedusing micro-CT.

Angiographic Imaging

In-vitro imaging using micro-CT resulted in astonishingresults of the vasculature and microvasculature of smallanimals. Often, these images were generated after casting thevasculature with contrast media such as silicon-basedcompound (Microfil MV-122, Flow Tech, Carver, MA,USA) [65].

For in-vivo imaging various contrast media have beendiscussed in Section. Due to the long scan times of micro-CTmainly blood-pool contrast agents have been used to imagethe vasculature (Fig. (4)) [37-39]. The detection of the aorta,

pulmonary trunk, inferior vena cava, renal artery and vein inmice and rats [35, 66] has been described.

Standard iodinated contrast media typically used forclinical examinations clear too fast from the vessels to beused in micro-CT scanners. However, mini-CT-scanning insuch short times became recently feasible.In-vivoimaging ofall central vessels of mice and even that of dilatedsubcutaneous vessels that drain implanted tumors as well assmaller vessels inside the tumor were demonstrated usingflat-panel mini-CT in combination with standard iodinatedcontrast media (Fig. (5)). Characteristics of tumor vesselarchitecture could be assessed [52] and their changes traced

Fig. (4).Blood-pool contrast media enhanced scan of a rat in an early (A) and later (B) phase as scanned by a micro-CT shown in Fig. ( 1) D.

The scanners spatial resolution is optimized for the dimensions of rats and mice. Due to the scan time of 180 s blood-pool contrast mediawere used (image courtesy Prof. W. Kalender, Institute of Medical Physics, Erlangen, Germany).

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in follow-up studies [20].

Assessment of Slow Contrast Agent Kinetics

To assess the slow time-resolved contrast mediaenhancement in micro-CT long circulating contrast mediawere used [39, 66]. Other authors assumed that even withiodinated contrast agents quantification of several aspects ofrenal function as shown in clinical CT imaging may besuccessful [1, 35]. The renal cortex, medulla, pelvicalycealsystem and ureter could be differentiated in contrast

enhanced studies. A hydronephrosis model has been charac-terized using this method successfully [35].

Through the use of faster flat-panel CT scanners, alsofast dynamic contrast-enhanced applications may become

broadly available [52] (Fig. (6)).

Cardiac and Thoracic Motion-Compensated Imaging

Thoracic imaging of living animals is challenged by heartand lung movements. These movements lead to imaging

artefacts, which reduce the accuracy of thoracic imaging ofsmall animals and make reliable measurements e.g. of lungtumors more problematic [57, 67] (Fig. (2)).

To compensate for such movements prospective as wellas retrospective gating and triggering methods can beapplied. In prospective methods, the imaging process ismodulated according to the movements of the small animal.Therefore only one animal can be imaged during one

acquisition. Prospective methods take longer scan times thanretrospective methods but are usually more dose efficientthan retrospective methods.

Regardless of whether prospective or retrospectivemethods are used, both methods add up projections acquiredduring defined phases of cardiac or respiratory cycles toform one complete dataset for CT image reconstruction.Therefore, position reproducibility between several breathingas well as cardiac cycles is crucial. However, this reprodu-cibility is not better than 100 m in small animals [68],limiting the maximum spatial resolution of thoracic imagingof free-breathing small animals to a resolution in this order.

So far predominantly prospective methods have beenimplemented in small animal motion compensated CT

imaging: To physically stop breathing excursions smallanimals can be intubated and ventilated. Breathing can bearrested for a short time period. Due to the long scan timesof micro-CT complete datasets can not be acquired duringone breath hold.

However, the short breathing movement arrests thatoccur during the physiological ventilation plateau phase can

Fig. (6).Perfusion CT using standard iodinated contrast media of a

bone metastasis model in rats [76] as scanned with flat-panel CT.The contralateral (black arrow) as well as the partially destroyed

ipsilateral tibia (white arrow) can be seen. Rotation time was 6 s,half-scan reconstruction was performed every 3 s using 180 data.

Injection of contrast media was initiated together with scanning. At

0 s (A) no contrast media was detected in the selected ROI, which

covered the bone metastasis, while at 24 s (B) pronounced rim andvessel enhancement can be found. The time-course of averaged CT

numbers within the ROI is given in (C) (unpublished data by Dr.

Tobias Baeuerle, German Cancer Research Center (DKFZ),

Heidelberg, Germany).

Fig. (5). Mini-CT angiography using standard, iodinated contrast

media of a mouse scanned using a flat-panel detector CT as in Fig.

(1) F [12]. 100 l contrast media were injected three seconds beforescan; data acquisition took three seconds, resulting in a late

arterial/venous phase. A volume rendering, in which the slice

positions of the axial slices (A-D) are indicated, is given in (A).

Furthermore, a coronal slice is given in (E), showing the kidneys

(long black arrow) and suprarenal glands (long white arrow). Thecontrast enhancement in the caval vein (short black arrow) is more

pronounced than in the aorta (short white arrow). Small structures

such as the splenal vein (white dashed arrow), the hepatic venous

confluence (white arrow head) and the suprarenal vessels (blackdashed arrow) can be detected. Regardless of cardiac motion the

right and left ventricle (black asterisk) can be distinguished clearly.

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be used to acquire at least a part (e.g. one projection) of theimaging dataset [4, 33, 57]. The trigger point can beoptimized so that data acquisition is performed in phases ofthe respiratory cycle that show least motion [69].

In another study, the animal could breathe freely. Itsbreathing excursions were tracked by a pneumatic cushionthat was rigidly attached to the animal. In a definable

breathing excursion position projections were acquired.

After each acquisition the animal was rotated to the nextprojection view [70]. Acquiring various datasets at variousbreathing excursion points allowed four-dimensionalimaging of the thorax and lung.

With a variation of the study design multiple projectionsat one angle were acquired and retrospectively sorted accor-ding to various breathing cycles. Intrinsic image informationwas used for the projection sorting, thereby not requiring anadditional device to collect gating information [71].

Cardiac motion compensation can be implemented usinga prospective ECG triggering [4] combined with breathingtriggering in intubated and ventilated mice. Repeating atvarious trigger points resulted in 4D imaging of a rodentheart [67].

The benefits of lung motion compensation are obvious.Lung parenchyma, diaphragm and vessels were displayedsharper than in non-gated scans [4, 70]. Respiratory gatingmade the measurement of lung tumor sizes more precise

because smearing caused by lung motion artefacts could bereduced or avoided. Smearing results in inaccurate sizemeasurements [57]. ECG-gating further improved thedisplay of lung and cardiac anatomy [4].

Soft Tissue Imaging

While micro- and mini-CT provide high resolution ofhigh contrast structures such as bone and contrast mediafilled vessels, the soft-tissue contrast is relatively low. Asmentioned, the soft-tissue contrast is strongly disturbed by

image noise and is therefore strongly depending on theapplied radiation dose.

To this point it is uncertain, how suited micro- and mini-CT are to differentiate soft-tissue structures because thedifferentiation of soft tissue in mini- and micro-CT is usuallynot as good as in clinical scale CT scanners.

However, differentiation of fat from other soft tissuestructures is possible. Measurement of fat volume as well asfat distribution in a mouse disease model has been demons-trated [35]. Fat measurements open promising perspectivesin imaging diabetic animal models or in studying kachexia(e.g. during neoplastic disease). Most likely, if fat can bedifferentiated from surrounding soft tissue, determination of

muscle volume also becomes feasible and can be used tocharacterize musculoskeletal disorders, e.g. in transgenicmice with muscle dystrophia.

Organs in the peritoneal cavity can already be differen-tiated in non-enhanced scans. Liver, kidney, spleen, heart,lung, stomach, adrenal gland, gut and bladder can beidentified and their borders be delineated.

To enhance the contrast between organs or to different-tiate tumors and organs in the peritoneal cavity, standard,

iodinated contrast media have been injected in to the perito-neal cavity, facilitating visualization of their boundaries.

In contrast-enhanced flat-panel micro-CT scans also thelongitudinal assessment of urethral tumors in the bladder

became feasible [72], where tumors were identified asnegative shapes in the contrast media filled bladder.

Image Registration with Other Modalities

The diagnostic value of studies can be improved byfusing and matching the data from different imagingmodalities thus directly combining morphological andfunctional features: registration is the method to align themodalities in space, while fusion is a function that combinesall or parts of the information of the modalities into a newdataset.

Potential complementary imaging data for registrationwith micro-CT can derive from radionuclide, optical and MRimaging as well as histology.

In general, registration methods can be characterized asfollows: intrinsic (registration based on the image infor-mation itself) or extrinsic (registration based on additionalmarkers or frames), retrospective (after scanning) or pros-

pective (before scanning, e.g. through reference frames orscan beds), manual (user based) or automatic (softwarealgorithm), rigid (no spatial distortion is allowed to register

both datasets) or non-rigid (deformation is allowed).

Small animal multimodality registration can be done bycombining both imaging modalities in one scanner and

perform simultaneous or near simultaneous data acquisitions[16, 73]. The advantage of combining two modalities in onescanner system is that there is no need to move the animal

between acquisitions, meaning both scanners can use thesame coordinate system, resulting in very good registrationaccuracies. The main disadvantage of simultaneous imagingdevices is a certain loss of the flexible choice of scannergeometry.

Another method is the use of a scan bed or a framework,to which the animal can be fixated. The scan bed can then beused to transfer the animal from one modality to the nextinsuring accurate animal positioning. This method has beensuggested as the most feasible method in small animalimaging that provides a very good accuracy [74].

If multi-modality imaging acquisition is not conducted ina combined scanner or using fixed reference frames, retros-

pective methods such as software-based or manual methodsneed to be applied. Multiple software-based approaches forretrospective, automatic methods are also described [75].Registration accuracy was in the order of 1 mm. The usedsoftware approaches still needed significant user interaction.

Alternatively, marker based approaches could be chosen,here registration markers that are visible in both modalitieswere attached rigidly to the animal. So far radionuclideimage registration with micro-CT has been published mostly.Registration between 18F-fluoride PET bone images andmicro-CT of the rat skull [75], 18F-FDG PET of the rat andmicro-CT (Fig. (7)) [74, 75] as well as SPECT and micro-CT[16, 74] was described.

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ACKNOWLEDGEMENTS

We thank Willi Kalender and Michael Grasruck for their

support by providing image material.

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Fig. (7).Micro-PET and micro-CT co-registration. Structural information of a mouse from micro-CT (A) is registered to 18F-FDG micro-PET

data (B), resulting in a combination of both modalities (C). The combination image provides anatomic information on the tumor location

(white arrows) and its metabolism (from Meei-Ling Jan et al., Institute of Nuclear Energy Research, Longtan, Taiwan, 2005 IEEE).

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Received: October 27, 2006 Revised: December 13, 2006 Accepted: December 14, 2006

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