three-dimensional noninvasive imaging of the vasculature ... · three-dimensional noninvasive...

Three-dimensional noninvasive imaging of thevasculature in the mouse brain using a high

resolution photoacoustic scanner

Jan Laufer,1 Edward Zhang,1 Gennadij Raivich,1,2 and Paul Beard1,*1Department of Medical Physics and Bioengineering, University College London,

Gower Street, London WC1E 6BT, UK2Centre for Perinatal Brain Research, University College, Gower Street, London WC1E 6BT, UK

*Corresponding author: [email protected]

Received 6 October 2008; revised 13 January 2009; accepted 16 January 2009;posted 22 January 2009 (Doc. ID 102329); published 30 March 2009

The application of a novel photoacoustic imaging instrument based on a Fabry–Perot polymer film sen-sing interferometer to imaging the small animal brain is described. This approach provides a convenientbackward mode sensing configuration that offers the prospect of overcoming the limitations of existingpiezoelectric based detection schemes for small animal brain imaging. Noninvasive images of the vas-culature in the mouse brain were obtained at different wavelengths between 590 and 889nm, showingthat the cerebral vascular anatomy can be visualized with high contrast and spatial resolution to depthsup to 3:7mm. It is considered that the instrument has a role to play in characterizing small animal mod-els of human disease and injury processes such as stroke, epilepsy, and traumatic brain injury. © 2009Optical Society of America

OCIS codes: 170.5120, 170.0110, 170.3880.

1. Introduction

Small animal models are widely used as a researchplatform to study basic brain function and diseaseand injury processes in an effort to improve the diag-nosis and treatment of neurological conditions suchas stroke, traumatic brain injury, and epilepsy [1].There is a consequent need to characterize thesemodels not only by revealing anatomical changesbut also by eliciting precursive functional informa-tion via the physiological and metabolic status oftissue. Imaging techniques such as x-ray CT, MRI,SPET, and PET have all been used for small animalneuroimaging. However, these modalities either pro-vide solely anatomical information (CT) or functionalinformation with relatively low spatial resolution(SPET/PET), or where they do provide both tend tobe limited in the range of physiological parameters,

and thus the specificity of the functional information,they can provide [2,3]. Optical imaging techniquescan provide a wide variety of functional parametersby exploiting changes in absorption, scattering, orfluorescence [3,4]. For example, by using absorptionspectroscopy, it is possible to monitor changes in oxy-hemoglobin and deoxyheomoglobin and cytochromeoxidase, all of which are important markers of brainfunction [5]. Furthermore there is a wide range ofoptical contrast agents available that can be targetedto specific biomolecules to reveal processes at a cel-lular or molecular level [6]. However, strong scatter-ing in most tissues limits penetration depth andspatial resolution. Thus, techniques such as opticalcoherence tomography and confocal and multiphotonmicroscopy that rely on predominantly unscatteredor singly backscattered light to provide spatial infor-mation are unable to provide penetration depthsexceeding ∼1mm [4]. This limits small animal opti-cal brain imaging to visualizing the surface of the cor-tex, and even then, to achieve the highest possible

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resolution, it is usually necessary to expose the cor-tex by removing the scalp and thinning or resectingthe skull [3]. Diffuse optical tomography can achievepenetration depths of several centimeters using nearinfrared (NIR) light but has found limited applica-tion in small animal brain imaging on account ofits relatively low spatial resolution [7].Photoacoustic imaging offers an alternative

approach that provides the spectroscopic based spe-cificity and functional capability of optical imagingmethods but without the attendant penetrationdepth/spatial resolutions limitations. The techniqueinvolves irradiating the tissue surface with nanose-cond NIR or visible laser pulses. Absorption of thelaser energy and subsequent thermoelastic expan-sion produces broadband pulses of ultrasound thatpropagate to the surface where they are detectedat different spatial locations. By measuring the timeof arrival of the acoustic pulses at each location, andknowing the speed of sound in tissue, the acousticsignals can be backprojected to reconstruct a 3D im-age of the absorbed optical energy distribution. Thus,while image contrast is based on tissue optical prop-erties, principally absorption, spatial resolution isdefined by the physics of ultrasound propagation.The fundamental spatial resolution limit is set bythe frequency dependent acoustic attenuation of softtissues and scales with penetration depth: for centi-meter depths, submillimeter spatial resolution ispossible, decreasing to sub-100 μm for millimeter pe-netration depths and sub-10 μm spatial resolution fordepths of a few hundred micrometers. For depthsbeyond 1mm or so, photoacoustic imaging thereforeprovides significantly higher spatial resolution thanpurely optical imaging methods—a consequence ofthe much lower scattering that ultrasound waves un-dergo in soft tissues compared to photons. Although,hard tissues such as bone strongly scatter ultra-sound, the skull of mice or rats, unlike larger animalssuch as primates or indeed humans, is usually suffi-ciently thin that this does not excessively compro-mise image quality.The most important source of endogenous photo-

acoustic image contrast is hemoglobin due to itsstrong optical absorption at visible and NIR wave-lengths. This makes photoacoustic imaging particu-larly well suited to visualizing the vasculature. Inaddition, quantitative functional information canbe potentially obtained via measurements of abso-lute blood oxygen saturation and flow—the formerby obtaining images at multiple wavelengths andapplying a spectroscopic analysis to exploit the dif-ferences in the spectral characteristics of oxyhemo-globin and deoxyhemoglobin [8], the latter byextracting the acoustic Doppler shift encoded ontophotoacoustic signals emitted by moving red bloodcells [9,10]. In common with other optical imagingtechniques there is the further prospect of using tar-geted exogenous contrast agents to image specificcellular, intracellular, and molecular receptors[11,12]. This ability to noninvasively image vascular

anatomy and function, as well as provide molecularinformation over the spatial scales of the mouse orrat brain, suggests photoacoustic imaging could bea powerful investigative tool for studying small ani-mal cerebral hemodynamics, oxygenation, and cellu-lar function. For example, the ability to localizedeficiencies in tissue perfusion and oxygenation sta-tus, as well as any ensuing neuronal cell damage,would be of major benefit in the study of strokeand other conditions where hypoxic ischemic damageoccurs.

Several experimental studies have now demon-strated the ability of photoacoustic imaging to non-invasively image the structure and function of thevasculature in the rat or mouse brain with a spatialresolution in the tens to hundreds of micrometersrange [13–20]. The most commonly used implemen-tation employs a circular scanning detection geo-metry. In this approach, a piezoelectric ultrasounddetector or array of detectors is mechanicallyscanned around the circumference of the head ofthe animal while the surface is illuminated withpulses of laser light. This method provides the neces-sarily large angular detection aperture for high qual-ity image reconstruction and allows the excitationlight to be delivered without it being obscured bythe detector. The disadvantage, however, lies in cou-pling the ultrasound to the detector: either the ani-mal has to be fitted with a breathing mask [13,17,19]and completely immersed in water or some otheracoustic couplant or only the top of the headis acoustically coupled via a thin walled helmet toa water bath containing the ultrasound receiver[14–16]. Neither is particularly convenient in termsof animal handling and achieving the correct orien-tation of the head with respect to the detection sys-tem. It would be significantly more convenient tomap the photoacoustic signals over the illuminatedtop surface of the head: the so-called backward or re-flection mode of detection. The challenge then is todeliver the laser light without it being obstructedby the detectors. One approach is to deliver the lightcircumferentially around a mechanically scanned fo-cused piezoelectric detector, so-called photoacousticmicroscopy [21]. Although high quality images ofthe superficial cortical vasculature in a mouse havebeen obtained with this method [22], the highest lat-eral spatial resolution is achieved only at a limitedrange of depths, those that coincide with the nondi-vergent region of the transducer focus. It also re-mains to be seen whether this sequential scanningapproach can be parallelized to overcome the in-evitable limited acquisition speed associated withmechanical scanning.

In this paper we describe the use of a novel highresolution photoacoustic scanner for small animalbrain imaging that can overcome these limitations.The distinguishing feature of the system is that itemploys an optical ultrasound imaging sensor basedon a Fabry–Perot polymer film interferometer (FPI)that is transparent. The sensor can therefore be

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placed on the surface of the animal head, theexcitation laser pulses transmitted through it, andthe photoacoustic signals mapped in 2D over the ir-radiated region. In this way it provides a truly back-ward mode planar detection configuration that doesnot require the animal to be wholly immersed inacoustic couplant and avoids the handling difficultiesassociated with the use of a helmet that cylindricaldetection schemes require. Since a tomographic im-age reconstruction approach is employed it does notsuffer from the depth limited spatial resolution lim-itations of using a fixed focus transducer, and thesensor read-out scheme can be readily parallelizedto achieve high frame rates.The design, operating principles, and performance

of the scanner are described in Ref. [23]. This studyshowed that the system could provide high fidelity3D images of tissue mimicking phantoms. In thispaper, the emphasis is on applying the instrumentto imaging the mouse brain. Section 2 provides abrief overview of the operation of the scanner andits performance; Section 3 presents images of thevascular anatomy in the mouse brain obtained at avariety of excitation wavelengths.

2. Photoacoustic Imaging System

A. Principles of Operation

A full description of the imaging system is describedin Ref. [23]. Briefly, it comprises a tunable pulsedexcitation laser system that is used to generatethe photoacoustic waves and a 2D planar backwardmode optical ultrasound mapping system for record-ing them over the surface of the skin. From the mea-sured photoacoustic signals, a 3D image is thenreconstructed.The excitation laser system is a fiber coupled type I

optical parametric oscillator (OPO) (GWU VisIR)pumped by the 355nm frequency tripled outputof a Q-switched Nd:YAG laser (Spectra-Physics,Quanta Ray LAB170). This system is capable of pro-viding 8ns optical pulses over the wavelength rangefrom 410 to 2100nm with end-of-fiber pulse energiesin the 12–36mJ range (depending on wavelength) ata pulse repetition frequency of 10Hz.The key component of the ultrasoundmapping sys-

tem is a sensor head that is placed in light contactwith the surface of the mouse head over the regionindicated by the dotted line in the photograph inFig. 1. A drop of water is sandwiched between thesensor head and the skin surface in order to acous-tically couple the photoacoustic signals to the sensor.The critical requirement in this respect is that thereis no air between the sensor and the skin, as thiswould act as a strong acoustic reflector and impedethe transmission of the acoustic waves. The sensorhead comprises a 10mm thick wedged poly(methylmethacrylate) substrate with a polymer film FPIformed on its lower side. The FPI is fabricated byvacuum depositing a multilayer structure consistingof a 38 μm thick Parylene C polymer film spacer

sandwiched between a pair of dichroic dielectric mir-rors. These mirrors are highly reflective (>95%) be-tween 1500 and 1650nm but highly transmissivebetween 600 and 1200nm. Excitation laser pulsesin the latter wavelength range can therefore betransmitted through the sensor head into the under-lying tissue. Absorption of the laser energy producesphotoacoustic pulses that travel back to the sensorhead where they modulate the optical thickness ofthe FPI and hence its reflectivity.

The sensor is read out by raster scanning a focusedinterrogation laser beam (at 1550nm where the FPImirrors are highly reflective) point-by-point over thesurface of the sensor using an x − y galvanometerbased scanner. At each point of the scan, the sensoris optimally biased by tuning the interrogation laserwavelength to the point of maximum slope on theinterferometer transfer function, the relationshipbetween reflected optical power and phase. Underthese conditions, a small acoustically inducedmodulation of the optical thickness of the FPI canbe regarded as being linearly converted, via the in-terferometer transfer function, to a corresponding re-flected optical power modulation. This time-varyingoptical power modulation, which represents thephotoacoustic waveform, is then detected by an In-Gas photodiode-transimpedance amplifier unit andrecorded using a digital storage oscilloscope (DSO).Once a waveform has been acquired at a particularspatial point on the sensor it is stored within the on-board memory of the DSO. The interrogation laser

Fig. 1. (Color online) Experimental arrangement. The FP sensorhead is placed on the surface of the mouse head over the regionindicated on the photograph. Nanosecond excitation laser pulsesemitted by a tunable OPO laser system are directed on to the sen-sor head, and transmitted through it into the underlying tissuethereby exciting acoustic waves. A second laser emitting at1550nm provides a focused interrogation laser beam that is rasterscanned over the surface of the sensor in order to map the distri-bution of the photoacoustic waves arriving at the sensor head.From the 2D distribution of the photoacoustic waves, a 3D imageis then reconstructed.


beam is then moved to the next scan point andthe process repeated. At the end of the 2D scan,the entire set of waveforms are downloaded fromthe DSO to a PC and input to a k-space backpropa-gation algorithm [24] in order to reconstruct a 3D im-age of the initial pressure distribution poðx; y; zÞ: thephotoacoustic image.

B. System Performance

The maximum area over which the interrogation la-ser beam can be scanned over the sensor is defined bya circle of diameter 50mm. The 1=e2 diameter of thelaser beam at its focus is 64 μm and represents, to afirst approximation, the dimensions of the acousti-cally sensitive region of the sensor. The minimumscan step size (limited by the 12 bit resolution ofthe PC digital-to-analog conversion card thatcontrols the optical scanner) is 10 μm. The peaknoise-equivalent pressure of the sensor is approxi-mately 0:2kPa over a measurement bandwidth of20MHz—comparable to that of a 1mm diameterpolyvinylidine fluoride piezoelectric receiver. The fre-quency response of the sensor is defined primarily bythe acoustic thickness of the polymer film spacer [23]and is characterized by a smooth roll-off from a max-imum value at 100kHz (the −3dB cutoff frequency ofthe ac-coupled photodiode) to zero at 100MHz witha −3dB bandwidth of 22:2MHz. A plot of thefrequency response of the sensor can be found inRef. [23].The point-to-point acquisition time of the system

described in Ref. [23] was approximately 1 s per scanstep. This was limited by the sequential nature of thewaveform downloading process in which each photo-acoustic waveform was downloaded to the PC via aGPIB interface immediately following its acquisi-tion. In the current study, a different acquisitionapproach was adopted whereby the complete set ofwaveforms acquired over the scan were stored withinthe onboard memory of the DSO and downloaded tothe PC in a single step. Although this enabled thepoint-to-point acquisition time of the scanner itselfto be reduced to 10ms [25], the average acquisitiontime achieved in this study was limited by the 10Hzpulse repetition frequency of the laser system to100ms per scan step.To obtain a measure of the spatial resolution that

the system can provide, measurements of the instru-ment line spread function (LSF) have been made andare reported in Ref. [23]. The lateral LSF depends onthe overall detection aperture (the scan area), the ef-fective acoustic element size, spatial sampling inter-val, and the bandwidth of the sensor. Furthermore itdepends on the angular aperture subtended by thedetection aperture and therefore the axial and lat-eral position of the source. Taking into account therelatively small scan areas (∼1 cm2) used in thisstudy, the data presented in Ref. [23] suggest a lat-eral LSF in the range from 50 to 100 μm for depthsup to 5:5mm at the center of the detection aperture.Unlike the lateral LSF, the vertical LSF is, to a first

approximation, spatially invariant and dependentlargely on the bandwidth, which is defined by thethickness of the polymer film spacer. Again, usingthe measurements obtained in Ref. [23], the verticalLSF is estimated at approximately 50 μm. In bothcases the estimated LSFs are measures of the instru-ment response alone. They do not account for fre-quency dependent acoustic attenuation and shouldtherefore be regarded as only approximate indicatorsof the spatial resolution that will be in achieved whenimaging tissue.

3. Noninvasive Imaging of the Mouse Brain

Ex vivo images of the cerebral vasculature of 3 monthold mice (CL57/BL6) were obtained. The time be-tween sacrifice and image acquisition was approxi-mately 90 min. The images were acquired at avariety of excitation wavelengths in order to identifywhich might be optimal in terms of image contrastand penetration depth. Three experiments wereundertaken. In the first, an excitation wavelengthof 590nm was used because it is strongly absorbedby hemoglobin and so offered the prospect of achiev-ing high contrast. However it was expected to providea relatively small penetration depth due to its ab-sorption by blood in the capillary bed and thus islikely to be best suited to visualizing superficial ves-sels. For this reason, the use of longer wavelengthswas explored. In the second experiment two wave-lengths, 637 and 800nm, were used. Both wave-lengths are attenuated less than 590nm, andcombining the two meant that a higher total fluencecould be obtained for increased penetration depth.Finally, in the third experiment an excitation wave-length of 889nm was used.

All the images were obtained noninvasivelythrough the intact skin and skull, the only prepara-tion being that the hair around the region of interestwas removed using VEET hair removal cream. In allcases, no signal averaging was used and the incidentlaser fluence was below the safe maximum permissi-ble exposure (MPE) for skin [26]. To aid the visuali-zation of deeply lying features, a depth dependentscaling factor was used to compensate for opticaland acoustic attenuation in a manner analogous totime-gain compensation used in diagnostic medicalultrasound imaging. The scaling factor was obtainedempirically from the 3D reconstructed data poðx; y; zÞset by extracting the maximum value of po from eachx − y plane in the data set as a function of z to givepoðmÞðzÞ. An exponential function of the form e−αz

was then fitted to poðmÞðzÞ by varying α, and it is thisfunction that was used to compensate for attenuationin the reconstructed image. Apart from this and theuse of an interpolation algorithm, no other imageprocessing was employed. The reconstructed imagesin this paper are presented as maximum intensityprojections or volume rendered representations.The latter are also available as animated volumerendered images and can be viewed online asMedia 1, Media 2, Media 3 or the University College


London website [27]. These animations provide themost compelling demonstrations of the 3D imagingcapability of the system.

A. λ¼590nm

In the first experiment, the excitation laser wave-length was set to 590nm. The incident fluence was6:8mJ=cm2 and thus significantly below the MPEof 20mJ=cm2 for skin at this wavelength. The sensorinterrogation beam was scanned over an area of di-mensions 9mm × 10mm in steps of 100 μm, and asingle photoacoustic waveform was acquired at eachstep. The total number of waveform acquisitions wastherefore 9000, and the time taken to acquire thesewas 15 min. A 3D image data set of dimensions9mm× 10mm× 3:8mm was then reconstructedfrom the detected signals using the k-space backpro-pagation algorithm referred to in Subsection 2.A.Lateral (x − y) and vertical (y − z) maximum intensityprojections (MIP) were computed from the 3D recon-structed data set and are shown in Fig. 2. From theseimages, the principal cerebral cortical vessels, the sa-gittal sinus, transverse sinus, and other superficialveins can be identified. At 590nm, penetration depthis relatively low due to strong optical absorption byblood in the capillary bed. As a consequence only ves-sels on the surface of the cortex are evident. Notethat the y − z MIP shows only the vasculature onthe cortical surface around the side of the head,not vessels extending vertically into the underlyingwhite matter. This can be seen more clearly in theanimated volume rendered representation availableonline (Media 1) from which a sequence of four stillsat different angles is shown in Fig. 3.

B. λ¼637 and 800nm

A second experiment was undertaken using the samemouse. This time, however, the OPO signal and idlerbeams at 637 and 800nm, respectively, were com-bined to provide a single excitation beam at thesetwo wavelengths. The incident fluence was thereforehigher than in the previous example at 10mJ=cm2

but still below theMPE. The spatial scan parametersand acquisition time were the same as the previousexperiment. Figure 4 shows the lateral and verticalMIPs. Figure 5 (Media 2) shows four stills from theanimated volume rendered representation availableonline. The superficial cerebral vessels are againevident on the x − y MIP. In addition, absorptionby the retina has produced features correspondingto the location of the eyes. However, unlike the imageobtained at 590nm (Fig. 3), the penetration depth issignificantly greater due to the reduced optical ab-sorption of blood at 637 and 800nm and the higherfluence. The most apparent features, which arenot visible in the image obtained at 590nm, arethe sinus rectus and inferior sagittal sinus. The lat-ter can be seen to extend deep into the brain to adepth of approximately 3:1mm Again this is mostapparent in the online animated volume renderedrepresentation.

C. λ¼889 nm

A third experiment was undertaken using a differentmouse and a wavelength of 889nm. To ensure thatthe whole of the cortical vasculature could be imaged,the photoacoustic signals were recorded over a largerarea (12mm × 14mm) than in the previous experi-ments. This required reducing the step size to

Fig. 2. Photoacoustic image of the vasculature in themouse brainobtained using an excitation wavelength of 590nm. Left,schematic of superficial cerebral vascular anatomy: A, superior sa-gittal sinus; B, transverse sinus; C, inferior cerebral vein. Topright, x − ymaximum intensity projection (MIP) of 3D photoacous-tic image; bottom right, y − z MIP.

Fig. 3. (Color online) Volume rendered representation of imagesof the mouse brain vasculature at different viewing angles ob-tained using an excitation wavelength of 590nm (Media 1). Thisimage was reconstructed from the data set used to form the MIPsshown in Fig. 2.


150 μm so as not to exceed the maximum array sizethat the acquisition software could handle. The x − y,x − z, and y − zMIPs are shown in Fig. 6 (Media 3). In

common with the images shown in Figs. 2–5, theprincipal cerebral cortical vessels and the sagittalsinus and transverse sinus can be seen, albeit withreduced contrast. The sinus rectus and inferiorsagittal sinus are also visible, but the latter can nowbe seen extending to a greater depth (∼3:7mm) thanobserved in Fig. 4. In addition, the branching of theinferior sagittal sinus at its distal end can be seen.This latter feature is most evident in the online ani-mated volume rendered representation, which alsoreveals an earlier branching that is not discernibleon the MIPs in Fig. 6 (Media 3).

4. Conclusion

The use of a novel photoacoustic imaging instrumentfor small animal brain imaging has been demon-strated. It has been shown that 3D images of thecerebral vascular anatomy can be obtained noninva-sively through the intact skin and scalp to depths ofseveral millimeters with high contrast and spatialresolution. It is noted that, as well as revealing

Fig. 4. Photoacoustic image of vasculature in mouse brain ob-tained using excitation wavelengths of 637nm and 800nm simul-taneously. (top) x-y maximum intensity projection (MIP) of 3Dphotoacoustic image and (bottom) y − z MIP. A, superior sagittalsinus B, transverse sinus; C—inferior cerebral vein; E, eyes; S,sinus rectus; D, inferior sagittal sinus.

Fig. 5. (Color online) Volume rendered representation of imagesof mouse brain vasculature at different viewing angles obtainedusing excitation wavelengths of 637nm and 800nm simulta-neously (Media 2). This image was reconstructed from the dataset used to form the MIP images shown in Fig. 4.

Fig. 6. Photoacoustic image of the vasculature in mouse brain ob-tained using an excitation wavelength of 889nm (Media 3). (top)x − y maximum intensity projection (MIP), (middle) x − z MIP, and(bottom) y − z MIP. A, superior sagittal sinus; B, transverse sinus;E, eyes; S, sinus rectus; D, inferior sagittal sinus; D0, lateralbranching at the distal end of the inferior sagittal sinus.


the superficial cortical vasculature, deeply lyingfeatures such as the inferior sagittal sinus and thebranching at its distal end can be visualized.Such features have not generally been observed inprevious photoacoustic imaging studies of the smallanimal brain [13–19].There are several advantages of the system over

other photoacoustic imaging instruments for ima-ging the small animal brain. These include the easeand convenience with which the instrument can beinterfaced to the animal. This is a consequence ofthe transparent nature of the sensor head, whichallows it to be placed on the surface of the head ofthe animal and the photoacoustic signals recordedover a planar surface coincident with the excitationlight. It thus obviates the need for fully immersingthe animal in acoustic couplant or the use of complexholders for the head, which schemes based on cylind-rical detection geometries often require. The sensoralso provides a level of acoustic performance that canbe difficult to achieve with piezoelectric based detec-tion methods, particularly in relation to the spatialsampling of the incident acoustic field as discussedin Ref. [23]. Because the sensor is optically addressedusing a focused laser beam, the incident acousticfield can be spatially sampled with significantlyhigher (potentially optical diffraction limited) resolu-tion than can be achieved with piezoelectric recei-vers. It is this along with the broadband uniformacoustic frequency response of the sensor and itshigh detection sensitivity that are responsible forthe high contrast and spatial fidelity of the recon-structed images. Although the penetration depth isnot currently as great as some photoacoustic imagingsystems [20], there is, as noted in Ref. [23], signifi-cant potential to increase it by increasing detectionsensitivity. In addition there is the prospect of in-creasing the sensor bandwidth and optimizing thespatial sampling parameters to obtain higher spatialresolution. While the current acquisition speed maybe acceptable for studying slow long term structuralchanges such as the evolution of the vasculature of atumor over a period of days it would be insufficientfor studying more rapid events such as the hemody-namic response to functional activation in the brain,which may occur on a subsecond time scale. However,there is significant scope to increase acquisitionspeed, potentially obtaining real time image acquisi-tion rates, especially if 2D images are acceptable,through the use of higher repetition rate laser sys-tems [28,29] or parallelizing the sensor read-outscheme using a photodetector array as described inRef. [30]. Finally, the optical nature of the sensorlends itself to multimodal imaging studies. For exam-ple, the system can readily be made compatible withMRI as the sensor head is fabricated from nonmag-netic materials and its transparent nature wouldallow fluorescence and bioluminescence emissionsto be imaged through it.In summary, this type of instrument may provide a

useful tool that is complementary to other preclinical

imaging techniques for characterizing small animalmodels of brain injury and disease processes,particularly those that require studying vascularanatomy and function such as stroke, epilepsy, andtraumatic brain injury.

This work was supported by the UK Engineeringand Physical Sciences Research Council.

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