Video rate optoacoustic tomography ofmouse kidney perfusion

Andreas Buehler, Eva Herzog, Daniel Razansky, and Vasilis Ntziachristos*Institute for Biological and Medical Imaging, Technische Universität München und Helmholtz

Zentrum München, Ingoldstädter Landstrasse 1, 85764 Neuherberg, Germany*Corresponding author: v.ntziachristos@tum.de

Received April 26, 2010; revised June 16, 2010; accepted June 24, 2010;posted June 30, 2010 (Doc. ID 127592); published July 15, 2010

Optoacoustic tomography can visualize optical contrast in tissues while capitalizing on the advantages of ultra-sound, such as high spatial resolution and fast imaging capabilities. We report a novel multispectral optoacoustictomography system for deep tissue small animal imaging. The previously undocumented capacity of whole-bodyoptoacoustic tomography at a video rate has been demonstrated by visualizing mouse kidney perfusion usingIndocyanine Green in vivo. © 2010 Optical Society of AmericaOCIS codes: 170.3880, 170.5120, 170.6280, 170.6920.

Macroscopic optical imaging plays an increasingly impor-tant role in biomedical studies, as it can noninvasivelyexamine structural, physiological, and molecular tissuefeatures in vivo [1]. In response, small animal opticalimaging has emerged as an important tool in basic re-search, drug discovery, and clinical translation due tothe widespread use of animal models of human diseasein understanding systemic responses. An emerging ap-proach for high-resolution volumetric imaging of opticalcontrast in tissues is optoacoustic (photoacoustic) tomo-graphy, which can visualize tissue chromophores withultrasonic resolution [2]. Using multispectral methods,imaging of extrinsic optical agents, such as fluoro-chromes and fluorescent proteins, has also been shownto be possible [3,4]. Optoacoustic imaging correspond-ingly offers important advantages in small animalimaging, including the use of nonionizing radiation, theversatile sensing of chromophoric molecules for probingcellular and subcellular function, ultrasonic spatial reso-lution, and, in analogy to ultrasound imaging, real-timeoperation.Nevertheless, the majority of current small animal

optoacoustic systems considered for small animal imag-ing utilize scanning arrangements that are typically in-appropriate for real-time whole-body imaging. Some ofthe existing systems scan a single transducer [3–6]around the sample, yielding long acquisition times (min-utes to hours). Control of animal physiological param-eters, motion, and anesthesia during these extendedmeasurement periods can present a significant challengefor obtaining high-quality images. The use of ultrasoundtransducer arrays from clinical systems [7,8] has alsobeen considered for increasing the imaging speed; how-ever, commercial arrays, optimized for clinical imaging,do not have the geometric arrangement and broadbandfrequency characteristics that would make them appro-priate for optoacoustic small animal imaging. Dedicatedsmall animal optoacoustic scanners, which utilize a highnumber of detection elements, have also been developed[9,10]. One of these approaches utilizes a 512-elementring-shaped focused ultrasound transducer array, custommade for small animal imaging, and was successfully ap-plied for anatomical imaging of cerebral blood vessels aswell as for imaging of cortical hemodynamics [10,11].

However, none of these approaches has been, so far,shown suitable for whole-body small animal visualizationin real time.

We report herein on the development of a high-performance optoacoustic imaging system capable offast volumetric multispectral optoacoustic tomography(MSOT). The system optimizes light energy delivery thatavoids the need for data averaging and offers parallel de-tection of optoacoustic signals geometrically arrangedaround the area of maximal light deposition for datacollection. Using this implementation, we demonstratethe previously undocumented ability of optoacousticimaging to tomographically visualize in real-time three-dimensional (3D) structures in the animal body,demonstrated by monitoring kidney perfusion after theintravenous injection of Indocyanine Green (ICG).

The system developed is based on a custom-made64-element curved transducer array Imasonic SaS, Voray,France) covering a solid angle of 172° around the imagedobject, as shown in Fig. 1. The individual elements aremanufactured using piezocomposite technology with acentral frequency of 5 MHz, bandwidth of more than 50%,and sensitivity of ∼18 μV=Pa and are shaped to create acylindrical focus at 40 mm. According to the ultrasonic

Fig. 1. (Color online) (a) Schematic representation of theMSOT system. A curved array of wideband and cylindrically fo-cused ultrasound transducers enables parallel data acquisition.Optical fibers are used to homogeneously illuminate the object.A special animal holder with a transparent plastic membrane isused for animal positioning. (b) Picture of the mouse during ascan, showing the position of the mouse and illuminationwith respect to the array of focused wideband ultrasoundtransducers.

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0146-9592/10/142475-03$15.00/0 © 2010 Optical Society of America

diffraction limit and focal width at the central frequency,the effective spatial resolution of the array is estimated at150 μm (in plane) and 800 μm (elevational) [12].Excitation light originates from a tunable (680–

950 nm) optical parametric oscillator laser (Phocus,Opotek Incorporated, Carlsbad, California), delivering<10 ns duration pulses with repetition frequency of10 Hz. The beam is guided into a silica fused-end fiberbundle (Model PowerLightGuide, CeramOptec GmbH,Bonn, Germany) consisting of 630 fibers partitioned into10 arms. The arms are positioned 3 cm from the animaland create a ring-shaped illumination pattern of ∼7 mmwidth upon the surface of the animal, coinciding with theultrasound detection plane [Fig. 1(b)]. The beam is suffi-ciently broadened to keep the laser pulse fluence on thesurface of the imaged objects under 20 mJ=cm2. Thedetected signals are digitized at 60 megasamples=s fre-quency by eight multichannel analog-to-digital conver-ters (Model PXI5105, National Instruments, Austin,Texas) having a noise floor of∼3:8 nV=

ffiffiffiffiffiffi

Hzp

. To facilitatein vivo measurements, the system contains an imagingchamber that holds animals inside a water-impenetrablemembrane that avoids animal contact with water whileproviding a wide tomographic view of ∼180°. A linearstage (NRT150, Thorlabs GmbH, Karlsfeld, Germany) al-lows linear translation of the animal holder in the axial zdirection over a 150 mm range with a minimal step size of2 μm for acquisition of 3D data sets. The data acquisitionis synchronized so that the signals are acquired onlywhen the stage comes to a complete rest. Image recon-struction is based on the interpolated-matrix-modelinversion method [13], where a 200 × 200 grid is usedwith a pixel size of 110 μm.For imaging, a female, eight-week-old CD1 mouse was

anesthetized with a mixture of ketamine and xylazine,and hair was removed with a hair removing lotion. Theanimal was then placed in the supine position in the imag-ing chamber and 300 nmol of ICG solution was injectedinto its tail vein. Following single-pulse measurements ofkidney perfusion at a constant position, a multispectraldataset of the kidney was also acquired at nine wave-lengths (750, 770, 790, 810, 830, 850, 870, 890, and910 nm). In contrast to the video rate perfusion imagingthat was performed at 10 Hz without averaging, here 50

averages were acquired per wavelength and the per-wavelength energy variations were further correctedby powermeter readings (Model FieldMaxII-TOP, Coher-ent GmbH, Dieburg, Germany) for quantitative multi-spectral reconstructions. After imaging, the animal waseuthanized, cryosliced, and photographed for confirma-tion of the noninvasive optoacoustic images.

Figure 2 depicts results from the real-time imagingstudies and corresponding confirmatory ex vivo ima-ging. Figure 2(a) shows a stack of subsequent single-wavelength (800 nm)optoacoustic slices obtained in vivofrom the pelvis and the kidney region. Figure 2(b) showsthe slice that was imaged during the injection of ICG,while the photograph of a cryoslice through the mousekidneys at a similar height is presented in Fig. 2(c). Thereis accurate congruence between features in the noninva-sive optoacoustic imaging and the anatomical photographof the cryosliced mouse. Beside the kidneys, the spine,surroundingmuscle, and vasculature are also clearly iden-tified on the optoacoustic image. Differences observed inthe images can be explained by slight organ movementsbetween in vivo and ex vivo imaging and differences inthe exact height that the in vivo images and the cryoslicedimages were acquired from.

Figure 3 shows a time series of images acquired in realtime after ICG injection as (a) single-wavelength imagesand as (b) an overlay of the contrast enhancement ob-served (color) superimposed on the image before injec-tion of ICG. Frames at different time points were selectedto show the dynamics of ICG kidney perfusion. About16 s after injection of ICG, there is the first appearance

Fig. 2. (Color online) (a) Stack of representative slices of a3D dataset of the pelvis and kidney region of a female eight-week-old CD1 mouse. (b) Cross-sectional optoacoustic imagebefore injection of the slice in which the kinetics of ICG wasmonitored. (c) Photograph of a cross-sectional cryoslice at aheight corresponding to that shown in (b): 1, vena cava; 2, por-tal vein; 3, kidneys; 4, spinal cord; 5, backbone muscles; and 6,spleen.

Fig. 3. (Color online) (a) Cross-sectional optoacoustic imagesat different time points of the kidneys of a female CD1 mouseilluminated at 800 nm after having injected 300 nmol of ICG.(b) Superposition of the absorption difference and the single-wavelength image before injection.

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of ICG in the imaging plane, showing as a localized activ-ity in the center of the image. In the subsequent slices,a gradual spread of the dye into the kidneys can be ob-served, while an increasing number of vascular struc-tures enhance their contrast (optical absorption) dueto the presence of ICG.Finally, Fig. 4 shows a MSOT image, which detects the

ICG presence based on its unique spectral signature in-stead of resolving it as an absorption difference over abackground image. ICG has a characteristic absorptionspectrum [14] in the 750–900 nm spectral window thatdiffers significantly from the spectra of blood [4] andoverall background tissue [Fig. 4(b)], and it was resolvedherein with per-pixel multispectral decomposition[3,4]. Considering an approximate total blood volumeof about 2 ml in the mouse, a single resolution-limited voxel (i.e., 150 μm × 150 μm × 800 μm ¼ 18 nl)contained 300 nmol × 18 nl=2 ml ¼ 2:7 pmol of ICG.In conclusion, we have shown a novel optoacoustic

imaging scanner capable of volumetric real-time animalimaging in vivo. The scanner provides 3D datasets byanimal translation along its vertical axis while simulta-neously acquiring optoacoustic data at multiple wave-lengths for visualization of functional and molecularcontrast with MSOT. The system has proven to be cap-able of cross-sectional animal imaging with real-time dataacquisition and the reconstructed 2D images delivered at10 Hz rates; i.e., each slice shown in Figs. 2 and 3 is col-lected and processed within less than 100 ms. Imagingperformance was demonstrated by resolving mouse kid-ney anatomy and perfusion using ICG, an exogenously

introduced contrast agent. Images produced were con-gruent with corresponding photographs of the mouseanatomy, based on cryoslices obtained from the sameanimal imaged with MSOT. Spectral unmixing furtherdemonstrated the ability of the system to visualize exter-nally administered contrast based on its unique spec-tral signature without using background measurementsmade prior to the probe administration.

V. Ntziachristos acknowledges support from the Eur-opean Research Council through an Advanced Investiga-tor Award. D. Razansky acknowledges support from theGerman Research Foundation (DFG) under researchgrant RA 1848/1.

References

1. V. Ntziachristos, J. Ripoll, L. H. V. Wang, and R. Weissleder,Nat. Biotech. 23, 313 (2005).

2. L. V. Wang, Med. Phys. 35, 5758 (2008).3. D. Razansky, M. Distel, C. Vinegoni, R. Ma, N. Perrimon,

R. W. Koster, and V. Ntziachristos, Nat. Photon. 3, 412(2009).

4. D. Razansky, C. Vinegoni, and V. Ntziachristos, Opt. Lett.32, 2891 (2007).

5. R. Ma, A. Taruttis, V. Ntziachristos, and D. Razansky, Opt.Express 17, 21414 (2009).

6. X. D. Wang, Y. J. Pang, G. Ku, X. Y. Xie, G. Stoica, andL. H. V. Wang, Nat. Biotech. 21, 803 (2003).

7. S. A. Ermilov, T. Khamapirad, A. Conjusteau,M. H. Leonard, R. Lacewell, K. Mehta, T. Miller, andA. A. Oraevsky, J. Biomed. Opt. 14, 024007 (2009).

8. B. Z. Yin, D. Xing, Y. Wang, Y. G. Zeng, Y. G. Zeng, Y. Tan,and Q. Chen, Phys. Med. Biol. 49, 1339 (2004).

9. H. P. Brecht, R. Su, M. Fronheiser, S. A. Ermilov,A. Conjusteau, and A. A. Oraevsky, J. Biomed. Opt. 14,064007 (2009).

10. J. Gamelin, A. Maurudis, A. Aguirre, F. Huang, P. Y. Guo,L. V. Wang, and Q. Zhu, Opt. Express 17, 10489 (2009).

11. C. Li, A. Aguirre, J. Gamelin, A. Maurudis, Q. Zhu, andL. V. Wang, J. Biomed. Opt. Lett. 15, 010509 (2010).

12. B. A. J. Angelsen,Ultrasound Imaging (Emantec AS, 2000).13. A. Rosenthal, D. Razansky, and V. Ntziachristos, IEEE

Trans. Med. Imaging 28, 1997 (2009).14. M. L. J. Landsman, G. Kwant, G. A. Mook, and W. G. Zijlstra,

J. Appl. Physiol. 40, 575 (1976).

Fig. 4. (Color online) (a) Superposition of a single wavelengthimage (890 nm) and the unmixed component corresponding tothe ICG signal. (b) Molar extinction coefficient of ICG [14],oxygenized and deoxygenized hemoglobin [4].

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