three-dimensional endomicroscopy using optical coherence ... · 2007 nature publishing group...

© 2007 Nature Publishing Group

Three-dimensional endomicroscopyusing optical coherence tomography

DESMOND C. ADLER1, YU CHEN1, ROBERT HUBER1,2, JOSEPH SCHMITT3, JAMES CONNOLLY4

AND JAMES G. FUJIMOTO1*1Department of Electrical Engineering and Computer Science and Research Laboratory of Electronics, Massachusetts Institute of Technology, Cambridge,

Massachusetts 02139, USA2Lehrstuhl fur BioMolekulare Optik, Fakultat fur Physik, Ludwig-Maximilians-Universitat Munchen, 80538 Munchen, Germany3Lightlab Imaging, Westford, Massachusetts 01886, USA4Department of Pathology, Beth Israel Deaconess Medical Center, Boston, Massachusetts 02215, USA

*e-mail: [email protected]

Published online: 25 November 2007; doi:10.1038/nphoton.2007.228

Optical coherence tomography enables micrometre-scale, subsurface imaging of biological tissue by measuring the magnitude andecho time delay of backscattered light. Endoscopic optical coherence tomography imaging inside the body can be performed usingfibre-optic probes. To perform three-dimensional optical coherence tomography endomicroscopy with ultrahigh volumetricresolution, however, requires extremely high imaging speeds. Here we report advances in optical coherence tomographytechnology using a Fourier-domain mode-locked frequency-swept laser as the light source. The laser, with a 160-nm tuning rangeat a wavelength of 1,315 nm, can produce images with axial resolutions of 5– 7 mm. In vivo three-dimensional optical coherencetomography endomicroscopy is demonstrated at speeds of 100,000 axial lines per second and 50 frames per second. This enablesvirtual manipulation of tissue geometry, speckle reduction, synthesis of en face views similar to endoscopic images, generation ofcross-sectional images with arbitrary orientation, and quantitative measurements of morphology. This technology can be scaledto even higher speeds and will open up three-dimensional optical-coherence-tomography endomicroscopy to a wide range ofmedical applications.

Optical coherence tomography (OCT) performs micrometre-scale,cross-sectional imaging by measuring the echo time delay ofbackscattered light1. It is a fibre-optic technique, enablinginternal body imaging with fibre-optic probes2. Two-dimensionalendoscopic OCT has been demonstrated with axial resolutions of2.4–3.7 mm using femtosecond lasers3,4. Three-dimensionalendoscopic OCT (3D-OCT) promises to provide more completecharacterization of tissue structure, but requires extremely highimaging speeds and data processing rates.

Time-domain OCT (ref. 5) uses low-coherence reflectometry,with imaging speeds of 4,000–8,000 axial lines per second(4–8 kHz) achieved using reference-arm grating phase delaylines6–10. Optical frequency-domain reflectometry (OFDR) usingfrequency-swept lasers is well established for measuringreflections in photonic devices11–15. The use of swept lasers forOCT was demonstrated a decade ago16–18; however, it was onlyrecently recognized that Fourier-domain detection using sweptlasers or spectrometers dramatically improves sensitivity andimaging speed19–21.

Frequency-swept lasers are a key technology for high-speed‘swept source’ OCT imaging. The sweep repetition rate, tuningrange, and instantaneous linewidth of the laser determine theimaging speed, axial resolution and ranging depth of the OCTsystem, respectively. Swept lasers using a tunable filter composedof a diffraction grating and a rotating polygon mirror haveachieved sweep rates of 15.7–115 kHz over tuning ranges of74–80 nm (refs 22,23). Using a similar laser with a tuning range

of 111–125 nm, endoscopic 3D-OCT imaging has beendemonstrated in the porcine oesophagus and coronary artery atsweep rates of 10 kHz and 54 kHz, respectively24,25.

Conventional swept lasers suffer fundamental performancelimitations as sweep rate is increased, because lasing mustrepeatedly build up from spontaneous emission during thesweep. This degrades the power, tuning range, and instantaneouslinewidth26. Recently, Fourier-domain mode-locked (FDML)lasers were demonstrated to achieve high performance at highsweep rates27–29. Fourier-domain mode-locked lasers use a longfibre cavity with a semiconductor amplifier and fibreFabry–Perot tunable filter (FFP-TF). The filter is tunedsynchronously to the cavity round trip time in a quasi-stationaryoperating regime, giving sweep rates up to 370 kHz over tuningranges of 100 nm (ref. 28).

This paper describes 3D-OCTendomicroscopy using an FDMLlaser operating at a sweep rate of 100 kHz with a 160-nm tuningrange. Axial resolutions of 5–7 mm are achieved, representing thehighest endoscopic speed and resolution obtained so far with aswept-source OCT imaging system. The system uses opticalclocking to provide real-time image capture at 50 frames persecond with 2,000 axial lines per frame. In vivo imaging isdemonstrated in the rabbit colon using a spiral-scanning fibreendoscope probe. Over an 8.8-mm length of colon, 1.1 gigavoxelswith a 3D resolution of 9 mm � 20 mm � 7 mm (1,260 mm3) areacquired in 17.7 s. Epithelial structures linked to neoplasticchanges are detected. A variety of visualization techniques are

ARTICLES

nature photonics | VOL 1 | DECEMBER 2007 | www.nature.com/naturephotonics 709

mailto:[email protected]

http://www.nature.com/doifinder/10.1038/nphoton.2007.228

www.nature.com/naturephotonics


demonstrated, including en face surface views, generation ofcross-sectional images with arbitrary orientation, and quantitativefeature measurement using image co-registration. Three-dimensional endoscopic OCT promises to enable high-resolutionendomicroscopy techniques for visualizing tissue pathologies.

RESULTS

FOURIER-DOMAIN MODE-LOCKED LASER

One key requirement for swept source 3D-OCT endomicroscopyis a laser that can provide a high sweep rate simultaneouslywith a large tuning range, high output power and narrowinstantaneous linewidth. Conventional frequency-swept laserssuffer a fundamental trade-off between sweep rate andperformance, limiting their utility above several tens ofkilohertz22,24–26. Fourier-domain mode-locked lasers remove thesefundamental limitations to sweep rate27–29. In FDML lasers, thefilter tuning period is matched to the optical roundtrip time ofthe laser cavity, enabling a quasi-stationary operating regimewhere each wavelength propagating in the cavity perceives thefilter to be stationary. Fourier-domain mode-locked lasers enableOCT imaging at superior speeds with axial resolutions andimaging depths not possible with conventional swept lasers. Asperformance does not degrade as the sweep rate is increased,FDML technology is inherently scalable to higher speeds.

For this study, an FDML laser was developed for high-speedimaging at the highest axial resolution achieved so far with afrequency-swept laser. Figure 1a shows a schematic diagram ofthe laser. A piezoelectrically actuated FFP-TF (LambdaQuest)with a free spectral range of 170 nm and a finesse of 500 is usedas a tunable narrowband filter. A semiconductor optical amplifier

(SOA) (Covega) at 1,310 nm provides the gain. Two lengths ofsingle-mode optical fibre (Corning), each 2 km long, form thelaser cavity. A second SOA externally amplifies the output. Thisdesign contains no free-space components, resulting in goodoperational stability.

The FFP-TF is driven sinusoidally, resulting in an alternatingseries of low-to-high and high-to-low frequency sweeps. Athigh sweep rates, the high-to-low sweep exhibits increasednoise. Therefore the FDML laser uses a buffered geometry,where two time-delayed copies of the low-to-high frequencysweep are extracted from the cavity28. The SOA is modulatedoff during the high-to-low sweep. This design enables the sweeprate to be multiplied and unidirectional sweeps created withoutsacrificing duty cycle.

Figure 1b shows the time-averaged spectrum of the laser outputmeasured with an optical spectrum analyser. The tuning range is160 nm, with a full-width at half-maximum (FWHM) of 97 nmaround a centre wavelength of �1,315 nm. The broad tuningrange is achieved primarily by the use of newly developed,polarized, broadband SOAs with high saturation output powers.The increased gain bandwidth of the intracavity SOA directlyimproves the FDML tuning range. The high-power booster SOAenables a decrease in the cavity output coupling ratio, raisingintracavity power levels and further extending the tuning range.The average output power is 35 mW and the sweep rate is100 kHz. Axial resolution and imaging depth are criticalparameters for OCT imaging. Figure 1c shows OCT point spreadfunctions measured with a Michelson interferometer at differentimaging depths, with the vertical axis showing system sensitivity.At a depth of 0.5 mm in air, the sensitivity is 98 dB, anddecreases by 6.5 dB at a depth of 2 mm, reflecting the low noise

FFP-TF

2 km2 km

ISO

ISO

SOA

80%

70%

ISOSOA

ISO

To O

CT

1,250 1,300 1,350 1,4000

0.2

0.4

0.6

0.8

1.0

Wavelength (nm)

Ampl

itude

(a.u

.)

40

60

80

100

Sens

itivi

ty (d

B)

Imaging depth (mm)

440 460 480 500 520 5400

0.2

0.4

0.6

0.8

1.0

Imaging depth (μm)Am

plitu

de (a

.u.)

0 1 2 3 4 5 6 7 8

Figure 1 FDML laser schematic and performance. a, Laser schematic. ISO: isolator. b, Time-averaged spectrum of laser output. c, Logarithmic OCT point spread

functions at various imaging depths. d, Linear point spread function at an imaging depth of 490 mm, demonstrating an axial resolution of 7.1 mm (FWHM) in air or

5.1 mm in tissue.

ARTICLES

nature photonics | VOL 1 | DECEMBER 2007 | www.nature.com/naturephotonics710



and narrow instantaneous linewidth of the laser. Fourier-domainmode-locking achieves a large imaging depth without requiringFourier-transform complex-conjugate image removal30–34. Theenlarged linear point spread function in Fig. 1d demonstratesthat the laser achieves an axial resolution of 7.1 mm (FWHM) inair or 5.1 mm in tissue.

OCT IMAGING ENGINE

To be practical in a clinical setting, a 3D-OCT system must providecontinuous real-time data capture and display with dense spatialsampling. Extremely large data sets must be digitized, transferredto computer memory, processed, and displayed in milliseconds.The OCT imaging engine developed for this study operates atsustained speeds compatible with FDML lasers by using anoptical frequency clock (OFC) integrated with a200 megasamples per second (MS s21) 12-bit analog-to-digitalconverter (ADC).

The formation of OCT images from frequency-domain datarequires the interference signal to be evenly spaced in opticalfrequency n before Fourier transformation35,36. As the frequencysweeps of both conventional and FDML lasers are not linear intime, it is necessary to correct for this effect. One commonapproach is to acquire a reference interference fringesimultaneously with the OCT signal. The n evolution of thereference signal can be analysed and used to later correct the n

spacing of the OCT signal22,35. This has the disadvantages ofdoubling the required data acquisition (DAQ) bandwidth becausetwo signals must be captured, requiring signal post-processing tocorrect the OCT data and limiting real-time operation.

The OCT imaging engine developed here uses an opticaltriggering technique to automatically correct for nonlinear n

spacing during data acquisition without storing a separatereference signal. Figure 2 shows a schematic of the system and aphotograph of the fibre probe tip. Of the laser output, 5% isrouted to an asymmetric Mach–Zehnder interferometer (MZI)that produces interference fringes with zero crossings evenlyspaced in n. This concept is shown in the left inset of Fig. 2. The

MZI fringes are detected by a dual-balanced photoreceiver andthe zero crossings are identified by an analogue voltagecomparator in the clock generator, creating a digital pulse trainthat functions as an OFC. The OFC is synchronized to the startof each laser sweep by a trigger signal from the FDML laser.

A dual-balanced Michelson interferometer consisting of a pairof optical circulators, and a 50/50 fibre-optic splitter is used togenerate the OCT interference fringes. Dual-balanced detectioncancels excess noise, reducing the dynamic range requirements ofthe DAQ system and improving sensitivity. The same frequencysweep from the laser generates both the MZI and OCT fringes, sothe spacing of the OFC pulses corresponds to evenly spacedn intervals in the OCT signal. The OFC triggers a 12-bit,200 MS s21, circularly buffered ADC that samples the OCTfringes. As the resulting signal is evenly spaced in n, only onesignal needs to be digitized and stored, and no post-processing isrequired for n correction.

After digitization, the OCT signal is continuously streamed tothe computer RAM over a PCI-X link at an average of46 Mbyte s21 and a peak of 150 Mbyte s21. Hammingwindowing and a fast Fourier transform are then performed tosynthesize the image. The axial resolution of the total systemdegrades from �5 mm to �7 mm in tissue, primarily because ofbroadening from the Hamming window and residual phase jitterin the OFC. Each frame is interpolated into polar coordinatesand displayed as a radial image in real time at .20 frames persecond. During 3D acquisition, sustained acquisition rates of100,000 axial lines per second are achieved while maintainingreal-time display.

3D-OCT IMAGING STUDIES

In vivo volumetric 3D-OCT data sets of the rabbit colon wereacquired with dense spatial sampling to demonstrate the abilityto detect microscopic, clinically relevant structures in thegastrointestinal (GI) tract. A spiral-scanning fibre probe with adiameter of 0.83 mm and a focal spot size of 9 mm was used forintraluminal imaging. An 8.8-mm segment of colon was imaged

FDML laser

DisplayFFTRAMPCI-X

200 MS s–1

ADC

OCT

OFC

Computer system

P

MZI

P

Ampl

itude

Time

Δν Δν

95/5C PIU

1 2

SampleRM

P P

C1 2

3

3

Probe

+ DA – + DA –

95/5

Clockgenerator

MZI input

TRIG

50/5

0

Figure 2 3D-OCT set-up. Volumetric data is acquired at 100,000 axial lines per second and 50 frames per second using a rotary fibre-optic endoscope probe

(C: circulator; RM: reference mirror; MZI: Mach–Zehnder interferometer; P: photodetector; DA: differential amplifier; TRIG: sweep trigger input; OCT: OCT signal input;

RAM: random access memory; FFT: fast Fourier transform; 95/5: 95%/5% coupler; 50/50: 50%/50% coupler. ‘1’, ‘2’ and ‘3’ refer to the port number of the

circulator.) Left inset: principle of OFC generation using the MZI output. Zero crossings are unevenly spaced in time, but evenly spaced in optical frequency n.

Right inset: photograph of the distal tip of the fibre probe, with the scale in mm.

ARTICLES




in 17.7 s, with 3 mm � 10 mm (XY) interval spatial sampling.The probe optics were rotated at 50 revolutions per second,generating images with 2,000 transverse pixels, and pulled back at0.5 mm s21. Figure 3a shows a single radial frame from near thecentre of the colon segment. The inset shows an enlarged view ofthe epithelium. Colonic crypts are clearly distinguishable as darkregions surrounded by bright filament-like bands of laminapropria, consistent with previous in vivo studies in smallanimals3,4. An example of a crypt is indicated by the red arrow.The ability to detect and analyse crypt structures in threedimensions is important, because abnormal crypts are signaturesof early neoplastic change37.

The complete volumetric data set is composed of 885 radialframes, which can be processed and displayed in threedimensions. Figure 3b shows a cutaway view of the rendered dataset to allow visualization of the tissue morphology. Groups ofcolonic crypts are visible on the luminal surface as well as in theepithelial wall. It is possible to perform a virtual incision byunfolding the tissue from a cylindrical form into a rectangularform, as shown in Fig. 3c. A flattening algorithm is used to

improve the quality of the two-dimensional (2D) cross-sectionalimages obtained from the 3D data set. The unfolded visualizationenables clearer appreciation of tissue morphology andcomparison to techniques such as endoscopy or biopsy-basedhistology. (See Supplementary Information, Videos 1–5, forvirtual fly-throughs of the rendered data sets, with downsamplingapplied to reduce file sizes.) After unfolding, cross-sectional OCTimages with arbitrary orientations can be generated that areprecisely registered to the surface of the tissue.

Figure 4a shows a single longitudinal (YZ) slice through themiddle of the unfolded volume. Because the data set is sampledwith a high spatial density, consecutive slices can be averaged toreduce speckle noise without significantly blurring imagefeatures. Figure 4b shows the mean of seven consecutive slices,equivalent to averaging a 21-mm-thick section. This dimensioncorresponds to less than two epithelial cells, so tissue structure islargely constant and minimal blurring is observed. As the specklesize is approximately equal to the 9-mm focal spot size, thespeckle pattern is decorrelated over the section and imageaveraging enhances tissue contrast. Figure 4c shows an enlarged

Figure 3 Volumetric renderings of in vivo rabbit colon. a, Single radial frame acquired in 20 ms. Inset shows enlarged view of epithelium, with crypt indicated by

a red arrow. b, Cutaway view of the rendered volume. c, Unfolded data set showing the cylindrical volume as a rectangular tissue slab. The entire volume was

acquired in 17.7 s.

Absorbing cellsColonic mucosaLamina propria

Muscularis mucosaSub-mucosa Circular muscleMyenteric plexus

Longitudinal muscle

Serosa

Fatty tissue

ACCMLP

MMSMCiMP

LM

Se

FT

AC

CM

LPMMSM

Ci

MP

LM

500 μm

200 μm 200 μm

500 μm

Z

Y

Z

Y

Z

Y

Figure 4 Longitudinal (YZ) images through the centre of the colon volume. a, Single YZ slice. b, Average of seven slices spanning a width of 21 mm, showing

reduced speckle noise. c, Enlarged view of b showing visibility of layered structure. d, Representative histology with haematoxylin and eosin stain.

ARTICLES




view of the region marked in Fig. 4b. The OCT image correlates wellwith representative histology of colonic tissue from the sameanimal (Fig. 4d), although the mucosal surface is flattened in theOCT images as a result of compression by the imaging probe.The dark vertical bands are caused by faecal material on theluminal surface.

En face (XY) OCT images can also be generated, similar to thoseobtained with a magnifying endoscope or microscope, except thatthe OCT images can be viewed at arbitrary tissue depths.Figure 5a shows an en face OCT image of the colon averaged overan 18-mm section at a depth of 144 mm, corresponding to thecentre of the colonic mucosa. In comparison with human tissue,crypts in the rabbit colon are smaller, less organized, and moretightly packed. Rabbit crypts are often separated by only a fewmicrometres of lamina propria, making it difficult to identifysingle crypts in the XY and YZ images owing to the 10-mm spatialsampling in Y. Nevertheless, groups of crypts are visible as darkregions between bright white bands of thicker lamina propria inFig. 5a. The three dark bands extending along the Y dimension arepolarization artifacts caused by probe rotation. Figure 5b shows anenlarged view of a section of Fig. 5a with groups of crypts markedby arrows, and Fig. 5c shows an en face histological section fromapproximately the same depth. The histology appears differentfrom the OCT data as a result of tissue shrinkage, which occursduring histology processing, and differences in crypt orientation inthe OCT image compared with the histological plane.

To verify that the 3D-OCT system is capable of resolvingindividual human crypts, an excised sample of formalin-fixedhuman colonic tissue was also imaged. An en face OCT imageaveraged over the depth of the mucosa is shown in Fig. 5d, with

representative histology shown in Fig. 5e. The optical propertiesare different from the rabbit, with ex vivo human tissue showinglow-scattering lamina propria and high-scattering crypts38.Excellent correlation is observed between the OCT images andhistology. With 3D-OCT, crypts are measured to be �100 mm indiameter with a separation of 30–60 mm.

Densely sampled volumetric 3D-OCT data sets containcomprehensive information about tissue microstructure andenable the generation of any desired cross-sectional image.Figure 6a shows an unfolded cross-sectional (XZ) image from the3D in vivo rabbit data set averaged over a 20-mm length.Figure 6b shows an enlarged view of a segment of Fig. 6a. As aresult of the higher spatial sampling density in X compared withY, it is easier to identify large individual crypts, as shown by theblack arrows. The dashed line defines a plane, YN, normal to thelong axis of the crypts. Figure 6c shows this plane averaged alonga 140-mm section parallel to the long axis. The crypts, markedwith black arrows and circumscribed with a black circle, are ovalin cross-section, probably due to compression of the tissue fromthe probe. Based on the size of the low-scattering region withinthe lamina propria bands, the crypts are �50 � 130 mm incross-section. From Fig. 6b, the crypts extend �350 mm indepth. The ability to quantitatively measure individual cryptscould be important in recognizing clusters of aberrant cryptfoci37, which are markers of early colon cancer.

DISCUSSION

The objective of 3D-OCT endomicroscopy is to visualize themicroscopic architectural morphology of tissue, so high spatial

X

Y

500 μm

500 μm

500 μm500 μm

1,000 μm

Figure 5 En face (XY ) images from colonic mucosa layer. a, Average of seven slices spanning a depth of 18 mm, showing crypt structure and luminal folds in

in vivo rabbit colon. b, Enlarged view showing details of crypt distribution. c, Representative rabbit colon histology with hematoxylin and eosin stain. d, En face

image of ex vivo human colonic tissue sample, averaged over the mucosal layer. e, Representative human colon histology with hematoxylin and eosin stain.

ARTICLES




resolution is essential. As an example, aberrant crypt foci areabnormal colonic crypts of 40–100 mm in diameter withsubtle structural variations linked to early colon cancer37.To visualize these structures, a 3D resolution of better than10 mm�10 mm�10 mm is required.

It is important to differentiate between optical resolution,spatial sampling density and true resolution. Optical resolution,determined by the focal spot size in XY and the width of thepoint spread function in Z, defines the best theoretical resolution.Spatial sampling density is the distance between consecutive axialscans in X and between consecutive frames in Y. Spatial samplingdensity in Z can be set arbitrarily by selecting the Fourier-transform length. True resolution is determined by acombination of optical resolution and spatial sampling density,and defines the smallest feature size that can be visualized. Itshould be noted that optical resolution is not constant withimaging depth, because the beam diverges away from the focalposition. In addition, the ability to resolve structures depends onfeature contrast.

The Nyquist criterion requires at least two spatial samples ineach dimension for every optical resolution element. For theOCT system reported here, the optical resolution in tissue is9 mm � 9 mm � 7 mm, so spatial samples should be acquiredevery 4.5 mm � 4.5 mm � 3.5 mm. The XY spatial samplingdensity is fundamentally limited by imaging speed, tissue surfacearea and maximum imaging duration. As surface area andimaging duration are determined by anatomy and physiology, theonly parameter available for increasing spatial sampling density isimaging speed. Therefore high imaging speeds, as determined bythe laser sweep rate and DAQ capacity, are required formicroscopic 3D-OCT resolution.

Previous ultrahigh-resolution OCTendoscopic imaging studiesachieved 3.7-mm axial resolutions at 3,125 axial lines per second inthe rabbit GI tract using time-domain detection and a femtosecondCr:forsterite laser with a 150-nm FWHM bandwidth at 1,250 nm(ref. 3). Record endoscopic axial image resolutions of 2.4 mm

at 20,000 axial lines per second were recently achieved in themouse colon using spectral/Fourier-domain detection and afemtosecond Ti:sapphire laser with a 160-nm FWHM bandwidthat 800 nm (ref. 39). Both studies demonstrated ultrahigh axialresolution, but were focused on 2D imaging. Recent 3D-OCTendoscopy studies in the pig using polygon-mirror-based sweptlasers achieved optical resolutions of 15 mm �15 mm � 7 mm,but had XY spatial sampling densities of only 25 mm � 33 mm asa result of restricted imaging speeds24,25. Consequently, the trueresolution was 50 mm � 66 mm � 7 mm (23,100 mm3) at imagingspeeds of 10 kHz in the oesophagus and 54 kHz in the coronaryartery. This study also demonstrated virtual unfolding of luminalstructures and en face viewing, but true resolutions wereinsufficient to resolve microscopic structures in three dimensions.

The current study uses a 100 kHz FDML laser andoptically clocked DAQ system to enable XY spatial sampling at3 mm � 10 mm intervals over a surface area of �55 mm2 withan imaging time of 17.7 s. The true resolution is therefore9 mm � 20 mm � 7 mm (1,260 mm3), optically limited in XZ andsampling limited in Y. This represents an 18 times improvementin true volumetric resolution, a 10 times improvement in GIimaging speed, and an improvement of about a factor of two inimaging speed compared with arterial imaging in recentendoscopic 3D-OCT studies.

The current results suggest the feasibility of performingquantitative measurements of structures such as crypts. The mainsources of measurement error are the limited tissue contrast,uncertainty in the tissue refractive index, and rotational probejitter. Contrast is the most significant source of error because thecrypt outline is subjectively determined by a human reader,which could lead to a 10–20% error. Uncertainty in therefractive index leads to a fixed error in all measurements, butwould not affect comparative measurements between differentsites in the same organ. Probe jitter is also a source of error, withconsecutive frames registered to within a few micrometres of eachother, and can be reduced with different probe designs.

Figure 6 Cross-sectional (XZ ) images with quantitative feature measurements. a, Average of two slices spanning a length of 20 mm, showing crypts extending

from mucosa to luminal surface. b, Enlarged view showing details of the crypt structure. The dashed line indicates the YN cut plane perpendicular to the long axis of

the crypts. Black arrows indicate the location of two specific crypts. c, Average image formed by projecting the YN plane defined in b over the 140 mm length

parallel to the crypt axis. The same two crypts are visible and can be measured as 50mm � 130mm � 350 mm oval cylinders.

ARTICLES




The high resolution of the system reported here enables in vivo3D-OCT imaging and measurement of small features such ascolonic crypts, which has not been possible using other methods.However, further increases in imaging speed are required to fullymake use of the optical resolution. Imaging speed is at presentlimited by the DAQ electronics. As imaging speed increases, theinterference-fringe frequency associated with a given imagingdepth increases at the same rate. Therefore the ADC andcomputer data bus bandwidths must increase to maintaincontinuous transfer to the system RAM or hard drives. Opticalcoherence tomography imaging of the skin at rates of 370 kHzhas been demonstrated using FDML lasers28, but was limited to asmall spatial volume as the DAQ system could not continuouslytransfer data to the computer at this high speed.

In the future, improved ADC hardware and data busarchitectures will enable higher imaging speeds with continuousdata transfer. Future 12-bit ADC hardware is expected to operateat 800 MHz, and would support an imaging depth .2 mm usinga 370-kHz FDML laser. Additional advances are expected toimprove the system axial resolution to the 5-mm level supportedby the laser’s tuning range. The system should achieve a trueresolution of 9 mm � 9 mm � 5 mm while reducing totalacquisition time by a factor of 2.5. Analog-to-digital conversionhardware using new PCI Express bus architecture will supporttransfer rates up to 16 Gbyte s21, removing another keybottleneck. Real-time data compression can further aid intransferring and managing the data sets. Fourier-domain mode-locked technology has been proven to scale to speeds not possiblewith conventional swept lasers while maintaining a broad tuningrange and narrow instantaneous linewidth, which will enablenext-generation DAQ systems to be harnessed for high-performance 3D-OCT endomicroscopy. Finally, FDML lasers alsoprovide improved phase stability compared with conventionalwavelength-swept lasers29. This feature, combined with the highsweep rates, promises to enable high-sensitivity 3D-OCT mappingof microvasculature using Doppler techniques40–43.

In conclusion, we have demonstrated a new imagingplatform for in vivo 3D-OCT endomicroscopy. High imagingspeeds and spatial resolutions were achieved, enabling thevisualization and measurement of microscopic features assmall as colonic crypts. Three-dimensional endoscopic OCT datasets enable powerful visualization techniques includingmanipulation of tissue geometry, speckle averaging, generationof en face views similar to endoscopic images, and generationof cross-sectional images with arbitrary orientations. Futureimprovements in data acquisition technology will allowvolumetric imaging at even higher rates, further enhancingmicroscopic resolutions and enabling a wide range of clinical3D-OCT endomicroscopy applications.

METHODS

FIBRE-OPTIC ENDOSCOPE PROBE

The fibre-optic probe is a 0.83-mm diameter spiral-scanning device thatcombines rapid rotary motion (up to 50 revolutions per second) with linearpullback (0.5–5 mm s21) to image 3D volumes. A photograph of the distal tip ofthe probe is shown in the right inset of Fig. 2. The probe is composed of a single-mode optical fibre inside a flexible polymer tube that provides mechanicalstability for manipulation of the device. A metal torque cable is attached to thefibre to transfer rotary and linear motion from the proximal end of the probe tothe distal tip. The use of a torque cable results in highly uniform probe rotation,which is required to ensure that successive frames are registered to within a fewmicrometres of each other. The proximal end of the fibre is attached to a patientinterface unit (PIU) containing the rotary and push/pull actuators. The distalend is attached to an angle-polished microlens, providing a 2v0 focal spot of9 mm as measured with a resolution test target. The microlens is 1 mm long,

125 mm wide, and polished at a 408 angle to direct the imaging beam through thepolymer tube. This design reduces specular reflections from the probe surfacesbecause the beam exits the sheath on a non-normal trajectory. The probe can beflushed with water to wash material away from the imaging site and furtherreduce specular reflections.

ANIMAL PREPARATION

The animal used in this study was an adult female New Zealand white rabbit.Imaging was performed in accordance with a protocol approved by the MITCommittee on Animal Care. The animal was fed a liquid diet for 24 h prior toimaging to clear the colon of most faecal material. Anaesthesia was administeredusing an intramuscular injection of ketamine (35 mg kg21) and xylazine(5 mg kg21). Maintenance doses of ketamine (8 mg kg21) and xylazine(1 mg kg21) were administered approximately every 20–30 min.

Before imaging, the probe was disinfected with 90% isopropanol. A flexibleintroducer was placed in the rectum to reduce pressure on the probe. The probewas advanced until a light resistance to motion was detected, indicating thepresence of a bend in the colon. This location was noted to prevent over-insertion. Imaging was conducted at multiple locations prior to the point of lightresistance with periodic water flushing. At the conclusion of the experiment, theanimal was killed with pentobarbitol sodium (250 mg kg21). Samples of colontissue were harvested, fixed in formalin and processed using standardhistological techniques.

SIGNAL PROCESSING

A flattening algorithm to correct for distortions caused by the non-uniformcentration of the probe sheath was applied as part of the virtual incision andunfolding process. The algorithm operates on each frame independently. First, a15 � 15 pixel gaussian filter was applied to reduce speckle noise. Next, a hardthreshold was applied to remove low-amplitude features. A Sobel edge detectorwas then used to locate the outer surface of the probe sheath. A polynomial curvefit to the detected edge was then used to shift each axial line in the original,unfiltered image such that the probe sheath appears flat within the frame. Noother processing or corrections for motion artifacts were applied.

Received 6 July 2007; accepted 2 October 2007; published 25 November 2007.

References1. Huang, D. et al. Optical coherence tomography. Science 254, 1178–1181 (1991).2. Tearney, G. J. et al. In vivo endoscopic optical biopsy with optical coherence tomography. Science 276,

2037–2039 (1997).3. Herz, P. R. et al. Ultrahigh resolution optical biopsy with endoscopic optical coherence tomography.

Opt. Express 12, 3532–3542 (2004).4. Tumlinson, A. R. et al. In vivo ultrahigh-resolution optical coherence tomography of mouse colon

with an achromatized endoscope. J. Biomed. Opt. 11, 06092RR (2006).5. Swanson, E. A. et al. In vivo retinal imaging by optical coherence tomography. Opt. Lett. 18,

1864–1866 (1993).6. Tearney, G. J., Bouma, B. E. & Fujimoto, J. G. High-speed phase- and group-delay scanning with a

grating-based phase control delay line. Opt. Lett. 22, 1811–1813 (1997).7. Rollins, A. M., Kulkarni, M. D., Yazdanfar, S., Ung-arunyawee, R. & Izatt, J. A. In vivo video rate

optical coherence tomography. Opt. Express 3, 219–229 (1998).8. Hsiung, P.-L. et al. High-speed path-length scanning with a multiple-pass cavity delay line. Appl. Opt.

42, 640–648 (2003).9. Yang, V. X. D. et al. High speed, wide velocity dynamic range Doppler optical coherence tomography

(Part I): System design, signal processing, and performance. Opt. Express 11, 794–809 (2003).10. Oldenburg, A. L., Reynolds, J. J., Marks, D. L. & Boppart, S. A. Fast-Fourier-domain delay

line for in vivo optical coherence tomography with a polygonal scanner. Appl. Opt. 42,4606–4611 (2003).

11. Barfuss, H. & Brinkmeyer, E. Modified optical frequency-domain reflectometry with high spatial-resolution for components of integrated optic systems. J. Lightwave Technol. 7, 3–10 (1989).

12. Glombitza, U. & Brinkmeyer, E. Coherent frequency-domain reflectometry for characterization ofsingle-mode integrated-optical wave-guides. J. Lightwave Technol. 11, 1377–1384 (1993).

13. Passy, R., Gisin, N., Vonderweid, J. P. & Gilgen, H. H. Experimental and theoretical investigations ofcoherent OFDR with semiconductor-laser sources. J. Lightwave Technol. 12, 1622–1630 (1994).

14. Vonderweid, J. P., Passy, R. & Gisin, N. Midrange coherent optical frequency-domain reflectometrywith a DFB laser-diode coupled to an external cavity. J. Lightwave Technol. 13, 954–960 (1995).

15. Passy, R., Gisin, N. & Vonderweid, J. P. High-sensitivity-coherent optical frequency-domainreflectometry for characterization of fiberoptic network components. IEEE Photon. Technol. Lett. 7,667–669 (1995).

16. Fercher, A. F., Hitzenberger, C. K., Kamp, G. & Elzaiat, S. Y. Measurement of intraocular distances bybackscattering spectral interferometry. Opt. Commun. 117, 43–48 (1995).

17. Chinn, S. R., Swanson, E. A. & Fujimoto, J. G. Optical coherence tomography using a frequency-tunable optical source. Opt. Lett. 22, 340–342 (1997).

18. Golubovic, B., Bouma, B. E., Tearney, G. J. & Fujimoto, J. G. Optical frequency-domain reflectometryusing rapid wavelength tuning of a Cr4þ:forsterite laser. Opt. Lett. 22, 1704–1706 (1997).

19. Choma, M. A., Sarunic, M. V., Yang, C. H. & Izatt, J. A. Sensitivity advantage of swept source andFourier domain optical coherence tomography. Opt. Express 11, 2183–2189 (2003).

20. de Boer, J. F. et al. Improved signal-to-noise ratio in spectral-domain compared with time-domainoptical coherence tomography. Opt. Lett. 28, 2067–2069 (2003).

21. Leitgeb, R., Hitzenberger, C. K. & Fercher, A. F. Performance of Fourier domain vs. time domainoptical coherence tomography. Opt. Express 11, 889–894 (2003).

ARTICLES




22. Yun, S. H., Tearney, G. J., de Boer, J. F., Iftimia, N. & Bouma, B. E. High-speed optical frequency-domain imaging. Opt. Express 11, 2953–2963 (2003).

23. Oh, W. Y., Yun, S. H., Vakoc, B. J., Tearney, G. J. & Bouma, B. E. Ultrahigh-speed optical frequencydomain imaging and application to laser ablation monitoring. Appl. Phys. Lett. 88, 103902 (2006).

24. Yun, S. H. et al. Comprehensive volumetric optical microscopy in vivo. Nature Med. 12,1429–1433 (2006).

25. Vakoc, B. J. et al. Comprehensive esophageal microscopy by using optical frequency-domainimaging. Gastrointestinal Endoscopy 65, 898–905 (2007).

26. Huber, R., Wojtkowski, M., Taira, K., Fujimoto, J. G. & Hsu, K. Amplified, frequency swept lasers forfrequency domain reflectometry and OCT imaging: design and scaling principles. Opt. Express 13,3513–3528 (2005).

27. Huber, R., Wojtkowski, M. & Fujimoto, J. G. Fourier domain mode locking (FDML): A newlaser operating regime and applications for optical coherence tomography. Opt. Express 14,3225–3237 (2006).

28. Huber, R., Adler, D. C. & Fujimoto, J. G. Buffered Fourier domain mode locking (FDML):Unidirectional swept laser sources for OCT imaging at 370,000 lines per second. Opt. Lett. 31,2975–2977 (2006).

29. Adler, D. C., Huber, R. & Fujimoto, J. G. Phase-sensitive optical coherence tomography atup to 370,000 lines per second using buffered Fourier domain mode locked lasers. Opt. Lett. 32,626–628 (2007).

30. Wojtkowski, M., Kowalczyk, A., Leitgeb, R. & Fercher, A. F. Full range complex spectral opticalcoherence tomography technique in eye imaging. Opt. Lett. 27, 1415–1417 (2002).

31. Choma, M. A., Yang, C. H. & Izatt, J. A. Instantaneous quadrature low-coherence interferometrywith 3 � 3 fiber-optic couplers. Opt. Lett. 28, 2162–2164 (2003).

32. Davis, A. M., Choma, M. A. & Izatt, J. A. Heterodyne swept-source optical coherence tomography forcomplete complex conjugate ambiguity removal. J. Biomed. Opt. 10, 064005 (2005).

33. Yun, S. H., Tearney, G. J., de Boer, J. F. & Bouma, B. E. Removing the depth-degeneracy in opticalfrequency domain imaging with frequency shifting. Opt. Express 12, 4822–4828 (2004).

34. Zhang, J., Nelson, J. S. & Chen, Z. P. Removal of a mirror image and enhancement of the signal-to-noise ratio in Fourier-domain optical coherence tomography using an electro-optic phasemodulator. Opt. Lett. 30, 147–149 (2005).

35. Brinkmeyer, E. & Ulrich, R. High-resolution OCDR in dispersive wave-guides. Electron. Lett. 26,413–414 (1990).

36. Wojtkowski, M., Leitgeb, R., Kowalczyk, A., Bajraszewski, T. & Fercher, A. F. In vivo human retinalimaging by Fourier domain optical coherence tomography. J. Biomed. Opt. 7, 457–463 (2002).

37. Takayama, T. et al. Aberrant crypt foci of the colon as precursors of adenoma and cancer. New Engl. J.Med. 339, 1277–1284 (1998).

38. Hsiung, P. L. et al. Ultrahigh-resolution and 3-dimensional optical coherence tomography ex vivoimaging of the large and small intestines. Gastrointestinal Endoscopy 62, 561–574 (2005).

39. Tumlinson, A. R. et al. Endoscope-tip interferometer for ultrahigh resolution frequency domainoptical coherence tomography in mouse colon. Opt. Express 14, 1878–1887 (2006).

40. Li, H. et al. Feasibility of interstitial Doppler optical coherence tomography for in vivo detection ofmicrovascular changes during photodynamic therapy. Lasers Surg. Med. 38, 754–761 (2006).

41. Aalders, M. C. G. et al. Doppler optical coherence tomography to monitor the effect ofphotodynamic therapy on tissue morphology and perfusion. J. Biomed. Opt. 11, 044011 (2006).

42. Hanna, N. M. et al. Feasibility of three-dimensional optical coherence tomography andoptical Doppler tomography of malignancy in hamster cheek pouches. Photomed. Laser Surg. 24,402–409 (2006).

43. Seki, J., Satomura, Y., Ooi, Y., Yanagida, T. & Seiyama, A. Velocity profiles in the ratcerebral microvessels measured by optical coherence tomography. Clin. Hemorheol. Microcirc. 34,233–239 (2006).

AcknowledgementsWe gratefully thank Bob Shearer from Lightlabs Imaging for his contributions. This research wassponsored by the National Institutes of Health (grants R01-CA75289-09 and R01-EY011289-20), AirForce Office of Scientific Research (grants FA9550-040-1-0046 and FA9550-040-1-0011), NationalScience Foundation (grants BES-0522845 and ECS-0501478), the Natural Sciences and EngineeringResearch Council of Canada (supporting D.C.A.), and the German Science Foundation DFG (supportingR.H.). R.H. was visiting from the Ludwig-Maximilians-Universitat Munchen, Munchen, Germany.Correspondence and requests for materials should be addressed to J.G.F.Supplementary information accompanies this paper on www.nature.com/nature photonics.

Author contributionsD.C.A. and Y.C. were responsible for project planning, experimental work and data analysis. R.H. wasresponsible for experimental work. J.S. was responsible for experimental work, designed the equipmentand provided technical material. J.C. was responsible for project planning and provided biologicalmaterials. J.G.F. obtained support, administered the project and was responsible for project planning.

Competing financial interestsThe authors declare competing financial interests: details accompany the full-text HTML version of thepaper at www.nature.com/nature photonics.

Reprints and permission information is available online at http://npg.nature.com/reprintsandpermissions/

ARTICLES


www.nature.com/nature photonics

http://npg.nature.com/reprintsandpermissions/


three-dimensional endomicroscopy using optical coherence ... · 2007 nature publishing group...

Documents