British Journal of Radiology (2005) 78, 976-988
© 2005 British Institute of Radiology
doi: 10.1259/bjr/55735832
Optical coherence tomography
A Gh Podoleanu, PhD
School of Physical Sciences, University of Kent, Canterbury CT2 7NR, UK
 |
Abstract
|
|---|
A review is presented of different scanning, acquisition and processing techniques used to obtain depth-resolved information in optical-coherence tomography (OCT). The principles and performances of different OCT techniques are discussed and images from different types of tissue are presented. Special attention is devoted to the progress in using the time-domain flying spot OCT technique and combination of the en face OCT imaging with confocal microscopy. Although OCT is based on white light interferometry, which is a well established and an old technology, the quest for higher resolution and faster acquisition of in vivo images has ensured OCT a rapid evolution in the last decade. Highly adventurous avenues to expand the OCT capabilities and trends are presented at the end of the review.
|
Introduction
|
|---|
Optical-coherence tomography (OCT) is a non-invasive high-resolution imaging modality which employs non-ionizing optical radiation. OCT derives from low-coherence interferometry. This is an absolute measurement technique which was developed for high-resolution ranging and characterization of optoelectronic components [1, 2]. The first application of low-coherence interferometry in the biomedical optics field was for the measurement of eye length [3]. Adding lateral scanning to a low-coherence interferometer, allows depth resolved acquisition of three-dimensional (3D) information from the volume of biological material [4]. The concept was initially employed in heterodyne scanning microscopy [5]. OCT has the potential of achieving high-depth resolution, which is determined by the coherence length of the source [6]. This is the length over which a process or a wave maintains strict phase relations; an ideal laser source, for instance, emits light with more than a few kilometres coherence length, while the coherence length of light emitted by a tungsten lamp could be as short as 1 µm. Interference takes place only between events that happen within the coherence length. Optical sources are now available with coherence lengths below 1 µm [7]. When combined with confocal microscopy (CM), OCT adds improved depth resolution and sensitivity.
Figure 1
illustrates the unique capabilities of the OCT technology, a cross sectional image from the skin on the tip of a finger. Within a depth of 700 µm, back-scattering structures are resolved and clearly displayed such as the stratum corneum, sweat ducts and epidermis. The image provides information on subsurface structure otherwise obtainable only by histology. This explains why OCT is referred to as an optical biopsy method. Similar images, from different types of tissue, are obtainable with modern OCT tools in fractions of a second and with optical powers well below the maximum safety level. Different OCT methods and scanning procedures have been devised to generate images such as that in Figure 1, as presented in this review.
View larger version (69K):
[in this window]
[in a new window]
|
Figure 1. Cross section optical-coherence tomography (OCT) image from the skin on the human finger tip. Image obtained in 0.5 s using a superluminescent diode of central wavelength 793 nm and delivering 600 µW optical power to skin. Depth resolution is 15 µm.
|
|
Using sources with extremely short coherence length, submicron depth resolution is achievable even when the microscope objective is far away from the investigated target. This is one of the most important features of OCT, which explains the high level of interest for OCT in ophthalmology. Confocal imaging was initially applied for high resolution imaging of the eye. In confocal microscopy, the depth resolution is inversely proportional to the square of the numerical aperture of the microscope objective. Because the retina is 2 cm away from the eye lens, which takes the role of the microscope objective, for a non dilated pupil size the numerical aperture is below 0.1. Therefore, the depth resolution in imaging the retina with confocal laser scanning ophthalmoscopy is limited to >0.3 mm by the combined effect of a low numerical aperture and aberrations of the anterior chamber. Using OCT however, a depth resolution better than 3 µm [7] is achievable and, when imaging the cornea, OCT does not require contact in contrast to confocal microscopy.
OCT has also been extended to image higher scattering tissue, such as skin, collagen, dentin and enamel [8, 9]. An increased interest is manifested in the applications of OCT in endoscopy, with the development of specialized catheters to accommodate different internal organs [10].
As explained below, the main characteristic of OCT, the depth resolution, derives from the manipulation of wave-trains of finite length emitted by a low-coherence light source. The principle of operation is different from that of other medical imaging technologies. OCT employs optical and infrared waves and therefore is dominated by diffraction which precludes algorithms for image reconstruction used in X-ray or MRI. Sometimes analogies are made of OCT cross section images with B-scan ultrasound images. However, ultrasound beams are longitudinal waves, whereas the waves in OCT are transverse. It is true that similarity does exist between the time taken for the ultrasound to propagate back and forth to the probe head (giving distance for a known ultrasound velocity in tissue) and the time taken by the optical waves in OCT to travel over a certain path length. However, whereas ultrasound imaging is a time of flight technique, where time gating is used to display ordered time events, in OCT the gating process operates in space, based on interferometry, as explained below.
In terms of the product of depth resolution and penetration depth, OCT fills the gap between confocal microscopy and ultrasound imaging. This product is approximately 0.1 µm x 500 µm in confocal microscopy [11], 1 µm x 3000 µm in OCT and 50 µm x 5000 µm in high frequency ultrasound [12].
|
Principle of operation
|
|---|
As mentioned above, interference lies at the core of the OCT imaging method. One of the two signals which interfere is internal, the reference beam, and the other, the signal reflected by the target, the object beam. Different interferometer configurations can be used in OCT. Let us consider a Michelson interferometer as shown in Figure 2a, illuminated by an optical source (OS). Light from the OS is divided into two beams by using a plate beam-splitter, BS. Figure 2b shows a different set-up, where the bulk beam-splitter, BS, is replaced by an in-fibre beam-splitter, termed a directional coupler, DC. It has all the properties of a bulk beam-splitter plus the advantage that the input and the output ports can easily be altered by moving the fibre ends. (This is exploited in delivering light to the tissue and collecting the back-reflected light via an optical fibre, which can be as thin as 0.125 mm). The two beams, reference and object, are recombined by the beam-splitter BS in Figure 2a or by the directional coupler DC in Figure 2b onto a photodetector. From interference theory, the photodetected signal is:
where 
is the photodetector responsivity, O the target reflectivity, R the reference mirror reflectivity, 
the central wavelength of the optical source and P0 the power incident on the object. The optical path difference (OPD) between the two optical paths in the two arms is d=2|lr–lo| (Figure 2b). The interference amplitude depends also on the degree of polarization, described by the factor 
, i.e. the degree of similarity of the orientation of the electric fields in the two optical beams.
View larger version (13K):
[in this window]
[in a new window]
|
Figure 2. Optical-coherence tomography set-up. (a) Michelson interferometer; (b) in-fibre equivalent of the configuration in (a); OS, optical source; BS, beam-splitter; OUT, object under test; DC, single mode directional coupler.
|
|
The first two terms represent time non-varying components and determine noise, usually O<<R. The third term describes the interference; this is periodic and dependent on d and . Each time the reference mirror is moved by an extra /2 (which determines a round trip OPD of ), the photodetected signal strength repeats itself. Hence, maxima and minima are recorded as the reference mirror is moved.
If the optical source is purely monochromatic (an ideal laser source), its spectrum is reduced to a narrow component of a well defined wavelength. In this case, Equation (1)
represents a fringe signal, with an infinite number of peaks as the OPD is changed, as shown at the top of Figure 3. This system cannot be used for depth selection. However, if the optical spectrum contains a spread of components, 
, then configurations such as presented in Figure 2 can select in depth. This arises from the fact that Equation (1) has to be superposed for each wavelength within the range . This superposition essentially involves the cosine terms and will lead to cancellation of the signal at the photodetector for most OPD values. Clearly however, for an OPD=0 most wavelengths will superpose constructively giving a signal at the photodetector. The coherence length lc of the optical source determines the width of the envelope of the fringe pattern (Figure 3 bottom) and hence the depth resolution. The coherence length lc and , are related by:
|
and thus the larger , the shorter lc and hence the better the OCT depth resolution. The temporal extension of the superposed waves determines the strength of the signal. As can be seen in Figure 4, by delaying one of the waves, such as that reflected by the mirror, the amount of superposition of the two waves (giving a constructive interference signal) is correspondingly reduced. This explains why the envelope of the third term in Equation (1) goes to zero for |OPD|>lc.
View larger version (31K):
[in this window]
[in a new window]
|
Figure 3. Comparison between the photodetector output in Figure 2 when an ideal laser (top) or a broadband source (bottom) is used. OPD, optical path difference.
|
|
View larger version (23K):
[in this window]
[in a new window]
|
Figure 4. Superposition of two wavetrains (interference of two beams generated by a source with a large bandwidth).
|
|
Different depths into tissue can be probed by changing the position of the reference mirror (using stepper motors).
If the reference mirror is moved with a known speed then the third term (cosine) in Equation (1) will vary with time and allows the dc components of the signal to be filtered out and the time-varying signal (ac component) amplified.
As an illustration of the technique, two interfaces are considered in Figure 5. At the air–glass interface, 
1, the wave crosses from index of refraction n1 (air) to n2 (glass) and at the glass–air interface, 2, the wave crosses from index of refraction n2 to n1. The photodetected signal is amplified, rectified and low pass filtered to retain its envelope only. Two peaks in the rectified photodetected signal (Figure 5 bottom) result, at positions z1 and z2 of the reference mirror. In this way, the depth positions of the interfaces 1 and 2 are detected. The amplitudes of the envelopes, A1 and A2 are proportional to the square root of each interface reflectivity, as determined by the third term in Equation (1) (i.e. with the index of refraction variation n between the two interfaces). The profile of the square root of the reflectivity versus depth is termed an A-scan. These profiles are shown at the bottom of Figure 5, represented against time obtained by dividing the position of the reference mirror, z, by its velocity, v. As mentioned above, the cosine term shown in the middle of Figure 5 oscillates with a periodicity of /2. Consequently, if the /2 spatial interval is scanned with a velocity v, then in time the peaks are separated by (/2)/v which is the inverse of a frequency f. This means that by moving the reference mirror, the photodetected signal oscillates with a frequency f=2v/. This is called Doppler frequency because f is identical in value with the Doppler shift experienced by the frequency of a wave reflected from a moving mirror.
|
Sources for OCT
|
|---|
The larger the line-width, , the smaller the coherence length lc (Equation (2)). This is a property of Fourier transformation, as the process of superposing the cosine terms in Equation (1) can be described by a Fourier integral. The larger the bandwidth, the shorter the wave-train lengths and smaller the superposition length of the two wave-trains for an OPD different from zero. For an ideal laser source, the wave-trains in Figure 4 would be infinitely long and delaying them relatively would not change the strength of the superposition signal. Let us consider a central wavelength =1 µm. A laser could have <0.01 nm, in which case, using Equation (2), lc
9 cm. In comparison, a tungsten bulb has a 300 nm which gives lc3 µm. However, light from a bulb is difficult to confine to a fibre, therefore special devices have been developed using laser diodes, e.g. light emitting diodes (LED) and superluminiscent diodes (SLD). A typical SLD has a bandwidth =20 nm, i.e. lc44 µm and since lc is defined for a round trip, this leads to a depth pixel size of 22 µm in air, which in tissue, considering a refractive index n of 1.4 gives 16 µm. Modern sources for OCT use Kerr-lens mode-locked lasers and photonics crystal fibres to achieve submicron coherence length [7, 13].
Apart from the wide line-width, the sources for OCT have to exhibit a smooth Gaussian spectrum profile. Therefore, different procedures have been devised to spectrally smooth the output spectrum from sources which manifest a high level of spectral ripple.
|
Main characteristics of OCT
|
|---|
1. OCT uses wavelengths within the band 600 nm to 2000 nm where the main constituents of the tissue, water, pigments, etc. exhibit low absorption [14].
2. The main driving force behind OCT development is its high depth resolution. The wider the optical spectrum line-width, the smaller the coherence length and the better the depth resolution.
3. For interference to take place, a strict phase relationship is required between the interfering waves. Multiple scattered events lose the phase information. Therefore, only single scattered photons contribute to the interference. Consequently, the maximum penetration depth in OCT is the depth from which single scattered photons still originate. This depth is about 1.5 mm in skin using wavelengths within the 800 nm band and about 2 mm when using 1300 nm [15] due to lower scattering at longer wavelengths.
4. Photodetection at the interferometer output involves multiplication of the two optical waves, therefore the weak signal in the object arm, backscattered or transmitted through the tissue, is amplified by the strong signal in the reference arm. This explains the higher sensitivity of OCT when compared with confocal microscopy, which for instance in skin can produce images only to a depth of 0.5 mm [11].
5. Since all OCT systems are built around a confocal microscope, the transverse resolution is determined by diffraction.
|
Different scanning procedures
|
|---|
To obtain 3D information about an object, the imaging system is equipped with two scanning means; one to scan the object in depth and another to scan the object transversally, usually composed of two orthogonal scanners. Depending on the order in which these scans are operated and on the scanning direction associated with the line displayed in the raster of the final image delivered, different scan planes are possible. OCT systems using charge-coupled device (CCD) cameras or arrays of sensors or arrays of emitters eliminate the need of scanning. The scanning terminology is illustrated in Figure 6.
View larger version (25K):
[in this window]
[in a new window]
|
Figure 6. Relative orientation of the axial scan (A-scan), en face scan (T-scan), longitudinal slice (B-scan) and en face or transverse slice (C-scan).
|
|
A-scan based B-scan
B-scan images, analogous to ultrasound B-scans are generated by collecting many A-scans (Figure 6) for different and adjacent transverse positions, as shown in Figure 7 top. The lines in the raster generated correspond to A-scans, i.e. the lines are orientated along the depth coordinate. The transverse scanning means (operating along X or Y, or along the angle 
with 
constant in polar coordinates in Figure 6, with X shown in Figure 7 top) advances at a slower pace to build a B-scan image. The majority of reports in the literature [7, 8] refer to this method of operation.
View larger version (38K):
[in this window]
[in a new window]
|
Figure 7. Different modes of operation of the three scanners in a flying spot optical-coherence tomography system.
|
|
T-scan based B-scan
In this case, the transverse scanners (or scanner) determine(s) the fast lines in the image [16, 17]; each image line is a T-scan (Figure 6). This can be produced by controlling either the transverse scanner along the X-coordinate, or along the Y-coordinate with the other transverse scanner fixed, or controlling both transverse scanners along the polar angle for a given , with the axial scanner fixed. The example in the middle of Figure 7 illustrates the generation of a B-scan using several T-scans, where the X-scanner produces the T-scans and the axial scanner advances slower in depth, along the Z-coordinate. This procedure has a net advantage in comparison with the B-scan generated from several A-scans as it allows production of C-scans, i.e. of OCT transverse (or en face) images (Figure 6) for a fixed reference path, as presented below.
C-scan
C-scans are made from many T-scans along either of X, Y, or coordinates, repeated for different values of the other transverse coordinate, Y, X, or , respectively, in the transverse plane. The repetition of T-scans along the other transverse coordinate is performed at a slower rate (Figure 7 bottom), which determines the frame rate. In this way, a complete raster is generated. Different transversal slices are collected for different depths Z, either by advancing the optical path difference in the OCT in steps after each complete transverse (XY) or (,) scan, or continuously at a much slower speed than the frame rate.
|
Different OCT versions
|
|---|
There are three main OCT methods:- Time domain OCT operates as explained in Figure 5, where an A-scan is produced by varying the OPD in the interferometer to output a reflectivity profile in depth. En face flying spot OCT belongs to the same category where a T-scan is produced by transversally scanning the beam over the target maintaining the reference mirror fixed to generate a reflectivity profile versus angle or lateral position;
- Spectral or Fourier domain OCT, where the interferometer output is sent to an optical spectrometer. The Fourier transform of the spectrometer signal returns the A-scan;
- Swept source OCT, where a laser source is used in contrast to time domain OCT and Fourier domain OCT which both employ a wideband source. In order to obtain similar depth resolution as in the time domain OCT or Fourier domain OCT, the laser frequency needs to be swept within an equivalent band to that of the broadband source used in the time domain OCT or Fourier domain OCT.
Time domain OCT
An in-fibre interferometer (an extension of that shown in Figure 2b) equipped with all three scanners for time domain OCT is shown in Figure 8.
View larger version (20K):
[in this window]
[in a new window]
|
Figure 8. An in-fibre interferometer equipped with all three scanners for time domain optical coherence tomography. SLD, superluminescent diode; DC, directional coupler; C1, C2, microscope objectives; M, mirror; SXY, galvanometer scanning mirror head; MX, MY, scanner mirrors; L1, L2, lenses; PD, photodetector; ASO, analogue storage oscilloscope; TX, TY, triangle waveform generators.
|
|
Longitudinal OCT
This is the main OCT technology used to image tissues and refers to cross-section images (B-scan sections) in the plane: (depth axis) x (lateral or angular axis), where the slice is constructed from many A-scans repeated for subsequent transverse pixels. The signal is processed as explained in Figure 5. Development of the longitudinal OCT based on A-scans was facilitated by a technical advantage: when moving the mirror in the reference path of the interferometer, not only is the depth scanned, but a carrier (in the form of an oscillating signal) is also generated. As explained above, the carrier signal is the oscillation inside the envelope of the two pulses in the middle of Figure 5. Its frequency is the Doppler shift produced by the axial scanner itself (Figure 8, moving along the axis of the system, Z, to explore the tissue in depth). In longitudinal OCT, the axial scan is the fastest (Figure 7, top) and its movement is synchronous with displaying the pixels along the line in the raster, while the lateral scanning determines the frame rate.
En face OCT
When using the set-up in Figure 8 for en face OCT, lateral scanning is the fastest and determines the line rate while the depth scanning to produce a B-scan image determines the frame rate. In the C-scan regimen, one of the transverse scanners determines the line rate and the other is slower, operating at the frame rate. The depth scanning is the slowest in this case. It is more difficult to generate en face OCT images than longitudinal OCT images as the reference mirror is fixed and no carrier is generated. In T-scan based OCT images, a phase modulator is needed [18]. It has been shown that the X or Y-scanning device itself introduces a path modulation which plays a similar role to the path modulation created by the longitudinal scanner employed to produce A-scans or B-scans based on A-scans. When the incident beam falls on the axis of rotation of the transverse scanner mirror, the signal collected from a mirror looks like that in Figure 9 top right. Frequencies from zero up to a maximum value are generated, shown in Figure 9 top left. If the beam is placed off-axis a small distance, the oscillation in the photodetector signal becomes denser and an equivalent carrier to that in Figure 5 middle is generated as shown by the spectrum of the photodetected signal in Figure 9 left bottom.
View larger version (29K):
[in this window]
[in a new window]
|
Figure 9. Spectra (left) and the temporal evolution (right) obtained with the set-up in Figure 8 from a mirror as target [16]. If the beam is shifted by 3 mm from the rotation axis (bottom), higher frequency oscillations are generated, concentrated around a carrier, as shown in the left bottom figure [17]. a.u. represents arbitrary units for the strength of the photodetected signal.
|
|
C-scan images and 3D reconstruction are shown in Figure 10 for images collected from the finger tip of a volunteer [19], using the phase modulation introduced by the scanner only. A glass window was used as support for the finger tip. 40 OCT transversal images were collected by moving the glass plate support along with the finger towards the OCT system in steps of 25 µm measured in air. The SLD used operated at a central wavelength of 850 nm and delivered a power to the finger of 0.27 mW. The finger-tip ridges are visible touching the glass plate interface at the top of the 3D volume in Figure 10 left, in the plane "a". The longitudinal cuts show the stratum corneum, epidermis and dermis. The thickness of different layers can be easily determined in the faces "b" and "c". The longitudinal cut in the face "b" in Figure 10 right shows the spiraled structure of the sweat ducts.
View larger version (32K):
[in this window]
[in a new window]
|
Figure 10. 3D display of in vivo optical-coherence tomography image of normal human skin from a volunteer's finger tip, [19]. Volume size: 5 mm x 4 mm x 1 mm (depth measured in air). The arrow ED shows the direction of exploration of the 3D reconstructed volume made from 40 en face slices acquired at 25 µm depth interval.
|
|
|
Simultaneous en face OCT and confocal imaging
|
|---|
The depth resolution in confocal microscopy when using low numerical aperture microscope objectives is limited while the transversal resolution in OCT is affected by random interference effects from different scattering centres (speckle). Therefore there is scope in combining CM with OCT. A separate confocal receiver can be added to the OCT system [20] based on a beam-splitter which reflects a percentage of the returned light from the object to a separate photodetector via a lens and a pinhole. An example of images provided by such a system in the B- and C-scan regimens of operation are shown in Figure 11, from a healthy eye. The system outputs pairs of OCT and confocal images. In Figure 11a, C-scan OCT images are shown at different depths from the fovea region and the confocal corresponding image. Figure 11b shows the pair of images in the B-scan regimen, confocal image (top) and a B-scan OCT image (bottom). The layers clearly discernible in the OCT image bear strong resemblance to histology [21].
View larger version (57K):
[in this window]
[in a new window]
|
Figure 11. (a) C-scan images of the fovea [20] obtained with the dual channel optical-coherence tomography (OCT)/confocal instrument in the C-scan imaging regimen. CSO: confocal scanning ophthalmoscopy image produced by the confocal channel; all the other images were provided by the OCT channel at depths shown below each image. Lateral size: 3 mm x 3 mm. (b) Large angle pair of images produced with the OCT/confocal instrument in the B-scan imaging regimen, showing the fovea and the optic nerve. 28° lateral size (peak to peak angular extension of the fan of scanned rays at the pupil). Top: CSO image; Bottom: B-scan OCT image (depth 2 mm in air). RNFL, retinal nerve fibre layer; PL, photoreceptor layer; RPE, retinal pigment epithelium [21]. Depth resolution in the OCT images: 12 µm.
|
|
Simultaneous OCT and fluorescence imaging
Once a confocal channel is added to the OCT channel, all known applications of confocal microscopy can be implemented on the confocal channel while benefiting from the simultaneous information offered by the OCT channel. One such possibility is to tailor the confocal channel to provide information on the fluorescence of the tissue under investigation. A simultaneous OCT/fluorescence imaging instrument has been produced which provides a fluorescence image of indocyanine green (ICG) at the same time with an en face OCT image from the eye fundus [22]. The same optical source, a SLD at 0=793 nm centre wavelength was employed to produce the OCT signal as well as to excite the ICG. A pair of OCT/ICG images is shown in Figure 12. Aside from the obvious advantages of cost savings and reduction of optical power incident on the patient's retina, the acquisition of simultaneous corresponding images leads to an unique correlation of anatomical features with vascular functional changes.
View larger version (103K):
[in this window]
[in a new window]
|
Figure 12. Simultaneous indocyanine green (ICG) (left) and en-face optical-coherence tomography (right) image from an healthy eye. Lateral size: 26°. Images collected at 40 s after ICG was released into the body [22]. RPE, retinal pigment epithelium.
|
|
|
En face non-scanning systems
|
|---|
Using parallel collection of rays within the system and interfering them with a cluster of reference rays, allows the generation of B-scan and C-scan images with no need for lateral scanning, a method called Coherence Radar [23]. For each pixel in the transverse coordinates of the object, a pair of object beam and reference beam can be identified, using telecentric optics. The simplest implementation set-up employs a CCD photodetector array (the information along a line of pixels is equivalent to a T-scan in Figure 6). The amplitude of the interference signal is recovered using phase-stepping algorithms. Phase shifts are introduced by exact path difference steps, which in total add up to a wavelength, or by continuous change of the OPD and comparing the sequences obtained. This means that real time processing is not possible. The detection of reflective interfaces in a multilayer object using the Coherence Radar method is limited by the dynamic range of the analogue to digital (A/D) converter of the combination CCD-frame grabber system.
For examining micrometre size features, a system based on a Linnik interference microscope with high-numerical-aperture objectives has been reported [24]. Lock-in detection of the interference signal is achieved in parallel on a CCD by use of a phase modulator (such as a photoelastic birefiringence modulator) and full-field stroboscopic illumination. C-scan images are obtained in real time at 1 Hz with better than 80 dB dynamic range, which enables tomography to be generated in scattering media such as biological tissue.
A faster processing method uses an array of photodetectors in a smart chip [25]. A photodetector is employed for each pixel in the en face image, followed by a processing electronics channel (demodulation, rectifier, amplifier, conditioning). One pixel consists of a 35 µm x 35 µm silicon photodiode coupled to a complementary metal-oxide semiconductor electronic circuit. The smart chip is composed of 58 x 58 smart pixels. Because there is no transverse scanning to alter the OPD, a phase modulator is used. The reading is sequential, similar to the reading of a CCD camera in a coherence radar system but the output is a fully demodulated OCT signal. The amplitude of the signal provided by each channel is proportional to the envelope of the OCT interference signal. The smart chip can also operate in the C-scan and B-scan regimen.
Spectral or Fourier domain OCT
In the last 5 years considerable research has been devoted by different groups developing OCT for tissue imaging into the spectral OCT method. This refers to Fourier transformation of the optical spectrum of a low coherence interferometer [26]. This method is an extension of the work on white light interferometry with initial applications in absolute ranging and sensing [27]. Fourier domain OCT is attractive because it eliminates the need for depth scanning in time domain OCT, performed usually by mechanical means. The operation of Fourier domain OCT is based on the demodulation of the optical spectrum output of an interferometer. The spectrum exhibits peaks and troughs, and the period of such a modulation is proportional to the OPD in the interferometer [28]. If multi-layered objects are imaged, such as tissue, each layer imprints its own modulation periodicity, depending on its depth (OPD), with the amplitude of the spectrum modulation proportional to the square root of the reflectivity of that layer. A CCD camera can be used to transform the optical spectrum into an electrical signal which exhibits ripples of different frequencies. A fast Fourier transform (FFT) of the spectrum of the CCD signal translates the periodicity of the channelled spectrum into peaks of different frequency, related to the path imbalance. Such a profile is essentially the A-scan profile of the reflectivity in depth.
Recent studies [29] have shown that Fourier domain OCT can provide a signal to noise ratio that is more than 20 dB better than the conventional time domain OCT and sufficient sensitivity was demonstrated by displaying video-rate images from the retina [30]. Initially, Fourier domain OCT lagged behind the coherence time domain OCT due to the low speed and limited dynamic range of CCDs. Progress in the field of high-speed high-dynamic range CCDs and photodetector arrays now makes Fourier domain OCT attractive. Fourier domain OCT has two disadvantages: (i) focusing point by point in depth (a procedure called dynamic focus, often utilized in the time domain OCT) is not possible, and therefore the interface optics is devised with a large depth of focus, to accommodate the entire range of the A-scan, usually 1–2 mm; this precludes the possibility of using a high numerical aperture objective to enhance the transverse resolution; (ii) the optical spectrum of the interferometer output consists of symmetric spectral terms, i.e. the same image results for positive and negative OPDs. For the latter, an initial adjustment of the OPD=0 outside the range of interest is required. This is not possible all the time, especially when imaging moving thick organs or tissue. Different methods have been devised to attenuate the symmetric terms in order to obtain a correct image such as phase-shifting interferometry, or complex signal processing [31].
Swept source OCT
Recent progress in the fast tunability of laser sources has revived interest in swept source OCT. The achievable signal to noise ratio is similar to that of Fourier domain OCT [29]. Swept source OCT suffers from the same disadvantage as Fourier domain OCT, as the modulation frequency of the interferogram is proportional to the absolute value of the OPD. Therefore, the origin of OPD=0 must be placed outside the depth range of interest. The time required to tune the wavelength determines the time to produce an A-scan. Tuning speeds of a few kHz have been achieved, which allows swept source OCT to compete with time domain OCT and Fourier domain OCT in terms of speed.
|
Functional OCT
|
|---|
OCT provides depth resolved information of reflectivity, phase, and polarization of the backscattered signal. This signal is intimately related to functional disturbances, which usually precede morphological changes. Different versions of OCT systems provide functional information, such as polarization sensitive OCT, spectrometric OCT, differential absorption OCT and Doppler OCT.
Polarization sensitive OCT
Polarization sensitive optical-coherence tomography takes into account the vectorial nature of light waves (state of polarization). It can detect and quantify the polarization properties of the tissue by analysing changes in the polarization state of the backscattered probe light beam. The information provided by polarization sensitive OCT images can be used to identify birefringent structural constituents in the target tissue that are otherwise invisible to conventional OCT or other imaging techniques. A change in the polarization sensitive OCT images of a tissue can be related to a change in the structure, functionality or integrity of the target. For instance, thermal injury denatures collagen in skin and polarization sensitive OCT can sense changes in the collagen birefringence. Retinal nerve fibre layer, cornea and dentin are birefringent. The retinal nerve fibre thickness is an essential parameter in the diagnosis of glaucoma and instruments based on retardation measurements infer the thickness by assuming that the double-pass phase retardation per unit of depth is the same throughout the eye fundus. Polarization sensitive OCT studies have shown that the retinal nerve fibre layer double-pass phase retardation per unit of depth (measured in degrees of rotation of the linear polarization vector/mm) is not the same but varies with values between 0.18 and 0.37 degree/µm [32].
Since the first report of a functional polarization sensitive OCT system [33], a diversity of polarization sensitive OCT configurations has been investigated. They all differ in complexity, capabilities and signal processing schemes. The most complete information about the polarization properties of a biological target is given by systems capable of producing depth resolved Mueller matrix elements [34]. These configurations can account for depolarization as well as changes in the total, linear and circular degree of polarization of the probe beam during propagation in tissue.
If the prevailing polarization characteristic of the target is linear birefringence, more basic configurations can be used. Such a configuration processes two interference signals corresponding to the vertical, V, and horizontal, H, linear polarization components (with respect to the plane of the system) in the reference and sample arms. The rectified envelopes H and V are fed into the signal inputs of a frame grabber to display either the V and H images or the reflectivity and birefringence retardation maps.
Figure 13 displays C-scan OCT images of an extracted human tooth, obtained using a SLD operating at a central wavelength of 850 nm and delivering less than 250 µW to the tooth. Dental tissue structures visible in the individual H and V images are different and indicate the presence of birefringence. The retardation map at the bottom of the right column clearly displays the banded structure characteristic of birefringent tissue. The darkest pixel level corresponds to a phase retardation of 0° whereas the brightest corresponds to a value of 90° retardation.
View larger version (157K):
[in this window]
[in a new window]
|
Figure 13. En face polarization sensitive optical-coherence tomography images of an extracted human tooth: linear output images of the V and H channels (top), net reflectivity and birefringence retardation images (bottom). The size of each image was 2 mm x 2 mm. The images were acquired from a depth corresponding to a 300 µm optical path in dental tissue and their thickness is less than 15 µm.
|
|
Oximetry
Information on the oxygenation or on the concentration of specific constituents of the tissue can be obtained by exploiting their spectral absorption behaviour. The larger the number of interrogating wavelengths the better the quantitative analysis. Multiple wavelength systems were constructed for oximetry before the advent of OCT. However, OCT has the ability to provide additional information on the depth distribution of the oxyhaemoglobin and deoxyhaemoglobin in the tissue [35]. The availability now of large bandwidth sources allows simultaneous OCT measurements in multiple band windows, not technologically possible using the same number of discrete sources. For instance, conventional sources for OCT are SLD, which in the 800 nm range have 20 nm bandwidth. A wide band source [7] with 200 nm band can provide the equivalent of 10 such SLD sources.
Differential absorption OCT
This can be implemented using two channels OCT, each channel operating on a different wavelength. One wavelength is chosen close to the absorption peak of the constituent to be measured with the other wavelength exhibiting low absorption. Such a technique was used to obtain depth resolved concentration of water in the cornea, by operating simultaneously at 1331 nm (where the water has relatively low absorption) and at 1448 nm (where the water exhibits an absorption peak) [36].
Doppler OCT
Returning to Equation (1), if the back-scattering centre moves with constant velocity, then the OPD varies linearly in time: d=vt. Introducing this equation in Equation (1) leads to a frequency oscillation f=2v/, where v is the projection of the velocity vector along the interrogating beam direction. This frequency is the Doppler shift suffered by the frequency of the object beam and appears as a frequency bit when the object beam is mixed with the reference beam on the photodetector in the OCT interferometer. Hence, Doppler OCT can be used to measure or monitor Brownian motion and flows of biological liquids [37, 38]. In addition to laser anemometry, Doppler OCT provides a depth resolved profile of the flow velocity in the vessel, with the resolution determined by the coherence length of the source. Due to the fact that the scanning itself shifts the frequency of the OCT signal, a challenging avenue in research is to produce the OCT image and velocity map simultaneously [39].
|
Tissue clearing and contrast media for OCT
|
|---|
OCT operates with ballistic photons, i.e. photons which have been scattered only once. Therefore, in highly scattering tissue, OCT exhibits a short penetration depth. The depth of light penetration into highly scattering tissue can be improved by the application of biocompatible and osmotically active chemical agents [40]. Two active agents have been investigated on porcine stomach tissue, glycerol and dimethyl sulphoxide. It was demonstrated that glycerol could enhance both OCT imaging depth and contrast while dimethyl sulphoxide enhanced the penetration depth only. Similarly, to address the main source of scattering in blood, which is the difference in refractive index between the cytoplasm of erythrocytes and serum, dextran and an intravenous contrast agent have been investigated in vitro [41]. Improvements in the penetration depth of 69% and 45%, respectively, were obtained.
Contrast enhancing mechanisms for OCT are needed to allow differentiation of tissue types that are morphologically or optically similar. A pump-probe OCT method was proposed to enhance the contrast of OCT images by transient absorption in the tissue induced by an external optical pump beam [42]. When pumped by a Q-switched Nd-Yag laser (532 nm, 
50 ns pulse width, 10–50 J pulse–1, 1 kHz repetition rate), a sample target of methylene blue exhibited attenuation of the OCT signal measured at 830 nm. Similarly, a protein, phytochrome A was spatially located in intralipid [43] by comparing A-scans produced by an OCT system working at 750 nm with and without pumping light at 660 nm. If such a technique can be sensitized to operate with safe levels of optical power, molecular contrast imaging may become feasible.
Non-linear optical effects have also been investigated. Interference between two coherent anti-Stokes Raman scattering and second harmonic generation signals in separate benzene samples shows potential for heterodyne detection with extremely molecular specific contrast [44]. The "fingerprint" nature of Raman spectroscopy could permit selective detection and imaging of different molecules in the sample arm by using similar molecules in the reference arm of the non-linear interferometer.
Incorporating melanin, gold and carbon nanoparticles [45] into the shells of oil-filled microspheres, contrast agents can be engineered. Intravenous injection of such an agent into a mouse demonstrated enhancement of the OCT images from the liver.
|
Progress in imaging different types of tissue
|
|---|
OCT has provided ophthalmologists with depth resolutions in imaging the posterior and anterior segment of the eye previously only achievable with histology. The recent developments of very short coherence optical sources has impacted the in vivo imaging of microstructural morphology. Ultrahigh resolution in vivo tomograms of 1 µm axial x 3 µm lateral resolution of half of a Xenopus laevis cell have been generated [7]. The ultrahigh resolution OCT will benefit the early diagnosis of cancer, although standard OCT has already proved capable of identifying transitional cell carcinoma in a rabbit bladder [46]. Although scattering in the tissue affects the quality of OCT images, a large variety of tissue types have been investigated, such as skin, larynx, colon, heart, brain and breast, and the list is under continuous expansion [8, 13]. Considerable progress has been recorded in OCT based endoscopy [10]. Special catheters have been devised delivering A-scan based B-scans or small needles [47].
In addition to instruments restricted to imaging only, versatile complex systems are designed where OCT is combined with other technologies. A hybrid catheter integrated an OCT probe head and a scintillating fibre for simultaneous OCT structural imaging and detection of beta particles for molecular imaging of vascular diseases [48]. Vulnerable plaques are metabolically active and can be detected at an early stage by intravascular detection of beta radiation emitted by a suitable radionuclide injected into the body. It is hoped that coregistered OCT and total radiation counts may enhance the contrast of detecting small atherosclerotic plaques.
OCT also proved useful in guiding arterial ablation [49], in surgery of breast cancer [50] and in evaluating the effect of Photo Dynamic Therapy [51].
|
Conclusions
|
|---|
No other technology can offer non-invasive non-contact in vivo real time subsurface images with such a high depth resolution. In ophthalmology, OCT has allowed imaging of the back of the eye with 100 times better depth resolution than that using confocal scanning laser ophthalmoscopy. In endoscopy, OCT prevails as the only technology capable of high resolution imaging, with better than 15 times depth resolution than high frequency ultrasound. The efforts of joint teams of hardware developers and clinicians has pushed the performance of OCT from initial applications restricted to imaging the eye with 20 µm at 1 Hz frame rate, to a wide range of tissue types, more than 20 times better depth resolution, operating at video rate and generating not only morphology imaging but providing functional information as well. We expect that, in a few years, broadband sources essential for ultrahigh depth resolution will become more available and more compact. Combination of imaging OCT technology with other imaging modalities should evolve beyond the dual OCT-confocal microscopy presented here, towards adding curing tools such as laser scalpels and ablation lasers to OCT imaging channels. New imaging technology brings not only new information to the clinician, but with it the requirement of interpretation. OCT is no exception in this respect; curvature of the tissue, coupled with the variation of different optical properties from layer to layer, organ movement and the specific way of ray scanning make the correct representation of the OCT image challenging. Specific algorithms for carrying out corrections are being developed to operate either post acquisition or in real time [52]. As the technology matures, we expect more dedicated processing protocols to be developed in the near future, to account for the sophistication of novel hardware techniques and tissue particularities.
|
Acknowledgments
|
|---|
The author acknowledges the support of EPSRC, Swindon, UK; New York Eye and Ear Infirmary, NY, USA; Ophthalmic Technology Inc., Toronto, Canada; Pfizer Sandwich, UK; Leverhulme Trust and Superlum, Moscow, Russia.
Received for publication March 16, 2004.
Revision received March 3, 2005.
Accepted for publication March 14, 2005.
|
References
|
|---|
- Al-Chalabi SA, Culshaw B, Davies DEN. Partially coherent sources in interferometric sensors. First Internat. Conf. Optical Fibre Sensors 26–28 April 1983, I.E.E. London, 132–135.
- Youngquist RC, Carr S, Davies DEN. Optical coherence-domain reflectometry: a new optical evaluation technique. Opt Lett 1987;12:158–60.
- Fercher AF, Mengedoht K, Werner W. Eye length measurement by interferometry with partially coherent light. Opt Lett 1988;13:186–9.[CrossRef]
- Huang D, Swanson EA, Lin CP, Schuman JS, et al. Optical coherence tomography. Science 1991;254:1178––81.
- Sawatari T. Optical heterodyne scanning microscope. Appl Optics 1973;12:2766–72.
- Mandel L, Wolf E. Optical coherence and quantum optics. Cambridge University Press: Cambridge, 1995:148–50.
- Drexler W. Ultrahigh-resolution optical coherence tomography. J Biomed Opt 2004;9:47–71.[CrossRef][Medline]
- Bouma BE, Tearney GJ, editors. Handbook of optical coherence tomography. New York, Basel: Marcel Dekker Inc., 2002.
- Masters BR, Thompson BJ, editors. Selected papers on optical low-coherence reflectometry & tomography. SPIE Milestone Series, MS 165. Washington: SPIE Optical Engineering Press, 2000.
- Tearney GJ, Brezinski ME, Bouma BE, Boppart SA, Pitris S, Southern JH, et al. In vivo endoscopic optical biopsy with optical coherence tomography. Science 1997;276:2037–9.[Abstract/Free Full Text]
- Rajadhyaksha R, Anderson R, Webb R. Video-rate confocal scanning laser microscope for imaging human tissues in vivo. Appl Opt 1999;38:2105–15.[CrossRef][Medline]
- Shammas HJ, Dunne S, Fisher YL. Three-dimensional ultrasound tomography of the eye. Ontario, Canada: Nova Coast Publishing Eden Mills, 1998.
- Fercher A, Drexler W, Hitzenberger CK, Lasser T. Optical coherence tomography – principles and applications. Rep Progr Phys 2003;66:239–303.[CrossRef]
- Lim HW, Sotter NA. Clinical photomedicine. New York: Marcel Dekker, 1993.
- Pan Y, Farkas DL. Noninvasive imaging of living human skin with dual-wavelength optical coherence tomography in two and three dimensions. J Biomed Opt 1998;3:446–55.[CrossRef]
- Podoleanu AGh, Dobre GM, Webb DJ, Jackson DA. Coherence imaging by use of a Newton rings sampling function. Opt Lett 1996;21:1789–91.
- Podoleanu AGh, Dobre GM, Jackson DA. En-face coherence imaging using galvanometer scanner modulation. Opt Lett 1998;23:147–9.
- Hitzenberger CK. Three-dimensional imaging of the human retina by high-speed optical coherence tomography. Opt Express 2003;11:2753–61.[Medline]
- Podoleanu AGh, Rogers JA, Jackson DA, et al. Three dimensional OCT images from retina and skin. Opt Express 2000;7:292–8.[Medline]
- Podoleanu AGh, Jackson DA. Noise analysis of a combined optical coherence tomography and confocal scanning ophthalmoscope. Appl Opt 1999;38:2116–27.[Medline]
- http://www.udel.edu/Biology/Wags/histopage/colorpage/cey/cey.htm [Accessed 25 April 2005].
- Dobre G, Rosen RB, Podoleanu AGh. Simultaneous OCT/ICG fluorescence imaging system for investigations of the eye fundus. Opt Lett 2005;30:58–60.[CrossRef][Medline]
- Dresel T, Hausler G, Venzke H. Three-dimensional sensing of rough surfaces by Coherence Radar. Appl Optics 1992;31:919–25.
- Dubois A, Vabre L, Boccara AC, et al. High-resolution full-field optical coherence tomography with a Linnik microscope. Appl Optics 2002;41:805–12.
- Bourquin S, Monterosso V, Seitz P, et al. Video-rate optical low-coherence reflectometry based on a linear smart detector array. Opt Lett 2002;25:102–4.
- Hausler G, Lindner MW. Coherence radar and spectral radar - new tools for dermatological diagnosis. J Biomed Optics 1998;D3:21–31.
- Smith LM, Dobson CC. Absolute displacement measurements using modulation of the spectrum of white light in a michelson interferometer. Appl Opt 1981;28:3339–42.
- Jenkins FA, White HE. Fundamentals of optics. New York: McGraw-Hill, 1957:284.
- Choma MA, Sarunic MV, Yang C, Izatt JA. Sensitivity advantage of swept source and Fourier domain optical coherence tomography. Opt Express 2003;11:2183–9.
- Nassif NA, Cense B, Park BH, Pierce MC, et al. In vivo high-resolution video-rate spectral-domain optical coherence tomography of the human retina and optic nerve. Opt Express 2003;12:367.[CrossRef]
- Targowski P, Wojtkowski M, Kowalczyk A, et al. Complex spectral OCT in human eye imaging in vivo. Opt Commun 2004;(1–6):79–84.
- Cense B, Chen TC, Park BH, Pierce MC, de Boer JF. In vivo birefringence and thickness measurements of the human retinal nerve fiber layer using polarization-sensitive optical coherence tomography. J Biomed Opt 2004;9:121–5.[CrossRef][Medline]
- Hee MR, Huang D, Swanson EA, Fujimoto JG. Polarization-sensitive low-coherence reflectometer for birefringence characterization and ranging. JOSAB 1992;9:903–8.
- de Boer JF, Milner TE. Review of polarization sensitive optical coherence tomography and Stokes vector determination. J Biomed Opt 2002;7:359–71.[CrossRef][Medline]
- Faber DJ, Mik EG, Aalders MCG, van Leeuwen TG. Light absorption of (oxy-) hemoglobin assessed by spectroscopic optical coherence tomography. Opt Lett 2003;28:1436–8.[CrossRef][Medline]
- Pircher M, Götzinger E, Leitgeb R, Fercher AF, Hitzenberger CK. Measurement and imaging of water concentration in human cornea with differential absorption optical coherence tomography. Opt Express 2003;11:2190–7.[Medline]
- Yazdanfar S, Kulkarni MD, Izatt JA. High resolution imaging of in vivo cardiac dynamics using color Doppler optical coherence tomography. Opt Express 1997;1:424–31.[Medline]
- Zhao Y, Chen Z, Saxer C, Xiang S, de Boer JF, Nelson JS. Phase-resolved optical coherence tomography and optical Doppler tomography for imaging blood flow in human skin with fast scanning speed and high velocity sensitivity. Opt Lett 2000;25:114–6.[CrossRef][Medline]
- Rollins AM, Yazdanfar S, Barton JK, Izatt JA. Real-time in vivo color Doppler optical coherence tomography. J Biomed Opt 2000;7:123–9.
- He YH, Wang RKK. Dynamic optical clearing effect of tissue impregnated with hyperosmotic agents and studied with optical coherence tomography. J Biomed Opt 2004;9:200–6.[Medline]
- Brezinski M, Saunders K, Jesser C, Li X. Fujimoto J. Index matching to improve optical coherence tomography imaging through blood. Circulation 2001;103:1999–2003.[Abstract/Free Full Text]
- Rao KD, Choma MA, Yazdanfar S, Rollins A, Izatt JA. Molecular contrast in optical coherence tomography using a pump-probe technique. Opt Lett 2003;28:340–2.[Medline]
- Yang C, Choma MA, Lamb LA, Simon JD, Izatt JA. Protein-based molecular contrast optical coherence tomography with phytochrome as the contrast agent. Opt Lett 2004;29:1396–9.[CrossRef][Medline]
- Vinegoni C, Bredfeldt JS, Marks DL, Boppart SA. Nonlinear optical contrast enhancement for optical coherence tomography. Opt Express 2004;12:331–40.[Medline]
- Lee TM, Oldenburg AL, Sitafalwalla S, Marks DL, Luo W, Toublan FJ-J, et al. Engineered microsphere contrast agents for optical coherence tomography. Opt Lett 2003;8:1546–8.
- Jesser CA, Boppart SA, Pitris C, Stamper DL, Nielsen GP, Brezinski ME, et al. High resolution imaging of transitional cell carcinoma with optical coherence tomography: feasibility for the evaluation of bladder pathology. Br J Radiol 1999;72:1170–6.[Abstract]
- Li X, Chudoba C, Ko T, Pitris C, Fujimoto JG. Imaging needle for optical coherence tomography. Opt Lett 2000;25:1520–2.[CrossRef][Medline]
- Zhu Q, Piao D, Sadeghi MM, Sinusas AJ. Simultaneous optical coherence tomography imaging and beta particle detection. Opt Lett 2003;28:1704–6.[Medline]
- Patel NA, Li XD, Stamper DL, Fujimoto JG, Brezinski ME. Guidance of aortic ablation using optical coherence tomography. Int J Cardiovascular Imag 2003;19:171–8.
- Boppart SA, Luo W, Marks DL, Singletary KW. Optical coherence tomography: feasibility for basic research and image-guided surgery of breast cancer. Breast Cancer Res Treatment 2004;84:85–97.[CrossRef][Medline]
- Yang VXD, Gzarnota GJ, Vitkin IA, Kolios M, Sherar M, de Boer J, et al. Ultrasound backscatter microscopy/spectroscopy and optical coherence (Doppler) tomography for mechanism-specific monitoring of photodynamic therapy in vivo and in vitro. Proc SPIE 2002;4612:128–35.[CrossRef]
- Westphal V, Rollins AM, Radhakrishnan S, Izatt JA. Correction of geometric and refractive image distortions in optical coherence tomography applying Fermat's principle. Opt Express 2002;10:397–404.
This article has been cited by other articles:
|
|
|
|
 
R. B. Rosen, M. Hathaway, J. Rogers, J. Pedro, P. Garcia, G. M. Dobre, and A. Gh. Podoleanu
Simultaneous OCT/SLO/ICG Imaging
Invest. Ophthalmol. Vis. Sci.,
February 1, 2009;
50(2):
851 - 860.
[Abstract]
[Full Text]
[PDF]
|
|
|
|
|
|
|
M. G. Ghosn, V. V. Tuchin, and K. V. Larin
Nondestructive Quantification of Analyte Diffusion in Cornea and Sclera Using Optical Coherence Tomography
Invest. Ophthalmol. Vis. Sci.,
June 1, 2007;
48(6):
2726 - 2733.
[Abstract]
[Full Text]
[PDF]
|
|
|
|
|
|
|
K A Goatman
A reference standard for the measurement of macular oedema.
Br. J. Ophthalmol.,
September 1, 2006;
90(9):
1197 - 1202.
[Abstract]
[Full Text]
[PDF]
|
|
|
Top
Abstract
Introduction
