OCT has features which are common to both ultrasound and microscopy. The resolution of clinical ultrasound imaging is typically 0.1 to 1 mm and depends on the sound wave frequency (3 to 40 MHz) used for imaging. High frequency ultrasound has been developed for research and clinical applications such as intravascular imaging.
Resolutions of 15 to 20 μm and finer have been achieved with frequencies of ~100 MHz. However, these high frequencies are strongly attenuated in biological tissues and imaging depths are limited to only a few millimeters. Microscopy and confocal microscopy are examples of imaging techniques which have extremely high resolutions, approaching 1 μm. Imaging is typically performed in an en face plane and resolutions are determined by the diffraction limit of light. Imaging depth in biological tissue is limited because image signal and contrast are significantly degraded by optical scattering. In most biological tissues, imaging can be performed to depths of only a few hundred microns.
OCT fills a gap between ultrasound and microscopy. Current OCT technologies have axial resolutions ranging from 1 to 15 μm, approximately 10 to 100 times finer than standard ultrasound imaging. The high resolution of OCT imaging enables the visualization of tissue architectural morphology. The principal disadvantage of OCT is that light is highly scattered by most tissues and attenuation from scattering limits the image penetration depths to ~2 mm. However, because OCT is an optical technology, it can be integrated with a wide range of instruments such as endoscopes, catheters, laparoscopes, or needles which enable internal body imaging.
The performance of an OCT system is mainly determined by its longitudinal (axial) resolution, transverse resolution, dynamic range (i.e. sensitivity) and data acquisition specifications, including digitization resolution and speed. For application in medical diagnosis, additional factors, e.g. non-contact vs. contact applicability, possible penetration into the investigated tissue, image contrast as well as extraction of functional or biochemical information in addition to the visualization of micro-structural morphology have to be considered. In addition, for clinical applications, compactness, user-friendliness, robustness, flexibility, overall costs of the OCT system, as well as the possibility to interface it to existing diagnostic technology are decisive factors.
In contrast to conventional and confocal microscopy, OCT achieves very high axial image resolutions independent of focusing conditions. The axial and transverse resolutions of OCT are decoupled:
- Axial (depth) resolution – defined by the coherence length of the light source (rather than the depth of field as in microscopy)
- Transverse resolution – defined by the focal spot size
As in conventional microscopy the transverse resolution and the depth of focus are determined by the focused transversal spot size. Increasing the numerical aperture of the objective increases the transverse resolution by reducing the focal spot size, but it decreases the depth of field, quantified by the confocal parameter b (cf. Figure below). Thus, improving the transverse resolution can be accomplished by increasing the numerical aperture (NA) of the objective, but at the same time decreasing b. A solution to this limitation is the use of a dynamic focus tracking system. Especially for ophthalmic retinal OCT imaging, low numerical aperture focusing is employed, because it is desirable to have a large depth of field and to use OCT to achieve high axial resolution.
The interference signal detected at the output of the interferometer is the electric field autocorrelation of the light source. As mentioned before, the full width at half maximum (FWHM) of this autocorrelation is the coherence length lc, which gives the axial resolution Δz and is inversely proportional to the width of the power spectrum. The envelope of this field autocorrelation is equivalent to the Fourier transform of the power spectrum. For a source with a Gaussian spectral distribution, the axial resolution Δz is primarily determined by the coherence length of the optical light source.
Hence high axial resolution may be achieved even with low numerical aperture (NA) beam delivery optics. Since the coherence length of a light source is inversely proportional to its spectral bandwidth, broad-bandwidth optical sources are required to improve the axial resolution in OCT, which are important to detect early changes of various diseases occurring at cellular level. To improve axial OCT resolution, the spectral bandwidth must be either increased and or the center wavelength decreased.
Improving the axial resolution in OCT is challenging and requires the use of highly sophisticated ultrabroad bandwidth light sources. The figure above depicts ‘iso-resolution lines’, i.e. the optical bandwidth (at full-width-at-half-maximum (FWHM)) for a given central wavelength necessary to achieve a desired axial OCT resolution. The iso-resolution lines range from 16 µm to 0.125 µm, measured in free space. For the standard wavelength region used for OCT retinal imaging (800 nm), assuming a Gaussian spectrum of the light source as well as non-dispersive imaging medium, this figure shows that for 1 µm axial resolution, an optical bandwidth of ~200 nm is needed. It also illustrates that in order to achieve the same axial resolution at 1300 nm, more than 500 nm of optical bandwidth (FWHM) is needed, i.e. more than two times more, compared to what is needed at 800 nm. Finally, in the visible wavelength region (500 nm), only 75 nm is need to achieve 1 µm axial OCT resolution. This demonstrates the significant wavelength dependence and the challenge to balance resolution versus penetration/contrast depending on the particular OCT application.
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