Optical coherence tomography (OCT) [1
] is a non-invasive imaging modality that provides micrometer scale resolution of tissue structures over depth ranges of a few millimeters. The technique has found a number of biomedical applications, most notably in ophthalmic [2
] and cardiovascular [3
] imaging. Fourier-domain OCT (FDOCT) is an improvement to OCT that provides a dramatic sensitivity advantage over traditional time domain techniques [4
]. In FDOCT, the reference arm is held stationary and a spectrally resolved interferometric signal is acquired as function of wavenumber. The sample’s depth can then be retrieved from the Fourier transform of this spectral interferogram. FDOCT can be realized in two ways, either through the use of a broadband source and spectrometer (spectral-domain, or SDOCT) or a frequency swept laser and high bandwidth detector (swept-source, or SSOCT). Both FDOCT techniques suffer from an inherent (sample independent) reduced imaging depth range, typically limited to between 1 and 5mm. Optical attenuation from absorption and scattering in tissue typically limit how much light is recovered from depths beyond a few millimeters, and thus for many applications this inherent reduced depth range is not the limiting factor in determining the practical imaging depth. However, several important OCT applications would benefit from extended imaging depths, including ophthalmic imaging of the anterior segment, intrasurgical imaging, small animal imaging, and catheter imaging of coronary arteries.
Extending the imaging range of FDOCT has thus been an area of interest for which a number of techniques have been developed [7
]. These techniques include phase shifting using a PZT-mounted reference arm [7
] or electro-optic phase modulator [8
], heterodyne SSOCT [9
], instantaneous acquisition of phase separated interferograms using 3x3 interferometers [13
] or polarization encoding [14
], harmonic lock-in detection of phase modulation [15
], imparting a phase ramp across a B-scan with B-M mode scanning [16
] and pivot-offset scanning [17
], sinusoidal phase modulation [20
], rapidly switching between multiple reference arms [21
] and dispersion encoding [22
]. Unfortunately, all of these techniques are accompanied by drawbacks in the form of reduced sensitivity, reduced axial resolution, reduced imaging speed, required lateral oversampling, increased system complexity, increased cost and/or increased signal processing overhead. In addition, most of these techniques produce incomplete suppression of the complex conjugate artifact, resulting in distracting “ghost” images.
Arguably the most effective of these methods is heterodyne complex-conjugate resolved SSOCT (HCCR-SSOCT), which resolves the ambiguity by shifting the peak sensitivity position away from electronic DC, such that positive and negative displacements from that position can be discerned. As this technique shifts, rather than suppresses, the complex conjugate, it completely resolves the artifact. In addition, HCCR-SSOCT does not result in any reduction in imaging speed or require lateral oversampling. In this method, one or two active elements acting as frequency shifters, usually acousto-optic modulators (AOM’s) [9
] (though electro-optic modulators, EOM’s, have been used [11
]) are used to apply a differential modulation frequency between the sample and reference arms. While effective, this technique is limited in that such modulators are expensive and require careful alignment. More significantly, active frequency shifters tend to have appreciable insertion losses, resulting in reduced sensitivity, and restricted optical bandwidth, resulting in spectral distortion and broadening of the axial point-spread function. In addition, processing of the acquired data requires either hardware demodulation [10
] or significant post-processing [9
In this work, we present a method of realizing HCCR-SSOCT using a dispersive optical delay line (D-ODL) as an alternative to AOM’s or EOM’s. This technique confers the same advantages as traditional HCCR-SSOCT in that it doubles the inherent imaging range without sacrificing imaging speed or requiring lateral oversampling. Furthermore, this technique bears no resolution penalty, incurs little to no sensitivity loss, is low-cost and easy to implement, requires no additional signal processing and can be designed to support broad wavelength ranges and arbitrary imaging speeds. As an additional benefit, the D-ODL also allows for hardware dispersion management, which reduces dependence on software dispersion compensation algorithms and adds flexibility to system design.