# Endoscopic Optical Imaging Technologies and Devices for Medical Purposes: State of the Art
## 2 Medical Imaging Applications
### Optical Coherence Tomography  
The working principle of OCT is analogous to that of an ultrasound, with the difference that low coherent light is used instead of sound. Laser light is transmitted to the semi-transparent tissue surface using an optical fiber or other means discussed later in the paper. The light backscattered from the tissue is monitored to reconstruct a cross-sectional image of the targeted tissue. The differential backscattered intensity from different tissue interfaces (refractive index changes) provides the contrast in OCT [7]. The speed of light in tissue (225 m/\(\upmu\)s) is about five orders of magnitude larger than the speed of sound in tissue (150 \(\times\) 10\({}^{-5}\) m/\(\upmu\)s). The resolution of an image is directly proportional to the wavelength of the used signal. Smaller wavelength light provides a finer resolution (10 to 100 times) image as compared to that of ultrasound [8]. On the other hand, the time delay between the source signal and the reflected signal is much smaller and cannot be measured directly. Thus, an interference pattern must be measured either using a Michelson or a Mach-Zehnder interferometer [9]. The light coming from the laser source is split into two paths: one going towards the sample (called sample arm), and the other going towards a mirror (reference arm). The backscattered light from the tissue is compared with the light from the reference arm to form an interference pattern if the two light components are within the coherence length of the light source [10]. A schematic diagram of an OCT scanner is shown in Figure 1.  
Scanning can be performed in a time domain (TD) or in a Fourier domain (FD). In TD-OCT, a low-coherence broadband light source provides light to the two arms through an optical splitter. The path length of the reference arm is changed in time by moving the reference mirror in the axial direction. The interference pattern between the reflected light from the reference arm mirror (with distance travelled variable in time) and the backscattered light from sample provides a depth profile of the sample tissue. The two-or three-dimensional structure image is obtained by transversally scanning the light beam along the sample arm [10; 11]. In FD-OCT, the reference arm is static. Axial scanning can be performed in time using a frequency sweep of the light source to cover the broad bandwidth (swept-source OCT, or SS-OCT), or in space using a spectrometer to separate the different wavelengths on the detection side (spectral domain OCT or SD-OCT). Depth information of the sample is obtained using the inverse Fourier transformation. In SS-OCT, the light source is a rapidly tunable laser, and the detector is a simple photodiode. In SD-OCT,  
Figure 1: The schematic diagram of an optical coherence tomography (OCT) scanner. the light source is a low-coherence laser, and the detector is a line scan camera. The elimination of the mechanically moving mirror in FD-OCT allows the sample to image with higher acquisition speeds, which is the main reason that most of the current OCT systems are FD-OCT [10; 11]. Analogous to ultrasound, it is also possible to use a doppler OCT (DOCT) to measure the velocity of the scattering materials flowing through a channel, such as a blood vessel [10]. In OCT, the lateral resolution depends on the lens system as:  
\[(\delta x,\ \delta y)=0.56\frac{\lambda}{NA}, \tag{1}\]  
with \(\lambda\) being the central wavelength of the light spectrum, and \(NA\) is the numerical aperture. The lateral field of view (FOV) is given by:  
\[FOV_{x,y}=\frac{f\ tan\theta}{2}, \tag{2}\]  
where \(f\) is the focal length of the lens, and \(\theta\) is twice the scanning angle [10]. The axial resolution depends on the light source used by:  
\[\delta z=\frac{2 ln2}{\pi}\frac{\lambda^{2}}{\Delta\lambda}, \tag{3}\]  
where \(\Delta\lambda\) is the full width half maximum of the light spectrum.