114 3 Sight Measurement
Figure 3.13. Ophthalmic biometry device
Figure 3.14. Ultrasound a-scan from a normal subject
optic disc, the retinal vasculature and a deeper pigmented area known as the fovea in which the concentration of photoreceptors is greatest and this is the area of fine acuity.
3.7 Optical Coherence Tomography
Optical coherence tomography (OCT) is a new imaging technique which is capable of producing cross sectional images of the retina or cornea with a resolution that surpasses that of conventional imaging techniques.
In principle, OCT is very similar to ultrasonic imaging (Section 3.5), the fundamental difference being that electro-magnetic radiation in the visible or infra-red
3.7 Optical Coherence Tomography |
115 |
Figure 3.15. Ophthalmoscope
Figure 3.16. Digital fundus camera
Figure 3.17. Normal retina
portion of the spectrum is used instead of sound waves. An OCT ‘A-scan’ is obtained by shining a beam from a super-luminescent diode onto the tissue to be imaged. As in ultrasound, the principle of echo delay is then employed to calculate the depth of interfaces within the tissue. A series of adjacent A-scans can be combined to form a B-scan, a two-dimensional cross sectional view through the structure. Successive B-scans may be combined to provide tomographic information on the structure.
116 3 Sight Measurement
3.7.1 Echo Delay
In both ultrasound and OCT, when the incident sound or light wave encounters an interface within the tissue, it is partly reflected back towards the emitter and partly transmitted into the tissue. In ultrasound, the depth of an interface is calculated by measuring the time taken for the reflection, or echo, arising from that interface to travel back to the source. The depth of the interface can then be calculated, provided that the speed of sound in the medium is known. Since the speed of light is several magnitudes greater than sound, measuring the echo time delay of a reflected light wave would require ultra-fast resolution, of the order of femtoseconds. In OCT imaging, this problem is overcome by using low coherence interferometry for high resolution measurements.
3.7.2 Low Coherence Interferometry
Figure 3.18 shows a schematic diagram of an optical interferometer. The light source of wavelength λ directs the incident optical wave Ei(t) towards a beamsplitter. At the beamsplitter, one portion of the beam ES(t) is transmitted into the tissue sample and another portion ER(t) is reflected towards the reference mirror which is at a known spatial position. The beam which is incident on the tissue sample undergoes partial reflection whenever it encounters a structure or surface within a tissue. Thus, the reflected beam travelling back towards the beamsplitter contains multiple echoes from the interfaces within the tissue. The beam incident on the reference mirror is reflected back towards the beamsplitter. These two reflected beams are recombined in the beamsplitter and the resulting beam ER(t) + ES(t) is analysed by the detector.
Figure 3.18. Schematic diagram of an optical interferometer
3.7 Optical Coherence Tomography |
117 |
The intensity IO(t) measured by the detector is proportional to the square of the electromagnetic field. Because of interference effects, the intensity of the output from the interferometer will oscillate as a function of the difference between the path lengths of the reference and specimen beams. From electromagnetic theory, the intensity of the combined beam arriving at the detector is given by
IO(t) = 1/4 [ER]2 + 1/4 [ES]2 + 1/2ERES cos [2 [2π/λ ]δl]
where lR is the distance that light travels in the reference path of the interferometer, lS is the distance that light travels in the measurement path (reflected from the specimen) and δl = lR − lS.
Varying the position of the reference mirror changes the value of lR and hence δl and will cause the two beams to interfere constructively or destructively. The intensity will oscillate between maximum and minimum each time the path length between reference and measurement arms changes by one optical wavelength (as the position of the reference mirror changes by half a wavelength).
If the light beam is coherent, constructive interference will be observed for a wide range of relative path lengths of the reference and measurement arms. However, in optical imaging it is important to know the position of the structures within the specimen precisely, thus light with a short coherence length must be used. With short coherence length light, constructive interference is seen only when the paths travelled by the reference and measurement beams are nearly equal. If the paths are mismatched by more than the coherence length of the light, no interference effects are observed. The coherence length of light therefore determines the spatial resolution of the imaging system. Since short coherence length light is composed of several frequencies of light, it can be characterized by a frequency or wavelength bandwidth. It can be shown from electromagnetic and optical theory that the range resolution can be related to the bandwidth by the following equation:
L = 2 ln2/π λ2/ λ
where L is the ranging resolution, λ is the operating wavelength of the light source and λ is the full width half maximum of the spectral bandwidth.
A typical OCT system for ophthalmic applications uses a superluminescent diode operating with a wavelength of around 850 nm and an optical bandwidth of 30 nm. From the above equation this gives an estimated range resolution of approximately
10μm.
3.7.3An OCT Scanner
In practice, OCT scanners are built using fibre optics. The light source is a superluminescent diode operating in the near infrared region of the electromagnetic spectrum and a short coherence length of around 10 μm. This is coupled directly into an optical fibre leading to the optical fibre coupler which functions as a beamsplitter. The resultant intereference beam is analysed by a photodiode together with signal-processing electronics and computer data acquisition. The resultant