Optical coherence tomography with the "Spectral Radar" - Fast optical analysis in volume scatterers by short coherence interferometry

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1 Optical coherence tomography with the "Spectral Radar" - Fast optical analysis in volume scatterers by short coherence interferometry M. Bail, G. Häusler, J. M. Herrmann, M. W. Lindner, R. Ringler Physics Institute, University of Erlangen-Nürnberg, Staudtstrasse 7/B2, Erlangen, Germany ABSTRACT We present a sensor for acquisition of cross-sectional images of volume scatterers (e.g. biological tissue), we call it "spectral radar". Medical and technical applications are possible. The sensor is a modified Michelson-interferometer, with a broad bandwidth light source. The scattering amplitude a(z) along one vertical axis from the surface into the bulk can be measured with-in one exposure. No reference arm scanning is necessary. Measurement results of stationary and non stationary scattering phantoms, human skin and of a fish eye in vitro are shown. Optical tomography, white light interferometry, scattering media, cross-sectional imaging, biological tissue. 1. INTRODUCTION An important medical aim is the early diagnosis of pathological tissue alterations (e.g. skin cancer). High resolution imaging methods not harmful to the healthy tissue are necessary. During the last years noninvasive cross-sectional optical imaging methods under the heading "optical tomo graphy" have been developed. An overview about perspectives of optical tomography is given in Ref. 1. The sensing methods are manifold, but are based on one common feature: the tis sue (which is a volume scatterer) is illuminated, and it is measured how many photons are scattered back to the detector, dependent on the path length in the tissue. This gives, to a certain extent, access to the scattering amplitude a(z), which is the key to exa mine the local scattering and absorption behavior in the tissue. The hope is that pathological tissue displays significant scattering properties and can therefore be separated from healthy tissue. Measurement of path length distribution of the photons can be done directly by time of flight measurements or utilizing short coherence interferometry. Optical coherence tomography 2 (OCT) uses a broad bandwidth light source in an interferometrical setup, interference is detected only, if the object path length equals the reference path length. The path length of the photons to be detected (backscattered from the tissue) can be adjusted by the reference path length. The main advantage of OCT is, that the path length resolution is roughly given by the coherence length of the light, which can be in the µm range. However, the spatial resolution of internal structure in strongly scattering biological tissue will be worse, due to multiple scattering of the photons 3. One dis advantage of the conventional OCT-setup is that the reference path has to be scanned through the depth range. This is often time consuming. But there are new OCT-setups 4, which use piezo-electric fiber stretching and obtain fiber length variations at rates of some m/s. During the last few years we developed several interferometrical methods for shape measurements of rough surfaces, called coherence radar 5,6,7. These methods were based on a modified Michelson interferometer. On basis of these sensors we developed the "spectral radar" 8,9, which is specifically suited for morphological analysis of volume scatterers. We use the spectral radar for cross-sectional imaging of the internal structure of strongly scattering phantoms and biological tissue. A similar principle has been suggested for measurements in the weakly scattering eye 10.

2 2. MEASUREMENT PRINCIPLE The spectral radar measures the scattering amplitude a(z) along one axis from the surface into the bulk within one detector exposure, no scanning in depth is necessary, hence a short measurement time is possible. Non stationary processes can be monitored. For two dimensional imaging a transverse scanning is necessary. The sensor is a modified Michelson-Interferometer. A broad bandwidth superluminescent diode (SLD) is used as a light source (fig.1), instead of a SLD also a wavelength-scanning narrow band light sourse could be used. The SLD is imaged onto the object surface and onto a reference mirror. Elementary waves are scattered back from different depths of the object. These waves interfere with the plane reference wave. Light is spectrally separated by a grating spectrograph and detected with a high dynamic linear detector (1024 photodiodes). Fig.1 Spectral radar, basic setup. The measured detector signal I, a function of wave number k, can be calculated from a simplified approach: we assume that the photons, we measure, are scattered only once. Only then we can be sure that the photons are really scattered in the depth z (the half of their total path length). This is a good approximation for scatterers, which exhibit a forward peaked scattering characteristic (gfactor near +1). The depth should be only a few mm and a the scattering coefficient should be not much larger than 10mm -1. Then the object signal o (i.e. the signal re-emitted from the scatterring media) can be regarded as a sum of elementary waves emerging from depth z with (scattering) amplitude a(z). s(k) is the spectral amplitude dis tribution of the light source (k=2π/λ, λ is the wavelength), n is the refractive index

3 of the scatterer. iknz o = s( k ) a( z ) e dz ( 1) 0 The reference wave is a plane wave: r = s( k ) e ikr ( 2) R is the reference path length. Reference and object wave are superimposed, and at the detector we get the intensity I(k): 2 I( k) = s( k) a( z)cos( 2knz) dz +... ( 3) 0 The valuable term of equation (3) describes a Fourier transformation of a symmetrized version ã(z): ã(z) =½(a(z)+a(-z)). Hence, we can retrieve a(z) by inverse Fourier transformation of I(k). This can be compared with the Fourier diffraction theorem described by Wolf 11. The equation is valid for single scattering and a constant refractive index. As in the experiment, the reference is placed virtually onto the object surface (R=0). I(k) is a sum of cosine functions with varying frequency and amplitude a(z). For example, if all light is backscattered from the same depth in the sample, I(k) would contain only one frequency, in interferometry this is known as Müller fringes 12. An analysis of the frequency content of I(k) is possible by Fourier transformation. Inverse Fourier transformation of I(k) delivers the scattering amplitude ã(z). The high frequency content of the detected signal represents larger depth. One technical problem is that already for a depth of only a few mm the frequency is rather high. For z=2mm one period in the spectrum is about 0.2nm (λ=850nm). Hence, a spectrograph with high resolution is required. z max = ( 4 ) 4n λ What is the maximum depth zmax, that can be measured with a spectrometer resolving λ? From Eq.(3) it can be seen that for z=zmax, the period of the cosine fringes is: k=π/nzmax. The spectrometer has to resolve at least k/2. With k = 2π λ/λ 2, we get: zmax is proportional to the "coherence length" of a virtual source with spectral width λ. Our SLD has a spectral half width of 30nm and a mean wavelength of 853nm. The light from the interferometer exit illuminates via a fibre the entrance slit of a high resolution grating spectrograph. The spectrograph resolves λ=0.05nm, therefore zmax=2.4mm (n=1.5). The spectrum is detected by a cooled photodiode array with 16bit dynamical range (SNR is about 10 4 ). 3. EXPERIMENTAL RESULTS Figure 2 shows measurement results of a dynamic phantom: two metal pins (diameter 160µm, distance center-center 320µm) in a vessel filled with an emulsion of milk/water.

4 Fig.2 Measurement of two metal pins (diameter 160µm, distance 320µm, depth 670µm ) in milk/water emulsion (25% milk) with the spectral radar. The concentration of the milk is about 25%, the scatterting properties of this dynamic phantom are similar to skin. We focus onto the pins into the emulsion and shoot several depth scans, while moving the object laterally in steps of 20µm. The spectrally separated interferograms are evaluated by Fourier transformation. The measured scattering amplitude a(z) versus depth z and lateral coordinate x is displayed in Fig. 2. The signal from the metal pins (depth 670µm) can be clearly detected and the pins are resolved. Despite of the strong scattering medium the resolution of the sensor is high. This is surprising. But the scatterers around the pins are moving. Therefore scattered light becomes incoherent over the integration time of our detector. Then the scattered light contributes only to frequency zero after Fourier transformation of the detector signal. It is possible to separate the multiple scattered light of the milk/water emulsion from the light scattered from the stationary metal pins. In a second experiment we measure the thickness of layers in a multilayer phantom (given to us for experimental use from Prof. Rinneberg, Physikalisch Technische Bundesanstalt, Berlin). The phantom is built up from 3 homogeneous 0.5mm layers of different scattering coefficients µs: 21cm -1, 86cm -1, 151cm -1 (850nm), g 0.8 (800nm). A lateral scan across the surface shows the profile of the surface and of the boundary layer 1 / layer 2 (fig.3). This boundary in the depth of 0.5mm was measured with a longitudinal uncertainty of about 13µm. The longitudinal rms measurement uncertainty of the surface is 3µm. This quite good accuracy is also interesting for many technical applications, where a rough surface of a diaphanous object has to be measured. The ratio of the scattering amplitudes for the surface peak and for the boundary layer 1 / layer 2 is 500:1. This shows that high SNR is absolutely necessary.

5 Fig. 3 Measurement of a multilayer phantom. The boundary between the first two layers with different scattering coefficients could be accurately localized; the longitudinal rms measurement uncertainty is σ= 13µm. The uncertainty for the surface is σ=3µm. As a biological sample we measured human skin (from hand and back) in vitro. The fresh skin was placed in a vessel filled with an isotonic NaCl solution. We call the obtained cross-sectional images optograms, because the sensor is sensitive to the optical parameters inside the measured sample. A comparison (fig. 4) of the optograms, of the spectral radar with the histological images of the skin shows that the thickness of the stratum corneum could be measured (thickness at the hand about 160µm, at the back about 50µm). Small differences are caused by the histological preparation, which leads to a skrinking of the sample of 10%-30%.

6 Fig. 4 Comparison of "optograms", measured with the spectral radar, of hand 4a) and back 4c) together with the histological images of hand 4b) and back 4d). The thickness of the stratum corneum was measured. As another biological object, we measured the small fish eye (brocade-barbel) in vitro (fig.5). The eye was embedded in isotonic NaCl solution. Cornea and iris can be clearly seen.

7 4. CONCLUSIONS For the spectral radar we expect the same resolution as for reference arm scanning OCT. An accurate comparision is difficult, because the results depend stronly upon the scattering coefficient of the sample and the measured depth. The main advantage of the spectral radar is that no scanning in depth is necessary, measurement time can be short, in principle. With the first laboratory setup of the spectral radar we could detect and measure structures in different samples: a dynamic object (pins in milk), a stationary multilayer phantom (variation of scattering coefficient) and biological samples (skin, fish eye). Unfortunately, present SLDs deliver only a few mw. This limits the SNR. With a stronger light source we expect to measure deeper into the media. More medical applications are being investigated. 5. ACKNOWLEDGEMENTS The research reported here was supported by the BMBF, registration 13N6301. Fig. 5 Optogram of a fish eye (brocade-barbel). 6. REFERENCES 1. D. Benaron, G. Müller, B. Chance, A Medical Perspective at the Threshold of Clinical Optical Tomography, in Medical optical tomography: functional imaging and monitoring, G. J. Müller, ed., 3-9, SPIE (1993) 2. D. Huang, E. A. Swanson, C. P. Lin, J. S. Schuman, W. G. Stinson, W. Chang, M. R. Hee, T. Flotte, K. Gregory, C. A. Puliafito, J. G. Fujimo to, Optical Coherence Tomography, Science 254, (1991) 3. G. Häusler, J. M. Herrmann, R. Kummer, M. W. Lindner, "Observation of light propagation within volume scatterers with fold slow motion", Opt. Lett. 21 (1996) 4. S. A: Boppart, G. J. Tearney, B. Bouma, J. G. Fujimoto, M. E. Brezinsky, Optical coherence tomography of developing embryonic morphology CLEO 96, (1996) 5. G. Häusler, German Patent DE (1991) 6. T. Dresel, G. Häusler, H. Venzke, 3D-sensing of rough surfaces by coherence radar, Appl. Opt. 31, (1992) 7. G. Häusler, J. Neumann, Coherence radar - an accurate 3D-sensor for rough surfaces, in Optics,

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