Patent Application: US-58749005-A

Abstract:
an arrangement and method for imaging and / or measuring tissue qualities , such as tissue thickness , tissue surface roughness and degree of tissue fiber linearization . the arrangement includes at least one light generating element , at least one light detecting element , a probe with an extension , and possibly a control apparatus including a signal processor for processing the detected signals and / or images . the extension is designed to convey light from the light generating element to the tissue for visualization and / or measurement . the extension is also designed to convey light back - scattered from the tissue to the light detecting element . the detecting element is designed to measure the intensity and / or spatial distribution of light back - scattered from the tissue .

Description:
twelve hip joint condyles from bovine calves were obtained from a local slaughterhouse less than 24 hours after sacrifice . two of the condyles were used for reference measurements and the other ten for thickness experiments . the condyles were stored in saline in a refrigerator and prepared for cartilage measurements through the removal of soft tissues and tendons surrounding the joint . three sites on each condyle surface were used for the measurements . a handheld , rotating , grinding machine was used to reduce the cartilage layer thickness . sandpaper with the roughness p100 was used for grinding . care was taken to grind in short episodes ( 5 - 15 s ) so as not to increase the temperature of the cartilage . thickness measurement of the cartilage layer was done with a high - resolution ultrasound scanner ( b - mode 20 mhz , dermascan 3v3 , cortex technology , hadsund , denmark ). the probe scanned over the measurement site and an image of the cartilage / bone interface was presented on the computer screen . optical reflection spectra were recorded by using an oriel instaspec iv ccd spectrometer equipped with an ocean optics broad spectrum tungsten lamp hl 2000 ( spectral range 360 - 2000 nm ). the light was guided by optical glass fiber bundles ( na = 0 . 35 ) and the measurements were taken at a small distance ( 2 - 5 mm ) to the condyle surface . the bundles were arranged in a probe head ( diameter 4 mm ) with the emitting fiber bundle encircling the detecting bundle . the reflection spectra were calculated according to the formula : i ⁡ ( λ ) = i tissue - i background i reference - i background ( 1 ) where i tissue is the raw spectrum of the examined tissue , i background the detector background signal , and i reference the diffuse reflectance spectrum taken from a white reference ( baso 4 ). for each measurement position , spectra and ultrasound images were recorded from the intact cartilage layer , for 4 - 5 intermediate cartilage thicknesses ( obtained by grinding ) and when bone level had been reached . twenty - three pieces of pure cartilage , about 1 mm thick , were removed from the two joints by using a sharp knife . remains of subchondral bone were carefully removed to secure a pure cartilage sample . reflection spectra were measured , with the equipment described above , for each piece of cartilage placed on a black plastic sheet . the joints were cut in half , washed in saline and stored ( in saline ) for a few days to remove remains of blood . finally , reflection spectra were taken from 10 positions on the exposed bone samples . mean reference spectra for cartilage ( n = 23 ) and bone ( n = 10 ) were calculated and will be referred to as s cartilage and s bone , respectively . furthermore , the reference spectrum of blood ( s blood ) was estimated as the inverse absorption spectrum of oxyhemoglobin , taken from the literature ( takatani and graham , 1979 ). to decrease the influence of remains of blood in the bone , s bone was adjusted by subtracting s blood until the characteristic hemoglobin peaks could no longer be distinguished . each measured reflectance spectrum was matched to the true cartilage thickness ( d ), as determined from the stored ultrasound images . for the thickness determination , the manufacturer &# 39 ; s software including a cursor system was used . the resolution of the ultrasound image was 0 . 06 mm . as a measure of cartilage thickness , determined from the spectroscopic data , d spec was defined as the percentage contribution (%) of cartilage spectrum in the measured reflectance spectrum : d spec = a a + b + c ( 2 ) where a , b and c are the coefficients for optimal match ( least square fitting ) between the measured spectrum ( s measured ) and the reference spectra according to : an exponential regression model ( 4 ) was used for statistical comparison between reference cartilage thickness and d spec . d spec = k 1 ( 1 − e − k 2 · d ) ( 4 ) in order to test the assumption of an exponential relation between the cartilage thickness d and the corresponding spectroscopic data , a monte carlo model ( de mul et al ., 1995 ) was used ( fig1 ). the model has a cartilage layer of known thickness ( d ) and diffusion theory optical properties ( μ ac , μ sc , g c ) positioned on top of a semi - infinite layer of bone containing blood ( μ ab , μ sb , g b ). the optical properties of cartilage , bone and blood were taken from the literature ( beek et al ., 1997 ; firbank et al ., 1993 ; tuchin , 2000 ) for the single wavelength 633 nm . the optical properties of bone containing blood were set to the bone coefficients increased by those of blood at a selected perfusion level ( 10 %) according to : all optical parameters are presented in table 1 . refractive indices of both layers were set to 1 as specular effects were not of interest . the pathways of 10 6 photons , incident in a point at the cartilage surface , were calculated ( de mul et al ., 1995 ). the back - scattered photons reaching a ring shaped detector ( outer radius 3 mm , inner radius 1 mm ) were counted . simulations were performed for cartilage thicknesses d = 0 - 3 mm in steps of 0 . 1 mm . as a model , we used collagen on glass plates . two samples were polymerized , one inside a strong magnetic field ( giving rise to a linearized sample ) and another outside of the field ( reference sample with randomized fiber directions ). the irradiance transmitted through the samples was studied . a polarized hene laser ( 633 nm ) was used as a light source . transmitted irradiance was recorded during rotation of the linearized sample in 10 ° steps between 0 - 360 °. this gave rise to a variation in transmitted intensity due to the relative difference between the collagen fiber direction of the sample and the laser polarization plane . the procedure was repeated using the reference sample . the mean reflection spectra of the reference material are presented in fig2 . the cartilage spectrum appears relatively “ white ” with a shift towards the blue region , whilst the bone spectrum appears distinctly “ red ” and includes the characteristic absorption peaks of hemoglobin at 542 nm and 576 nm . in the thickness analysis , the latter spectrum is divided into s bone and s blood ( see equation 3 ). the mean (± sd ) thickness of intact cartilage was 1 . 21 ± 0 . 30 mm ( n = 30 ). a typical example of spectra from a grinding sequence is shown in fig3 . as the cartilage layer gets thinner a clear influence of bone can be seen . the spectroscopic estimation of cartilage thickness ( d spec ) is plotted against the ultrasound reference cartilage thickness in fig4 . the regression model of equation 4 is used ( r = 0 . 69 , p & lt ; 0 . 000001 , s = 0 . 167 , k 1 = 0 . 75 , k 2 = 3 . 81 , n = 182 ). for thinner cartilage layers ( d & lt ; 0 . 5 mm ), the model mean error is 0 . 19 ± 0 . 17 . the monte carlo simulation results are presented in fig5 . this result is similar to the experimental results in fig4 , with more photons reaching the detector at a thicker cartilage layer . the monte carlo simulation results support the assumption of an exponential relationship between cartilage thickness and the spectroscopic data . in the measurements on polymerized collagen , polarized laser light transmission for the linearized sample showed a clear relationship to the angle between laser light polarization plane and collagen fiber direction , with maximum transmission when the two were in parallel . the reference sample showed no such relationship . the possibility to extract objective information about cartilage thickness by studying the reflectance spectrum from the cartilage surface , and the possibility to obtain information on collagen fiber linearization by studying polarized light transmission . these results implicate that it is possible to use a minimally invasive technique to characterize cartilage in connection with in situ diagnosis . the reflectance spectrum from a condyle surface can be seen as a sum of spectra from cartilage and subchondral bone ( containing blood ). typical spectra from cartilage and bone can be seen in fig2 . cartilage contains relatively few cells , which occupy 10 - 20 % of its volume . the remainder is extracellular material , which is highly hydrated and contains up to 80 % water by weight . the material consists primarily of large hydrated proteoglycan aggregates , entrapped within a matrix of collagen fibrils . this fiber structure and the fact that absorption of water in the investigated wavelength region is low , gives reason to believe that the character of the cartilage spectrum is an effect of reduced scattering at longer wavelengths according to the mie theory . consequently , the cartilage layer acts as a diffuse reflector for photons , impeding them from reaching the highly absorbing subchondral bone . we intended to assess cartilage thickness by studying the relative content of cartilage and bone components in the combined reflectance spectrum . by using this approach , the accuracy of cartilage thickness determination only becomes dependent upon the variability of the optical properties of these components . this variability remains to be investigated for a larger human material , after which it also will be possible to model the behavior for the complete spectrum and not just for a single wavelength . the exponential relationship between cartilage thickness and diffuse reflectance was expected ( nötzli et al ., 1989 ). a tentative source of error could be the variability of perfusion in the underlying bone . flow rates in ( rabbit ) tibial and femoral cortical bone vary in a physiological range of 1 . 6 - 7 . 0 ml · min − 1 · g − 1 ( shepherd and öberg , 1990 ). increased perfusion / blood content probably leads to increased absorption by the hemoglobin of the bone . some of the data scattering in fig4 may be due to the resolution of the ultrasound reference system ( 0 . 06 nm according to the manufacturer &# 39 ; s specifications ). this value can be compared to the model mean error for thinner cartilage layers 0 . 19 ± 0 . 17 mm . the accuracy of the ultrasound reference method depends on the ultrasound speed in cartilage ( jurvelin et al ., 1995 ). without a calibration of the device to cartilage ultrasound speed data we may have some spreading in data . grinding causes roughening of the cartilage surface , affecting the degree of specular reflection . however , the specular reflection can be considered wavelength independent , not affecting the thickness estimation , solely based on spectral distribution . for the same reason , the measurement distance was not precisely controlled . to achieve maximal penetration depth , we chose to record the spectrum between 330 - 835 nm . it is reasonable to assume that a few characteristic wavelengths can be found that can be used for thickness calculations , thereby eliminating the need for recording the complete spectrum . such an approach can facilitate the design of an instrument based on the presented principle . an attractive feature of this principle is that it is based on fiber optics . thus , it can easily be “ imbedded ” in an arthroscope , which , in addition to the visual assessment of the cartilage surface can give quantitative information about the thickness of the cartilage layer under study . based on the polarized light transmission results , other important aspects of cartilage quality , such as the linearization and surface roughness , seem possible to extract using optical sensing in an arthroscope . we found a large variation in the thickness of the bovine hip joint cartilage ( 0 . 67 - 1 . 98 mm ). the same variation can be found in human hip joint cartilage ( 1 . 14 - 2 . 84 mm ), depending on where on the joint the cartilage is measured ( nakanishi et al ., 2001 ). there are reasons to believe that hip joints or smaller joints can be assessed through the sterile introduction of a fiber optic bundle but the most interesting application for this new principle may be the assessment of the knee via arthroscopy . the cartilage thickness of healthy and osteoarthritic human knee joints varies in the range of 0 . 5 to 7 . 4 mm ( kladny et al ., 1999 ). with the present method the hemoglobin absorption peaks could often be seen for thicker cartilage layers ( fig3 ) but a clear spectral effect occurred at cartilage thicknesses below 0 . 5 mm . in a well - perfused bone the variation and sensitivity of the method may be improved . penetration depth may also be improved by focusing on the diffuse reflection component , by using more efficient optical components and by geometrical separation of light source and detector . blood perfusion of bone has , for instance , been measured at 3 . 5 mm penetration depth , including penetration of a 1 mm thick cartilage layer , using a 632 . 8 nm laser doppler technique ( nötzli et al ., 1999 ). after studying bovine hip joint condyles and polymerized collagen , it was found that information about cartilage thickness and collagen fiber structure can be extracted using optical reflectance spectroscopy and polarized light transmission , respectively . for thicker cartilage layers , a high reflection for the wavelengths 400 - 600 nm is seen , and for thinner cartilage layers , the characteristic spectra of blood and bone dominate . the opposite is seen for wavelengths in the near infrared region . linearized collagen shows higher transmission when the polarization plane of the incoming light is parallel with the fiber direction . consequently , the optical reflectance spectrum may be used to characterize cartilage , and specifically cartilage thickness , in connection with in situ diagnosis . furthermore , polarization measurement can be used for studying cartilage collagen linearization . imaging by using white light illumination and signal processing for creating images with enhanced contrast between tissues with different optical properties . the solution is presented in fig6 and fig7 . white light is supplied by light source 1 ( 8 ) emitting light into fiber bundle 1 ( 7 ). the light source is driven by a light source driver unit ( 11 ), controlled by a control unit ( 10 ) in a control apparatus ( 6 ). light source 1 ( 8 ) could be a broad band tungsten lamp and fiber bundle 1 ( 7 ) could consist of high aperture optical glass fibers . fiber bundle 1 ( 7 ) is passing through a channel in the extension ( 3 ) of the probe ( 1 ) and supplies illumination of the measurement object via a lens ( 2 ). this light serves both as illumination for investigation and measurement . in another solution , the fiber bundle 1 ( 7 ) is mounted on the extension ( 3 ) in a removable fashion . light reflected from the measurement object is collected via the lens ( 2 ) and projected onto a ccd camera ( 5 ) via an adapter ( 4 ). a ccd camera with enhanced sensitivity in the near - infrared wavelength region could be used . after manual input via an input device ( 9 ) in the control apparatus ( 6 ), requesting measurement to start , the ccd image is read by the control apparatus ( 6 ) and processed in a signal processor ( 13 ). the reading could be made continuous . according to the theory presented above , the image should be processed as to increase wavelength information where differences in absorption between the tissue components are seen . these wavelengths could correspond to hemoglobin absorption peaks ( 5 - 10 nm in the vicinity of 425 , 542 or 576 nm ) or water absorption ( 1000 - 2000 nm , preferably 1300 nm ). contrast enhanced images of the measurement object are presented on a display unit ( 12 ) in the control apparatus ( 6 ). the input device ( 9 ), control unit ( 10 ), signal processor ( 13 ) and / or display unit ( 12 ) can be that of a computer . imaging by discrete wavelength and / or polarized light illumination for creating images with enhanced contrast between tissues with different optical properties and / or between tissues with a linearized and normal fiber structure . the solution is presented in fig7 and fig8 . light source 2 ( 14 ) emits white light for illumination for investigation and a number of discrete wavelengths for measurement . the white light source could be a broad band tungsten lamp and the discrete wavelength light sources could be light emitting diodes . wavelengths emitted from said discrete wavelength light sources could correspond to hemoglobin ( 5 - 10 nm in the vicinity of 425 , 542 or 576 nm ) or water absorption ( 1000 - 1600 nm , preferably 1300 nm ). in one solution , light source 2 ( 14 ) is replaced by a single white light source . an optical filter extracts discrete wavelengths for measurement and the unfiltered white light is used for illumination for investigation . the optical filter could be a liquid crystal filter . light source 3 ( 15 ) with a polarization filter ( 16 ) is used for a separate channel . this light source could be a broad band tungsten lamp . in one solution light source 2 ( 14 ) is used for both channels and light source 3 ( 15 ) becomes redundant . both light source 2 ( 14 ) and light source 3 ( 15 ) are driven by a light source driver unit ( 11 ), controlled by a control unit ( 10 ) in a control apparatus ( 6 ). the light from light source 2 ( 14 ) is emitted into fiber bundle 1 ( 7 ). this fiber bundle could consist of high aperture optical glass fibers . the light from light source 3 ( 15 ) is emitted into fiber bundle 2 ( 17 ). this fiber bundle could consist of polarization maintaining optical glass fibers . in another solution high aperture optical glass fibers are used and the polarization filter ( 16 ) is positioned at the distal end of the extension ( 3 ) of the probe ( 1 ). both fiber bundle 1 ( 7 ) and fiber bundle 2 ( 17 ) pass through a channel in the extension ( 3 ) of the probe ( 1 ) and supply illumination of the measurement object via a lens ( 2 ). the light from fiber bundle 1 ( 7 ) serves both as illumination for investigation and measurement . in another solution , fiber bundle 1 ( 7 ) and / or fiber bundle 2 ( 17 ) is mounted on the extension ( 3 ) of the probe ( 1 ) in a removable fashion . light reflected from the measurement object is collected via the lens ( 2 ) and projected onto a ccd camera ( 5 ) via an adapter ( 4 ). a ccd camera with enhanced sensitivity in the near - infrared wavelength region could be used . after manual input via an input device ( 9 ) in the control apparatus ( 6 ), requesting measurement to start , the ccd image is read by the control apparatus ( 6 ) and processed in a signal processor ( 13 ). the reading could be made continuous . according to the theory presented above , images enhancing contrast between tissue components with different optical properties are obtained during discrete wavelength illumination by light source 2 ( 14 ), and images enhancing contrast between tissues with a linearized and normal fiber structure are obtained during polarized light illumination by light source 3 ( 15 ). contrast enhanced images of the measurement object are presented on a display unit ( 12 ) in the control apparatus ( 6 ). the input device ( 9 ), control unit ( 10 ), signal processor ( 13 ) and / or display unit ( 12 ) can be that of a computer . visualization based on discrete wavelength and / or polarized light illumination for enhancing contrast between tissue components with different optical properties and / or between tissues with a linearized and normal fiber structure . the solution is presented in fig7 and fig9 . light source 2 ( 14 ) emits white light for illumination for investigation and a number of discrete wavelengths for measurement . the white light source could be a broad band tungsten lamp and the discrete wavelength light sources could be light emitting diodes . wavelengths emitted from said discrete wavelength light sources could correspond to hemoglobin ( 5 - 10 nm in the vicinity of 425 , 542 or 576 nm ) or water absorption ( 1000 - 1600 nm , preferably 1300 nm ). in one solution , light source 2 ( 14 ) is replaced by a single white light source . an optical filter extracts discrete wavelengths for measurement and the unfiltered white light is used for illumination for investigation . the optical filter could be a liquid crystal filter . light source 3 ( 15 ) with a polarization filter ( 16 ) is used for a separate channel . this light source could be a broad band tungsten lamp . in one solution light source 2 ( 14 ) is used for both channels and light source 3 ( 15 ) becomes redundant . both light source 2 ( 14 ) and light source 3 ( 15 ) are driven by a light source driver unit ( 11 ), controlled by a control unit ( 10 ) in a control apparatus ( 6 ). the light from light source 2 ( 14 ) is emitted into fiber bundle 1 ( 7 ). this fiber bundle could consist of high aperture optical glass fibers . the light from light source 3 ( 15 ) is emitted into fiber bundle 2 ( 17 ). this fiber bundle could consist of polarization maintaining optical glass fibers . in another solution high aperture optical glass fibers are used and the polarization filter ( 16 ) is positioned at the distal end of the extension ( 3 ) of the probe ( 1 ). both fiber bundle 1 ( 7 ) and fiber bundle 2 ( 17 ) pass through a channel in the extension ( 3 ) of the probe ( 1 ) and supply illumination of the measurement object via a lens ( 2 ). in another solution , fiber bundle 1 ( 7 ) and / or fiber bundle 2 ( 17 ) is mounted on the extension ( 3 ) of the probe ( 1 ) in a removable fashion . light reflected from the measurement object is collected via the lens ( 2 ) and viewed by the observer through an eye - piece ( 18 ) via an adapter ( 4 ). measurement is started via an input device ( 9 ) in the control apparatus ( 6 ). the reading could be made continuous . according to the theory presented above , images enhancing contrast between tissue components with different optical properties are obtained during discrete wavelength illumination by light source 2 ( 14 ), and images enhancing contrast between tissues with a linearized and normal fiber structure are obtained during polarized light illumination by light source 3 ( 15 ). the input device ( 9 ) and / or the control unit ( 10 ) can be that of a computer . in this solution ( technical solution 3 ) the signal processor ( 13 ) and display unit ( 12 ) of the control apparatus ( 6 ) become redundant . imaging by discrete wavelength and / or polarized light illumination for creating images with enhanced contrast between tissue components with different optical properties and / or between tissues with a linearized and normal fiber structure , combined with point - wise measurement of tissue layer thickness . the solution is presented in fig7 and fig1 . light source 2 ( 14 ) emits white light for illumination for investigation and a number of discrete wavelengths for measurement . the white light source could be a broad band tungsten lamp and the discrete wavelength light sources could be light emitting diodes . wavelengths emitted from said discrete wavelength light sources could correspond to hemoglobin ( 5 - 10 nm in the vicinity of 425 , 542 or 576 nm ) or water absorption ( 1000 - 1600 nm , preferably 1300 nm ). a reference wavelength where similar absorption between the tissue components is seen should be used . said reference wavelength could be in the 600 - 800 nm region , preferably 630 nm . in one solution , light source 2 ( 14 ) is replaced by a single white light source . an optical filter extracts discrete wavelengths for measurement and the unfiltered white light is used for illumination for investigation . the optical filter could be a liquid crystal filter . light source 3 ( 15 ) with a polarization filter ( 16 ) is used for a separate channel . this light source could be a broad band tungsten lamp . in one solution light source 2 ( 14 ) is used for both channels and light source 3 ( 15 ) becomes redundant . both light source 2 ( 14 ) and light source 3 ( 15 ) are driven by a light source driver unit ( 11 ), controlled by a control unit ( 10 ) in a control apparatus ( 6 ). the light from light source 2 ( 14 ) is emitted into fiber bundle 1 ( 7 ). this fiber bundle could consist of high aperture optical glass fibers . the light from light source 3 ( 15 ) is emitted into fiber bundle 2 ( 17 ). this fiber bundle could consist of polarization maintaining optical glass fibers . in another solution high aperture optical glass fibers are used and the polarization filter ( 16 ) is positioned at the distal end of the extension ( 3 ) of the probe ( 1 ). both fiber bundle 1 ( 7 ) and fiber bundle 2 ( 17 ) pass through a channel in the extension ( 3 ) of the probe ( 1 ) and supply illumination of the measurement object via a lens ( 2 ). in another solution , fiber bundle 1 ( 7 ) and / or fiber bundle 2 ( 17 ) is mounted on the extension ( 3 ) in a removable fashion . the light from fiber bundle 1 ( 7 ) serves both as illumination for investigation and measurement . light reflected from the measurement object is collected via the lens ( 2 ) and projected onto a ccd camera ( 5 ) via an adapter ( 4 ). a ccd camera with enhanced sensitivity in the near - infrared wavelength region could be used . light reflected from a sub - section of the measurement object is collected by fiber bundle 3 ( 19 ). this fiber bundle could consist of low aperture optical glass fibers . in another solution , fiber bundle 3 ( 19 ) is mounted on the extension ( 3 ) of the probe ( 1 ) in a removable fashion . the radiation intensity emerging from fiber bundle 3 ( 19 ) is measured by a detection unit ( 20 ), preferably containing photo diodes . after manual input via an input device ( 9 ) in the control apparatus ( 6 ), requesting measurement to start , the ccd image is read by the control apparatus ( 6 ) and processed in a signal processor ( 13 ). the reading could be made continuous . according to the theory presented above , images enhancing contrast between tissue components with different optical properties are obtained during discrete wavelength illumination by light source 2 ( 14 ), and images enhancing contrast between tissues with a linearized and normal fiber structure are obtained during polarized light illumination by light source 3 ( 15 ). contrast enhanced images of the measurement object are presented on a display unit ( 12 ) in the control apparatus ( 6 ). following sequential illumination by light source 2 ( 14 ), the signal processor ( 13 ) calculates quotas between the intensities of measurement and reference wavelengths as measured by the detection unit ( 20 ). the calculated quotas are translated by the signal processor ( 13 ) to tissue layer thicknesses according to the theory presented above , and the tissue layer thickness is presented on the display unit ( 12 ) in the control apparatus ( 6 ). the input device ( 9 ), control unit ( 10 ), signal processor ( 13 ) and / or display unit ( 12 ) can be that of a computer . miniaturized measurement system for point - wise measurement of tissue layer thickness . the solution is presented in fig7 and fig1 . light source 2 ( 14 ) emits a number of discrete wavelengths for measurement . the light sources could be light emitting diodes . wavelengths emitted from said discrete wavelength tight sources could correspond to hemoglobin ( 5 - 10 nm in the vicinity of 425 , 542 or 576 nm ) or water absorption ( 1000 - 1600 nm , preferably 1300 nm ). a reference wavelength where similar absorption between the tissue components is seen should be used . said reference wavelength could be in the 600 - 800 nm region , preferably 630 nm . in one solution , light source 2 ( 14 ) is replaced by a single white light source . an optical filter extracts discrete wavelengths for measurement . the optical filter could be a liquid crystal filter . light source 2 ( 14 ) is driven by a light source driver unit ( 11 ), controlled by a control unit ( 10 ) in a control apparatus ( 6 ). the light from light source 2 ( 14 ) is emitted into fiber bundle 1 ( 7 ). this fiber bundle could consist of high or low aperture optical glass fibers . fiber bundle 1 ( 7 ) passes through a protective tube ( 21 ) and supply illumination of the measurement object . the protective tube ( 21 ) is preferably manufactured in a tenable plastic material . light reflected from the enlightened section of the measurement object is collected by fiber bundle 3 ( 19 ). this fiber bundle could consist of high or low aperture optical glass fibers . the radiation intensity emerging from fiber bundle 3 ( 19 ) is measured by a detection unit ( 20 ), preferably containing photo - diodes . the reading could be made continuous . after manual input via an input device ( 9 ) in the control apparatus ( 6 ), requesting measurement to start , sequential illumination by light source 2 ( 14 ) is performed . the signal processor ( 13 ) calculates quotas between the intensities of measurement and reference wavelengths as measured by the detection unit ( 20 ). the calculated quotas are translated by the signal processor ( 13 ) to tissue layer thicknesses according to the theory presented above . the tissue layer thickness is presented on the display unit ( 12 ) in the control apparatus ( 6 ). the input device ( 9 ), control unit ( 10 ), signal processor ( 13 ) and / or display unit ( 12 ) can be that of a computer . 1 probe 2 lens 3 extension 4 adapter 5 ccd camera 6 control apparatus 7 fiber bundle 1 8 light source 1 9 input device 10 control unit 11 light source driver unit 12 display unit 13 signal processor 14 light source 2 15 light source 3 16 polarization filter 17 fiber bundle 2 18 eye - piece 19 fiber bundle 3 20 detection unit 21 protective tube armstrong , c . g ., and mow , v . c . 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