WO2025184677A1 - Apparatus for confocal endomicroscopy of a sample - Google Patents

Abstract

An apparatus for three-dimensional confocal endomicroscopy of a sample, in particular of a luminal organ, without any moving parts in the distal end of the apparatus, comprises: a light source (1) for emitting light to the sample, an optical sensor comprising a capsule (6) that can be swallowed or inserted into the sample and a bundle (3) of optical fibres for guiding light from the light source (1) to the capsule (6), the bundle (3) with a distal end (3a) thereof being introduced into the capsule (6) in a light-conducting manner, light detecting means for detecting fluorescent light emitted by the sample, image generating means for generating an image on the basis of signals obtained by the light detecting means, and is characterized in that the apparatus further comprises an optical focusing unit (13) arranged at the distal end (3a) of the bundle (3) inside the capsule (6) and adapted to direct light introduced into the optical fibres and emitted from them inside the capsule (6) in a converging manner onto a reflector (14) arranged in the capsule (6) and adapted to guide fluorescent light impinging onto the reflector (14) back into the optical fibres in a focused manner and the reflector (14) has the shape of a concavely curved surface of revolution on its side facing the focusing unit (13), the curved surface of revolution being arranged coaxially with the bundle (3) of optical fibres.

Description

Apparatus for confocal endomicroscopy of a sample

The present invention relates to an apparatus for confocal endomicroscopy of a sample , in particular of a luminal organ, comprising : a light source for emitting light to the sample , an optical sensor comprising a capsule that can be swal lowed or inserted into the sample and a bundle of optical fibres for guiding light from the light source to the capsule , the bundle with a distal end thereof being introduced into the capsule in a light-conducting manner, light detecting means for detecting fluorescent light emitted by the sample , and image generating means for generating an image on the basi s of signals obtained by the light detecting means .

The introduction of white light endoscopy has improved health care outcomes over a long period of time . However, white light endoscopy is not sensitive to early diseases in gastrointestinal tracts and has a limited diagnostic accuracy . Studies have shown that endoscopists , for example , only correctly identi fy 40-50% of precancerous conditions in the esophagus . As a result , for diagnostic purposes , gastroenterologists extract random biopsies in the suspect region . However, biopsy only samples a fractional area of tissue , thus the protocol is associated with signi ficant sampling errors . It is estimated that the miss rate of neoplasia in the esophagus is 30% to 50% using white light endoscopy and subsequent biopsies .

Moreover, due to patients ' discomfort , the endoscopic procedures are usually performed under sedation, which is an important contributor to the high cost of endoscopy . As an example , sedation accounted for 30 - 50% of the total procedural cost of upper endoscopy . Since recovery from the sedation often requires additional nursing and monitoring, this further confines the normal endoscopic procedures to speciali zed facilities .

To improve diagnostic accuracy, di f ferent imaging technologies are proposed . Among them, confocal laser endomicroscopy ( CLE ) is the only 3D modality that provides diagnostic information at cellular levels in real-time . A known commercial CLE system (Mauna Kea Technologies ) uses a minimi zed probe through the working channel of a standard endoscope to perform optical biopsy of gastrointestinal tracts . Clinical studies have shown that these probes of fer a promising approach for detecting early cancer in the gastrointestinal tract .

However, known probes have severely limited field of view ( FOV) , usually less than 500 pm x 500 pm (micrometer ) . Thus , they are used as the adj unct imaging method to white light endoscopy . There is also a lack of 3D imaging capability for the known probes , since they only image a fixed subsurface plane within the tissue , failing to deliver full cellular structures in depth .

Another form of CLE employs the spectrally encoded confocal microscopy ( SECM) to image gastrointestinal tracts using a tethered capsule . The capsule can be swallowed by a patient and the tether enables the movement of the capsule along gastrointestinal tracts . During movement , a motor inside the capsule rotates an imaging head to circumferentially image the inner surface of gastrointestinal tracts . The SECM capsule does not suf fer from limited field of view as it can image large area of gastrointestinal tracts with the movement of the capsule . However, as SECM capsule relies on the motor rotation to reali ze cross-sectional imaging, the imaging speed is slow . The SECM capsule also fails to of fer 3D imaging capability .

Confocal microscopy is a powerful optical imaging technology that of fers high contrast imaging of cellular features . Confocal laser endomicroscopy is the endoscopic embodiment of confocal microscopy, and it of fers cellular resolution microscopic views of mucosa in the gastrointestinal tract to allow real-time hi stological diagnosis . There is hope that confocal laser endomicroscopy can one day replace traditional biopsy and provide a non-invasive , comprehensive optical biopsy of large segments of gastrointestinal tracts to detect diseases at an early stage .

Apparatus as mentioned above are , for example , disclosed in US 2013 / 0310643 Al . It describes an apparatus with a tethered catheter with an encapsulated optical sensor on a distal end of a bundle of optical fibres . As another example , WO 2019/ 140174 A2 discloses an apparatus for capsule endomicroscopy comprising a tether having a proximal end and a distal end, and an optical fibre disposed within the tether . The optical element is optically coupled to a distal end of the optical fibre and is configured to direct light received from the optical fibre to a perimeter of the housing .

However, current state of the art confocal laser endomicroscopy technologies still suf fer, among others , from very limited fields of view of the known optical sensors , lack of 3D imaging capability and slow imaging speed for long gastrointestinal tracts . As a result , the commercially avai lable confocal laser endomicroscopy-systems are mainly used as an adj unct modality for traditional white light endoscopy .

The present invention is , therefore , directed to providing an apparatus of the initially mentioned kind that overcomes the above shortcomings . The inventive apparatus combines fibreoptic technology and confocal microscopy within a tethered capsule for imaging gastrointestinal tracts in a minimally- invasive manner .

In order to achieve this goal , the apparatus according to the invention comprises : a light source for emitting light to the sample , an optical sensor comprising a capsule that can be swal lowed or inserted into the sample and a bundle of optical fibres for guiding light from the light source to the capsule , the bundle with a distal end thereof being introduced into the capsule in a light-conducting manner, light detecting means for detecting fluorescent light emitted by the sample , and image generating means for generating an image on the basi s of signals obtained by the light detecting means and is characteri zed in that it further comprises an optical focusing unit arranged at a distal end of the bundle inside the capsule and adapted to direct light introduced into the optical fibres and emitted from them inside the capsule in a converging manner onto a reflector arranged in the capsule and adapted to guide fluorescent light impinging onto the reflector back into the optical fibres in a focused manner .

Preferably, the optical fibres are arranged around an axis and the reflector is designed to reflect light emitted from the optical fibres onto a region of the sample extending over an angle of 360 ° in the circumferential direction around the axis . In this way, the light beams emitted from the individual optical fibres are reflected into di f ferent angular directions with respect to the axis towards the sample in order to cover a region of the sample extending over an angle of 360 ° . In other words , together with the focusing unit , the reflector is able to generate focus points in the sample that are distributed over an angle of 360 ° in a circumferential direction with respect to said axis . On the other hand, the reflector, together with the optical focusing unit is adapted to guide fluorescent light generated at each focal point of the sample and impinging onto the reflector back into the optical fibres in a focused manner .

According to the invention, the reflector has the shape of a concavely curved surface of revolution on its side facing the optical focusing unit . Due to the curvature of the reflector surface , focus of light emanated from a fibre core closer to the bundle centre will be proj ected at a deeper depth in the tissue and the focus of light emanated from a fibre core further away from the bundle center will be proj ected at a shal lower depth inside the tissue . In other words , the reflector is not only able to generate optical focus points in the sample that are distributed over an angle of 360 ° in a circumferential direction with respect to the axis , these focus points are also distributed at di f ferent radial positions with respect to said axis . At the same time , fluorescent light emitted from aforementioned optical focus points at various depths inside the tissue will be reflected by the reflector through the light paths of the optical focusing unit back into the corresponding bundle of optical fibres . The small diameter of the cores of optical fibres acts as the confocal gating to reali ze optical sectioning . Preferably, the optical f ibres are arranged at various radial distances from the axis and the reflector therefore reflects light emitted from the optical fibres onto focal points that are arranged in a plane extending orthogonal to the axis and at di f ferent radial distances from the axis . In this way, it is possible to obtain fluorescent light signals from di f ferent depths of the sample and thereby obtain a three-dimensional image of the sample .

By employing a reflector as herein defined, it becomes feas ible to reali ze 3D imaging by pulling back the capsule from the gastrointestinal tract without any moving parts inside the capsule , as its geometry will divert axially impinging beams of light to the outside , i . e . into the tissue of the organ to be examined . The surrounding perimeter of the capsule can therefore be scanned by emitting light from the bundle of optical fibres and by registering the fluorescence coming back from the tissue .

The present invention, thus , provides for a small and robust optical sensor that is relatively cheap to manufacture as it does not have any moving parts . The rotational geometry of the reflector allows for a 360 ° field of view which, together with a high sampling frequency of the optical fibres in the bundle , allows for rapid and complete 3D-scanning of the inner tissue of the respective organ .

With a known peripheral position of an illumination spot emerging from the bundle of optical fibres at the distal end of the bundle , with an illumination spot being comprised of light emerging at the distal end from a multitude of neighboring optical fibres and the known geometry of the reflector it will be possible to attribute fluorescence light fed back from the tissue into said neighboring optical fibres of said illumination spot via the reflector to a speci fic peripheral positions around the circumference of the capsule and to a certain depth within the tissue at the respective peripheral position . The reflector will only feed light emerging from the optical focusing unit ' s and the reflector' s focusing depth back into the respective neighboring optical fibres of the respective illumination spot . Light emitted from other depths of the tissue of the sample will not reach the axial end of the bundle in a focused manner so that it will be excluded from travelling back through the optical fibres due to the confocal gating provided by the multitude of fibres in an illumination spot . The exclusion of unfocussed fluorescence light provides the capsule endoscopic imaging with all the bene fits associated with confocal microscopy .

In accordance with a preferred embodiment of the present invention, the capsule can be filled with at least one liquid or a mixture of liquids , wherein the at least one liquid preferably is selected from the group consisting of water and oils , in particular mineral oils . This brings about an increased numerical aperture of the optical focusing units inside the capsule , better aberration correction and refractive index matching between the capsule and the surrounding tissue and will , thus , enhance the optical performance of the inventive apparatus .

In order to make the bundle more easily manageable and to protect the fibre optics in the bundle , the invention is preferably devised such, that the fibre bundle is enclosed in a sheath to form a tether for the capsule . The tether also allows placing the capsule inside the organ as it of fers the necessary rigidity .

According to a preferred embodiment of the present invention, the capsule is displaceable in an axial direction of the bundle by a motori zed drive unit for the bundle . This allows for an automated scanning process of the tissue to be examined and for more consistent fluorescence data over a multitude of examinations performed .

Preferably, the bundle is fed to a light beam dividing unit with a proximal end of the bundle facing away from the capsule pointed onto a dichroic mirror arranged at an angle with respect to the axis of the bundle to direct fluorescent light out of the axis of the bundle and onto the light detecting means , in particular a photo detector array . This allows for separation of the fluorescence light returning from the tissue and for its positional registration . This is helpful for any method of calculating the position and depth of origin of the fluorescence light and for this , the recorded fluorescence light will be at least temporarily stored in a respective digital memory . The dichroic mirror will typically be arranged at an angle of 45 ° with respect to the axis of the bundle . The axis of the bundle , with the bundle being rather flexible , is deemed to be in the middle of the bundle and in the direction of the axis of the fibre optics within the bundle at its end facing the mirror .

In order to be able to ef fectively scan the surrounding of the capsule and to yield fluorescence light data with the light detecting means , in particular the photo detector array that can be attributed to a certain position and depth in the tissue of the organ to be examined, preferably, the dichroic mirror is arranged between a first Nipkow disk ( scanning disk) arranged between a light inj ection unit and the dichroic mirror and a second Nipkow disk ( scanning disk) arranged between the dichroic mirror and the proximal end of the bundle , wherein the first and the second Nipkow disks have pinholes in registration with each other and are drivable to synchroni zed rotation . This allows for the inj ection of light in periodically varying peripheral positions corresponding to the respective peripheral positions of the illumination spots comprised of a multitude of optical fibres as described above as the Nipkow disks of the present invention have pinholes arranged in a nested cluster of Archimedean spirals of constant pitch . Rotating the disks with the light inj ection unit turned on will , therefore , impinge light onto speci fic regions of the bundle creating an illumination spot at these speci fic regions . The di f ferent illumination spots at the distal end of the bundle are assigned to speci fic regions on the reflector by the optical focusing unit . Thus , the inj ected light will be focused into the tissue at speci fic depths inside the tissue and will , thus , evoke fluorescent light phenomena in those areas , i . e . depths of focus within the tissue of the organ . As explained above , due to the properties of the reflector together with the properties of the optical focusing unit , only the fluorescent light from the depths of focus of the respective illumination spot returns through the very same optical fibres of the illumination spot within the bundle and reaches the dichroic mirror through the pinholes in the second disk in registration with the pinholes of the first disk and will , thus , be registered in an assigned position on the light detecting means , in particular the photo detector array . Mapping of the actual position of the fluorescence phenomenon can then be carried out from the stored image data by polar coordinate conversion . In order to optimi ze the yield of image data from a certain flux of light inj ected into the fibre optics of the bundle , the present invention is preferably characteri zed in that the pinholes of the first Nipkow disk are installed with microlenses adapted to focus a collimated light beam onto the pinholes of the second Nipkow disk and that a lens arrangement is arranged between the second Nipkow disk and the proximal end of the fibre bundle adapted to focus light impinging onto it through the pinholes o f the second Nipkow disk onto the bundle . By these groups of lenses it is assured that the light entering a pinhole in the first Nipkow disk will more or less completely reach the fibre optics in the bundle . Thus , high yields of fluorescence light can be achieved in the illumination spots .

The present invention will now be exempli fied in more detail with reference to the attached drawings . In the drawings , Figure 1 shows a simpli fied representation of the arrangement of various modules of the inventive apparatus , Fig . 2 shows a detailed sectional view of the capsule , Figure 3 shows a more detailed schematic view o f the light inj ection unit with the dichroic mirror and the Nipkow disks and Figure 4 illustrates the processing of original data to a cross-sectional imaging frame .

In Figure 1 , a light source is denoted by reference numeral 1 . The light source 1 provides light in order to excite fluorescence from endogenous or exogenous fluorophores inside the biological tissue 7 . The fluorescence light is collected using an optical detection apparatus 8 to create images . The light source 1 may comprise laser diodes at , for example , 405 nm, 488 nm or 683 nm depending on the targeted fluorophores . Other sources such as solid lasers or gas lasers may also be suitable to excite the f luorophores .

A swallowable/ insertable capsule 6 is provided to image the inner surfaces of biological tissues 7 , especially gastrointestinal tracts . The capsule 6 is connected with a tether 4 and movable along the tissue 7 by the tether 4 . The movement of the capsule 6 can be performed either manually or automatically using a motori zed drive unit 5 , such as a motori zed stage . The dimension of the capsule 6 should allow the surface of the tissue 7 , especially gastrointestinal tracts such as esophagus , intestines or colons , to be closely attached to the capsule 6 surface for imaging . The capsule 6 has imaging optics to image tissues 7 in a cross-sectional manner .

The excitation light from the light source 1 is transmitted by coherent optical fibres in a bundle 3 of optical fibres into the capsule 6 . The bundle 3 of optical fibres also transmits the emitted fluorescence from tissue 7 to the optical detection apparatus 8 for detection . The bundle 3 of optical fibres contains a multitude of individual optical fibres or fibre optics . Commercially available fibre bundles , such as Fuj i kura products can provide 3000 to 100000 optical fibres for one bundle 3 , though higher number of cores can be provided upon customi zation . These fibres typically have a core-to-core distance of 3 . 7 pm and a core diameter of 2 . 5 pm, which are ideal for the inventive apparatus . With each core working as a channel to visuali ze the tissue 7 , the number of optical fibres should allow a comprehensive sampling of the tissue cross-section . A light inj ection unit 2 simultaneously illuminates multiple fibre cores at a high scanning rate creating an illumination spot to allow fast imaging . The collected data from the optical detection apparatus 8 is first stored in a memory 9 and a coordinate trans former 10 needs to trans fer the original data from memory 9 from its imaging plane to the tissue crosssection plane for visuali zation . The need of trans formation is due to the optics design in the capsule 6 , which will be explained in more detail below . The resulting cross-sectional imaging frames after the coordinate trans former 10 are stored in another memory 11 and combined for 3D visuali zation . A controller 12 is needed to synchroni ze the light inj ection unit 2 , motori zed drive unit 5 and optical detection apparatus 8 .

In Figure 2 , it can be seen that the bundle 3 with a distal end 3a thereof is introduced into the capsule 6 in a lightconducting manner and it is firmly connected to it . A plurality of illumination spots 13a at the distal end 3a of the bundle 3 inside the capsule 6 is directed by the optical focusing unit 13 onto a reflector 14 arranged in the capsule 6 and is further adapted to guide fluorescent light impinging from the tissue 7 onto the reflector 14 into the optical fibres of the fibre bundle 3 in a focused manner . The optical focusing unit in the example depicted in Fig . 2 is comprised of a plurality of lenses 13b . The reflector 14 has the shape of a concavely curved surface of revolution on its side facing the optical focusing unit 13 . The fibre bundle 3 is enclosed in a sheath 4a to form a tether 4 for the capsule 6 . A central axis of the bundle 3 is denoted by reference numeral 15 .

As can be seen in Figure 2 , illumination spots or light beams emerging from the bundle 3 at di f ferent radial positions also impinge on the reflector 14 at different locations and will therefore be focused into the tissue 7 at different depths. With scanning of the fibre optics within the bundle 3, a 360° image will be obtained with precise depth resolution as described above.

By designing the curvature of the freeform cone reflector 14 as a continuous function of the radial distance 16, beams from different neighboring fibre cores are focused at different depths. For example, the fibre cores at 25 are closer to the fibre bundle center 26, corresponding to the central axis 15 and their beam is reflected by the freeform cone reflector 14 at a deeper position 27 inside the tissue 7. The fibre cores at 28 are further away from the fibre bundle center 26, and their beam is reflected by the freeform cone reflector 14 at a shallower position 29 inside the tissue 7. Therefore, the lenses 13b, the freeform cone reflector 14 and the distance between the cone reflector 14 and the lenses 13b determines a one-to-one mapping 30 between fibre core, i.e. illumination spot radial position to the depth of the focal point inside the tissue 7. Illumination of multiple fibre cores at the same time means fluorescence will emit simultaneously from multiple points at different circumferential positions and depths within the tissue cross-section. The emitted fluorescence from the tissue 7 is reflected by the cone reflector 14 and focused into corresponding fibre cores by lenses 13b. Thus, the small core diameter serves as the confocal gating by rejecting the out-of-focus fluorescence. The fibre bundle 3 with its numerous fibre cores and the freeform cone reflector 14 enables 3D confocal imaging, without active moving elements inside the capsule 6. Figure 3 shows in a schematic manner the composition of optical inj ection apparatus 2 and the optical detection apparatus 8 . The excitation light from the light source 1 is coll imated by lens 31 and expanded by lenses 32 and 33 . A reflection mirror 34 reflects the excitation beam in the direction of arrow 18 to two coaxially aligned scanning disks designed as Nipkow Disks 20 and 21 . Disks 20 and 21 are identical Nipkow disks with pinholes arranged in a nested cluster of Archimedean spirals of constant pitch . Pinholes 35 of disk 20 are installed with microlenses 35a . Pinholes of disk 21 are empty pinholes 36 . The surfaces of disks 20 and 21 preferably contain a polished black chrome coating to reduce optical ref lection . A motor 37 rotates the disks 20 and 21 simultaneously . Typical pinhole si zes on disks 20 and 21 are 50 pm, with 250 pm pinhole spacing . A dichroic mirror 38 passing the excitation light and reflecting the fluorescence from the tissue is mounted at 45 ° to the excitation beam . The dichroic mirror 38 is also mounted between disks 20 and 21 . The microlenses 35a focus the excitation light through the dichroic mirror 38 into corresponding pinholes 36 of disk 21 . Accurate alignment of disks 20 and 21 is required for the transmission of the excitation light from disks 20 to 21 . The rest of the excitation light is blocked by disk 20 thereby creating a beam of light with dimensions corresponding to the si ze and shape of the pinholes 35 and 36 . After disk 21 , the excitation beams are focused by lenses 39 and 40 into the fibre bundle 3 . Lenses 39 and 40 should produce focus spots on the fibre bundle 3 with a spatial intensity point spread function equaling the fibre core diameter to optimi ze the coupling ef ficiency . Lenses 39 and 40 should correct spherical aberrations and field curvatures to achieve a wavefront qual ity with wavefront error of the order of X/ 8 . The motor 37 rotates disks 20 and 21 at speeds up to several thousand rounds per minute . The fluorescence emitted from the tissues is focused by lenses 39 and 40 through pinholes 36 in the direction of arrow 19 onto the dichroic mirror 38 . The fluorescence is then reflected through a barrier filter 41 . The barrier filter 41 blocks the excitation light from entering the optical detection apparatus 8 . Transmitted fluorescence signals are focused by lenses 42 and 43 on a photo detector array 44 , such as a charge-coupled device ( CCD) sensor . The controller 12 should strictly synchroni ze the rotation of optical disks 20 and 21 with the exposure time of the photo detector array 44 . The exposure time must be an integer multiple of the time for a full scan of the fibre bundle surface .

Figure 4 symbolically illustrates the processing of the original data from the optical detector array 44 to a cross- sectional imaging frame . Assuming using a CCD as the optical detector array 44 , the controller 12 controls the rotational speed of motor 37 and the exposure time of CCD so that the exposure time equals one full scan of the fibre bundle 3 . A full scan of the fibre bundle 3 corresponds to a full scan of the tissue 7 cross-section . It is evident then that the cross- sectional imaging frame rate of the apparatus equals to the CCD frame rate . A realistic cross-sectional imaging frame rate of the device may be 1000 Hz for longer CCD exposure and higher imaging signal to noise ratio .

The resulting original data 45 from the optical detector array 44 is the Hadamard product of the image of the fibre bundle surface and the spatial distribution of fluorescence intensity in a tissue cross-section . Every pixel of the original data 45 is represented in Polar coordinates ( r, 0 ) , with the fibre bundle center 26 as the origin . The pixel intensity reflects the fluorescence intensity emitted from a unique point source in the tissue cross-section . This unique one-to-one correspondence between a point source in the tissue cross- sectional plane 48 and a pixel in the CCD plane 44 is due to mapping 30 of Figure 2 . With the mapping 30 , a point source in the tissue cross- sectional plane 48 and the corresponding pixel in the CCD plane 44 have the same angle 0 , but di f ferent axial lengths in their respective coordinate systems . The change of radial lengths is known and determined by lenses 13b, the curvature of the freeform cone reflector 14 , and spacings between these components . Thus , the coordinate trans former 10 does the inverse trans form of mapping 30 and maps each pixel in the CCD image plane 44 back into respective points in the tissue cross-sectional plane 48 . Examples are the trans formation of 46 and 47 in the CCD plane 44 to point sources 50 and 51 in the tissue cross-sectional plane 48 , respectively . Another example shows that features such as concentric lines 52 and 53 in the CCD image plane 44 will be converted as respective concentric lines 54 and 55 in the tissue cross-sectional plane 48 , respectively . The tissue cros s-sectional image obtained after the coordinate trans former 10 is further processed with Gaussian low pass filter and spline interpolations to remove the pixelated fibre core pattern and improve resolutions .

Claims

Claims :

1. Apparatus for confocal endomicroscopy of a sample, in particular of a luminal organ, comprising: a light source (1) for emitting light to the sample, an optical sensor comprising a capsule (6) that can be swallowed or inserted into the sample and a bundle (3) of optical fibres for guiding light from the light source (1) to the capsule (6) , the bundle (3) with a distal end (3a) thereof being introduced into the capsule (6) in a light-conducting manner, light detecting means for detecting fluorescent light emitted by the sample, and image generating means for generating an image on the basis of signals obtained by the light detecting means, wherein the apparatus further comprises an optical focusing unit (13) arranged at the distal end (3a) of the bundle (3) inside the capsule (6) and adapted to direct light introduced into the optical fibres and emitted from them inside the capsule (6) in a converging manner onto a reflector (14) arranged in the capsule (6) and adapted to guide fluorescent light impinging onto the reflector (14) back into the optical fibres in a focused manner, characterized in that the reflector (14) has the shape of a concavely curved surface of revolution on its side facing the focusing unit (13) , the curved surface of revolution being arranged coaxially with the bundle (3) of optical fibres.

2. Apparatus according to claim 1, characterized in that the optical fibres are arranged around an axis (15) and the reflector (14) is designed to reflect light emitted from the optical fibres onto a region of the sample extending over an angle of 360° in the circumferential direction around the axis .

3. Apparatus according to claim 2, characterized in that the optical fibres are arranged at various radial distances from the axis (15) and the reflector (14) is designed to reflect light emitted from the optical fibres onto focal points that are arranged in a plane extending orthogonal to the axis and at different radial distances from the axis.

4. Apparatus according to any one of claims 1 to 3, characterized in that the capsule (6) is filled with at least one liquid or a mixture of liquids, wherein the at least one liquid preferably is selected from the group consisting of water and oils, in particular mineral oils.

5. Apparatus according to any one of claims 1 to 4, characterized in that the bundle (3) is enclosed in a sheath to form a tether for the capsule.

6. Apparatus according to any one of claims 1 to 5, characterized in that the capsule (5) is displaceable in an axial direction of the bundle (3) by a motorized drive unit (5) for the bundle (3) .

7. Apparatus according to any one of claims 1 to 6, characterized in that the bundle (3) is fed to a light beam dividing unit with a proximal end (3b) of the bundle (3) facing away from the capsule (6) pointed onto a dichroic mirror (38) arranged at an angle with respect to the axis (15) of the bundle (3) to direct fluorescent light out of the axis (15) of the bundle (3) and onto the light detecting means, in particular a photo detector array (44) .

8. Apparatus according to claim 7, characterized in that the dichroic mirror (38) is arranged between a first Nipkow disk

(20) arranged between a light injection unit (2) and the dichroic mirror (38) and a second Nipkow disk (21) arranged between the dichroic mirror (38) and the proximal end (3b) of the bundle (3) , wherein the first and the second Nipkow disks (20, 21) have pinholes (35, 36) in registration with each other and are drivable to synchronized rotation.

9. Apparatus according to claim 7 or 8, characterized in that the pinholes (35) of the first Nipkow disk (20) are installed with microlenses (35a) adapted to focus a collimated light beam onto the pinholes (36) of the second Nipkow disk

(21) and that a lens arrangement (39, 40) is arranged between the second Nipkow disk (21) and the proximal end (3b) of the fibre bundle (3) adapted to focus light impinging onto it through the pinholes (36) of the second Nipkow (21) disk onto the bundle (3) .