the present invention is directed to a method and system for the in vivo , noninvasive detection and mapping of the biochemical and or functional alterations of tissue , e . g ., tissue within a subject . upon selection of the appropriate pathology differentiating agent that enhances the optical contrast between normal and pathologic tissue ( depending on the pathology of the tissue ), this agent is administered , e . g ., topically , to the tissue . as used herein a pathology differentiating agent is any agent capable of altering the optical property of a tissue , e . g ., an agent capable of altering the reflection characteristics or the fluorescence characteristics of a tissue . the pathology differentiating agent may be an acidic solution , a basic solution , a porphyrine solution or a porphyrine precursor solution . preferred examples of a pathology differentiating agent for use in the methods of the invention include an acetic acid solution , e . g ., a weak acetic acid solution , or 5 - amino luvelinic acid . in fig1 , the tissue ( t ), is sprayed using an atomizer ( a ), which contains the agent , e . g ., acetic acid . at the same time , the tissue is illuminated with a source that emits light having a frequency within a specific spectral band that depends on the optical characteristics of both the agent and the tissue . the characteristics of the light emitted from the source can be controlled by choosing particular sources ( ls ), and optical filters ( ofs ). sources of light for illuminating the tissue include light emitting diodes , and lasers . for imaging the area of interest , light collection optics ( l ) may be used , which focus the image onto a two - dimensional optical detector ( d ). the output signal of the latter is amplified , modulated and digitized with the aid of appropriate electronics ( eis ) and finally the image is displayed on a monitor ( m ) and stored in the data - storing means of a personal computer ( pc ). between tissue ( t ) and detector ( d ), optical filters ( ofi ) can be interposed . the filter can be interposed for tissue ( t ) imaging in selected spectral bands , at which the maximum contrast is obtained between areas that are subjected to different grade of alterations in their optical characteristics after administering the appropriate agent . before administration of the latter , images can be obtained and used as references . after the agent has been administered , the detector ( d ) helps to capture images of the tissue , in successive time instances , which are then stored in the computer &# 39 ; s data - storage means . the measuring rate is proportional to the rate at which the tissue &# 39 ; s optical characteristics are altered , following the administration of the agent . as used herein , an optical property , p , is a property that arises from the interaction of electromagnetic waves and a material sample , e . g ., a tissue , such as a tissue within a subject . for example , the property can be the intensity of light after it interacts with matter , as manifested by an absorption , emission , or raman spectrum . a dynamic optical property is a property that is obtained from a time - dependent optical property , p ( t ), and is determined from the measurement of p ( t ) at more than one time . for example , a dynamical optical property can be a relaxation time , or a time integral of p ( t ). in fig1 , images of the same tissue area are schematically illustrated , which have been stored successively before and after administering the agent ( sti ). in these images , the black areas represent tissue areas that do not alter their optical characteristics ( nat ), while the gray - white tones represent areas that alter their optical characteristics ( at ), following the administration of the agent . the simultaneous capture of the intensity of the light re - emitted from every spatial point of the tissue area under analysis and in predetermined time instances , allows the calculation of the kinetics of the induced alterations . in fig1 , two curves are illustrated : pixel value at position xy ( pvxy ), versus time t . the curve atc corresponds to an area where agent administration induced alterations ( at ) in the tissue &# 39 ; s optical characteristics . the curve ( natc ) corresponds to an area where no alteration took place ( nat ). each pixel , ( x , y ), can be associated with a pixel value , such as intensity i , which generally depends on time . for example , at time ti and pixel ( x , y ), the pixel value can be denoted by pv xy ( ti ). one useful dynamical spectral property , which can be obtained by measuring pixel value versus time at a particular pixel ( x , y ), is the relaxation time t ret ( x , y ). letting the maximum of a pv time curve be denoted by a , then t ret ( x , y ) satisfies pv xy ( t ret )= a / e , where e is the base of the natural logarithm . for example , if the pixel value versus time curve can be approximated by an exponential with relaxation rate r , pv xy ( t )= a exp (− rt ), where r & gt ; 0 , then t rel ( x , y )= 1 / r . the calculation of these parameters ( p ) at every spatial point of the area under analysis allows kinetic information ( ki ) to be obtained , with pixel values that are correlated with these parameters . these values can be represented with a scale of pseudocolors ( p min , p max ), the spatial distribution of which allows for immediate optical evaluation of the intensity and extent of the induced alterations . depending on the correlation degree between the intensity and the extent of the induced alterations with the pathology and the stage of the tissue lesion , the measured quantitative data and the derived parameters allow the mapping , the characterization and the border - lining of the lesion . the pseudocolor image of the phenomenon &# 39 ; s kinetics ( ki ), which expresses the spatial distribution of one or more parameters , can be superimposed ( after being calculated ) on the tissue image , which is displayed in real - time on the monitor . using the superimposed image as a guide facilitates the identification of the lesion &# 39 ; s boundaries , for successful surgical removal of the entire lesion , or for locating suspicious areas to obtain a biopsy sample ( s ). furthermore , based on the correlation of the phenomenon &# 39 ; s kinetics with the pathology of the tissue , the measured quantitative data and the parameters that derive from them can provide quantitative clinical indices for the in vivo staging of the lesion or of sub - areas of the latter . in some cases it is necessary to capture the kinetics of the phenomenon in more than one spectral band . this can help in the in vivo determination of illumination and / or imaging spectral bands at which the maximum diagnostic signal is obtained . furthermore , the simultaneous imaging in more than one spectral band can assist in minimizing the contribution of the unwanted endogenous scattering , fluorescence and reflection of the tissue , to the optical signal measured by the detector . the measured optical signal comprises the optical signal generated by the marker - tissue interaction and the light emitted from the endogenous components of the tissue . in many cases , the recorded response of the components of the tissue constitutes noise since it occludes the generated optical signal , which carries the diagnostic information . therefore , separation of these signals , based on their particular spectral characteristics , results in the maximization of the signal - to - noise ratio and consequently in the improvement of the obtained diagnostic information . fig2 illustrates a method for measuring in two spectral bands simultaneously and in any spatial point of the area under analysis , the kinetics of the alterations in the characteristics of the light emitted from the tissue , before and the after the administration of the contrast enhancing agent . the light emitted from the tissue is collected and focused by the optical imaging module ( l ) and allowed to pass through a beam splitting ( bsp ) optical element . thus , two identical images of the tissue ( t ) are generated , which can be captured by two detectors ( d 1 , d 2 ). in front of the detector , appropriate optical filters ( of λ1 ), ( of λ2 ) can be placed , so that images with different spectral characteristics are captured . besides beam splitters , optical filters , dichroic mirrors , etc ., can also be used for splitting the image of the object . the detectors ( d 1 ), ( d 2 ) are synchronized so that they capture simultaneously the corresponding spectral images of the tissue ( ti λ1 ), ( ti λ2 ) and in successive time - intervals , which are stored in the computer &# 39 ; s data storage means . generalizing , multiple spectral images can be captured simultaneously by combining multiple splitting elements , filters and sources . fig3 illustrates another method for capturing in different spectral bands simultaneously and in any spatial point of the area under analysis , the kinetics of the alterations in the characteristics of the light emitted from the tissue , before and the after the administration of the contrast enhancing agent . with the aid of a special prism ( mip ) and imaging optics , it is possible to form multiple copies of the same image onto the surface of the same detector ( d ). various optical filters ( of λ1 ), ( of λ2 ), ( of λ3 ), and ( of λ4 ), can be interposed along the length of the optical path of the rays that form the copies of the object &# 39 ; s image , so that the multiple images correspond to different spectral areas . for the clinical use of the methods of the invention , the different implementations of imaging described above can be integrated to conventional optical imaging diagnostic devises . such devises are the various medical microscopes , colposcopes and endoscopes , which are routinely used for the in vivo diagnostic inspection of tissues . imaging of internal tissues of the human body requires in most cases the illumination and imaging rays to travel along the same optical path , through the cavities of the body . as a result , in the common optical diagnostic devices the tissue &# 39 ; s surface reflection contributes substantially to the formed image . this limits the imaging information for the subsurface characteristics , which is in general of great diagnostic importance . this problem becomes especially serious in epithelial tissues such as the cervix , larynx , and oral cavity , which are covered by fluids such as mucus and saliva . surface reflection also obstructs the detection and the measurement of the alterations in the tissue &# 39 ; s optical properties , induced after the administration of agents , which enhance the optical contrast between normal and pathologic tissue . more specifically , when an agent alters selectively the scattering characteristics of the pathologic tissue , the strong surface reflection that takes place in both pathologic ( agent responsive ) and normal ( agent non responsive ) tissue areas , occludes the diagnostic signal that originates from the interaction of the agent with the subsurface features of the tissue . in other words , surface reflection constitutes optical noise in the diagnostic signal degrading substantially the perceived contrast between agent responsive and agent non - responsive tissue areas . for accurate diagnoses using the aforementioned imaging devices , appropriate optics can be used to eliminate noise arising from surface reflection . fig4 illustrates a schematic diagram of a medical microscope that includes a light source ( ls ), a magnification selection mechanism ( ms ), an eyepiece ( ep ) and a mount for attaching the image capturing module ( ca ), ( detector ( s ), readout electronics etc ). to eliminate surface reflection , a pair of linear polarizers is employed . light from the source passes through a linear polarizer ( lpo ) with the resulting linearly polarized light ( ls ) then impinging on the tissue . the surface reflected light ( ts ) has the same polarization plane as the incident light ( fresnel reflection ). by placing another linear polarizer ( ipo ), oriented at a right angle with respect to the first , in the path of the light emitted from the tissue , the contribution of the surface reflected light is eliminated . the light that is not surface reflected enters the tissue , where due to multiple scattering , light polarization is randomized . thus , a portion of the re - emitted light passes through the imaging polarization optics , carrying improved information for the subsurface features . fig5 illustrates an endoscope that includes an eyepiece ( ep ), which can be adapted to an electronic imaging system , and optical fibers or crystals for the transmission of both illumination and image rays . the endoscope also includes a first linear polarizer ( lpo ), disposed in the optical path of the illumination rays ( le ), and a second polarizer ( ipo ), oriented at right angles to the first , disposed in the path of the light emitted by the tissue ( ii ). the polarizer ( lpo ) can be disposed as shown in the figure , or , alternatively , where the light enters the endoscope ( il ). in the latter case , the endoscope has to be constructed using polarization preserving crystals or fiber optics for transferring the light . if polarization preserving light transmission media are used , then the polarizers for the imaging rays can be disposed in their path , in front or in back of the eyepiece ( ep ). a problem for the effective clinical implementation of the method described above involves the micro - movements of the patient , which are present during the snapshot imaging of the same tissue area . this problem is eliminated when the patient is under anesthesia ( open surgery ). in most cases , however , the movements of the tissue relative to the image capturing module , occurring during the successive image capturing time - course , result in image pixels , with the same image coordinates , which do not correspond to exactly the same spatial point x , y of the tissue area under examination . this problem is typically encountered in colposcopy . a method for eliminating the influence to the measured temporal data of the relative movements between tissue and image capturing module is presented below . a colposcopic apparatus , illustrated in fig6 , includes an articulated arm ( aa ), onto which the optical head ( oh ) is affixed . the head ( oh ) includes a light source ( ls ), an objective lens ( obj ), an eyepiece ( ep ) and optics for selecting the magnification ( ms ). the image - capturing module is attached to the optical head ( oh ), through an opto - mechanical adapter . a speculum ( k . d ), which is used to open - up the vaginal canal for the visualization of the cervix , is connected mechanically to the optical head ( oh ), so that its longitudinal symmetry axis ( la ) is perpendicular to the central area of the objective lens ( obj ). the speculum enters the vagina and its blades are opened up compressing the side walls of the vagina . the speculum ( i ( d ), being mechanically connected with the optical head ( oh ), transfers any micromovement of the patient to the optical head ( oh ), which , being mounted on an articulated arm ( aa ), follows these movements . thus the relative position between tissue and optical head remains almost constant . an important issue that must also be addressed for the successful clinical implementation of the diagnostic method described herein is the synchronization of the application of the pathology differentiating agent with the initiation of the snapshot imaging procedure . fig6 , illustrates an atomizer ( a ) attached to the optical head of the microscope . the unit ( mic ) is comprised of electronics for controlling the agent sprayer and it can incorporate also the container for storing the agent . when the unit ( mic ) receives the proper command from the computer , it sprays a predetermined amount of the agent onto the tissue surface , while the same or another command initiates the snapshot image capturing procedure . the diagnostic examination of non - directly accessible tissues located in cavities of the human body ( ear , cervix , oral cavity , esophagus , colon , stomach ) is performed with the aid of common clinical microscopes . in these devices , the illumination - imaging rays are near co - axial . more specifically , the line perpendicular to the exit point of light into the air , and the line perpendicular to the objective lens , form an angle of a few degrees . as a result , these microscopes operate at a specific distance from the subject ( working distance ), where the illuminated tissue area coincides with the field - of view of the imaging system . these microscopes are found to be inappropriate in cases where tissue imaging through human body cavities of small diameter and at short working distances is required . these technical limitations hinder the successful clinical implementation of the method described herein . as discussed above , elimination of surface reflection results in a substantial improvement of the diagnostic information obtained from the quantitative assessment of marker - tissue interaction kinetics . if a common clinical microscope is employed as the optical imaging module , then as a result of the above - mentioned illumination - imaging geometry , multiple reflections occur in the walls of the cavity before the light reaches the tissue under analysis . multiple reflections are more numerous in colposcopy because of the highly reflective blades of the speculum , which is inserted into the vagina to facilitate the inspection of the cervix . if the illuminator of the imaging apparatus emits linearly polarized light , the multiple reflections randomize the polarization plane of the incident light . as discussed above , if the light impinging on the tissue is not linearly polarized , then the elimination of the contribution from the surface reflection to the image can not be effective . fig7 illustrates an optical imaging apparatus that includes a light source located at the central part of its front - aperture . with this arrangement , the central ray of the emitted light cone is coaxial with the central ray of the light beam that enters the imaging apparatus . this enables illumination rays to directly reach the tissue surface under examination before multiple reflections occur with the wall of the cavity or speculum . a reflective - objective lens is used , which includes a first reflection ( 1rm ) and a second reflection ( 2rm ) mirror . a light source ( ls ) is disposed at the rear of the second reflection mirror ( 2rm ), together with , if required , optics for light beam manipulation such as zooming and focusing ( so ). the reflective - objective lens ( ro ), by replacing the common refractive - objective used in conventional microscopes , provides imaging capability in cavities of small diameter with the freedom of choosing the working distance . the zooming and focusing optics of the light beam can be adjusted simultaneously with the mechanism for varying the magnification of the optical imaging system so that the illumination area and the field - of - view of the imaging system vary simultaneously and proportionally . thus , image brightness is preserved regardless of the magnification level of the lens . the imaging - illumination geometry embodied in this optical imaging apparatus , along with the light beam manipulation options , helps to eliminate the surface reflection contribution to the image and consequently helps to efficiently implement the method described herein . those skilled in the art will recognize , or be able to ascertain using no more than routine experimentation , many equivalents to the specific embodiments of the invention described herein . such equivalents are intended to be encompassed by the following claims .