KR0145114B1 - Phase modulated spectrophotometry - Google Patents

Abstract

The present invention relates to a method and apparatus for measuring the concentration of absorbent components in a scattering medium.

The light source 10 is connected to the scattering medium 20, moved through the medium, and detected by the detector 16.

The reference signal 18 is mixed with a detection signal that produces a phase shift detection signal from the signal whose concentration is measured.

Description

[Name of invention]

Phase Modulated Spectrophotometry

[Related Applications]

This application discloses pending applications, Britton Chane, filed Nov. 2, 1988, filed No. 266,166. Optical coupling system used to monitor oxidation status in living tissue, filed Nov. 2, 1998, filed No. 266,019. The tissue hemoglobin concentration of known absorbents was measured using the attenuation characteristics of a portable hemoglobinometer for measuring metabolic status in a subject and an electron attenuation scattered in vivo, filed December 21, 1998, application No. 287,847. An application related to a method and apparatus for measuring.

[Background of invention]

The application of the basic double-wavelength principle to detect hemoglobin and myoglobin changes in tissues has been described in G.A. It was written by Millikan and by Millikan and Pappenheimer, who detected the reduction of hemoglobin in the earlobe of the human body.

Multiwavelength instruments have been developed.

These instruments use either a multi-wavelength laser diode light source or one of time-division filter technology, and through several algorithms that deconvulte background signals, oxidized and reduced cytochrome signals, and oxidized and reduced hemoglobin signals. Results with high precision have been found.

These techniques are oxygen-composites and have low light levels that make it difficult to obtain a light source with a wavelength suitable for the algorithm developed or that can count photons.

It is typically priced at $ 80,000 and provides a large amount of experimental data on newborns and adults.

The fundamental problem with this method is that the optical path length is not known from the outset, but rather is measured by reference to an animal model from which hemoglobin is removed and the chitochrome can be studied immediately.

The transfer of such data from animals to humans is one difficult problem to be solved before the invention of time-resolved spectroscopy in which path lengths are directly measured.

See the optical coupling system used to examine the oxidation state in living tissue above.

Continuous wave spectroscopy (CWS) of hemoglobin in tissues has the advantage of being very simple and sensitive as well as early warning of hypoxia in tissues.

Applying picosecond time-analytical spectroscopy (TRS) in tissues to measure the optical path length and the amount of change in hemoglobin concentration and the actual concentrations of hemoglobin and chitochrome is well suited for clinical studies of hypoxia in tissues. It can be said.

Moreover, used in conjunction with continuous spectrophotometry, time-resolved spectroscopy provides a means of measuring the optical path length that a photon travels through tissue.

As long as these trends are of great value, the ability to measure hemoglobin levels with continuous light and pulsed light technology dramatically increases the applicability to clinical research.

US Patent Application No. 266,166, filed November 2, 1988, optical coupling system used to examine the oxidation state in living tissue, and electromagnetic radiation emission attenuation characteristics in vivo, US Patent Application No. 287,847, Dec. 21, 1988 See Methods and Apparatus for Determining Concentrations of Tissue Pigments of Known Absorbents Using "

[Summary of invention]

Principles of dual wavelength spectroscopy have been found that can be applied to time-resolved polarization spectroscopy in which the carrier frequency is selected from any value of the time characteristic that is compatible with the time delay of photon transfer from input to output through the scattering medium.

The present invention provides a method and apparatus for transmitting a modulated waveform to a scattering medium and detecting it after being moved through it.

The detected waveform can be modified and compared with the initial waveform.

For example, the waveform is phase shifted by a delay while passing through the scattering medium.

Thus in a preferred embodiment the phase of the waveform is modulated and phase shift is detected. The concentration of an absorbent component, such as hemoglobin, can be measured by the difference in the wavelength of the emitted electromagnetic waves having two different wavelengths.

Accordingly, it is an object of the present invention to provide a method and apparatus for irradiating photon shift using signal modulation techniques such as time, frequency and phase modulation.

It is a particular object of the present invention to provide methods and apparatus in which phase modulated spectroscopy (PMS) can be used in conjunction with continuous wave spectroscopy (CWS) to measure the threshold of absorbent pigments such as hemoglobin at the point where PCr / Pr ratio begins to decrease To provide.

It is another object of the present invention to provide a dual wavelength phase modulation system for clinically applying the advantages of time resolved spectroscopy in a specific embodiment economically and commercially.

[Brief Description of Drawings]

Figure 1a shows a simplified single wavelength phase modulated spectrophotometer according to the present invention.

FIG. 1B is a block diagram of a specific example of a dual wavelength phase modulation spectrophotometer according to the present invention.

2 is a block diagram of another embodiment of a phase modulation spectrophotometer according to the present invention.

3 is a block diagram of a preferred embodiment of the spectrophotometer of the present invention.

4 is a block diagram of another embodiment of the spectrophotometer of the present invention.

Detailed description of the invention

The time, frequency or phase of a signal can be modulated.

Phase modulation can be conveniently performed in the above-described time-resolved spectroscopy (TRS) technique.

In Fig. 1b, a single wavelength spectrophotometer using the principle of phase modulation is shown.

In this embodiment, the 4 mW laser diode 11 is excited by the frequency generator 17 operating at 200 MHz and emits light of wavelength 760 nm.

Light is conducted to the test subject 20 through the optical fiber 15.

Light is detected after it is moved through tissue.

This detector is preferably composed of a photomultiplier tube and a voltage supply device 16.

This device is a Hamamatus R928.

The frequency generator 17 also receives an input signal from the 50 KHz oscillator 19 and inputs a reference waveform of 200.05 MHz to the detector 16.

Therefore, the output waveform 22 of the detector 16 is a carrier frequency with a frequency difference of 50 KHz, for example.

The waveform 22 of the detector 16 and the reference waveform of the oscillator 19 are supplied to the phase and amplitude detector 24.

In this embodiment, phase and amplitude detector 24 is a lock-in amplifier.

The outputs of the lock-in amplifiers are signals representing the phase shift and amplitude of the detected signal.

These signals are then analyzed and absorbent fathers such as hemoglobin are associated with relative concentrations.

In FIG. 1a a broadband acoustic optical modulator 12 operating at 200 MHz is connected to the helium-neon laser light source 10.

The frequency of the acoustic optical modulator 12 modulates the light emitted by the laser 10.

Light is conducted through the optical fiber guide 14 to the forehead of the illustrated subject 20 or another portion to be irradiated.

A signal about 3-6 cm from the position of the input waveform is exceptionally received by the Hamamatus R928 detector 16.

Since the dynodes are modulated with a 220.050 MHz signal, a heterodyne signal of 50 Hz is obtained and supplied to a lock-in amplifier 24 such as the PAR SR510.

As above, the reference frequency for the lock-in amplifier is obtained from 50 Hz, the difference between the two frequencies.

The phase shift between the transmission waveform and the detection waveform is measured with high precision and the output waveform shown at 26 is shown as an analog signal to the strip chart recorder so that the user can track the change in the transmission of light through the brain or other tissue. .

The signal is then converted algebraically.

The result has a linear relationship with the change in concentration of absorbent components such as hemoglobin.

Referring to FIG. 2, a block diagram of a simplified embodiment of a dual wavelength phase modulated spectrophotometer made in accordance with the present invention is shown.

Unlike the single wavelength system of FIG. 1B, this embodiment enables the measurement of the concentration of absorbent components on an absolute basis.

The embodiment of FIG. 2 is similar to that shown in FIG. 1B except that light is transmitted to the subject at two discrete wavelengths.

2 shows a second embodiment of the apparatus of the present invention.

In this embodiment the laser diode light is amplitude modulated and the phase shift due to the photon's movement is measured by an optical detector, mixer and phase detector.

The dual frequency time division system consists of stable oscillators 30, 32, such as Kenwood Model # 321 for 220 Hz: If the oscillator system is preferably used, it may generate waveforms from 144 MHz to 440 MHz (Kenwood TM721A).

The above mentioned frequencies of 144,220 and 440MHz are suitable for application to other parts than the original purpose, but it is also possible to continuously change the frequency.

Oscillators 30 and 32 are set to 50 Hz difference and the frequency difference is detected by mixer 34 to obtain the reference phase signal 36 shown.

The 200 Hz electronic switch 38 excites laser diodes 40 and 42 which in turn operate between about 75-60 nm and 800-10 nm and 220 which are conducted through optical fiber guides 44 and 46 with a preferred diameter of about 3 mm. The MHz modulated light is emitted to the head surface of the test subject 20 or another part to be irradiated.

It has been found that the most effective detector 48 is Hamamatus R928 to achieve a satisfactory operating state at 220 MHz.

A more advantageous device, however, is the Hamamatus R1645u, a microchannel tube with a 120 picosecond transmission time range and a high gain as the miniature channel plate photomultiplier tube 48 in two states.

This tube with a current amplification capability of 5 x 105 (57 kW) is similar to those used for pulsed time measurements in time resolved spectroscopy (TRS) studies and is considered ideal for this purpose.

See US Patent Application No. 287,487 for a method and apparatus for measuring the concentration of known tissue pigments using the attenuation characteristics of electromagnetic radiation in vivo.

The photomultiplier tube 48 is connected to a high voltage source having an output voltage of about 3400V to obtain high gain.

The photomultiplier tube 48 may be located by being connected to a brain or other tissue part by the optical fiber guides 44 and 46 or directly connected to a housing isolated from the earth potential as shown.

As described above, the detector 48 is attached to the subject 20 and connected to the mixer 52, thereby mixing the 220 MHz output signal of the detector 48 with the 220.050 MHz signal of the oscillator 32 to downconvert to 50 KHz. do.

The lock-in amplifier 54 measures the phase of the output waveform and also obtains the logarithm of the signal.

This signal is then fed to a secondary phase detector / lock-in amplifier 56 that measures the difference between the signals at each of the two wavelengths, which is directly proportional to the concentration of absorbing pigment, such as hemoglobin.

This embodiment can be used in newborns as well as in the adult brain.

A preferred embodiment of a time division, dual wavelength laser diode phase modulation spectrophotometer is shown in FIG.

In this embodiment a pair of laser diodes 110, 102 are excited in parallel at 220 MHz by a stable frequency oscillator 104 (Kenwood 321).

Each diode 100, 102 emits electromagnetic waves of different wavelengths, preferably 760 nm and 800 nm.

The radiation of electromagnetic waves is time-divided by a vibration mirror 105 which illuminates a single optical fiber probe at a modulation frequency which is preferably about 60 Hz.

The motion of the mirror 105 and the synchronization of the 60 Hz phase detector 120 (described below) can be achieved by electrically coupling the reference voltage in the 60 Hz lock-in amplifier 120.

Thus, electromagnetic radiation at each wavelength is synchronized between radiation and detection.

The latter embodiment in the spectrophotometer of FIG. 3 and the embodiment of FIG. 2 uses a carrier modulation system that encodes the excitation power of one laser, while the embodiment of FIG. 3 outputs from two lasers excited at the same frequency. Those skilled in the art will appreciate that they switch continuously between lights.

Time-divided 760/800 nm light is applied to the subject 20 through the optical fiber 106.

The output probe 108, which preferably includes secondary fibers that are relatively centimeters away, captures light passing through the subject and is a suitable photoelectron detector (Hamamatsu R928) or a microchannel plate detector (Hamamatus R1645U). ).

The focused light is phase shifted from the input oscillation by time delay when the photons move between the input and output.

A secondary oscillator 114 generating a 220.030 MHz waveform is coupled to the mixer 112.

The 220.000 MHz output of the detector 110 is also connected to the mixer.

As a result, the phase modulation frequency is shifted downward to 30KHz, which is a convenient frequency for lock-in detection.

This signal is input to phase detector 116, preferably a lock-in amplifier.

The secondary input signal of the phase detector 116 can be obtained by connecting the input signals of the 220 MHz oscillator 104 and the 220.030 MHz oscillator 114 to a mixer 118 for obtaining a non-moved 30 KHz signal used as the reference phase.

The lock-in amplifier 116 thus operates on a reference phase obtained directly from the frequency generators 104 and 114 and a phase modulated input obtained by photon shift through the test subject 20.

The weft of the output signal changes between the phase according to the light wave at 800 nm and the phase according to the light wave at 760 nm.

The output of the lock-in amplifier 116 is thus a waveform of 60 Hz and its amplitude carries phase information at two wavelengths.

The output signal of the phase difference detector 116 is connected to the same waveform as driving the 60 Hz vibration mirror 105.

The output of the phase detector can be obtained by connecting a switch to a vibration modulation that sequentially connects the opposite phase of the 60 Hz waveform to the integrated network at the peak of the phase detector output waveform.

The output is input to a differential amplifier that records the amplitude difference between two parts of the 60 Hz waveform corresponding to 760 nm and 800 nm phase shifts.

This retardation output is properly filtered at 0.05 to 1 Hz and records the execution time of the change in hemoglobin concentration by double wavelength time resolution spectroscopy.

The advantage of the system illustrated in Figure 3 is to allow a single light guide input signal to the subject to be operated from two laser diodes operating continuously at the same oscillation frequency.

Therefore, the counterfeit phase difference of the frequency related to excitation is minimized.

That is, no differential phase shift is expected between the 760nm and 800nm signals. Thus, a signal with a difference of 30 KHz will indicate the actual phase delay between these two wavelengths.

Moreover, phase noise in this region is minimized by the differential detector 116.

The photoelectron tube detector 110 can be sufficiently high speed as the mixing function is separated from the detector.

Lock-in amplifier technology, which derives the difference in phase and amplitude of the two signals, has the highest possible signal-to-noise ratio for this type of device.

The principle of time-division dual wavelength spectrophotometry combined with lock-in techniques generally follows the principle adopted for dual wavelength spectrophotometry.

However, the present invention provides a very improved device which is fast enough to measure the photon travel time between the input and the output having a characteristic time where a carrier frequency of 220.000 MHz is observed at about 5 nanoseconds.

Therefore, as observed in the experiments, the sensitivity of the published system is high, about 70 ° per nanosecond or 3 ° per centimeter change in path length.

The application of the principle of dual wavelength spectrophotometry to time resolved spectrophotometry allows the carrier frequency to be selected from a value of time characteristic compatible with the time delay of photon transfer from input to output.

The disclosed apparatus achieved a very precise measurement of the change in absorptivity in photon travel at a particular distance of approximately 1 m, compared to the continuous light method where photon travel was measured for all possible path lengths as contrasted.

The path length of approximately 1 m is preferably selected to allow scanning all parts of the brain for brain hemorrhage studies.

Obviously high frequencies will choose a small portion of the brain that is more granular with respect to input and output appearance.

The only known method of measuring the path length of a photon transmitted in a multilayer scattering medium such as human tissue is to measure flight time and refractive index and calculate the travel distance from it.

The path length in the brain is centimeters, so travel time is nanoseconds or less.

There are several fundamental shortcomings in measuring this period directly in this time domain.

The clearer the required time resolution, the greater the detection bandwidth should be.

Signal power is constant at best, while noise power is proportional to the increase in bandwidth.

For sources such as laser diodes, where the average output voltage is approximately equal for both pulsed and continuous operation, the signal power actually drops as the pulse width decreases.

The duty cycle of the pulse train is practically low because the time between the probes must be sufficient to reduce the reflected light to about zero.

This implies the use of low average signal power or high peak power, which can damage the skin surrounding the irradiated tissue.

Finally, the cost and difficulty of constructing a suitable electronic circuit is significantly greater in pulsed systems than in continuous wave systems.

As an alternative to time-domain measurements, the CW system can adopt phase measurements instead of time intensity, and the characteristic decay time can be calculated by simply calculating based on the phase shift measurements between the probe and the reflected light at a single frequency.

Such a system has the advantage of the high average power of narrowband modulation, detection and probe circuits and a significant advantage in signal-to-noise ratio and data recognition time.

Considerable literature on these timing techniques exists, particularly for radar, time-base and spectroscopy.

Perhaps most closely related to this application is the literature on the phase resolution measurement of fluorescence attenuation kinetics.

Another embodiment of the apparatus of the present invention is shown in FIG.

This system relies more on communication technology than on NMR technology and is essentially a single sideband system in which sidebands are modulated in proportion to the required modulation frequency shift.

This design relies heavily on radio frequency transmitters / receivers, which currently cost about $ 300 per transmit / receive frequency.

A block diagram of such a system is shown in FIG.

In this embodiment the first standard communication transmitter / receiver (transceiver) operating at 220 MHz is used in the transmission mode for waveform generation that excites the laser diode 202.

The transceiver 200 is used in single sideband mode, which provides single side band modulation of 3KHz.

This carrier signal is fed back to the transceiver 200 and fed to a phase detector / filter 208 that receives an input from a second transceiver 204.

The foregoing laser diode 202 emits light conducted to the test subject 20 through the optical fiber 216.

Single-sideband modulated signals pass through the brain and are phase-modulated by motion delays.

Light travels through the subject, is scattered, absorbed and received by the optical coupler / fiber assembly 218.

The received light is then sent to a detector 220 which is either a photomultiplier tube or a microchannel plate type as described above in another embodiment.

The output of the detector 220 is connected to the RF input of the J transceiver 204.

For example, the transceiver is used in SSB reception mode and a phase modulated 3KHz pass is obtained and connected to the phase detector filter 208.

The output is a 3KHz phase shift signal input to the second SSB transceiver.

The first transceiver 200 and the second transceiver 204 form a phase locked loop to ensure phase coupling.

The 3 KHz carrier waveform is also synchronized to 220 MHz by the frequency divider 206, thereby synchronizing the phases of 220 MHz and 3 KHz and accurately measuring the phase shift.

As shown in FIG. 4, the axial force of the transmitter oscillator 200 is divided into about 7 × 10 5 to generate a 3 KHz signal.

Thus, the output of the phase detector / filter 208 is associated with phase shift and represents absorption in the subject.

The carrier frequency is initially selected at 200 MHz.

It is high enough to detect phase shift during a reduction time of several nanoseconds.

However, it is small enough to be commercially available and present within the mixer's bandwidth.

Diode-ring mixers can easily be as high as 36GHz but have a significantly lower dynamic range than dynamic (transistor bridge or linear multiplier) designs.

Large dynamic ranges are difficult for this type of spectrophotometer system.

A heterodyne system is selected that transmits and retrieves multiple optical wavelengths in parallel at individual subcarrier frequencies, and performs phase detection within the frequency range of a commercially available phase sensitive detector, such as, for example, a lock-in amplifier.

Such devices have excellent noise figure, linearity, dynamic range, and accuracy of phase and amplitude.

Performance is also superior to other phase detectors operating directly at the RF carrier frequency.

Generation of the reference signal for the lock-in amplifier by frequency division from the RF frequency oscillator provides adequate phase coherence for all subcarriers and all modulated signals for the carrier.

No phase measurement between wavelengths is required.

Frequency generation by division also provides the possibility of minimum phase noise for a given main oscillator.

If additional carrier frequencies are required, such as measuring multiple exponential decays, it is sufficient for this design to add 1 to N RF switches and additional RF oscillators.

Laser diodes are chosen as heat sources because of their high radiation, ease of coupling with optical fibers, narrow output spectrum and wavelength stability, long lifetime and ease of modulation at RF frequencies.

Single sideband suppression carrier modulation is used to maximize the signal-to-noise ratio of the system and to prevent intermodulation distortion due to the nonlinearity of the laser.

The intermediate frequency is selected in the range of 10 to 100 kHz.

It must be high enough to realize the Q of the SSB filter. However, it should be small enough to be within the range of commercially available lock-in amplifiers.

The heat absorption source of the laser is preferably a Peltier cooler used for temperature control and feedback control.

Temperature control is necessary to stabilize the laser wavelength and to sufficiently tune the output wavelength, including the tongue usage of the diode (almost ± 10 nm).

However, post-demodulation detection of phase shift is excluded because no constant wavelength or amplitude is required.

The optical system includes a laser, a lens assembly that combines the laser light with the optical fiber, an optical fiber bundle (s) that transmits light from the subject to the subject, a fiber-to-test seat coupler at the ends of the optical fiber deval (s), and a photodetector assembly. The photohague consists of an isolator.

The isolator is necessary to prevent the optical feedback flowing into the laser cavity by reflection from the optical instrument or the test subject.

This estimate is very low at -60 Hz, but it is well known to cause amplitude and phase noise in the laser source.

The fiber for anchoring the bundles should be sufficiently small in dispersion since the uncertainty of the phase induced at the nominal modulation frequency of 100 MHz is much smaller than the phase shift of interest.

At the same time, it is desirable that the diameter of the core and the size of the legs be as large as possible.

This makes the laser-fiber coupling simple and tight and greatly increases the reflected light signal from the subject.

For this reason, single-mode fibers are excluded despite their unusual bandwidth-length.

For multimode fibers, formal dispersion is sufficient for the source, length and wavelengths considered here.

Therefore, waveguides and dispersions are ignored.

Considering the primary step-index fibers, simple radiation optics show that determining formal dispersion is the numerical value of the opening, not the aperture of the core.

For an acceptable time uncertainty of 100-picoseconds and a total step-dess fiber length of 2 m, an aperture of approximately 0.17 or less is required.

All commercially available Step-Indes multimodal fibers have large openings. However, the ideal modified index multimodal fiber has no formal dispersion of maximum radiation dose according to Fermat's principle, and the actual bandwidth-length of commercially available improved index fiber is 100 MHz at wavelengths for which this is the main concern for this study. exceeds -km Therefore, an improved index fiber with a core size of 100 microns, an aperture of 0.3 and a bandwidth-length of 100 MHz-km is chosen.

This fiber is also inexpensive enough to install bundles ($ 0.5 / m).

The detection optics consists of an optical bandpass or comb filter and the detector itself that only pass the fiber bundle, the laser wavelength at the intersection coupled to the active region of the detector and prevent saturation of the detector by room light.

Initially photomultipliers with GaAs (Cs) photocathodes are used.

This detector has a gain of about 30 compared to a PMT with a silicon photocathode and about 300 compared to a silicon yielding photodiode and does not include an area that is much smaller than the non-turbulent region of the yielding photodiode. However, photomultipliers also have extra bandwidth for this application and require the extraction of signals from the middle of the dynode row, thus reducing the gain.

If the signal-to-noise ratio is sufficient, the system can be easily modified by replacing the breakdown photodiode instead of PMT, reducing the allocation and increasing bandwidth, reliability and robustness.

The breakdown photodiode detector is also small and can be used for simulation because of its low cost.

In other words, ultra-small channel plate photomultipliers can be used if larger detection bandwidth and higher gain are required simultaneously.

The disadvantage here is that the cost is increased and the photocathode sensitivity of about 30 is lower than that of GaAs (Cs).

After a variable gain, the detector's output signal is fed back to the IF frequency by generating a heterodyne and fed to a two phase lock-in amplifier.

One lock-in amplifier is used per optical wavelength (IF frequency).

The cost increase increases the data recognition period by about 1.41 (2). And avoid unnecessary assumptions about relative absorption kinetics at each wavelength.

The entire system can be controlled using IBM-compatible computers with IEEE 4.888, analog-digital, digital-analog, and bidirectional digital interfaces.

All of this can be installed using low-cost plug-in cards for IBM.

Adjustments using a computer have important advantages in the possibility of automatic operation for testing and simulation, ease of reconstruction, speed, accuracy and post-processing of data as the system is improved, especially in static analysis.

By selecting a simple device, the clinical experiment of the present invention can be simplified.

The multi-wavelength phase modulation technique, each protected by its subcarrier, can be easily performed as illustrated in the preferred embodiment of FIG.

The output of the system then replaces the continuous wave technology of the system and at the same time has the advantage of demodulation logarithms such as the hemoglobin and the various states of the chitochrome.

The big advantage is that the optical path length is known and not assumed. Therefore, phase modulation is a convenient means of implementing TRS technology, and it can emphasize the 5 nanosecond delay time that the attenuation is exponential and includes long path shifts.

Claims (32)

Light source means coupled to an optical input port for inputting light into biological tissue and coupled to modulating means for emitting modulated light, optical detecting means for detecting light received by the optical detector, and spectroscopic examination of biological tissue A spectroscopic apparatus having a signal processing means comprising a phase detecting means provided for the light source, wherein the light source means and the modulating means are provided to emit light of a wavelength selected from a modulation waveform having a first modulation frequency of 100 MHz, The first modulation frequency has a time characteristic compatible with the time delay of photon migration from the optical input port to the optical detection port in a biological tissue; The optical input port is located a few centimeters away from the optical detection port; The optical detection means is provided to move from the input port to the detection port through a movement path in the inspection tissue to detect photons of the selected wavelength and provide a detected waveform; The phase detection means is provided to receive the detected waveform from the optical detection means, and compares the detected waveform with the modulated waveform to determine phase shift at the selected wavelength; Syrup phase shifting is indicative of the optical path length of photon shift in the examined tissue, and indicates scattering and absorption of the examined tissue. 2. The apparatus of claim 1, wherein said modulating means comprises a single sideband modulator and said phase detecting means comprises a lock-in amplifier. 3. The apparatus of claim 2, further comprising second oscillation means coupled to a wing single sideband modulator for generating a second waveform at a second frequency, wherein the single side modulator and the wing optical detection means detect the detected edge at an intermediate frequency. Apparatus, characterized in that it is provided to generate a waveform. 2. The apparatus of claim 1, further comprising: the wing modulator comprises a sideband modulator and the phase detection means comprises a lock-in amplifier. 5. The apparatus of claim 4, further comprising: second oscillating means coupled to said bilateral band modulator for generating a second waveform at a second frequency, wherein the two-wing bimodal modulator and said optical detection means are detected at an intermediate frequency. Apparatus, characterized in that it is provided to generate a waveform. 6. The apparatus of claim 3 or 5, wherein the intermediate frequency is in the range of 3 KHz to 100 KHz. 2. The apparatus according to claim 1, wherein said modulating means and said phase detecting means further receive said detected waveform from said first transceiver means and said detecting means for generating said waveform of said first frequency and said first transceiver means. And second transceiver means for receiving a waveform from the waveform to provide a phase lock loop. 8. The apparatus of claim 7, wherein the first transceiver means and the second transceiver means operate as a single sideband modulator. 9. An apparatus according to claim 8, further comprising a frequency separator which is page locked to said first transceiver means and said phase detecting means generates an intermediate frequency for detecting said phase shift. 2. The apparatus of claim 1, wherein the light source means is further provided to emit a second wavelength in the visible to ultraviolet region, wherein the emitted light is modulated in a waveform having the first modulation frequency; The optical detecting means is further provided to move from the input port to the detection port through a movement path in the inspection tissue to detect photons of the selected second wavelength and provide the detected second waveform; And said phase detecting means is further configured to compare said detected waveform with said modulated second waveform to determine phase shift at a selected second wavelength. The optical modulation system of claim 1, wherein the modulator is configured to generate a second modulation waveform having a second modulation frequency, the second modulation frequency being from the optical input port to the optical detection port in a biological tissue. It has a time characteristic that is compatible with time delay; The optical means emits light modulated by the waveform of the second modulation frequency; The optical detecting means is provided to detect photons at the second frequency moved down from the input port to the detection port at a second modulation frequency through a movement path in the test tissue; And said phase detecting means is adapted to compare the detected second waveform with said modulated second waveform to dig a phase at said second modulation frequency. 11. The apparatus of claim 10, wherein the wing wavelength is sensitive to absorbing pigments, and the apparatus further comprises a processor for determining the ratio of phase shift determined at the two wavelengths. The device of claim 12, wherein the pigment is hemoglobin and the processor further determines the hemoglobin concentration in the tissue based on the rate of phase shift. 8. The apparatus of any one of claims 1, 2, 4 and 7, wherein the first modulation frequency is in the range of 144 MHz to 440 MHz. 8. The apparatus of any one of claims 1, 2, 4 and 7, wherein the first modulation frequency is about 220 MHz. 8. Apparatus according to any one of claims 1, 2, 4 and 7, wherein said processing means provides scattering and absorptive indications of the masonry inspected from said phase shift. 8. Apparatus according to any one of claims 1, 2, 4 and 7, wherein said processing means provides scattering and absorptive indications of the tissue examined from the amplitude of the detected signal. A method of spectroscopy in which the optical path length of a photon shift through a photon migration path in a biological tissue located between an optical input port and an optical detection port of the spectrometer is determined by scattering and absorption properties of a tissue part to be examined. Positioning the optical input port a few centimeters from the optical detection port; Generating a modulation waveform having a first modulation frequency of 100 MHz having a time characteristic compatible with the time delay of photon migration from the optical input port to the optical detection port in a biological tissue; Imposing the waveform on light source means to produce light of a selected wavelength modulated by the waveform; Injecting the modulated light of the wavelength into the tissue from the input port; Detecting, by the detection unit, photons of the wavelengths moving from the input port to the tissues through a movement path; Supplying a detected waveform corresponding to the detected photons of the wavelength to phase detection means; Comparing the detected waveforms with the modulated waveforms to determine a phase shift indicating a photon-shift optical path length in the examined tissue and indicating scattering and absorptivity of the examined tissue. 19. The method of claim 18, wherein said generating step uses a single side modulator to generate said waveform. 19. The method of claim 18, wherein said generating step uses a bilateral side modulator to generate said waveform. 21. The method of claim 19 or 20, further comprising generating at a second frequency, wherein the supplying step comprises generating a detected waveform at an intermediate frequency based on the second frequency, wherein the comparing step comprises: Characterized in that it is performed at the intermediate frequency. 19. The method of claim 18, wherein generating and supplying comprises using a transmitter to generate the waveform and a receiver to receive the detected waveform, wherein the transmitter and the receiver form a page look loop. Characterized in that the method. 22. The method of claim 21, further comprising generating a second waveform at a second frequency, wherein the supplying step comprises generating a waveform detected at an intermediate frequency, wherein the comparing step is performed at the intermediate frequency. Characterized in that the method. 19. The method of claim 18, further comprising: generating a modulation waveform having a second frequency in the biological tissue having a time characteristic compatible with the time delay of photon migration from the optical input port to the optical detection port; Imposing the modulation waveform on a light source means to generate light of a selected wavelength modulated by the waveform of the second modulation frequency; Injecting the modulated light of the wavelength having the waveform of the second modulation frequency into the tissue from the input port; Detecting, by the detection unit, photons of the wavelength moved from the input port through a movement path in the tissue; Supplying the detected waveform corresponding to the detected photons of the wavelength modulated by the waveform of the second modulation frequency to the phase detecting means; Comparing the detected waveform with the modulated waveform to determine a phase shift indicating an optical path length of photon shift in the tissue examined at the second modulation frequency and indicating scattering and absorptivity of the examined tissue; How to Bounty to Include. 19. The method of claim 18, further comprising: generating light of a second wavelength; Imposing the modulating waveform on the light of the second wavelength; Injecting light of the second wavelength into the tissue to be examined alternately with light of the first wavelength; Detecting, by the detector, a photon of the second wavelength that has moved within the tissue from the input port; Supplying a detected waveform corresponding to the detected photons of the second wavelength to the phase detecting means; A second phase indicating scattering and absorptivity of the examined tissue corresponding to the optical path length of the photon shift of the second wavelength in the examined tissue by comparing the detected waveform of the second wavelength with the modulated waveform Determining the movement. 23. The method according to any one of claims 18, 19, 20 and 22, wherein the vane wavelength is in the visible to ultraviolet wavelength range. 27. The method of claim 25, wherein said wavelength is sensitive to absorbing pigments, and said apparatus further comprises a processor for determining the ratio of said phase shift determined at said two wavelengths. The method of claim 27, wherein the pigment is hemoglobin and the processor further determines the hemoglobin concentration in the tissue based on the rate of phase shift. 23. The method of any of claims 18, 19, 20 and 22, wherein the first modulation frequency is in the range of 144 MHz to 440 MHz. 23. The apparatus of any of claims 18, 19, 20, and 22, wherein the first modulation frequency is about 220 MHz. 23. The method of any one of claims 18, 19, 20 and 22, wherein said processing means provides scattering and absorptive indications of tissue examined from said phase shift. 23. An apparatus according to any one of claims 18, 19, 20 and 22, wherein said processing means provides scattering and absorptive indications of tissue examined from the amplitude of said detected signal.