Patent Publication Number: US-2021161408-A1

Title: Scattered light signal measuring apparatus and information processing apparatus

Description:
TECHNICAL FIELD 
     The present technology relates to a scattered light signal measurement apparatus and an information processing apparatus that are capable of measuring information related to the velocity of a group of particles such as red blood cells in blood. 
     BACKGROUND ART 
     Various systems have been developed for a blood flow velocity sensor that measures the velocity of a blood flow flowing through a blood vessel. For example, Patent Literature 1 has disclosed a laser Doppler blood flow meter. 
     The laser Doppler blood flow meter emits laser light to a living body. The laser light is scattered by biological tissues. Scattering by particles such as red blood cells that move in a blood flow causes a shift (Doppler shift) of a wavelength due to the Doppler effect as compared to scattering by other tissues that do not move. 
     Therefore, it is possible to measure a motion velocity of particles in blood such as red blood cells, i.e., a blood flow velocity by detecting scattered light to obtain a signal change related to such a Doppler shift. Further, it is also possible to measure a velocity of a group of particles that move in a measurement target object other than the particles in blood by using this measurement principle. 
     CITATION LIST 
     Patent Literature 
     Patent Literature 1: Japanese Patent Application Laid-open No. 2017-192629 
     DISCLOSURE OF INVENTION 
     Technical Problem 
     However, in the laser Doppler blood flow meter as described in Patent Literature 1, a body motion (motion of a living body to which a sensor is attached) may cause a position shift or a relative velocity between the living body and the sensor. Since the Doppler shift caused by the motion of the particles in blood as described above is used for measurement in the laser Doppler blood flow meter, if a position shift or a relative velocity is caused between the living body and the sensor, this Doppler shift is affected and noise is generated in the measurement result. 
     In view of the above-mentioned circumstances, it is an object of the present technology to provide a scattered light signal measurement apparatus and an information processing apparatus that are capable of performing accurate measurement while preventing noise due to changes in relative velocity and relative position caused between a sensor and an observation target. 
     Solution to Problem 
     In order to accomplish the above-mentioned object, a scattered light signal measurement apparatus according to the present technology includes a light receiving element and an incident angle limiting unit. 
     The light receiving element receives scattered light obtained when coherent light emitted to a measurement target is scattered by the measurement target. 
     The incident angle limiting unit limits an incident angle of the scattered light that is incident on a single point of the light receiving element to be equal to or smaller than a predetermined angle and controls the single light receiving element to increase an aperture ratio. 
     In accordance with this configuration, the incident angle of scattered light of the scattered light scattered by the measurement target, which is incident on the single point of the light receiving element, is limited to be equal to or smaller than the predetermined angle. If a relative velocity caused by the body motion between the light receiving element and the measurement target is caused, noise can be generated due to the Doppler beat caused between scattered light beams having different incident angles on the light receiving element. However, since the incident angle of the scattered light that is incident on the single point of the light receiving element is limited to be equal to or smaller than the predetermined angle in the above-mentioned configuration, and thus the frequency of the Doppler beat between the scattered light beams having the incident angles different from each other is reduced. Accordingly, it is possible to detect only the Doppler beat caused by the motion of particles at the measurement target and to correctly calculate the motion of the particles at the measurement target. 
     The incident angle limiting unit may limit a light beam diameter of the scattered light that is incident on the single point of the light receiving element to be smaller than an entire light beam diameter that is incident on the light receiving element. 
     The incident angle limiting unit may concentrate the scattered light at a different point of the light receiving element in accordance with an incident angle. 
     The incident angle limiting unit may be a lens array in which a plurality of lenses is arranged. 
     The light receiving element may be at least one, and 
     the plurality of lenses may concentrate scattered light at an incident angle limited by each of a plurality of lenses to the single light receiving element. 
     It may further include a light shield that shields light incident on a portion between the plurality of lenses. 
     The incident angle limiting unit is a lens, and the scattered light signal measurement apparatus may further include a light shield that prevents mixing of light incident on the lens and light incident on a portion other than the lens. 
     The light receiving element may be at least one, and the incident angle limiting unit that sets a plurality of pin holes with respect to the single light receiving element, limits the incident angle of the scattered light that is incident on the single point of the light receiving element to be equal to or smaller than the predetermined angle, and concentrates scattered light having an amount larger than an amount of scattered light, which is concentrated by a single pin hole, to the light receiving element. 
     The scattered light signal measurement apparatus may further include: a coherent light source that emits the coherent light; and an irradiation light control unit that controls an irradiation diameter of the coherent light emitted from the coherent light source. 
     The irradiation light control unit may make the coherent light collimated, the coherent light being emitted from the light source. 
     The irradiation diameter may be 0.5 mm or more and 2 mm or less. 
     The scattered light signal measurement apparatus may further include: a first coherent light source that emits first coherent light having a first wavelength range to the measurement target; and a second coherent light source that emits second coherent light having a second wavelength range different from the first wavelength range to the measurement target. 
     The scattered light signal measurement apparatus may further include a light receiving element array in which a plurality of light receiving elements is arranged. 
     The incident angle limiting unit may include one louver layer or a plurality of louver layers, the louver layer may include a plurality of louvers extending in parallel to each other, and in a case of including the plurality of louver layers, the incident angle limiting unit may be an optical path limiting filter in which directions in which the respective louvers of the layers extend cross each other. 
     The scattered light signal measurement apparatus may further include a polarization filter disposed in any optical path between the measurement target and the light receiving element. 
     The coherent light source and the irradiation light control unit may make the coherent light incident on the measurement target in a direction inclined with respect to an optical path of scattered light travelling toward the light receiving surface from the measurement target. 
     The irradiation light control unit may shape the coherent light into an elliptical shape and emits the shaped coherent light such that an irradiation spot of the coherent light at the measurement target has a circular shape. 
     In order to accomplish the above-mentioned object, an information processing apparatus according to the present technology includes: a light receiving element that receives scattered light obtained when coherent light emitted to a measurement target is scattered by the measurement target; an incident angle limiting unit that limits an incident angle of the scattered light that is incident on a single point of the light receiving element to be equal to or smaller than a predetermined angle and controls the single light receiving element to increase an aperture ratio; and a particle velocity information calculation unit. 
     The particle velocity information calculation unit calculates particle velocity information that is information regarding velocity of a particle inside the measurement target on the basis of a signal output from the light receiving element. 
     The information processing apparatus may further include a noise cancellation unit that detects a relative motion between the sensor and the measurement target on the basis of an output of a light receiving element array in which a plurality of light receiving elements provided in the sensor is arranged, and cancels noise caused by the relative motion between the sensor and the measurement target on the basis of the particle velocity information. 
     The measurement target may be blood, and the particle velocity information calculation unit may calculate blood flow velocity information that is information regarding velocity of a red blood cell inside the measurement target. 
     The information processing apparatus may include a wearable device. 
     Advantageous Effects of Invention 
     As described above, in accordance with the present technology, it is possible to provide a scattered light signal measurement apparatus and an information processing apparatus that are capable of performing accurate measurement while preventing noise due to changes in relative velocity and relative position caused between a sensor and an observation target. 
    
    
     
       BRIEF DESCRIPTION OF DRAWINGS 
         FIG. 1  A schematic diagram showing a Doppler shift and a beat phenomenon caused by the Doppler shift. 
         FIG. 2  A schematic diagram showing the principle of a laser Doppler blood flow measurement apparatus. 
         FIG. 3  A schematic diagram showing a Doppler shift due to a relative motion between a light emitting unit of the laser Doppler blood flow measurement apparatus and a scatterer. 
         FIG. 4  A schematic diagram showing a Doppler shift due to a motion of the scatterer and a beat signal caused by it. 
         FIG. 5  A schematic diagram showing a Doppler shift due to a motion of the scatterer and a beat signal caused by it. 
         FIG. 6  A schematic diagram showing a Doppler shift due to a relative motion between the light receiving unit of the laser Doppler blood flow measurement apparatus and the scatterer. 
         FIG. 7  A schematic diagram showing a Doppler shift due to a relative motion between the light receiving unit of the laser Doppler blood flow measurement apparatus and two scatterers. 
         FIG. 8  A diagram showing a velocity distribution of red blood cells set by the laser Doppler blood flow measurement apparatus as a measurement target. 
         FIG. 9  A schematic diagram of a laser Doppler blood flow measurement apparatus having a general configuration. 
         FIG. 10  An example of a beat signal output from a light receiving unit of the blood flow measurement apparatus. 
         FIG. 11  An example of a signal processing result of the beat signal output from the light receiving unit of the blood flow measurement apparatus. 
         FIG. 12  A schematic diagram of the velocity distribution observed in a case where disturbance noise due to the body motion is not generated in the blood flow measurement apparatus. 
         FIG. 13  A schematic diagram showing a Doppler beat due to a relative velocity between a light emitting unit of the blood flow measurement apparatus and a living body. 
         FIG. 14  A schematic diagram showing the Doppler beat the relative velocity between the light receiving unit of the blood flow measurement apparatus and the living body. 
         FIG. 15  A schematic diagram showing a Doppler beat due to an angular velocity between the light emitting unit and the light receiving unit of the blood flow measurement apparatus and the living body. 
         FIG. 16  A schematic diagram of a blood flow measurement apparatus according to a first embodiment of the present technology. 
         FIG. 17  A block diagram showing functions of the blood flow measurement apparatus. 
         FIG. 18  A schematic diagram of a light emitting unit and a light receiving unit of the blood flow measurement apparatus. 
         FIG. 19  A schematic diagram of an incident angle limiting unit of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 20  A schematic diagram showing effects of the blood flow measurement apparatus. 
         FIG. 21  A schematic diagram showing effects of the blood flow measurement apparatus. 
         FIG. 22  A schematic diagram showing effects of the blood flow measurement apparatus. 
         FIG. 23  A schematic diagram showing effects of the blood flow measurement apparatus. 
         FIG. 24  A schematic diagram showing effects of the blood flow measurement apparatus. 
         FIG. 25  A schematic diagram showing another configuration of the light emitting unit of the blood flow measurement apparatus. 
         FIG. 26  A schematic diagram showing another configuration of the light emitting unit of the blood flow measurement apparatus. 
         FIG. 27  A schematic diagram showing another configuration of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 28  A schematic diagram showing another configuration of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 29  A schematic diagram showing another configuration of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 30  A schematic diagram showing another configuration of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 31  A schematic diagram showing another configuration of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 32  A schematic diagram showing another configuration of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 33  A schematic diagram showing another configuration of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 34  A schematic diagram showing another configuration of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 35  A schematic diagram showing another configuration of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 36  A block diagram showing a hardware configuration of an information processing apparatus of the blood flow measurement apparatus. 
         FIG. 37  A schematic diagram of a blood flow measurement apparatus according to a second embodiment of the present technology. 
         FIG. 38  A block diagram showing functions of the blood flow measurement apparatus. 
         FIG. 39  A cross-sectional view of a light emitting unit and a light receiving unit of the blood flow measurement apparatus. [ FIG. 40 ] A plan view of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 41  A schematic diagram showing another configuration of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 42  A block diagram showing a hardware configuration of an information processing apparatus of the blood flow measurement apparatus. 
         FIG. 43  A schematic diagram of a blood flow measurement apparatus according to a third embodiment of the present technology. 
         FIG. 44  A block diagram showing functions of the blood flow measurement apparatus. 
         FIG. 45  A cross-sectional view of a light emitting unit and a light receiving unit of the blood flow measurement apparatus. 
         FIG. 46  A plan view of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 47  A block diagram showing a hardware configuration of an information processing apparatus of the blood flow measurement apparatus. 
         FIG. 48  A schematic diagram of a blood flow measurement apparatus according to a fourth embodiment of the present technology. 
         FIG. 49  A block diagram showing functions of the blood flow measurement apparatus. 
         FIG. 50  A cross-sectional view of a light emitting unit and a light receiving unit of the blood flow measurement apparatus. 
         FIG. 51  A plan view of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 52  A schematic diagram showing another configuration of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 53  A block diagram showing a hardware configuration of an information processing apparatus of the blood flow measurement apparatus. 
         FIG. 54  A cross-sectional view of a light receiving unit of a blood flow measurement apparatus according to an embodiment of the present technology. 
         FIG. 55  A schematic diagram of a louver filter of a light receiving unit of the blood flow measurement apparatus. 
         FIG. 56  A schematic diagram of the louver filter of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 57  A cross-sectional view of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 58  A cross-sectional view of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 59  A cross-sectional view of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 60  A cross-sectional view of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 61  A schematic diagram of the louver filter of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 62  A schematic diagram of the louver filter of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 63  A schematic diagram of the louver filter of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 64  A cross-sectional view of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 65  A cross-sectional view of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 66  A cross-sectional view of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 67  A plan view of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 68  A cross-sectional view of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 69  A plan view of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 70  A cross-sectional view of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 71  A cross-sectional view of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 72  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 73  A schematic diagram of the laser light emitted from the light emitting unit of the blood flow measurement apparatus. 
         FIG. 74  A schematic diagram of the laser light emitted from the light emitting unit of the blood flow measurement apparatus. 
         FIG. 75  A schematic diagram of the light emitting unit of the blood flow measurement apparatus. 
         FIG. 76  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 77  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 78  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 79  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 80  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 81  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 82  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 83  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 84  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 85  A schematic diagram of the light receiving unit of the blood flow measurement apparatus. 
         FIG. 86  A schematic diagram of the laser light emitted from the light emitting unit of the blood flow measurement apparatus. 
         FIG. 87  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
         FIG. 88  A schematic diagram of the light emitting unit of the blood flow measurement apparatus. 
         FIG. 89  A schematic diagram of the light emitting unit and the light receiving unit of the blood flow measurement apparatus. 
     
    
    
     MODE(S) FOR CARRYING OUT THE INVENTION 
     The scattered light signal measurement apparatus according to the present technology is an apparatus that detects a velocity of a group of particles and a beat signal caused by interference between coherent light beams by using the laser Doppler phenomenon and measures a flow rate per unit volume of the group of particles and a velocity distribution by using the laser Doppler phenomenon. The measurement target particles are typically particles in blood, such as red blood cells flowing through a blood vessel. In a case where the scattered light signal measurement apparatus according to the present technology is used for observation of motions of particles in blood, an amount of moving particles in blood in an observation region proportional to the blood flow rate per hour of a tissue and a velocity distribution of particles in blood in the observation region (only a velocity along an observation direction axis) can be measured. Hereinafter, a blood flow measurement apparatus according to an embodiment of the present technology will be described. It should be noted that in this specification, red blood cells are used as a generic term for particles in blood that scatter laser light, the particles in blood including red blood cells. 
     [Regarding Principle of Laser Doppler Blood Flow Measurement Apparatus] 
     The laser Doppler blood flow measurement apparatus is a measurement apparatus using a Doppler shift of light. Hereinafter, the principle of the laser Doppler blood flow measurement apparatus will be described. It should be noted that regarding the Doppler shift phenomenon dealt with in the present invention, the relativistic effect is negligibly small because a relative velocity between target objects is sufficiently small with respect to a light velocity. Therefore, the relativistic effect is omitted from this principle description. Further, in reality, the refractive index differs from the refractive index of the sensor portion inside the human body. Therefore, a light velocity c, which determines an amount of Doppler shift, is different (or a wavelength A is different). However, the difference in refractive index has little effect on the principle of the present invention. Therefore, the effect of the refractive index is omitted from this principle description.  FIG. 1  is a schematic diagram showing a Doppler shift and a Doppler beat that is a phenomenon that occurs due to interference between Doppler-shifted waves and original waves. 
     It is assumed that after a wave at a frequency f 0  reaches a measurement point as shown in FIG.  1 ( a ), the frequency shifts due to the Doppler phenomenon and the frequency of the wave becomes f 0 +Δf as shown in  FIG. 1( b ) . 
     It is known that in this case, both the interfering waves have a beat at a frequency Δf as shown in  FIG. 1( c ) . This wave at the frequency Δf is referred to as a Doppler beat. Using such a beat phenomenon, a velocity of a movable object, which causes a Doppler shift, relative to a measurement point can be obtained by measuring a frequency Δf of a Doppler beat even without measuring a frequency f 0 +Δf at which the Doppler shift is caused. 
       FIG. 2  is a schematic diagram showing the principle of the laser Doppler blood flow measurement apparatus. As shown in the figure, laser light L is emitted from a light emitting unit  11 . When the laser light L is incident on a scatterer M, the laser light L is reflected by the scatterer M and scattered light S is generated. The scattered light S is incident on a light receiving unit  12  and is detected. 
       FIG. 3  is a schematic diagram showing a Doppler shift due to a relative motion between the light emitting unit  11  and the scatterer M. As shown in the figure, the scatterer M is irradiated with laser light L at a frequency f 0  from the light emitting unit  11 . The scatterer M moves at a velocity v in a direction to form an angle θ 1  with respect to an optical axis A of the laser light L. 
       FIG. 3  is a diagram in which the scatterer M is moving relative to the light emitting unit  11  that is stationary. In this case, irradiation light received by the moving scatterer M causes a Doppler shift to the scatterer M. In addition, when the scattered light emitted by the moving scatterer M is observed at a stationary measurement point, a Doppler shift occurs. In this case, a Doppler shift Δf, viewed from the scatterer M is expressed by (Expression 1-1) below. 
     
       
         
           
             
                 
             
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     In (Expression 1-1), (Expression 1-2), and (Expression 1-3), c is a light velocity, λ 0  is a wavelength of the laser light L, and α 0  is a proportional constant depending on the wavelength (λ 0 ) of the laser light L. As shown in Expression 1-1, the Doppler shift Δf is proportional to the relative velocity between the light emitting unit  11  and the scatterer M and varies in a manner that depends on the angle θ 1 . 
     Further, the Doppler shift Δf viewed from stationary coordinates depends on the emission direction of the scattered light S and varies in a manner that depends on an angle θ 2 . 
       FIGS. 4 and 5  are schematic diagrams each showing a Doppler shift due to a motion of a scatterer. As shown in  FIG. 4( a ) , it is assumed that a scatterer M A  is moving in light emitted from a light source. In the figure, a velocity vector of the scatterer M A  is shown as the arrow extended from the scatterer M A . 
       FIG. 4( b )  shows a power spectrum S A 2 of the light scattered by the moving scatterer M A  and a power spectrum S A 1 of the light scattered by the stationary scatterer M A . 
     As shown in the figure, the Doppler effect due to a motion of the scatterer M A  results in a Doppler shift Δf, and the frequency of the scattered light S A 2 increases from the original frequency f 0 . Further, when the scatterer M A  moves away from the light source, a Doppler shift occurs such that the frequency of the scattered light S A 2 decreases. 
       FIG. 4( c )  is a power spectrum S A 3 of a Doppler beat caused by interference between scattered light S A 1 and scattered light S A 2. A frequency of S A 3 is an absolute value of a difference between the frequency of 
     S A 1 and the frequency of S A 2. A motion velocity of the scatterer M A  can be determined on the basis of a peak frequency or mean frequency of this power spectrum S A 3. It should be noted that in order to simultaneously receive the scattered light S A 1 and the scattered light S A 2 for generating a Doppler beat, it is sufficient that a fixed scatterer is additionally prepared in order to direct the light from the light source to the light receiving unit. 
     Next, as shown in  FIG. 5( a ) , it is assumed that many scatterers M B  are moving in various directions in the light emitted from the light source. In the figure, velocity vectors of the scatterers M B  are shown as the arrows extended from the scatterers M B . The scatterers M B  with no arrows indicate that they are stationary. 
       FIG. 5( b )  shows a power spectrum S B 1 of light scattered by the scatterer M B  that is stationary and a power spectrum S B 2 of light scattered by the moving scatterer M B . For a blood flow measurement example, the stationary scatterer M B  is a body tissue or the like. The moving scatterer M B  is a red cell or the like. 
     In this case, a motion velocity of the respective scatterers M B  relative to the light source is different. Therefore, as shown in  FIG. 5( b ) , the power spectrum S B 2 of the scattered light is a composite wave of the Doppler shift due to the many scatterers M B . In a case where the ratio of the amount of stationary scatterers is much higher than the ratio of the amount of moving scatterers, the scattered light of the stationary scatterers M B  has much higher intensity than that of the scattered light of the moving scatterers M B . In this case, most of the interference between the scattered light causing the Doppler beat is interference between the scattered light of the stationary scatterers M B  and the scattered light of the moving scatterers M B . 
       FIG. 5( c )  is a power spectrum S B 3 of a Doppler beat caused by interference between scattered light S B 1 and scattered light S B 2. A frequency of S B 3 is an absolute value of a difference between the frequency of S B 1 and the frequency of S B 2. The shape of this power spectrum shows a velocity distribution of the scatterers M B  relative to the light source. Therefore, an actual velocity distribution of the scatterers M B  can be determined by considering a decrease in Doppler shift amount in a moving direction of each scatterer M B . 
     It should be noted that a technique of measuring this Doppler shift due to the many observed objects moving in the respective directions is called dynamic light scattering (DLS). 
     In the description so far, the Doppler shift caused by the relative motion of the scatterer M relative to the light emitting unit  11  has been described. Next, a Doppler shift caused by a relative motion of the light receiving unit  12  relative to the scatterer M will be described. 
       FIG. 6  is a schematic diagram showing a Doppler shift due to the relative motion of the scatterer M relative to the light receiving unit  12 . As shown in the figure, the scatterer M is irradiated with the laser light L at a frequency f 0  from the light emitting unit  11 . The light receiving unit  12  is moving at the velocity v in a direction to form an angle θ with respect to an optical axis B of the scattered light S. 
     A Doppler shift Δf caused in the scattered light S incident on the light receiving unit  12  in this case is expressed by (Expression 2) below. 
     
       
         
           
             [ 
             
               Formula 
                
               
                   
               
                
               4 
             
             ] 
           
         
       
       
         
           
             
               
                 
                   
                     
                       
                         
                           Δ 
                            
                           
                               
                           
                            
                           f 
                         
                          
                           
                         = 
                         
                           
                             v 
                             c 
                           
                            
                           
                               
                           
                            
                           f 
                            
                           
                               
                           
                            
                           cos 
                            
                           
                               
                           
                            
                           θ 
                         
                       
                     
                   
                   
                     
                       
                           
                          
                         
                           ≃ 
                           
                             
                               v 
                               c 
                             
                              
                             
                                 
                             
                              
                             
                               f 
                               0 
                             
                              
                             
                                 
                             
                              
                             cos 
                              
                             
                                 
                             
                              
                             θ 
                              
                             
                                 
                             
                              
                             
                               ( 
                               
                                 small 
                                  
                                 
                                     
                                 
                                  
                                 quadratic 
                                  
                                 
                                     
                                 
                                  
                                 term 
                                  
                                 
                                     
                                 
                                  
                                 is 
                                  
                                 
                                     
                                 
                                  
                                 ignored 
                               
                               ) 
                             
                           
                         
                       
                     
                   
                   
                     
                       
                           
                          
                         
                           = 
                           
                             
                               v 
                                
                               
                                   
                               
                                
                               cos 
                                
                               
                                   
                               
                                
                               θ 
                             
                             
                               λ 
                               0 
                             
                           
                         
                       
                     
                   
                   
                     
                       
                           
                          
                         
                           = 
                           
                             
                               α 
                               0 
                             
                              
                             
                                 
                             
                              
                             v 
                              
                             
                                 
                             
                              
                             cos 
                              
                             
                                 
                             
                              
                             θ 
                           
                         
                       
                     
                   
                 
               
               
                 
                   ( 
                   
                     Expression 
                      
                     
                         
                     
                      
                     2 
                   
                   ) 
                 
               
             
           
         
       
     
     In Expression 2, c is a light velocity, λ 0  is a wavelength of the laser light L, f is a frequency of scattered light emitted in a direction of the optical axis B, and α 0  is a proportional constant depending on the wavelength (λ 0 ) of laser light L. The scattered light S is emitted in all directions. The scattered light S along a straight line connecting the scatterer M and the light receiving unit  12  reaches the light receiving unit  12 . 
     The Doppler shift Δf is proportional to the relative velocity between the scatterer M and the light receiving unit  12  and varies in a manner that depends on the angle θ with respect to the optical axis B. 
     Next, a Doppler shift due to relative motions of two scatterers relative to the light receiving unit  12  will be described.  FIG. 7  is a schematic diagram showing a Doppler shift due to a relative pair motion the scatterer M 1  and the scatterer M 2  relative to the light receiving unit  12 . 
     It is assumed that the laser light L having the frequency f 0  from the light emitting unit  11  is emitted to the scatterer M 1  and the scatterer M 2  and the light receiving unit  12  is moving at the velocity v in a direction to form the angle θ with respect to a center axis (dashed line) of an incident angle difference ϕ between scattered light S 1  and scattered light S 2  on the light receiving unit  12 . It is assumed that the frequency of light of the scattered light emitted from the scatterer M 1 , which is emitted in the direction of the light receiving unit  12 , is f 1  and the Doppler shift amount is Δf 1  and the frequency of light of the scattered light emitted from the scatterer M 2 , which is emitted in the direction of the light receiving unit  12 , is f 2  and the Doppler shift amount is Δf 2 . 
     In this case, a frequency F 1  of the scattered light S 1  incident on the light receiving unit  12  from the scatterer M 1  is expressed by the following (Expression 3). 
     
       
         
           
             [ 
             
               Formula 
                
               
                   
               
                
               5 
             
             ] 
           
         
       
       
         
           
             
               
                 
                   
                     ( 
                     
                       
                         λ 
                         0 
                       
                       = 
                       
                         c 
                         
                           f 
                           0 
                         
                       
                     
                     ) 
                   
                    
                   
                     
 
                   
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                     ( 
                     
                       
                         α 
                         0 
                       
                       = 
                       
                         1 
                         
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                           0 
                         
                       
                     
                     ) 
                   
                    
                   
                     
 
                   
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                               1 
                             
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                                       ) 
                                     
                                   
                                 
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                                  
                                 
                                     
                                 
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                                   ( 
                                   
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                                     + 
                                     
                                       
                                         1 
                                         2 
                                       
                                        
                                       φ 
                                     
                                   
                                   ) 
                                 
                               
                             
                           
                         
                       
                     
                     
                       
                         
                             
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                             = 
                             
                               
                                 f 
                                 0 
                               
                               + 
                               
                                 Δ 
                                  
                                 
                                     
                                 
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                                   f 
                                   1 
                                 
                               
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                                   α 
                                   0 
                                 
                                  
                                 
                                     
                                 
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                                 v 
                                  
                                 
                                     
                                 
                                  
                                 cos 
                                  
                                 
                                     
                                 
                                  
                                 
                                   ( 
                                   
                                     θ 
                                     + 
                                     
                                       
                                         1 
                                         2 
                                       
                                        
                                       φ 
                                     
                                   
                                   ) 
                                 
                               
                             
                           
                         
                       
                     
                   
                 
               
               
                 
                   ( 
                   
                     Expression 
                      
                     
                         
                     
                      
                     3 
                   
                   ) 
                 
               
             
           
         
       
     
     In (Expression 3), Δf 1  is a Doppler shift caused in the scattered light S 1  travelling toward to the light receiving unit  12  and a is a proportional constant depending on the wavelength of the light source. (Expression 3) ignores the small quadratic term “Δf 1 (v/c)cos(θ+(½))” in the middle. 
     Further, a frequency F 2  of scattered light S 2  incident on the light receiving unit  12  from the scatterer M 2  is expressed by the following (Expression 4). 
       [Formula 6] 
         F   2   ≅f   0   +Δf   2 +α 0   v  cos(θ−½ϕ)  (Expression 4)
 
     In (Expression 4), Δf 2  is a Doppler shift caused in the scattered light S 2  travelling toward the light receiving unit  12  and α 0  is a proportional constant depending on the wavelength λ 0  of the light source. 
     From the above, a Doppler shift ΔF of the scattered light S 1  and the scattered light S 2  is expressed by (Expression 5) below. 
     
       
         
           
             [ 
             
               Formula 
                
               
                   
               
                
               7 
             
             ] 
           
         
       
       
         
           
             
               
                 
                   
                     β 
                     = 
                     
                       sin 
                        
                       
                         ( 
                         
                           
                             1 
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                         ) 
                       
                     
                   
                    
                   
                     
 
                   
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                         0 
                       
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                           = 
                           
                             
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                               1 
                             
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                               2 
                             
                           
                         
                       
                     
                     
                       
                         
                             
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                                   0 
                                 
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                                           1 
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                                       ) 
                                     
                                   
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                             = 
                             
                               
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                                   f 
                                   1 
                                 
                               
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                                   2 
                                 
                               
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                                   0 
                                 
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                                  
                                 
                                     
                                 
                                  
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                                  
                                 
                                     
                                 
                                  
                                 sin 
                                  
                                 
                                     
                                 
                                  
                                 
                                   ( 
                                   
                                     
                                       1 
                                       2 
                                     
                                      
                                     φ 
                                   
                                   ) 
                                 
                               
                             
                           
                         
                       
                     
                     
                       
                         
                             
                            
                           
                             = 
                             
                               
                                 ( 
                                 
                                   
                                     Δ 
                                      
                                     
                                         
                                     
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                                       f 
                                       1 
                                     
                                   
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                                      
                                     
                                         
                                     
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                                 ) 
                               
                               + 
                               
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                                  
                                 
                                     
                                 
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                                   f 
                                   3 
                                 
                               
                             
                           
                         
                       
                     
                   
                 
               
               
                 
                   ( 
                   
                     Expression 
                      
                     
                         
                     
                      
                     5 
                   
                   ) 
                 
               
             
           
         
       
     
     In (Expression 5), Δf 1 −Δf 2  is a frequency observed when the light receiving unit  12  is stationary and Δf 3  is a frequency that is added due to a motion of the light receiving unit  12 . In the expression showing Δf 3 , α 0 v is a fundamental frequency of the Doppler beat due to the relative motion velocity vector and β is a relative-velocity beat-frequency decrease factor determined in a manner that depends on an incident angle difference φ between the scattered light S 1  and the scattered light S 2  on the light receiving unit  12 . 
     In a case where the incident angle difference φ is sufficiently small, β is close to 0 and Δf 3  is sufficiently small as compared to (Δf 1 −Δf 2 ). At this time, a distribution of the frequency difference Δf 1 −Δf 2  caused by the blood flow velocity can be accurately measured. On the other hand, as the incident angle difference φ is larger, β is further from 0 and Δf 3  can include a frequency that cannot be distinguished from the frequency (Δf 1 −Δf 2 ). In particular, in interference between many stationary scatterers, (©f 1 −Δf 2 ) is 0. Therefore, when the light receiving unit is stationary, Δf 3  is 0 and thus no Doppler beat is caused. However, when the light receiving unit moves and φ is large, Δf 3  can take a value similar to the blood flow velocity. Therefore, when the light receiving unit moves and φ is large, the distribution of the frequency difference Δf 1 −Δf 2  caused by the blood flow velocity cannot be accurately measured. 
     [Regarding Velocity Distribution of Red Blood Cells] 
       FIG. 8  is a diagram showing a velocity distribution of red blood cells. In  FIG. 8( a ) , assuming that a direction perpendicular to the skin of a living body is a z direction and directions parallel to the skin are an x direction and a y direction, red blood cells H move in various directions at various velocities in a manner that depends on the direction of the blood vessel. In the figure, velocity vectors of the red blood cells H are shown as the arrows. 
       FIG. 8( b )  schematically shows a velocity distribution of red blood cells H at a particular moment. As shown in the figure, the velocity distribution of the red blood cells H spreads in various directions around the origin. A closed region K shown as the ellipse in  FIG. 8( c )  indicates the range of the velocity distribution of the red blood cells H described above. In the following description, the velocity distribution of the red blood cells H is indicated by such a closed region K. It should be noted that the velocity distribution of the red blood cells H increases/decreases in accordance with a change in pressure inside the blood vessel, which is caused by beating of the heart. Therefore, the elliptical shape K shown in  FIG. 8( c )  expands/contracts in accordance with beating of the heart. 
     [Regarding Configuration of General Laser Doppler Blood Flow Measurement Apparatus] 
     A configuration of a laser Doppler blood flow measurement apparatus having a general configuration will be described as a comparative of the laser Doppler blood flow measurement apparatus according to the present embodiment.  FIG. 9  is a schematic diagram of a laser Doppler blood flow measurement apparatus  500  having a general configuration. 
     As shown in  FIG. 5 , the laser Doppler blood flow measurement apparatus  500  (hereinafter, referred to as blood flow measurement apparatus  500 ) includes a sensor head  510  and an information processing apparatus  520 . The sensor head  510  and the information processing apparatus  520  are connected through a signal line  530 . 
     The sensor head  510  includes a light emitting unit  511  that emits laser light and a light receiving unit  512  that receives scattered light. The light emitting unit  511  is a laser light source, for example. The light receiving unit  512  is a photodiode, for example. The sensor head  510  is disposed close to or in close contact with a living body  600  that is a measurement target. The living body  600  has red blood cells  602  flowing through a blood vessel  601 , and a stationary tissue  603 . The stationary tissue  603  is a stationary living tissue other than blood. 
     In the blood flow measurement apparatus  500 , a laser beam (in the figure, L) is emitted to the living body  600  from the light emitting unit  511 . The laser light L is scattered by the red blood cells  602  and the stationary tissue  603 . The scattered laser light L is received by the light receiving unit  512  and is converted into an electrical signal. 
       FIG. 10  is an example of the electrical signal output from the light receiving unit  512 . The scattered light incident on the light receiving unit  512  is light scattered by the red blood cells  602  moving through the blood flow and light scattered by the stationary tissue  603  that does not move. Therefore, the signal shown in  FIG. 10  is an aggregate of many Doppler beats caused by the interference between both the scattered light beams. In this specification, a signal including such an aggregate of Doppler beats is referred to as a beat signal. 
     The information processing apparatus  520  obtains a beat signal from the sensor head  510  via the signal line  530  and performs signal processing such as Fourier transform.  FIG. 11  shows an example of the signal processing result.  FIG. 11( a )  is a graph obtained by converting a beat signal measured at a fingertip into a frequency domain.  FIG. 11( b )  is a graph obtained by converting a beat signal measured at a wrist into a frequency domain. 
     The power spectra shown in these graphs correspond to intensity for each beat frequency, i.e., distribution density of red blood cells moving at a velocity corresponding to the frequency. 
     For example, the flow rate of the blood flow at the fingertip is higher than the flow rate of the blood flow at the wrist. In  FIG. 11( a ) , the attenuation rate of the intensity with respect to the increase in beat frequency is lower than that in  FIG. 11( b ) . As a result, it can be seen that the ratio of red blood cells at a higher motion velocity in  FIG. 11( a )  is higher than in  FIG. 11( b ) . Further, the total area of the spectrum in  FIG. 11( a )  is wider than in  FIG. 11( b ) . Thus, it can be seen that the number of red blood cells moving within an observation area is larger at the fingertip. 
     The information processing apparatus  520  is further capable of calculating higher-order blood flow-related information such as a time change (pulse) of the mean velocity of the blood flow, a pulse rate, and a flow velocity of the blood flow on the basis of a time-series change of the power spectrum of this beat signal. 
     [Regarding Disturbance Noise Caused by Body Motion] 
     The disturbance noise caused by the body motion, which can be generated in the blood flow measurement apparatus  500 , will be described. 
       FIG. 12  shows a velocity distribution observed in the living body  600  when disturbance noise due to body motion is not generated in the blood flow measurement apparatus  500 . 
     As described above, the velocities of the red blood cells  602  spread around the origin inside the closed region K (see  FIG. 4( c ) ) in accordance with the direction and the flow velocity of the blood flow. 
     At the origin, the motion velocity is 0 with respect to each of the x, y, and z directions. This is a velocity distribution of the stationary tissue  603 . Here, the number of cells in the stationary tissue  603  is much greater than the number of red blood cells  602 . Therefore, light scattered by the stationary tissue is greater than light scattered by the red blood cells. The velocity distribution of zero velocity caused by the stationary tissue  603  is significantly greater than the velocity distribution of the red blood cells  602 . Therefore, a singular point T of the velocity distribution is formed at the origin. 
     When disturbance noise due to a body motion is not generated, the velocity distribution as shown in  FIG. 12  is obtained. When disturbance noise due to a body motion is generated, the velocity distribution of the living body  600  apparently changes from the state shown in  FIG. 12 . 
     The disturbance noise due to the body motion can be classified into the following three categories. 
     1. Change in Blood Flow due to Body Motion 
     2. Shift of Sensor Head due to Body Motion 
     3. Relative Velocity between Sensor Head and Living Body due to Body Motion 
     &lt;1. Change in Blood Flow Due to Body Motion&gt; 
     For example, when jogging is done in a state in which the sensor head  510  is mounted on the wrist, the blood flow is increased in velocity by centrifugal force of arm swing. Therefore, even if the measurement by the blood flow measurement apparatus  500  is correct, a measurement result that differs from the original blood flow velocity can be obtained. 
     This noise is overcome by the use of a low-pass filter that performs smoothing in a period longer than the frequency of the body motion or by a technique of obtaining the acceleration of the body motion through an acceleration sensor and eliminating components correlated to the acceleration, for example. 
     &lt;2. Shift of Sensor Head Due to Body Motion&gt; 
     When the mounting position of the sensor head  510  moves relative to the living body  600 , noise is generated due to a difference in blood flow velocity between the mounting positions, a difference in luminance distribution which is caused by spatial speckles due to interference of laser light, a difference in epidermis property, a difference in laser irradiation angle with respect to the blood vessel, and the like. 
     This noise can be overcome by increasing the size of the light receiving unit  512  or devising a method of fixing the sensor head  510  to the living body, for example. 
     &lt;3. Relative Velocity Between Sensor Head and Living Body Due to Body Motion&gt; 
     Regarding the relative velocity between the sensor head  510  and the living body due to the body motion, more specifically the following four points can be mentioned. 
     3a. Doppler Beat Due to Relative Velocity Between Light Emitting Unit and Living Body 
     3b. Doppler Beat Due to Relative Velocity between Light Receiving Unit and Living Body 
     3c. Doppler Beat Due to Relative Angular Velocity between Light Emitting Unit and Living Body 
     3d. Doppler Beat Due to Relative Angular Velocity between Light Receiving Unit and Living Body 
     {3a. Doppler Beat Due to Relative Velocity Between Light Emitting Unit and Living Body} 
       FIG. 13  is a schematic diagram showing a Doppler beat due to the relative velocity between the light emitting unit  511  and the living body  600 . As shown in  FIG. 13( a ) , it is assumed that the light emitting unit  511  moves at the angle θ and the velocity v with respect to the optical axis A of the laser light L. It should be noted that the irradiation angle of the laser light L is defined as an angle ϕ. 
       FIG. 13( b )  is a schematic diagram showing a velocity distribution of red blood cells in this case. When the light emitting unit  511  moves in a manner as described above, the relative velocity between the red blood cells and the light emitting unit  511  apparently moves to a closed region K′ at a position deviated from the closed region K by a velocity vector v. Further, the singular point T moves to a point T′ in a circle having a radius v. The angle formed by the optical axis of the laser light L and the velocity vector is distributed in a range of ±½ϕ having β as the center. Therefore, as shown in  FIG. 13( b ) , the closed region K′ and the singular point T′ of the red blood cell  602  and the stationary tissue  603  on each optical axis also move in the range of ±½ϕ in the circle having the radius v. For example, the singular point T of the stationary tissue  603  on the optical axis moves to T′ θ . The singular point T of the stationary tissue  603 , which is positioned on a line at an angle +½ϕ from the optical axis, moves to T′ θ+(½)ϕ . The singular point T of the stationary tissue  603 , which is positioned on a line at an angle −½ϕ from the optical axis, moves to T′ θ−(½)ϕ . 
     As described above, regarding the velocity of the stationary tissue  603  within the range of the angle ϕ, the singular point T moves to the singular point T′ having a spread at the angle θ. The stationary tissue  603  is irradiated with the laser light L from the light emitting unit  511 . The stationary tissue  603  at each location causes a Doppler shift corresponding to a different velocity. When they interfere upon light reception, they cause a Doppler beat, even though they are scattered light of the stationary tissues  603 . In  FIG. 13( b ) , a maximum velocity difference that causes a Doppler beat at the singular point T is denoted by E. Base on (Expression 5) above, a frequency corresponding to the velocity difference E is at most 2α 0 v sin θ((½)ϕ). 
     Therefore, the scattered light of the stationary tissue  603  which should not originally generate a beat signal generates a large beat signal caused by a motion of the light emitting unit  511 , and the beat signal generated by a motion of a red blood cell cannot be detected. 
     {3b. Doppler Beat Due to Relative Velocity Between Light Receiving Unit and Living Body} 
       FIG. 14  is a schematic diagram showing a Doppler beat due to the relative velocity between the light receiving unit  512  and the living body  600 . As shown in  FIG. 14( a ) , it is assumed that the light receiving unit  512  moves at the angle θ and the velocity v with respect to the optical axis B of the scattered light S. It should be noted that the incident angle of the scattered light S received by the light receiving unit  512  is an angle θ. 
       FIG. 14( b )  is a schematic diagram showing a velocity distribution of red blood cells in this case. Scattering by the living body  600  is highly uniform scattering close to Lambertian scattering. Therefore, scattering by the living body  600  is performed in a wide range of irradiation angles. Therefore, scattered light S emitted from the wide range of the living body  600  is incident on a single point of the light receiving unit  512 . That is, the angle θ formed by the scattered light S incident on the single point of the light receiving unit  512  is a large angle close to 180°. 
     Therefore, when the sensor head  510  moves in the manner as described above, cos θ in (Expression 2) above largely differs at each position of the stationary tissue  603 , and the apparent closed area moves to a wide area as shown as K′ in  FIG. 14( b ) . 
     Accordingly, the singular points T′ at the respective moving destinations, that is, scattered light beams of the stationary tissue  603  cause Doppler beats each having a higher frequency. In  FIG. 14( b ) , a maximum width of the velocity difference between the singular points T° is denoted by G. Based on (Expression 5) above, a frequency of a beat G ranges from 0 to α 0 v{1−sin(θ+(½)ϕ)}. 
     Therefore, the scattered light of stationary tissue  603 , which should not originally generate a beat signal, generates a large beat signal due to a motion of the light receiving unit  512 , and a beat signal generated by a motion of a red blood cell cannot be detected. 
     {3c. Doppler Beat Due to Relative Angular Velocity between Light Emitting Unit and Living Body/3d. Doppler Beat Due to Relative Angular Velocity between Light Receiving Unit and Living Body} 
       FIG. 15  is a schematic diagram showing a Doppler beat due to the angular velocity between the light emitting unit  511  and the light receiving unit  512  and the living body  600 . 
     As shown in  FIG. 15 , when an angular velocity indicated by an angle ω is generated in the sensor head  510 , a relative velocity is generated at one end of the irradiation spot of the laser light L such that the light emitting unit  511  and the light receiving unit  512  approach the living body  600 . A relative velocity is generated at the other end of the irradiation spot such that the light emitting unit  511  and the light receiving unit  512  moves away from the living body  600 . 
     Therefore, as in “3a. Doppler Beat Due to Relative Velocity Between Light Emitting Unit and Living Body” and “3b. Doppler Beat Due to Relative Velocity between Light Receiving Unit and Living Body” above, the scattered light of the stationary tissue  603  which should not originally generate a beat signal generates a large beat signal, and a beat signal generated by a motion of a red blood cell cannot be detected. 
     The blood flow measurement apparatus according to the present technology can prevent disturbance noise due to “3. Relative Velocity between Sensor Head and Living Body due to Body Motion”. 
     First Embodiment 
     [Configuration of Blood Flow Measurement Apparatus According to First Embodiment] 
     A configuration of a blood flow measurement apparatus  100  according to a first embodiment of the present technology will be described.  FIG. 16  is a schematic diagram of the blood flow measurement apparatus  100  according to the present embodiment.  FIG. 17  is a block diagram showing functions of the blood flow measurement apparatus  100 . 
     As shown in  FIG. 16 , the blood flow measurement apparatus  100  includes a sensor head  110  and an information processing apparatus  120 . The sensor head  110  and the information processing apparatus  120  are connected through a signal line  130 . The blood flow measurement apparatus  100  may be a wearable device worn by a user. The wearable device may be, for example, a head mounted display (HMD), smart ice glasses, a smart watch, a smart band, smart earphones, or the like. 
     Further,  FIG. 16  shows a living body  600  which is a blood flow measurement target. The living body  600  has red blood cells  602  flowing through a blood vessel  601 , and a stationary tissue  603 . 
     As shown in  FIG. 16 , the sensor head  110  is disposed close to or in close contact with the living body  600 . As shown in  FIGS. 16 and 17 , the sensor head  110  includes a light emitting unit  111 , a light receiving unit  112 , an analog signal processing unit  113 , and an ADO  114 . 
     The light emitting unit  111  irradiates the living body  600  with laser light L. The light receiving unit (sensor)  112  receives scattered light obtained when the laser light L is reflected by the living body  600 . The light receiving unit (sensor)  112  converts the received scattered light into a beat signal. The light emitting unit  111  and the light receiving unit  112  are mounted on the sensor head  110  and the relative position relative to the sensor head  110  is fixed. Details of the light emitting unit  111  and the light receiving unit  112  will be described later. It should be noted that laser light is included in the coherent light and examples of laser light are described as an example of coherent light in the respective embodiments. 
     The analog signal processing unit  113  performs signal processing such as amplification on the beat signal output from the light receiving unit  112  and supplies it to the ADC  114 . 
     The analog to digital converter (ADC)  114  converts the analog signal supplied from the analog signal processing unit  113  into a digital signal. The HMD  114  outputs the converted digital signal to the information processing apparatus  120  via the signal line  130 . 
     As shown in  FIG. 17 , the information processing apparatus  120  includes a power spectrum calculation unit  121 , a velocity pulse wave calculation unit  122 , a pulse rate calculation unit  123 , and a velocity distribution information calculation unit  124 . 
     The power spectrum calculation unit  121  performs arithmetic processing such as Fourier transform for a predetermined time interval of the beat signal obtained from the ADC  114 . Accordingly, the power spectrum calculation unit  121  calculates the power spectrum by converting the beat signal for each predetermined time into a frequency domain. The predetermined time interval is, for example, 1/300 to 1/10 seconds. The predetermined time is, for example, 1/300 to 1/10 seconds. 
     The velocity pulse wave calculation unit  122  calculates a velocity pulse wave on the basis of a time-series change of the power spectrum of the beat signal. The pulse rate calculation unit  123  calculates a pulse rate on the basis of the velocity pulse wave calculated by the velocity pulse wave calculation unit  122 . 
     The velocity distribution information calculation unit  124  performs an operation such as integration processing on the power spectrum of the beat signal to thereby calculate velocity distribution information indicating a velocity distribution of the red blood cells  602 . 
     The power spectrum calculation unit  121 , the velocity pulse wave calculation unit  122 , the pulse rate calculation unit  123 , and the velocity distribution information calculation unit  124  function as a blood flow-related information calculation unit that calculates information regarding the flow of red blood cells (hereinafter, referred to as blood flow-related information). The blood flow-related information calculation unit may have another configuration capable of calculating the blood flow-related information. 
     The blood flow measurement apparatus  100  has the above-mentioned configuration. It should be noted that sensor head  110  functions as an analog signal processing block of the blood flow measurement apparatus  100  and the information processing apparatus  120  functions as a digital signal processing block. 
     The configuration of the blood flow measurement apparatus  100  is not limited thereto. Digital signal processing may be performed at the sensor head  110  and analog signal processing may be performed at the information processing apparatus  120 . Further, not only a wired signal communication path but also a wireless signal communication path may be used for the signal line  130 . Further, the sensor head  110  and the information processing apparatus  120  may be integrated. 
     [Operation of Blood Flow Measurement Apparatus According to First Embodiment] 
     An operation of the blood flow measurement apparatus  100  will be described. 
     As shown in  FIG. 16 , when the laser light L at the frequency f 0  is emitted from the light emitting unit  111  to the living body  600 , the laser light L is scattered by the living body  600  and the scattered light is received by the light receiving unit  112 . The scattered light includes scattered light at the frequency f 0  which is scattered by the stationary tissue  603  and scattered light at the frequency f 0 +Δf which is scattered by the red blood cells  602  moving through the blood flow and causes a Doppler shift. 
     The Doppler shift Δf varies in a manner that depends on the rate of motion of the red blood cell  602  relative to the stationary tissue  603  (see  FIG. 5 ). Therefore, the light receiving unit  112  outputs a beat signal that is an aggregate of many Doppler beats (see  FIG. 10 ). 
     This beat signal is supplied to the power spectrum calculation unit  121  via the analog signal processing unit  113  and the ADC  114 . The power spectrum calculation unit  121  converts the beat signal in the predetermined time interval for each predetermined time into a frequency domain and calculates a power spectrum of the beat signal (see  FIG. 11 ). 
     Each of the velocity pulse wave calculation unit  122 , the pulse rate calculation unit  123 , and the velocity distribution information calculation unit  124  calculates blood flow-related information based on the power spectrum of the beat signal. 
     The blood flow measurement apparatus  100  performs the above-mentioned operation. Here, when disturbance noise due to body motion as described above is generated in the beat signal output by the light receiving unit  112 , the blood flow cannot be accurately measured. However, in the blood flow measurement apparatus  100 , disturbance noise due to body motion is prevented from being generated as follows. 
     [Configurations of Light Emitting Unit and Light Receiving Unit] 
       FIG. 18  is a schematic diagram showing the light emitting unit  111  and the light receiving unit  112 . As shown in the figure, the light emitting unit  111  includes a laser light source  151  and an irradiation light control unit  152 . 
     The laser light source  151  is a light source of the laser light L. The laser light source  151  can be a common semiconductor laser device. 
     The irradiation light control unit  152  collimates the laser light L emitted from the laser light source  151  and controls the irradiation light diameter of the laser light L. The irradiation light diameter of the laser light L is equal to the diameter of a spot (R in the figure) of the laser light L on the surface of the living body  600 .  FIG. 18  shows an irradiation light diameter d controlled by the irradiation light control unit  152 . 
     The irradiation light control unit  152  is constituted by an optical member such as glass. As shown in  FIG. 18 , the irradiation light control unit  152  includes a concave lens  152   a  and a convex lens  152   b.    
     The concave lens  152   a  expands the laser light L incident from the laser light source  151  to the same diameter as the irradiation beam diameter d. The convex lens  152   b  collimates the laser light L enlarged by the concave lens  152   a.    
     Thus, the laser light L having the irradiation light diameter d is emitted by the irradiation light control unit  152  to the living body  600 . The irradiation light diameter d is favorably 0.5 mm or more and 2 mm or less. 
     It should be noted that the configuration of the light emitting unit  111  is not limited to that shown here. It is sufficient that the irradiation light control unit  152  can roughly collimate the laser light L and control the irradiation light diameter to a predetermined diameter. Various configurations that the light emitting unit  111  can take will be described later. 
     Further, as shown in  FIG. 18 , the light receiving unit  112  includes a measurement light receiving element  161  and an incident angle limiting unit  162 . 
     The measurement light receiving element  161  photoelectrically converts light incident through the incident angle limiting unit  162  and outputs a beat signal. The measurement light receiving element  161  can be a photodiode having a general configuration. In the present technology, a large-size photodiode whose light receiving surface has a larger area as compared to the configuration of the conventional technology is favorably used. 
     The incident angle limiting unit  162  makes light from the living body  600 , i.e., the scattered light S of the laser light L incident on the measurement light receiving element  161 . Here, the incident angle limiting unit  162  is configured to limit the incident angle of the scattered light S incident on a single point of the measurement light receiving element  161  to be equal to or smaller than a predetermined angle and control the measurement light receiving element  161  to increase the aperture ratio as compared to a simple incident angle limiting mechanism with a single pin hole. 
     Specifically, the incident angle limiting unit  162  concentrates the incident scattered light S to a different point of the measurement light receiving element  161  for each incident angle. 
       FIG. 19  is an enlarged view showing the incident angle limiting unit  162 . As shown in the figure, the incident angle limiting unit  162  is constituted by an optical member and can be a lens array in which a plurality of lenses  162   a  is arranged. The lens  162   a  is not limited to a convex lens. The lens  162   a  may be a hologram lens or a Fresnel lens. The incident angle limiting unit  162  is arranged such that the light receiving surface of the light receiving element  161  is located at a position apart from the main point of each lens by approximately a focal length. In such an arrangement, the parallel light emitted from below in  FIG. 19  is concentrated in the vicinity of the surface of the light receiving element  161 . 
     In  FIG. 19 , scattered light that is incident at a certain incident angle with respect to the light receiving unit  112  will be referred to as scattered light S 1 . That is, components of the parallel light of the total scattered light, which has the certain incident angle, is the scattered light S 1 . Further, scattered light that is incident at an incident angle different from that of the scattered light S 1  with respect to the light receiving unit  112  will be referred to as scattered light S 2 . In addition, scattered light that is incident at an incident angle different from those of the scattered light S 1  and the scattered light S 2  will be referred to as scattered light S 3 . The scattered light S 2  and the scattered light S 3  are also components of the parallel light, respectively. 
     As shown in the figure, the scattered light S 1  is concentrated by each lens  162   a  to a point P 1  of the measurement light receiving element  161 . Further, the scattered light S 2  is concentrated by each lens  162   a  to a point P 2  on the measurement light receiving element  161 . In addition, the scattered light S 3  is concentrated by each lens  162   a  to a point P 3  on the measurement light receiving element  161 . 
     The point P 1 , the point P 2 , and the point P 3  are different points from each other. Further, light incident on the incident angle limiting unit  162  at an incident angle different from those of the scattered light S 1  to S 3  is also concentrated by each lens  162   a  to a different point on the measurement light receiving element  161  in a manner that depends on the incident angle. 
     As an angle range of the scattered light incident on a single point of the measurement light receiving element  161  is narrower, β described above takes a smaller value, which is more favorable. More specifically, in a case where the angle between the scattered light S incident on a single point is limited to be equal to or smaller than 10°, β is smaller than 0.1, which can provide a sufficient effect. 
     In addition, by using the lens array for the incident angle limiting unit  162 , the light beam diameter of the scattered light S concentrated to the single point of the measurement light receiving element  161  is limited to be smaller than the entire light beam diameter of the light incident on the measurement light receiving element  161 . As shown in  FIGS. 18 and 19 , a light beam diameter D of the scattered light S concentrated to the single point of the measurement light receiving element  161  is equal to the diameter of the single lens  162   a  and is equal to a diameter of an observation region Ti facing this lens  162   a  on the living body  600 . 
     Although the number of lenses  162   a  of the incident angle limiting unit  162  is arbitrary, it is favorable that the lenses  162   a  are disposed over the entire light receiving surface of the measurement light receiving element  161 . Further, the diameters of the lenses  162   a  may be the same or may be different from each other. 
     It should be noted that the configuration of light receiving unit  112  is not limited to that shown here. It is sufficient that the incident angle limiting unit  162  can limit the incident angle of the scattered light incident on the single point of the measurement light receiving element  161  to be equal to or smaller than the predetermined angle. Various configurations that can be taken by the light receiving unit  112  will be described later. 
     [Effects of Light Emitting Unit and Light Receiving Unit] 
     Effects of the light emitting unit  111  and the light receiving unit  112  having the above-mentioned configurations will be described. 
     In the blood flow measurement apparatus  500  having the above-mentioned general configuration, “3a. Doppler Beat Due to Relative Velocity Between Light Emitting Unit and Living Body” caused by the relative motion between the sensor head  510  and the living body due to the body motion becomes a problem (see  FIG. 13 ). 
     In the light emitting unit  111  in the present embodiment, this “3a. Doppler Beat Due to Relative Velocity Between Light Emitting Unit and Living Body” can be overcome.  FIG. 20  is a schematic diagram showing this effect. 
     As described above, at the light emitting unit  111 , the laser light L emitted from the laser light source  151  is collimated by the irradiation light control unit  152 . Therefore, as shown in  FIG. 20( a ) , the incident angle of the laser light L at each position of the stationary tissue  603  is the same even when the light emitting unit  111  is moving at the angle θ and the velocity v. 
     Accordingly, in (Expression 1) above, cos θ takes the same value at each position of the stationary tissue  603  and the velocity distribution moves to the same position in accordance with the velocity v as shown in  FIG. 20( b ) . Therefore, there is no Doppler beat between the stationary tissues  603  indicated by the singular points T′ (having a velocity difference denoted by E in  FIG. 13  and the beat signal caused by the motion of the red blood cell to be detected is maintained. The generation of the beat signal between the stationary tissues  603 , which is noise, is suppressed, and the beat signal generated by the motion of the red blood cell to be detected which is the measurement target, is maintained. Therefore, sensing of the signal is maintained. 
     In addition, in the blood flow measurement apparatus  500  having the above-mentioned general configuration, “3b. Doppler Beat Due to Relative Velocity between Light Receiving Unit and Living Body” caused by the relative motion between the sensor head  510  and the living body due to the body motion becomes a problem (see  FIG. 14 ). 
     In the light receiving unit  112  in the present embodiment, “3b. Doppler Beat Due to Relative Velocity between Light Receiving Unit and Living Body” can also be overcome.  FIG. 21  is a schematic diagram showing this effect. 
     As shown in  FIG. 21( a ) , when the light receiving unit  112  is moving at the angle θ and the velocity v, the irradiation angle θ of the scattered light S is large. Therefore, the apparent velocity distribution of the stationary tissue  603  moves to each position (see  FIG. 14( b ) ). 
     Here, as described above, in the light receiving unit  112 , the incident angle limiting unit  162  limits the incident angle of the scattered light S incident on the single point of the measurement light receiving element  161  to be equal to or smaller than the predetermined angle. Therefore, scattered light S at a different incident angle with respect to the incident angle limiting unit  162  is concentrated at a different point on the measurement light receiving element  161  (see  FIG. 19 ). 
     Thus, scattered light beams S at different incident angles with respect to the incident angle limiting unit  162  do not interfere with each other. As shown in  FIG. 21( b ) , even if the apparent velocity distribution of the stationary tissue  603  is different, the Doppler beat (G in  FIG. 14 ) is not generated between the singular points T′. At each point, the Doppler beat is generated in a combination of each closed region K′ and the singular point T′. Therefore, the beat signal caused by the motion of the red blood cell to be detected is maintained. The beat signal between T′ at different incident angles, which is a disturbance signal, is suppressed and the beat signal caused by the motion of red blood cell that is the measurement target is maintained. Therefore, sensing of the signal is maintained. 
     Further, in the blood flow measurement apparatus  500  having the above-mentioned general configuration, “3c. Doppler Beat Due to Relative Angular Velocity between Light Emitting Unit and Living Body/3d. Doppler Beat Due to Relative Angular Velocity between Light Receiving Unit and Living Body” caused by a relative motion between the sensor head  510  and the living body due to the body motion becomes a problem. 
     In the light emitting unit  111  and the light receiving unit  112  in the present embodiment, “3c. Doppler Beat Due to Relative Angular Velocity between Light Emitting Unit and Living Body/3d. Doppler Beat Due to Relative Angular Velocity between Light Receiving Unit and Living Body” can be also overcome.  FIG. 22  is a schematic diagram showing this effect. 
     As shown in the figure, the irradiation light control unit  152  collimates the laser light L, such that laser beams at a plurality of relative velocities are not emitted to a particular point of the stationary tissue  603 . Further, the irradiation spot R of the laser light L has the irradiation light diameter d. Therefore, a maximum velocity difference in the irradiation spot R is dω. 
     Further, the incident angle limiting unit  162  limits the region in which light is concentrated to the single point of the measurement light receiving element  161  to an observation region N. The observation region N and the lens  162   a  have the diameter D. Therefore, maximum velocity difference of the scattered light concentrated to the measurement light receiving element  161  through the single lens  162   a  is Do). 
     Thus, the relative velocity distribution obtained by combining the factors of the light emitting unit  111  and the light receiving unit  112  falls within the range of (d+D)ω. Therefore, if the diameter D of the lens  162   a  is reduced, the Doppler beat due to the relative angular velocity of the sensor head  110  is determined in a manner that depends on the irradiation light diameter d. 
     By reducing the irradiation light diameter d through the irradiation light control unit  152  in this manner, “3c. Doppler Beat Due to Relative Angular Velocity between Light Emitting Unit and Living Body/3d. Doppler Beat Due to Relative Angular Velocity between Light Receiving Unit and Living Body” can be overcome. 
     Specifically, the irradiation light diameter d is favorably 2 mm or less. On the other hand, if the irradiation light diameter d is too small, an influence due to a change in position of the sensor head  110  becomes large. Therefore, the spot diameter is favorably 0.5 mm or more. From the above, it is favorable that the irradiation light diameter d is 0.5 mm or more and 2 mm or less. It should be noted that the irradiation shape is not limited to the circular shape. The irradiation shape may be a shape having a long-side direction and a short-side direction, for example, an ellipse or a rectangle. The relative angular velocity of the sensor head  110  may have axial anisotropy in a manner that depends on the shape and mounting position of the sensor device. For example, in a case of mounting the sensor head to a watch-type device, rotation along an axis perpendicular to the dial of the watch is unlikely to occur, and further, rotation along an axis in the 3 o&#39;clock-9 o&#39;clock direction on the dial of the analog watch is also unlikely to occur. On the other hand, regarding rotation along the 12 o&#39;clock-6 o&#39;clock direction on the dial of the analog watch, a large angular velocity motion is generated by arm swing in walking or the like. In such a case, by arranging the long-side direction of the irradiation shape in parallel with the rotation axis having a high angular velocity, it is possible to enhance the resistance to the relative angular velocity motion while maintaining the irradiation area and maintaining the resistance to the change in position. 
     In addition, by using the incident angle limiting unit  162 , the light receiving unit  112  can increase the area of the measurement light receiving element  161  while preventing disturbance noise caused by “Relative Velocity between Sensor Head and Living Body due to Body Motion” as described above. 
     With this enlargement of the measurement light receiving element  161 , it is also possible to reduce the disturbance noise caused by “2. Shift of Sensor Head due to Body Motion” above. 
       FIGS. 23 and 24  are schematic diagrams each showing this effect.  FIG. 23  shows a case where the area of the measurement light receiving element  161  is small.  FIG. 24  shows a case where the area of the measurement light receiving element  161  is large. 
     As shown in  FIG. 23( a ) , when the measurement light receiving element  161  having the small area vibrates due to the shift of the sensor head  110 , the observation region facing the measurement light receiving element  161  in the living body  600  moves between the region Q 1  and the region Q 2 . 
     In  FIG. 23( b ) , V 1  is an actual blood flow velocity at the region Q 1  and V 2  is an actual blood flow velocity at the region Q 2 . The region Q 1  has a higher blood flow velocity than the region Q 2 , and is, for example, a region including many blood vessels extending toward the epidermis. 
       FIG. 23( c )  shows the blood flow velocity V 3  measured on the basis of the Doppler beat in this case. As shown in the figure, the measured blood flow velocity V 3  is caused by the motion of the measurement light receiving element  161  between the region Q 1  and the region Q 2 . This is largely different from the blood flow velocity V 1  and the blood flow velocity V 2  which are the original measurement targets. 
     On the other hand, as shown in  FIG. 24( a ) , it is assumed that the measurement light receiving element  161  having a large area vibrates due to the shift of the sensor head  110  and the observation region facing the measurement light receiving element  161  in the living body  600  moves between a region W 1  and a region W 2 . 
     In  FIG. 24( b ) , V 1  is the actual blood flow velocity at the region W 1  and V 2  is the actual blood flow velocity at the region W 2 . As shown in  FIG. 24( a ) , even if the measurement light receiving element  161  vibrates, the measurement light receiving element  161  has the large area, such that the region W 1  and the region W 2  overlap. Therefore, in  FIG. 24( b ) , the blood flow velocity V 1  and the blood flow velocity V 2  have a small difference. 
     Accordingly, as shown in  FIG. 24( c ) , the blood flow velocity V 3  measured on the basis of the Doppler beat has a shape similar to those of the blood flow velocity V 1  and the blood flow velocity V 2 , and the pulse wave is measured in the same manner as the blood flow velocity V 1  and the blood flow velocity V 2  which are the original measurement targets irrespective of vibration of the measurement light receiving element  161 . 
     As described above, in the blood flow measurement apparatus  100 , the laser beam is collimated by providing the irradiation light control unit  152  in the light emitting unit  111 . Accordingly, the angles of the laser light beams entering the respective positions of the living body  600  are set to be the same, and a Doppler beat can be prevented from being generated between the stationary tissues  603 . 
     Further, by controlling the light diameter of the laser light L through the irradiation light control unit  152 , it is possible to achieve both the resistance to a shift of the light emitting unit  111  and the resistance to the relative angular velocity of the light emitting unit  111 . 
     In addition, the angle of the scattered light incident on the single point of the measurement light receiving element  161  is limited by providing the incident angle limiting unit  162  in the light receiving unit  112 . Accordingly, it is possible to prevent interference between the scattered light beams at different incident angle and to prevent a Doppler beat from being generated between the stationary tissues  603 . 
     Further, by providing the incident angle limiting unit  162 , it is possible to increase the size of the measurement light receiving element  161 , to reduce the power consumption by increasing the utilization efficiency of light, and it is possible to reduce the influence due to the change in position of the light receiving unit  112 . 
     [Variations of Light Emitting Unit] 
     The light emitting unit  111  is not limited to the above-mentioned configuration. It is sufficient that the configuration of the light emitting unit  111  is capable of substantially collimating the laser light L and controlling the spot diameter to the predetermined diameter.  FIGS. 25 and 26  are schematic diagrams each showing the light emitting unit  111  having another configuration. 
     As shown in  FIG. 25 , the light emitting unit  111  may include a laser light source  151  and an irradiation light control unit  153 . The irradiation light control unit  153  includes a convex lens  153   a  and a convex lens  153   b . The convex lens  153   a  magnifies the laser light L incident from the laser light source  151 . The convex lens  153   b  collimates the laser light L magnified by the convex lens  153   a.    
     Further, as shown in  FIG. 26 , the light emitting unit  111  may include a laser light source  151  and an irradiation light control unit  154 . The irradiation light control unit  154  can be a hologram lens capable of magnifying and collimating the laser light L incident from the laser light source  151 . 
     Further, the light emitting unit  111  is not limited to those including the laser light source and the irradiation light control unit. The light emitting unit  111  may include only the laser light source. For example, vertical cavity surface emitting LASER (VCSEL) devices are laser devices capable of emitting relatively parallel laser light beams. The irradiation angle of the laser light source itself is favorable as the laser light source of the light emitting unit  111  not provided with the irradiation light control unit. 
     [Variations of Light Receiving Unit] 
     The light receiving unit  112  is not limited to the above-mentioned configuration. The light receiving unit  112  only needs to have a configuration to limit the incident angle of the scattered light S incident on the single point of the measurement light receiving element  161  to be equal to or smaller than the predetermined angle and to control the measurement light receiving element  161  to increase the aperture ratio.  FIGS. 27  to  35  are schematic diagrams each showing the light receiving unit  112  having another configuration. 
     As shown in  FIG. 27 , the incident angle limiting unit  162  may further include light shields  163 . The light shields  163  are disposed between the lens array  162   a  and the measurement target and shield scattered light that enters the portions between the lenses  162   a . It should be noted that the lens  162   a  is not limited to the convex lens. The lens  162   a  may be a hologram lens or a Fresnel lens. The same applies to each of the following configurations. 
     Accordingly, scattered light can be prevented from entering the measurement light receiving element  161  from portions between the lenses  162   a  and a Doppler beat uncorrelated with the signal to be measured on the basis of the scattered light can be prevented from being generated. Further, by adjusting the angle of view overlapping the adjacent lenses  162   a , it is possible to suppress the variation of the amount of light projected on the light receiving surface when the living body  600  moves and to enhance the resistance to the position dependency of the blood flow rate. In addition, in a case of using a paved-type lens array with sufficiently small portions between the lenses  162   a , the light shields  163  and the lens array  162   a  do not need to be strictly aligned. In addition, if the height and the light shield spacing satisfy a predetermined condition, it is unnecessary to establish a one-to-one relationship with each lens. The predetermined condition is to limit the range of the incident angle such that the light that has passed through the lens  162   a  is not incident on the same point on the photosensitive surface as the light that has passed through another lens  162   a,    
     Further, as shown in  FIG. 28 , the incident angle limiting unit  162  may further include light shields  164 . The light shields  164  are disposed between the lens array and the measurement light receiving element  161  and can prevent scattered light from entering the measurement light receiving element  161  from portions between the adjacent lens  162   a  and prevent the Doppler beat from being generated by the scattered light. Further, it is possible to enhance the resistance to the position dependency of the blood flow rate as in the light shields  163 . 
     Further, as shown in  FIG. 29 , the incident angle limiting unit  162  may include both light shields  163  and light shields  164 . By providing both the light shields  163  and the light shields  164 , it is possible to further reduce stray light incident from portions between the lenses  162   a , to can more effectively prevent disturbance noise due to the Doppler beat, and to enhance the resistance to the position dependency of the blood flow rate. 
     Further, as shown in  FIG. 30 , the light receiving unit  112  may include a measurement light receiving element  161  and an incident angle limiting unit  165 . The incident angle limiting unit  165  is a lens array in which lenses  165   a  convex toward the measurement light receiving element  161  are arranged. The incident angle limiting unit  165  includes light shields  166 . 
     The light shields  166  are disposed between the lenses  165   a  between the incident angle limiting unit  165  and the measurement light receiving element  161  and shields scattered light that enters the portions between the lenses  165   a.    
     Also with this configuration, it is possible to more effectively prevent disturbance noise due to a Doppler beat and to enhance the resistance to the position dependency of the blood flow rate. Further, it is possible to use the lens array as a cover for the measurement light receiving element  161 , and to enhance the resistance to external light by providing the lens array with a characteristic of absorbing light other than the light source wavelength such as IR black. 
     Further, as shown in  FIG. 31 , the light receiving unit  112  may include a plurality of light receiving elements  167  and the incident angle limiting unit  162 . The incident angle limiting unit  162  is configured such that the respective lenses  162   a  collect light to the respective light receiving elements  167 . The incident angle limiting unit  162  includes light shields  163  and light shields  164 . 
     With this configuration, it is possible to more effectively prevent disturbance noise caused by the Doppler beat, to observe the position dependency of the blood flow rate, and to achieve both resolution and coverage between the light receiving elements  167 . It should be noted that a plurality of lenses  162   a  may be arranged for a single light receiving element  167 . Further, either the light shields  163  or the light shields  164  may be provided or neither of them may be provided. 
     Further, as shown in  FIG. 32 , the light receiving unit  112  may include a measurement light receiving element  161  and an incident angle limiting unit  168 . The incident angle limiting unit  168  is constituted by a single lens  168   a . Also with this configuration, the disturbance noise due to the relative velocity between the sensor head  110  and the living body  600  can be overcome. 
     Further, by reducing the area of the light receiving surface of the measurement light receiving element  161  and reducing the diameter of the light beam collected to a single point of the measurement light receiving element  161  (in  FIG. 22 , D), the disturbance noise due to the relative angular velocity between the sensor head  110  and the living body  600  can be overcome. In addition, this configuration can be realized at low costs. 
     Further, as shown in  FIG. 33 , the incident angle limiting unit  168  may include a light shield  169 . The light shield  169  is provided around the lens  168   a  to limit the diameter of the light beam incident on the lens  168   a,    
     With this configuration, the relative velocity between the sensor head  110  and the living body  600  and the disturbance noise due to the relative angular velocity can also be overcome. In addition, this configuration can also be realized at low costs. 
     Further, as shown in  FIG. 34 , the incident angle limiting unit  168  may include a light shield  170 . The light shield  169  is provided around the lens  168   a  between the lens  168   a  and the measurement light receiving element  161  and limits the diameter of the light beam incident on the measurement light receiving element  161 . 
     With this configuration, the relative velocity between the sensor head  110  and the living body  600  and the disturbance noise due to the relative angular velocity can also be overcome. In addition, this configuration can also be realized at low costs. 
     Further, as shown in  FIG. 35 , the light receiving unit  112  may include a measurement light receiving element  161  and an incident angle limiting unit  171 . The incident angle limiting unit  171  is a pin hole structure including a plurality of pin holes  171   a  made of a light-shielding material and extending in a direction perpendicular to the light receiving surface of the measurement light receiving element  161 . 
     With this pin hole  171   a , the scattered light S 1  is incident on the point P 1  on the measurement light receiving element  161 , the scattered light S 2  is incident on the point P 2  on the measurement light receiving element  161 , and the scattered light S 3  is incident on the point P 3  on the measurement light receiving element  161 . That is, the incident angle limiting unit  171  limits the incident angle of the scattered light S incident on the single point of the measurement light receiving element  161  to be equal to or smaller than the predetermined angle. More specifically, when the predetermined angle is 10 degrees or less, β becomes 0.1 or less and satisfactory effects can be obtained. 
     Therefore, in accordance with a principle similar to that in a case where the incident angle limiting unit is a lens, it is possible to prevent the disturbance noise due the Doppler beat. Further, by guiding the scattered light from the plurality of pin holes  171   a  to the single measurement light receiving element  161 , the utilization efficiency of light can be increased by the number of pin holes  171   a  as compared to a sensor head with a single pin hole despite the single measurement light receiving element. Accordingly, high-sensitivity can be realized at low costs. 
     On the other hand, the scattered light S that has not passed through the pin hole  171   a  becomes lost. Therefore, the utilization efficiency of light is better in a case where the incident angle limiting unit is a lens. However, there is an advantage that the problem of aberration and stray light due to the lens can be avoided. 
     [Hardware Configuration] 
     A hardware configuration of the information processing apparatus  120  will be described.  FIG. 36  is a schematic diagram showing the hardware configuration of the information processing apparatus  120 . As shown in  FIG. 1 , the information processing apparatus  120  includes an interface (I/F)  181 , arithmetic processing circuit  182 , a bus  183 , and a memory  184 . 
     The I/F  181  provides a beat signal (see FIG.  10 ) obtained from the sensor head  110  to the bus  183 . The arithmetic processing circuit  182  is a processor such as a central processing unit (CPU). The arithmetic processing circuit  182  obtains the beat signal via the bus  183  and executes various types of arithmetic processing while exchanging information with the memory  184 . 
     The power spectrum calculation unit  121 , the velocity pulse wave calculation unit  122 , the pulse rate calculation unit  123 , and the velocity distribution information calculation unit  124  having the above-mentioned functional configurations are realized by the cooperation of the arithmetic processing circuit  182  and the program. 
     It should be noted that some or all of the configurations of the information processing apparatus  120  may be implemented on a computer network. Further, some or all of the configurations of the information processing apparatus  120  may be realized by arithmetic circuits that perform particular processing such as a field-programmable gate array (FPGA) and an application specific integrated circuit (ASIC). 
     Second Embodiment 
     [Configuration of Blood Flow Measurement Apparatus According to Second Embodiment] 
     A configuration of a blood flow measurement apparatus  200  according to a second embodiment of the present technology will be described.  FIG. 37  is a schematic diagram of the blood flow measurement apparatus  200  according to the present embodiment.  FIG. 38  is a block diagram showing functions of the blood flow measurement apparatus  200 . 
     As shown in  FIG. 37 , the blood flow measurement apparatus  200  includes a sensor head  210  and an information processing apparatus  220 . The sensor head  210  and the information processing apparatus  220  are connected through a signal line  230 . The blood flow measurement apparatus  200  may be a wearable device. 
     The living body  600  that is the blood flow measurement target has red blood cells  602  flowing through a blood vessel  601 , and a stationary tissue  603 . 
     As shown in  FIG. 37 , the sensor head  210  is disposed close to or in close contact with the living body  600 . As shown in  FIGS. 37 and 38 , the sensor head  210  includes a light emitting unit  211 , a light receiving unit  212 , an analog signal processing unit  213 , and an ADC  214 . 
     The light emitting unit  211  and the light receiving unit (sensor)  212  are mounted on the sensor head  210  and the relative position relative to the sensor head  210  is fixed.  FIG. 39  is a cross-sectional view of the sensor head  210 .  FIG. 40  is a plan view of the sensor head  210  viewed from the living body  600  side. 
     The light emitting unit  211  alternately emits two types of laser light having different wavelength bands in a time division manner and irradiates the living body  600  with the laser light L. 
     As shown in  FIGS. 39 and 40 , the light emitting unit  211  may include a first laser light source  251  and a second laser light source  252 . The first laser light source  251  and the second laser light source  252  have emission wavelengths different from each other. It should be noted that in  FIG. 39 , the first laser light source  251  should originally be located at the same position as the second laser light source  252 , but it is shown, shift to the left because it would interfere with the comprehension of the drawing. 
     It should be noted that of the first laser light source  251  and the second laser light source  252  may be provided with the irradiation light control unit described in the first embodiment. 
     The light receiving unit  212  receives the laser light L obtained when the scattered light is scattered by the living body  600 . The light receiving unit  212  converts the received scattered light into a signal. 
     The light receiving unit  212  includes a measurement light receiving element  261  and an incident angle limiting unit  262 . The measurement light receiving element  261  and the incident angle limiting unit  262  can respectively have the same configuration (see  FIG. 18 ) as the measurement light receiving element  161  and the incident angle limiting unit  162  described in the first embodiment. 
     Further, the incident angle limiting unit  262  may have another configuration (see  FIGS. 27 to 35 ) of the first embodiment. In addition, the light receiving unit  212  may be configured to include only the measurement light receiving element  261  and not to include the incident angle limiting unit  262 . 
     The measurement light receiving element  261  receives scattered light (hereinafter, referred to as first scattered light) emitted from the first laser light source  251  and scattered light (hereinafter, referred to as second scattered light) emitted from the second laser light source  252  for each predetermined time. 
     The analog signal processing unit  213  performs signal processing such as amplification on the signal output from the light receiving unit  212  and supplies it to the ADC  214 . 
     The analog to digital converter (ADC)  214  converts the analog signal supplied from the analog signal processing unit  213  into a digital signal. The ADC  214  outputs the converted digital signal to the information processing apparatus  220  via the signal line  230  as a beat signal. 
     As shown in  FIG. 38 , the information processing apparatus  220  includes a power spectrum calculation unit  221 , a power spectrum normalization unit  222 , an adaptive filter unit  223 , a velocity pulse wave calculation unit  224 , a pulse rate calculation unit  225 , and a velocity distribution information calculation unit  226 . 
     The power spectrum calculation unit  221  performs arithmetic processing such as Fourier transform on the beat signal obtained from the ADC  214 . Accordingly, the power spectrum calculation unit  221  calculates the power spectrum of the signal that may be changed to the frequency domain for each predetermined time. The predetermined time is, for example, 1/300 to 1/100 seconds. 
     The power spectrum calculation unit  221  separates the beat signal into a beat signal  2  generated from the first scattered light and a beat signal  1  generated from the second scattered light and calculates a power spectrum for each of them. The power spectrum based on the beat signal  1  is defined as a first power spectrum. The power spectrum based on the beat signal  2  is defined as a second power spectrum. 
     A power spectrum normalization unit  222  normalizes the first power spectrum and the second power spectrum. Light having different wavelengths has different Doppler shift frequencies generated at the same velocity. Therefore, the power spectrum normalization unit  222  normalizes the frequency axis of the spectrogram such that those having the same velocity are made to correspond to each other. 
     The adaptive filter unit  223  inputs the normalized first power spectrum and second power spectrum to an adaptive filter. Accordingly, the Doppler beat (see  FIGS. 13 and 14 ) caused by the stationary tissue  603  is removed. 
     The velocity pulse wave calculation unit  224  calculates a velocity pulse wave on the basis of a time-series change of the power spectrum of the beat signal output from the adaptive filter unit  223 . The pulse rate calculation unit  225  calculates a pulse rate on the basis of the velocity pulse wave calculated by the velocity pulse wave calculation unit  224 . 
     The velocity distribution information calculation unit  226  calculates velocity distribution information by performing an operation such as integral processing on the power spectrum of the signal output from the adaptive filter unit  223 . 
     The power spectrum calculation unit  221 , the velocity pulse wave calculation unit  224 , the pulse rate calculation unit  225 , and the velocity distribution information calculation unit  226  function as a blood flow-related information calculation unit that calculates blood flow-related information. The blood flow-related information calculation unit may have another configuration capable of calculating the blood flow-related information. 
     Further, the power spectrum normalization unit  222  and the adaptive filter unit  223  function as a noise cancellation unit that cancels noise caused by the relative motion between the sensor head  210  and the living body  600  from the blood flow-related information. The noise cancellation unit may have another configuration capable of cancelling the noise. 
     The blood flow measurement apparatus  200  has the above-mentioned configuration. It should be noted that sensor head  210  functions as an analog signal processing block of the blood flow measurement apparatus  200  and the information processing apparatus  220  functions as a digital signal processing block. 
     The configuration of the blood flow measurement apparatus  200  is not limited thereto. Digital signal processing may be performed at the sensor head  210  and analog signal processing may be performed at the information processing apparatus  220 . 
     [Operation of Blood Flow Measurement Apparatus According to Second Embodiment] 
     An operation of the blood flow measurement apparatus  200  will be described. 
     As shown in  FIG. 39 , the first laser light source  251  and the second laser light source  252  of the light emitting unit  211  are alternately irradiated with the laser light L to the living body  600 . 
     First scattered light obtained when laser light emitted from the first laser light source  251  is scattered and second scattered light obtained when laser light emitted from the second laser light source  252  is scattered are respectively received by the measurement light receiving element  261 . 
     Here, when the wavelengths of the laser light are different, the depth penetration of the laser light in the living body  600  is different, the depth of the observation area is different, and the patterns of multiple scattering thereafter are different. Therefore, when the wavelengths of the irradiated laser light are switched, differences is caused in the observation target red blood cells  602  and the observation target stationary tissue  603 , and the detected beat signals are also mixed at different ratios. Therefore, in the first beat signal generated by receiving the first scattered light and the second beat signal generated by receiving the second scattered light, signals from a plurality of targets are mixed at different ratios. This difference in mixing ratio can be utilized to effectively cancel noise components. 
     In the blood flow measurement apparatus according to the first embodiment, the relative velocity and the relative angular velocity noise can be effectively reduced, but the beat signal of the scattered light of the stationary tissue may still remain. The mixing ratio of the beat signal between the scattered light of such stationary tissue and the beat signal by the red blood cells is also different between the first beat signal and the second beat signal. Thus, the ratio of the components derived from the beat signal due to the stationary tissue  603  to the components derived from the beat signal due to the red blood cells  602  differs between the first power spectrum generated from the first beat signal and the second power spectrum generated from the second beat signal. Therefore, the relative velocity components between the sensor head  210  and the living body  600  can be removed by normalizing both the power spectra by the power spectrum normalization unit  222  and then performing the processing by the adaptive filter unit  223 . 
     In the blood flow measurement apparatus  200 , the disturbance noise can be prevented from being generated in the following manner. 
     [Variations of Light Receiving Unit] 
     The light receiving unit  212  is not limited to the above-mentioned configuration.  FIG. 41  is a schematic diagram showing a light receiving unit  212  having another configuration. 
     As shown in the figure, the light receiving unit  212  includes another measurement light receiving element  263  in addition to the measurement light receiving element  261 . The measurement light receiving element  261  and the measurement light receiving element  263  are light receiving elements for blood flow measurement as described above. The sensor head  210  outputs the difference between the beat signals detected by the two pixels as a beat signal. By taking such a configuration, it is possible to effectively remove noise components that vary in the same manner in two pixels, such as external light noise. 
     Although a technique of using such a difference signal has been known for a long time, it is possible to use a similar difference signal by devising the arrangement of pixels and the arrangement of light sources as shown in  FIG. 41  in the blood flow measurement apparatus  200 . 
     [Hardware Configuration] 
     A hardware configuration of the information processing apparatus  220  will be described.  FIG. 42  is a schematic diagram showing a hardware configuration of the information processing apparatus  220 . As shown in  FIG. 2 , the information processing apparatus  220  includes an interface (I/F)  281 , arithmetic processing circuit  282 , a bus  283 , and a memory  284 . 
     The I/F  281  provides a beat signal obtained from the sensor head  210  (see  FIG. 10 ) to the bus  283 . The first laser light source  251  and the second laser light source  252  emit light alternately. Therefore, for the beat signal obtained from the sensor head  210 , the beat signal  1  and the beat signal  2  are alternately supplied in time division. The arithmetic processing circuit  282  is a processor such as a central processing unit (CPU). A spatial speckle signal is obtained via the bus  283  and various types of arithmetic processing are performed while exchanging information with the memory  284 . 
     The power spectrum calculation unit  221 , the power spectrum normalization unit  222 , the adaptive filter unit  223 , the velocity pulse wave calculation unit  224 , the pulse rate calculation unit  225 , and the velocity distribution information calculation unit  226  having the above-mentioned functional configurations are realized by the cooperation of the arithmetic processing circuit  282  and the program. 
     It should be noted that some or all of the configurations of the information processing apparatus  220  may be realized on a computer network. 
     Third Embodiment 
     [Configuration of Blood Flow Measurement Apparatus According to Third Embodiment] 
     A configuration of a blood flow measurement apparatus  300  according to a third embodiment of the present technology will be described.  FIG. 43  is a schematic diagram of the blood flow measurement apparatus  300  according to the present embodiment.  FIG. 44  is a block diagram showing functions of the blood flow measurement apparatus  300 . 
     As shown in  FIG. 43 , the blood flow measurement apparatus  300  includes a sensor head  310  and an information processing apparatus  320 . The sensor head  310  and the information processing apparatus  320  are connected through a signal line  330 . The blood flow measurement apparatus  300  may be a wearable device. 
     The living body  600  that is the blood flow measurement target has red blood cells  602  flowing through a blood vessel  601 , and a stationary tissue  603 . 
     As shown in  FIG. 43 , the sensor head  310  is disposed close to or in close contact with the living body  600 . As shown in  FIGS. 43 and 44 , the sensor head  310  includes a light emitting unit  311 , a first light receiving unit  312 , a second light receiving unit  313 , a first analog signal processing unit  314 , a second analog signal processing unit  315 , a first ADC  316 , and a second ADC  317 . 
     The light emitting unit  311 , the first light receiving unit (sensor)  312 , and the second light receiving unit  313  are mounted on the sensor head  310 , and the relative position relative to the sensor head  310  is fixed.  FIG. 45  is a cross-sectional view of a sensor head  310 .  FIG. 46  is a plan view of the sensor head  310  viewed from the living body  600  side. 
     The light emitting unit  311  irradiates the living body  600  and the second light receiving unit  313  with the laser light L. As shown in  FIG. 45 , the light emitting unit  311  may include a laser light source  351  and a reflector  352 . 
     The reflector  352  is provided between the laser light source  351  and the living body  600  and reflects part of the laser light L emitted from the laser light source  351  toward the second light receiving unit  313 . The reflector  352  may be, for example, a micro-mirror mounted by a Micro Electro Mechanical Systems (MEMS) technology. 
     It should be noted that the laser light source  351  may be provided with the irradiation light control unit described in the first embodiment. 
     The first light receiving unit  312  receives receives scattered light obtained when the laser light L is reflected by the living body  600 , and converts the received scattered light into a signal. The first light receiving unit  312  includes a measurement light receiving element  361  and an incident angle limiting unit  362 . 
     The measurement light receiving element  361  and the incident angle limiting unit  362  can respectively have the same configurations (see  FIG. 18 ) as the measurement light receiving element  161  and the incident angle limiting unit  162  described in the first embodiment. Further, the incident angle limiting unit  362  may have another configuration of the first embodiment (see  FIGS. 27 to 35 ). 
     The second light receiving unit  313  receives the laser light L reflected by the reflector  352  and scattered light obtained when the laser light L is scattered by the living body  600 . The second light receiving unit  313  converts the received scattered light into a signal. The second light receiving unit  313  includes a reflected light receiving element  363  and an incident angle limiting unit  364 . 
     The reflected light receiving element  363  can have the same configuration as the measurement light receiving element  361 . The incident angle limiting unit  364  can have the same configuration as the incident angle limiting unit  362 . 
     Further, the first light receiving unit  312  and the second light receiving unit  313  do not need to be provided with the incident angle limiting unit  362  and the incident angle limiting unit  364 . 
     The first analog signal processing unit  314  performs signal processing such as amplification on the signal output from the first light receiving unit  312  and supplies it to the first ADC  316 . 
     The second analog signal processing unit  315  performs signal processing such as amplification on the signal output from the second light receiving unit  313  and supplies it to the second ADO  317 . 
     The first analog to digital converter (ADC)  316  converts the analog signal supplied from the first analog signal processing unit  314  into a digital signal. 
     The first ADC  316  outputs the converted digital signal to the information processing apparatus  320  via the signal line  330 . 
     The second ADC  317  converts the analog signal supplied from the second analog signal processing unit  315  into a digital signal. The second ADC  317  outputs the converted digital signal to the information processing apparatus  320  via the signal line  330 . 
     As shown in  FIG. 44 , the information processing apparatus  320  includes a power spectrum calculation unit  321 , a relative velocity beat extraction unit  322 , a velocity pulse wave calculation unit  323 , a velocity correction unit  324 , a pulse rate calculation unit  325 , and a velocity distribution information calculation unit  326 . 
     The power spectrum calculation unit  321  performs arithmetic processing such as Fourier transform on a beat signal obtained from the first ADC  316  and the second ADC  317 . Accordingly, the power spectrum calculation unit  321  calculates the power spectrum of the signal that may be changed to the frequency domain for each predetermined time. The predetermined time is, for example, 1/100 seconds. 
     The power spectrum calculation unit  321  calculates a power spectrum for each of a beat signal  1  generated from the scattered light received by the measurement light receiving element  361  and a beat signal  2  generated from the laser beam and the scattered light received by the reflected light receiving element  363 . The power spectrum based on the beat signal  1  is defined as a first power spectrum. The power spectrum based on the beat signal  2  is defined as a second power spectrum. 
     The relative velocity beat extraction unit  322  extracts a Doppler shift component due to the relative velocity between the sensor head  310  and the living body  600  from the first power spectrum and the second power spectrum. An adaptive filter technique can be used for this extraction. The second power spectrum includes frequency components of the beat signal due to the relative velocity v, which are substantially removed from the first power spectrum. Therefore, the second power spectrum often has a large crest in the high frequency part of the power spectrum. This component is hardly included in the first power spectrum. Therefore, this component can be extracted on a basis of a difference between both. The extracted frequency component has a peak at a frequency position proportional to the relative velocity v. Therefore, the extracted frequency component is informative about the relative velocity v. 
     The velocity pulse wave calculation unit  323  calculates a velocity pulse wave on the basis of a time-series change of the power spectrum of the beat signal output from the power spectrum calculation unit  321 . 
     The velocity correction unit  324  corrects the velocity distribution calculated on the basis of the first power spectrum using the Doppler shift component extracted by the relative velocity beat extraction unit  322 . An adaptive filter method can be used for this correction. 
     The pulse rate calculation unit  325  calculates a pulse rate on the basis of the velocity pulse wave calculated by the velocity pulse wave calculation unit  323  and corrected by the velocity correction unit  324 . 
     The velocity distribution information calculation unit  326  calculates velocity distribution information by performing an operation such as integral processing on the power spectrum of the beat signal output from the relative velocity beat extraction unit  322 . 
     The power spectrum calculation unit  321 , the velocity pulse wave calculation unit  323 , the pulse rate calculation unit  325 , and the velocity distribution information calculation unit  326  function as a blood flow-related information calculation unit that calculates blood flow-related information. The blood flow-related information calculation unit may have another configuration capable of calculating the blood flow-related information. 
     Further, the relative velocity beat extraction unit  322  and the velocity correction unit  324  function as a noise cancellation unit that cancels noise caused by the relative motion between the sensor head  310  and the living body  600  from the blood flow-related data. The noise cancellation unit may have another configuration capable of cancelling the noise. 
     The blood flow measurement apparatus  300  has the above-mentioned configuration. It should be noted that sensor head  310  functions as an analog signal processing block of the blood flow measurement apparatus  300  and the information processing apparatus  220  functions as a digital signal processing block. 
     The configuration of the blood flow measurement apparatus  300  is not limited thereto. Digital signal processing may be performed at the sensor head  310  and analog signal processing may be performed at the information processing apparatus  320 . Further, the second light receiving unit  313 , the second analog signal processing unit  315 , and the second ADC  317  are not essential. Noise may be cancelled in signal processing similar to that of the blood flow measurement apparatus according to the second embodiment by placing the reflector  352  so as to be able to reflect the laser light L toward the first light receiving unit  312  and then controlling the reflector  352  to alternately take a state to reflect the laser light L toward the first light receiving unit  312  and a state in which the laser light L does not reach the first light receiving unit  312 . The control on the reflector  352  described above can be realized by constituting the reflector  352  by an MEMS mirror and then changing the angle of reflection by a drive mechanism for the MEMS, for example. 
     [Operation of Blood Flow Measurement Apparatus According to Third Embodiment] 
     An operation of the blood flow measurement apparatus  300  will be described. 
     As shown in  FIG. 45 , the living body  600  is irradiated with the laser light L from the laser light source  351  of the light emitting unit  311 . The scattered light S obtained when the laser light L is reflected by the living body is received by the measurement light receiving element  361 . 
     Further, part of the laser light L is reflected by the reflector  352  and is received by the reflected light receiving element  363  without entering the living body  600 . 
     Here, when the relative velocity is caused between the sensor head  310  and the living body  600 , a beat is caused between the frequency f 0  of the laser light L incident on the reflected light receiving element  363  and the frequency f+Δf of the scattered light modulated with the relative velocity, which is incident on the measurement light receiving element  361 . Where Δf is approximately 2αv cos θ. 
     The relative velocity beat extraction unit  322  extracts a Doppler shift Δf component due to the relative velocity between the sensor head  310  and the living body  600  from the first power spectrum based on the scattered light and the second power spectrum based on the laser light. The center of gravity of the obtained Δf component represents 2αv cos θ. The velocity correction unit  324  removes components in the vicinity of this 2αv cos θ from the velocity information. 
     In the blood flow measurement apparatus  300 , the disturbance noise can be prevented from being generated in the following manner. 
     [Hardware Configuration] 
     A hardware configuration of the information processing apparatus  320  will be described.  FIG. 47  is a schematic diagram showing a hardware configuration of the information processing apparatus  320 . As shown in the figure, the information processing apparatus  320  includes an interface (I/F)  381 , an arithmetic processing circuit  382 , a bus  383 , and a memory  384 . 
     The I/F  381  provides a beat signal obtained from the sensor head  310  (see  FIG. 10 ) to the bus  383 . The arithmetic processing circuit  382  is a processor such as a central processing unit (CPU). The beat signal is obtained via the bus  383  and various arithmetic processing is executed while exchanging information with the memory  384 . 
     The power spectrum calculation unit  321 , the relative velocity beat extraction unit  322 , the velocity pulse wave calculation unit  323 , the velocity correction unit  324 , the pulse rate calculation unit  325 , and the velocity distribution information calculation unit  326  having the above-mentioned functional configurations are realized by the cooperation of the arithmetic processing circuit  382  and the program. 
     It should be noted that some or all of the configurations of the information processing apparatus  320  may be realized on a computer network. 
     Fourth Embodiment 
     [Configuration of Blood Flow Measurement Apparatus According to Fourth Embodiment] 
     A configuration of a blood flow measurement apparatus  400  according to a fourth embodiment of the present technology will be described.  FIG. 48  is a schematic diagram of the blood flow measurement apparatus  400  according to the present embodiment.  FIG. 49  is a block diagram showing functions of the blood flow measurement apparatus  400 . 
     As shown in  FIG. 48 , the blood flow measurement apparatus  400  includes a sensor head  410  and an information processing apparatus  420 . The sensor head  410  and the information processing apparatus  420  are connected through a signal line  430 . The blood flow measurement apparatus  400  may be a wearable device. 
     The living body  600  that is the blood flow measurement target has red blood cells  602  flowing through a blood vessel  601 , and a stationary tissue  603 . 
     As shown in  FIG. 48 , the sensor head  410  is disposed close to or in close contact with the living body  600 . As shown in  FIGS. 48 and 49 , the sensor head  410  includes a light emitting unit  411 , a light receiving unit  412 , an analog signal processing unit  413 , an ADC  414 , a light receiving element array  463 , an image analog signal processing unit  464 , and an image ADC  465 . 
     The light emitting unit  411  and the light receiving unit (sensor)  412  are mounted on the sensor head  410  and the relative position relative to the sensor head  410  is fixed.  FIG. 50  is a cross-sectional view of a sensor head  310 .  FIG. 51  is a plan view of the sensor head  410  viewed from the living body  600  side. 
     The light emitting unit  411  irradiates the living body  600  with laser light L As shown in  FIGS. 50 and 51 , the light emitting unit  411  includes a laser light source  451 . It should be noted that the laser light source  451  may be provided with the irradiation light control unit described in the first embodiment. 
     The light receiving unit  412  receives scattered light obtained when the laser light L is reflected by the living body  600  and converts the received scattered light into a signal. The light receiving unit  412  includes a measurement light receiving element  461  and an incident angle limiting unit  462 . 
     The measurement light receiving element  461  and the incident angle limiting unit  462  can respectively have the same configuration as the measurement light receiving element  161  and the incident angle limiting unit  162  described in the first embodiment (see  FIG. 18 ). 
     Further, the incident angle limiting unit  462  may have another configuration of the first embodiment (see  FIGS. 27 to 35 ). In addition, the light receiving unit  412  may be configured to include only the measurement light receiving element  461  or not to include an incident angle limiting unit  462 . 
     The analog signal processing unit  413  performs signal processing such as amplification on the signal output from the light receiving unit  412  and supplies it to the ADC  414 . 
     The analog to digital converter (ADC)  414  converts the analog signal supplied from the analog signal processing unit  413  into a digital signal. The ADC  414  outputs the converted digital signal to the information processing apparatus  420  via the signal line  430 . 
     The light receiving element array  463  is provided in the sensor head  410  and is an array in which a plurality of light receiving elements  463   a  is arranged. The number of light receiving elements  463   a  constituting the light receiving element array  463  is not particularly limited. It should be noted that a speckle particle size of a captured spatial speckle image may be adjusted by placing an optical component such as a pin hole and a lens between the light receiving element array  463  and the living body  600 . Alternatively, a surface image of the living body  600  may be captured in place of the spatial speckle image by placing an optical component such as a lens between the light receiving element array  463  and the living body  600 . 
     The image analog signal processing unit  464  performs signal processing such as amplification on the signal output from the light receiving element array  463  and supplies it to the image ADC  465 . 
     The image ADC  465  converts the analog signal supplied from the image analog signal processing unit  464  into a digital signal. The image ADC  465  outputs the converted digital signal to the information processing apparatus  420  via the signal line  430 . 
     As shown in  FIG. 49 , the information processing apparatus  420  includes a power spectrum calculation unit  421 , a velocity pulse wave calculation unit  422 , a relative velocity detection unit  466 , a velocity correction unit  423 , a pulse rate calculation unit  424 , and a velocity distribution information calculation unit  425 . 
     The power spectrum calculation unit  421  performs arithmetic processing such as Fourier transform on the beat signal obtained from the ADC  414 . Accordingly, the power spectrum calculation unit  421  calculates the power spectrum of the signal that may be changed to the frequency domain for each predetermined time. The predetermined time is, for example, 1/300 to 1/10 seconds. 
     The power spectrum calculation unit  421  calculates a power spectrum for the signal beat generated from the scattered light received the measurement light receiving element  461 . 
     The velocity pulse wave calculation unit  422  calculates a velocity pulse wave on the basis of a time-series change of the power spectrum of the beat signal output from the power spectrum calculation unit  421 . 
     The relative velocity detection unit  466  detects the relative velocity between the sensor head  410  and the living body  600  in the xy direction (direction perpendicular to the optical axis direction (z direction) of the laser light L) on the basis of the output of the light receiving element array  463 . The relative velocity detection unit  466  is capable of detecting the relative velocity on the basis of the amount of deviation for each frame in the spatial speckle image with the laser beam, which is captured by the light receiving element array  463 . 
     The velocity correction unit  423  corrects the velocity distribution calculated on the basis of the Power spectrum using the relative velocity between the sensor head  410  and the living body  600  in the xy direction supplied from the relative velocity detection unit  466 . An adaptive filter technique can be used for this removal. 
     The pulse rate calculation unit  424  calculates a pulse rate on the basis of the velocity pulse wave corrected by the velocity correction unit  423 . 
     The velocity distribution information calculation unit  425  calculates velocity distribution information by performing an operation such as integral processing on the power spectrum calculated by the power spectrum calculation unit  421 . 
     The power spectrum calculation unit  421 , the velocity pulse wave calculation unit  422 , the pulse rate calculation unit  424 , and the velocity distribution information calculation unit  425  function as a blood flow-related information calculation unit that calculates blood flow-related information. The blood flow-related information calculation unit may have another configuration capable of calculating the blood flow-related information. 
     Further, the velocity correction unit  423  functions as a noise cancellation unit that cancels noise caused by the relative motion between the sensor head  410  and the living body  600  from the blood flow-related information. The noise cancellation unit may have another configuration capable of cancelling the noise. 
     The blood flow measurement apparatus  400  has the above-mentioned configuration. It should be noted that the sensor head  410  functions as an analog signal processing block of the blood flow measurement apparatus  400  and the information processing apparatus  420  functions as a digital signal processing block. 
     The configuration of the blood flow measurement apparatus  400  is not limited thereto. Digital signal processing may be performed at the sensor head  410  and analog signal processing may be performed at the information processing apparatus  420 . 
     [Operation of Blood Flow Measurement Apparatus According to Fourth Embodiment] 
     An operation of the blood flow measurement apparatus  400  will be described. 
     As shown in  FIG. 50 , the living body  600  is irradiated with the laser light L from the laser light source  451  of the light emitting unit  411 . The scattered light S obtained when the laser light L is reflected by the living body is received by the measurement light receiving element  461  and the light receiving element array  463 . 
     Here, the amount of deviation of the image captured by the light receiving element array  463  for each frame is proportional to the relative velocity between the living body  600  and the sensor head  410 . The relative velocity in the xy direction obtained on the basis of the deviation vector for each frame corresponds to v sin θ which determines the frequency of the relative velocity noise. The velocity correction unit  423  removes a v sin θ-correlated component included in the velocity signal calculated by the power spectrum calculation unit  421 . The adaptive filter technique can be applied to this removal. 
     In the blood flow measurement apparatus  400 , the disturbance noise can be prevented from being generated as described above. 
     [Variations of Light Receiving Unit] 
     The light receiving unit  412  is not limited to the above-mentioned configuration.  FIG. 52  is a schematic diagram showing a light receiving unit  412  having another configuration. 
     As shown in the figure, the light receiving unit  412  may include another measurement light receiving element  467  in addition to the measurement light receiving element  461 . As in the measurement light receiving element  461 , the measurement light receiving element  467  is a light receiving element for measuring a beat signal generated due to the motion of the red blood cell  602 . 
     By taking a difference between the beat signal received by the measurement light receiving element  461  and the measurement light receiving element  467 , it is possible to cancel noise of a noise source such as external light, which causes similar changes in both the pixels. 
     [Hardware Configuration] 
     A hardware configuration of the information processing apparatus  420  will be described.  FIG. 53  is a schematic diagram showing a hardware configuration of the information processing apparatus  420 . As shown in the figure, the information processing apparatus  420  includes an interface (I/F)  481 , arithmetic processing circuit  482 , a bus  483 , and a memory  484 . 
     The I/F  481  provides a beat signal obtained from the sensor head  410  (see  FIG. 10 ) to the buses  483 . Further, the image data of the light receiving element array  463  obtained from the sensor head  410  is also supplied to the bus  483 . The arithmetic processing circuit  482  is a processor such as a central processing unit (CPU). A beat signal and image data are obtained via the bus  483  and various types of arithmetic processing are performed while exchanging information with the memory  484 . 
     The power spectrum calculation unit  421 , the velocity pulse wave calculation unit  422 , the relative velocity detection unit  466 , the velocity correction unit  423 , the pulse rate calculation unit  424 , and the velocity distribution information calculation unit  425  having the above-mentioned functional configurations are realized by the cooperation of the arithmetic processing circuit  482  and the program. 
     It should be noted that some or all of the configurations of the information processing apparatus  420  may be realized on a computer network. 
     (Other Configurations) 
     The light emitting unit and the light receiving unit of the blood flow measurement apparatus according to each of the above-mentioned embodiments can have configurations shown below. In the following description, the same reference signs as those of the first embodiment are given to the same configurations as those of the first embodiment, but they can also be applied to the second to fourth embodiments. 
     (Use of Louver Filter) 
     The light receiving unit of the blood flow measurement apparatus according to each of the above-mentioned embodiments includes an incident angle limiting unit that limits the incident angle of light incident on the measurement light receiving element to a predetermined angle or less. Here, a louver-type optical path limiting filter (hereinafter, referred to as louver filter) may be used as the incident angle limiting unit. The louver filter is often used as a filter for peep prevention of a display. 
       FIG. 54  is a schematic diagram showing a light receiving unit  1100  including an incident angle limiting unit using a louver filter. As shown in the figure, the light receiving unit  1100  includes a measurement light receiving element  1110  and a louver filter  1120 . 
     The measurement light receiving element  1110  can have the same configuration as the measurement light receiving element  161  described in the first embodiment. The measurement light receiving element  1110  includes a light receiving surface  1110   a.    
     The louver filter  1120  functions as an incident angle limiting unit. The louver filter  1120  includes a light transmissive member  1121  and louvers  1122 . The thickness of the louver filter  1120  is, for example, 0.3 mm. 
     The light transmissive member  1121  is made of a material that transmits at least scattered light. Each of the louvers  1122  is a plate-like member made of a material that does not transmit light having a wavelength of scattered light, the main surface of which extends along a direction perpendicular to the light receiving surface  1110   a . The plurality of louvers  1122  is embedded in the light transmissive member  1121  in a predetermined arrangement. 
       FIG. 55  is a schematic diagram showing the arrangement of the louvers  1122 . As shown in the figure, the louver filter  1120  includes a first louver layer  1122   a  and a second louver layer  1122   b . The first louver layer  1122   a  includes a plurality of louvers  1122  extending parallel to an X direction with a predetermined spacing. The second louver layer  1122   b  includes a plurality of louvers  1122  extending parallel to a Y direction with a predetermined spacing. The X direction and the Y direction are directions crossing each other. Typically, the X direction and the Y direction are orthogonal to each other. Further, a direction orthogonal to the X and Y directions defined as a Z direction. 
     The height (Z direction) and spacing (X and Y directions) of the louvers  1122  is adjustable. The louvers  1122 , for example, may have a height of about 130 μm and a spacing of about 40 μm. 
     The louver filter  1120  limits the incident angle of light incident on the louver filters  1120  to be equal to or smaller than a predetermined angle through the louvers  1122 .  FIG. 56  is a schematic diagram showing limitation of the incident angle of the incident light by the louver filter  1120 . 
     As shown in the figure, components of light incident on the louver filter  1120 , the angle θ of which is equal to or higher than a certain angle, are absorbed by the louvers  1122  and do not transmit through the louver filter  1120 . That is, the louver filter  1120  limits the incident angle of light incident on the measurement light receiving element  1110  to be equal to or smaller than the predetermined angle. It should be noted that the figure shows limitation of the incident angle in one of the X direction and Y direction, though similar limitation of the incident angle is done for the other. Further, the louver filter  1120  may have one louver layer or may have three or more louver layers. For example, in the case of a single-layer louver in the X direction, the limitation in the Y direction in which the angle is not limited by the louver can be performed by providing a light shielding window on the living body  600  side of the louver filter  1120  or the like. 
     The louver filter can set the limiting angle θ in a wide range of about 5° to 30° by adjusting the height and spacing of the louvers  1122 . It is favorable for the angular limitation of the present technology. 
     By using the louver filter  1120  in the incident angle limiting unit, it is possible to extremely thin the thickness of the incident angle limiting unit unlike the general light shield and the pin hole. It is thus possible to increase the utilization efficiency of light. Further, the louvers  1122  can be arranged at very close pitches (&lt;40 μm), which is difficult with conventional light shields. 
     Further, the light transmissive member  1121  may be made of a material such as IR black that shields light other than infrared rays. Accordingly, the louver filter  1120  is capable of shielding external light (sunlight) or the like that affects blood flow measurement. The sensor head can be thus thinned. 
     In addition, the light receiving unit may include a louver filter and a lens array.  FIG. 57  is a schematic diagram of a light receiving unit  1200  including a measurement light receiving element  1210 , a lens array  1220 , and a louver filter  1230 . 
     The measurement light receiving element  1210  can have the same configuration as the measurement light receiving element  161  described in the first embodiment. 
     The lens array  1220  is constituted by an optical member and can be a lens array in which a plurality of lenses  1220   a  is arranged. By bringing the lens array  1220  into contact with the measurement light receiving element  1210 , it is possible to facilitate the fixing of the positional relationship between the lens array  1220  and the measurement light receiving element  1210 . 
     The louver filter  1230  is disposed in front of the lens array  1220  (opposite to the measurement light receiving element  1210 ) and has the same configuration as the louver filter  1120 . 
     Since the louver filter  1230  is disposed in front of the lens array  1220 , the louvers  1122  have narrow pitches. Therefore, it is possible to limit the incident angle to the measurement light receiving element  1210  even without adjusting the predetermined positional relationship by defining the positional relationship between the lens array  1220  and the louvers  1122 . Without limiting the positional relationship, moire is generated in light incident on the measurement light receiving element  1210  under the influence of two different periodic structures. However, since the blood flow measurement apparatus according to the present technology does not utilize an image of incident light, generation of moire can be tolerated. Further, since the incident angle limiting unit is separated from the lens array  1220 , the lenses can be densely paved in the lens array  1220 , and only light having an incident angle at which the light is mixed between the neighboring lenses  1220   a  is removed. Therefore, it is possible to improve the utilization efficiency of light. 
     It should be noted that the lens array  1220  and the louvers  1122  may also be configured to limit the positions to a particular relationship as described below. 
       FIGS. 58 to 60  are schematic diagrams each showing a variation of the light receiving unit  1200 . As shown in  FIG. 58 , the lens array  1220  may be spaced apart from the measurement light receiving element  1210  and an air layer may be provided between the lens array  1220  and the measurement light receiving element  1210 . With this configuration, the focal length of the lens  1220   a  can be shortened. 
     In addition, as shown in  FIG. 59 , the lens array  1220  and the measurement light receiving element  1210  may be spaced apart and a filler  1231  such as a mold resin may be disposed therebetween. With this configuration, it is possible to adjust the height and tilt of the lens array  1220 . The resistance to the flatness of a light receiving surface  1210   a  of the measurement light receiving element  1210  and a flat surface  1220   b  which is the back surface of the lens array  1220  can be improved. 
     In addition, as shown in  FIG. 60 , the lens array  1220  may be provided in contact with the louver filter  1120 . The louver filter  1120  may be disposed such that the louvers  1122  are opposite to the measurement light receiving element  1210  as shown in the figure. 
     With any configuration of the light receiving unit  1200 , one or both of the lens array  1220  and the light transmissive member  1121  may be made of a material that shields light other than infrared rays, such as IR black. 
       FIG. 61  is a schematic diagram showing a relationship between the opening of the lens  1220   a  and the aspect ratio of the louvers  1122  when the lens array  1220  is in contact with the measurement light receiving element  1210 . As shown in the figure, it is assumed that θ L2  is an allowable upper limit incident angle of the lens  1220   a  at which the incident light enters the adjacent lens  1220   a  at an angle deviating from the incident angle range of ±θ L2  in the lens  1220   a  and θ L1  is an emission upper limit angle when the louver filter  1230  limits the exit angle to the range of ±θ L1 . 
     In this case, a site (irradiation spot R, see  FIG. 18 ) of the observation target, which emits the scattered light S, is located in the front of the measurement light receiving element  1210  (emits light in the front of the measurement light receiving element  1210 ), it is possible to prevent mixing of light with the adjacent lens  1220   a  by satisfying tan θ L2 &gt;tan θ L1 . 
     Further, when a site of the observation target, which emits the scattered light S, is located in one direction of the measurement light receiving element  1210  (emits light in the front of the light emitting unit), it is possible to prevent mixing of light with the adjacent lens  1220   a  by satisfying tan θ L2 &gt;2 tan θ L1 . 
       FIG. 62  is a schematic diagram showing a relationship between the opening of the lens  1220   a  and the aspect ratio of the louvers  1122  in a case where an air layer exists between the lens array  1220  and the measurement light receiving element  1210 . In this case, as shown in the figure, it is sufficient that θ L2  may be considered as an angle in the air layer. Other points are similar to those in the case of  FIG. 61 . 
     Further, the light receiving unit may include louvers and a lens array that are integrated defining the positional relationship between the louvers and the lens array.  FIG. 63  is a schematic diagram of a light receiving unit  1300  including a measurement light receiving element  1310  and a louver (hereinafter, referred to as louver lens array)  1320  integrated with a lens array. 
     The measurement light receiving element  1310  can have the same configuration as the measurement light receiving element  161  described in the first embodiment. 
     The louver lens array  1320  is constituted by an optical member and can be a lens array in which a plurality of lenses  1320   a  is arranged. Louvers  1321  are disposed between the individual lenses  1320   a . Like the louvers  1122 , the louvers  1321  are arranged in parallel with a predetermined spacing in the X and Y directions. Alternatively, in a case where the lens array is arranged in a honeycomb shape, louvers are arranged on side portions of the hexagon, which correspond to the walls of the honeycomb. 
     Accordingly, individual lenses  1320   a  are separated from neighboring lenses  1320   a  by the louvers  1321 . 
     As described above, it is possible to remove only the light mixed with the adjacent lens  1320   a  and make the other light incident on the measurement light receiving element  1310  by integrating the louvers and the lens array. Maximum light utilization efficiency can be realized. The louver lens array  1320  may be made of a material that shields light other than infrared rays, such as IR black. 
     It should be noted that in the light receiving unit  1200  and the light receiving unit  1300 , the louvers  1122  and the louvers  1321  does not need to be perpendicular to the light receiving surface of the measurement light receiving element, and the louvers  1122  and the louvers  1321  may be inclined with respect to the light receiving surface. 
       FIG. 64  is a schematic diagram showing a louver filter  1230  in which one of two louver layers is inclined.  FIG. 65  is a schematic diagram showing a louver lens array  1320  in which one of two louver layers is inclined. 
     As shown in these figures, in a case where the irradiation spot R is present in the living body  600  in front of the light emitting unit  111  (see  FIG. 18 ), a large amount of scattered light S emitted from the irradiation spot R can be guided to the measurement light receiving element by tilting the louvers. 
     (Use of Polarization Filter) 
     The light receiving unit of the blood flow measurement apparatus according to each of the above-mentioned embodiments may include a polarization filter as an incident angle limiting unit. 
       FIG. 66  is a schematic diagram showing a light receiving unit  1400  including a polarization filter. As shown in the figure, the light receiving unit  1400  includes a measurement light receiving element  1410  and the polarization filter  1420 . 
     The measurement light receiving element  1410  can have the same configuration as the measurement light receiving element  161  described in the first embodiment. 
     The polarization filter  1420  is a filter that attenuates a component of light having a polarization direction that is a specific direction. As shown in  FIG. 66 , it is disposed between the measurement light receiving element  1410  and the measurement target (living body). The polarization filter  1420  can be oriented in accordance with the polarization direction of the laser light source. 
       FIG. 67  is a schematic diagram showing the polarization direction of the polarization filter  1420  and is a plan view of the sensor head as viewed from the living body side. The laser light source  151  emits linearly polarized laser light.  FIG. 67  shows the polarization direction of the laser light emitted from the laser light source  151  as a direction D 1 . It should be noted that the polarization direction of laser light is not limited to that shown in the figure. 
     As shown in the figure, the polarization filter  1420  is arranged such that the polarization direction is the direction D 1 . By making the polarization direction of the polarization filter  1420  the same as the polarization direction of the laser light source  151 , scattered light generated by single scattering in the living body passes through the polarization filter  1420  almost as it is. 
     On the other hand, the direction of polarization of the scattered light generated by multiplex scattering in the living body becomes random by multiple scattering. Therefore, the amount of light is halved when the light is transmitted through the Polarization filter  1420 . In addition, the amount of external light is also halved. 
     Accordingly, the resistance to external light is improved about twice and the brightness of the irradiation spot directly hit by the laser becomes conspicuous, resulting in the effect of limiting the incident angle. Accordingly, another incident angle limiting mechanism can be omitted or can have a simplified configuration, and the polarization filter  1420  functions as an incident angle limiting unit. 
     Further, the polarization filter may be combined with another incident angle limiting mechanism.  FIG. 68  is a schematic diagram showing a light receiving unit  1500  including a polarization filter and another incident angle limiting mechanism. As shown in the figure, the light receiving unit  1500  includes a measurement light receiving element  1510 , a lens array  1520 , and a polarization filter  1530 . 
     The measurement light receiving element  1510  can have the same configuration as the measurement light receiving element  161  described in the first embodiment. 
     The lens array  1520  serves as an incident angle limiting unit. The lens array  1520  is constituted by an optical member and is a lens array in which a plurality of lenses  1520   a  is arranged. Further, the lens array  1520  may be a pin hole structure (see  FIG. 35 ) and may be a louver-type privacy louver filter  1120  as described above. 
     The polarization filter  1530  is a filter that attenuates a component of light having a polarization direction that is a specific direction. As shown in  FIG. 68 , it is disposed between the lens array  1520  and the measurement target (living body). Further, the polarization filter  1530  may be disposed between the lens array  1520  and the measurement light receiving element  1510 . 
     The polarization filter  1530  can be oriented in accordance with the polarization direction of the laser light source.  FIG. 69  is a schematic diagram showing the polarization direction of the polarization filter  1530  and is a plan view of the sensor head as viewed from the living body side. The laser light source  151  emits linearly polarized laser.  FIG. 69  shows the polarization direction of the laser light source  151  as a direction D 2 . It should be noted that the polarization direction of laser light source  151  is not limited to that shown in the figure. 
     As shown in the figure, the polarization filter  1530  is arranged such that the polarization direction is a direction D 3 . The direction D 3  is a direction perpendicular to the direction D 2 . By making the polarization direction of the polarization filter  1530  orthogonal to the polarization direction of the laser light source  151 , light that has caused single scattering in the living body is blocked by the polarization filter  1530 . 
     Accordingly, the ratio of light from measurement target particles in blood, which is contained in the scattered light incident on the measurement light receiving element  1510 , is greatly increased. Only multiple scattered light is observed. Therefore, although the absolute amount of light incident on the measurement light receiving element  1510  is reduced, an S/DC ratio (ratio of the amount of beat to the amount of light) is improved. Therefore, in a case where the incident intensity of the coherent light is sufficiently high relative to the incident intensity of the external light, it is favorable when lowering the intensity of the coherent light by increasing the amplification factor of the electric circuit and performing a low-power operation. 
     It should be noted that in the above-mentioned configuration, the polarization direction of the polarization filter  1530  may be set in the direction D 2 . In that case, the performance of resistance to the external light is enhanced. 
     In addition, the polarization filter may be combined with a louver filter.  FIG. 70  is a schematic diagram showing a light receiving unit  1600  including a polarization filter and a louver filter. As shown in the figure, the light receiving unit  1600  includes a measurement light receiving element  1610 , a lens array  1620 , a polarization filter  1630 , and a louver-type privacy filter (hereinafter, referred to as louver filter)  1640 . 
     The measurement light receiving element  1610  can have the same configuration as the measurement light receiving element  161  described in the first embodiment. 
     The lens array  1620  serves as an incident angle limiting unit. The lens array  1620  is constituted by an optical member and is a lens array in which a plurality of lenses  1620   a  is arranged. 
     The polarization filter  1630  is a filter that attenuates a component of light having a polarization direction that is a specific direction. As shown in  FIG. 70 , it is disposed between a louver filter  1640  and the lens array  1620 . In this case, the polarization filter is configured to be incorporated in a part of the incident angle limiting unit. Further, the polarization filter  1630  may be disposed between the louver filter  1640  and the measurement target (living body). 
     The louver filter  1640  is disposed between the lens array  1620  and the living body and has the same configuration as the louver filter  1120 . 
     The polarization filter  1630  can be installed with S-polarized light (large wristwatch-type device or the like) for the purpose of reducing power consumption under favorable conditions or can be installed with P-polarized light for the purpose of resistance to high external light conditions (wristband of thin design or the like). 
     In addition, the polarization filter may be combined with a louver-type privacy filter and a liquid crystal plate.  FIG. 71  is a schematic diagram showing a light receiving unit  1700  including a polarization filter, a louver-type privacy filter, and a liquid crystal plate. As shown in the figure, the light receiving unit  1700  includes a measurement light receiving element  1710 , a lens array  1720 , a polarization filter  1730 , a liquid crystal plate  1740 , and a louver-type privacy filter (hereinafter, referred to as louver filter)  1750 . 
     The measurement light receiving element  1710  can have the same configuration as the measurement light receiving element  161  described in the first embodiment. 
     The lens array  1720  serves as an incident angle limiting unit. The lens array  1720  is constituted by an optical member and is a lens array in which a plurality of lenses  1720   a  is arranged. 
     The polarization filter  1730  is a filter that attenuates a component of light having a polarization direction that is a specific direction. As shown in  FIG. 71 , it is disposed between the lens array  1720  and the liquid crystal plate  1740 . 
     The liquid crystal plate  1740  has a polarization direction which is rotated by 90 degrees between a state in which a voltage is applied and a state in which a voltage is not applied. The liquid crystal plate  1740  is disposed between the polarization filter  1730  and a louver filter  1750 . 
     The louver filter  1750  is disposed between the liquid crystal plate  1740  and the living body and has the same configuration as the louver filter  1120 . 
     The polarization filter  1730  can be disposed in a direction in which the S-polarized light passes in a state in which no voltage is applied to the liquid crystal plate  1740 . Accordingly, under a favorable condition with less ambient light, the S-polarized light can be used to reduce the power consumption. In a condition in which there is a lot of ambient light, the P-polarized light can be used by applying a voltage to the liquid crystal plate  1740  to enhance the external light resistance. 
     (Regarding Polarization Direction of Laser Light) 
     When a laser light source having linearly polarized light is used in the light emitting unit of the blood flow measurement apparatus according to each of the above-mentioned embodiments, it is favorable to set the polarization direction in a predetermined direction. It should be noted that the light receiving unit may include the polarized filter as described above or does not need to include the polarization filter. 
       FIG. 72  is a schematic diagram showing an optical path of the laser light L in a light emitting unit  1800  where the laser beam has a specific polarization direction.  FIG. 72( a )  is a side view.  FIG. 72( b )  is a plan view. As shown in the figure, the light emitting unit  1800  includes a laser light source  1810 , a lens  1820 , and a prism  1830 . 
     The laser light source  1810  can have the same configuration as the laser light source  151  described in the first embodiment. More favorably, the laser light source  1810  is configured to emit laser light having a predetermined direction described below as the polarization direction. 
     The lens  1820  collimates the incident laser light and makes it incident on the prism  1830 . The prism  1830  refracts the incident laser light and tilts its path. 
     The laser light L emitted from the prism  1830  is incident on the living body  600  to form a spot R. The scattered light S emitted from the spot R is received by the measurement light receiving element  161 . It should be noted that in the figure, the incident angle limiting unit (see  FIG. 18 ) is not shown. 
     As shown in  FIG. 72( a ) , when a straight line perpendicular to a light receiving surface  161   a  of the measurement light receiving element  161  is defined as a perpendicular line Lh, a configuration in which an optical axis C 1  of the laser light L emitted from the prism  1830  is inclined with respect to the perpendicular line Lh is employed. It should be noted that the configuration in which the laser light L is inclined is not limited to the one shown here. The details will be described later. 
     Further, as shown in  FIG. 72( b ) , the optical axis of the scattered light S received by the measurement light receiving element  161  is defined as an optical axis C 2  and the plane formed by the optical axis C 1  and the optical axis C 2  of the laser light L is defined as a plane E. The light emitting unit  1800  is configured such that the polarization direction of the laser light L is a direction perpendicular to the plane E (in the figure, direction D 4 ). 
       FIGS. 73 to 75  are schematic diagrams each showing the action of the laser light emitted from the laser light source  1810  having such a configuration. 
       FIG. 73  shows the laser light L incident on a cover  1840 . The cover  1840  is a transparent cover that is attached to the sensor head and covers the light receiving unit and the light emitting unit. As shown in the figure, when the laser light L is incident on the cover  1840 , part of the laser light L is reflected by the cover  1840 . 
     The figure shows reflected light Lr reflected by the cover  1840 . Since the reflected light Lr is not emitted to the living body, it is favorable that the amount of the reflected light Lr is small. 
     Here, in a case where the laser light source  1810  has the direction D 4  perpendicular to the plane E as the polarization direction as described above, the reflectance is reduced by the cover  1840  and it is possible to reduce the amount of reflected light Lr. 
     Further,  FIG. 74  shows the laser light L incident on the living body  600 . The surface (skin) of the living body  600  is rough, but reflected components are more on the front side shown as S. This indicates that the surface (skin) of the living body  600  is rough, but there are many surface components oriented in the normal direction of the surface. Therefore, when the direction D 4  perpendicular to the plane E is used as the polarization direction, reflected components are reduced, that is, light (Li in the figure) entering a skin  610  is increased. 
     In addition,  FIG. 75  is a schematic diagram showing an optical path of the laser light L in the light emitting unit  1800  described above. As shown in the same figure, the laser light L is refracted by a refractive surface  1830   a  of the prism  1830  and travels obliquely. 
     Here, in a case where the laser light source  1810  sets the direction D 4  perpendicular to the plane E as the polarization direction, the reflected light on the refractive surface  1830   a  decreases and the amount of laser light L reaching the living body can be increased. It should be noted that although the optical system for making the optical axis of laser light L oblique is not limited to that shown in  FIG. 75 , it is possible to reduce the reflected light at the refracting surface by setting the direction D 4  perpendicular to the plane E as the polarization direction as long as the optical path of the laser light L is on the plane E. 
     (Regarding Incident Direction of Laser Light) 
     In the light emitting unit of the blood flow measurement apparatus according to each of the above-mentioned embodiments, it is favorable that the incident direction of the laser light is as follows.  FIG. 76  is a schematic diagram showing an optical path of the laser light L in a light emitting unit  1900 . As shown in the figure, the light emitting unit  1900  includes a laser light source  1910 , a lens  1920 , and a prism  1930 . 
     The laser light source  1910  can have the same configuration as the laser light source  151  described in the first embodiment. The laser light source  1910  may emit laser light whose polarization direction is a specific direction or may emit laser light whose polarization direction is not defined. 
     The lens  1920  collimates the incident laser light and makes it incident on the prism  1930 . The prism  1930  refracts the incident laser light and tilts its path. 
     The laser light L emitted from the prism  1930  is incident on the living body  600  to form a spot R. The scattered light S emitted from the spot R is received by the measurement light receiving element  161 . It should be noted that in the figure, the incident angle limiting unit (see  FIG. 18 ) is not shown. 
     As shown in the figure, a straight line perpendicular to the light receiving surface  161   a  of the measurement light receiving element  161  is defined as a perpendicular line Lh. The light emitting unit  1900  is configured such that the optical axis C 1  of the laser light L emitted from the lens  1920  is inclined with respect to the perpendicular line Lh and the optical axis C 1  crosses the perpendicular line Lh on the surface of the living body  600 . 
     Thus, it is possible to improve the utilization efficiency of the laser light L by making the laser light L obliquely incident to form the spot R in front of the measurement light receiving element  161  and combining it with the incident angle limiting unit of the light receiving unit. 
       FIG. 77  shows a case where the laser light L is emitted from a laser light source  1941  to the front as a comparative. The laser light L is collimated by a lens  1942  and the irradiation spot R is formed on the surface of the living body  600  in front of the laser light source  1941 . The measurement light receiving element  161  obliquely receives the scattered light emitted from the spot R. 
     With this configuration, the irradiation spot R is formed in front of the laser light source  1941  and the range is spread in the internal scattering, though the spot R is formed away from the measurement light receiving element  161 . Light is scattered on the tissue of the living body  600  in all directions. Therefore, the amount of light reaching the measurement light receiving element  161  is reduced in accordance with the distance-square law. Therefore, it is difficult to improve the utilization efficiency of light. In addition, if the light receiving direction is greatly inclined, a problem in that the aberration is deteriorated arises in a case where the incident angle limiting unit is constituted by a lens, it is difficult to accurately limit the incident angle. 
     In contrast, as shown in  FIG. 76 , the irradiation spot R is formed in front of the measurement light receiving element  161  by making the laser light L obliquely incident. Accordingly, the distance between the irradiation spot R and the measurement light receiving element  161  is shortened. Further, even when the laser light source  1910  and the living body  600  are spaced away from each other, the laser light L is collimated by the lens  1920 . Therefore, the light intensity does not decrease. In addition, the measurement light receiving element  161  is capable of limiting the incident angle of the scattered light S through the incident angle limiting unit in a direction closer to the front of the measurement light receiving element  161 . Therefore, the measurement light receiving element  161  is capable of limiting the incident angle without being affected by deterioration of the aberration of the lens. 
     (Specific Configuration for Making Laser Light Obliquely Incident) 
       FIGS. 78 to 85  are schematic diagrams each showing a specific configuration of a light emitting unit  1900  for making the laser light L obliquely incident. In each figure, the light emitting unit  1900  is configured such that the laser light L is emitted from the laser light source  1910 , is collimated, and obliquely enters the front of the measurement light receiving element  161 . It should be noted that in  FIGS. 78 to 81 , the light receiving unit shows only the measurement light receiving element  161  and does not show the incident angle limiting unit. 
     As shown in  FIG. 78 , the light emitting unit  1900  may include a free-form surface prism  1951 . The laser light L emitted from the laser light source  1910  is inclined and collimated by the free-form surface prism  1951 . 
     Further, as shown in  FIG. 79 , the light emitting unit  1900  may include a support member  1952  for tilting the laser light source  1910 . The laser light L is emitted from the inclined laser light source  1910  and is collimated by a lens  1953 . 
     Further, as shown in  FIG. 80 , the light receiving surface  161   a  does not need to be parallel to the surface of the living body  600 . By arranging the light receiving unit also inclined in this manner, it is possible to reduce an adverse effect caused when light reflected from the light receiving unit enters the living body  600  again. 
       FIG. 89  is a diagram showing a more specific example of the configuration shown in  FIG. 80 . As shown in the figure, a laser light source  1910  and the measurement light receiving element  161  may be disposed in a light shield  1991  provided with openings  1991   a  and  1991   b  through the substrate  1952  and a substrate  1992 . 
     The laser light emitted from the laser light source  1910  travels through the opening  1991   a , is collimated by the lens  1953 , and enters the living body. The figure shows a living body surface  600   a  that has approached the sensor head in a relative motion and a living body surface  600   b  moves away from the sensor head in a relative motion. 
     The scattered light S emitted from the living body travels through an opening  1991   b , passes through a cover  1993 , and enters the light receiving unit. A lens array  1994  disposed in front of the measurement light receiving element  161 . The incident angle of the light receiving unit is limited by the lens array  1994  and the wall surface of the opening  1991   b . Further, if a light shielding mask having the same diameter is provided at a portion indicated by “※”, an equivalent angle limitation is possible. 
     With this configuration, the efficiency can be improved by tilting the laser light source  1910 . Further, the wall surface of the opening  1991   b  prevents light beams in the adjacent lenses in the lens array  1994  from mixing with each other. In addition, since the light receiving unit does not include the louvers and the light shield for each lens, the utilization efficiency of light is increased. 
     Further, as shown in  FIG. 81 , the light emitting unit  1900  may include a plurality of divided prisms. The laser light L emitted from the laser light source  1910  is collimated by a lens  1956 . The parallelized laser light L is inclined by refraction by a prism  1957  and a prism  1958 . 
     Further, as shown in  FIG. 81 , the light emitting unit  1900  may include a lens-integrated prism  1959  in which a lens  1959   a  and a prism  1959   b  are integrated. The laser light L emitted from the laser light source  1910  is collimated by the lens  1959   a . The parallelized laser light L is inclined by refraction by the prism  1959   b.    
     As shown in the figure, the light emitting unit  1900  can be disposed adjacent to the light receiving unit  1980 . The light receiving unit  1980  includes the measurement light receiving element  161 , the incident angle limiting unit  162 , and a light shield  1981 . The light shield  1981  prevents light leaking from the lens-integrated prism  1959  from reaching the incident angle limiting unit  162 . 
     Further, as shown in  FIG. 83 , the light emitting unit  1900  may include a lens-integrated prism  1960  in which a lens  1960   a  and a prism  1960   b  are integrated. The laser light L emitted from the laser light source  1910  is collimated by the lens  1960   a . The parallelized laser light L is inclined by refraction by the prism  1960   b.    
     Further, as shown in  FIG. 84 , the light emitting unit  1900  may include a lens-integrated prism  1961  in which a prism  1961   a  and a lens  1961   b  are integrated. The laser light L emitted from the laser light source  1910  is inclined by the prism  1961   a  and is collimated by the lens  1961   b . With this configuration, the amount of reflection when the coherent light is incident on the prism  1961  can be reduced by setting the incident angle on the lens-integrated prism  1961  around the Brewster&#39;s angle corresponding to the refractive index of the optical member of the prism  1961 . 
     Further, with any of the above-mentioned configurations, the reflection on the surface of the living body  600  can be reduced by setting the emission angle of light to the living body  600  around the Brewster&#39;s angle corresponding to the refractive index of the living body  600 . The details thereof will be described later. 
     As described above, in the light emitting unit  1900 , the laser light L can be tilted by various configurations. Further, the light emitting unit  1900  may have a configuration other than that described above as long as it can collimate and tilt the laser light L. 
     In addition, in the light emitting unit of the blood flow measurement apparatus according to each of the above-mentioned embodiments, the light receiving unit is favorably configured as follows in a case where the incident direction of laser light is inclined.  FIG. 85  is a schematic diagram showing the light receiving unit  1980  and a cover  1982 . 
     As shown in the figure, in a case where the laser light L is reflected in the cover  1982  at an angle θ, the incident angle limiting unit  162  can be configured to shield the incident light at the incident angle θ or more. Accordingly, it is possible to prevent the reflected light by the cover  1982  from reaching the measurement light receiving element  161 . It should be noted that in the configuration shown in  FIG. 80 , the light receiving unit and the living body  600  do not face each other. Therefore, the cover  1982  may be disposed so as not to cover the optical path of the light source light. It is possible to prevent the light at the angle θ from entering the incident angle limiting unit. 
     (Regarding Advantage in Oblique Incidence of Laser Light) 
     When the laser light is made incident obliquely as described above and the polarization direction of the laser light L is made perpendicular to the plane E (see  FIG. 72 ), the following advantage occurs. 
       FIG. 86  is a schematic diagram showing the laser light L obliquely incident on the living body  600 . It is favorable to set an incident angle A 1  in  FIG. 86  to be the Brewster&#39;s angle corresponding to the refractive index of the surface of the living body  600  (incident angle at which the reflectance of incident light of p-polarized light becomes 0 at the interface between flat surfaces of transparent materials). In a case where the living body  600  is human skin, the refractive index of the skin stratum corneum is known to be approximately 1.5, with a corresponding Brewster&#39;s angle of 56.3°. That is, it is favorable to set the incident angle of the laser light L such that A 1  takes an angle in the vicinity of 56.3°. 
     By aligning the polarization directions and utilizing the Brewster&#39;s angle, it is possible to minimize the loss of light reflected on the stratum corneum and to maximize the amount of light reaching the target to be observed. It should be noted that at an angle shallower than Brewster&#39;s angle (more parallel to the plane), the reflectivity increases rapidly as it becomes further from the Brewster&#39;s angle. Therefore, it is favorable to make the light incident at an angle slightly deeper than the Brewster&#39;s angle (more perpendicular to the plane) because the resistance to the fabrication tolerance and the like is increased. 
       FIG. 87  is an example of a configuration of a light emitting unit  1900  for making light incident on the living body  600  at the Brewster&#39;s angle. As shown in the figure, laser light L at a Brewster&#39;s angle A 2  corresponding to the refractive index of the lens-integrated prism  1961  is configured to be incident on the lens-integrated prism  1961 . Accordingly, the laser light L can be made to enter at a Brewster&#39;s angle A 1 . 
     (Regarding Spot Shape of Laser Light) 
     When the laser light is made incident on the living body  600  in the light emitting unit of the blood flow measurement apparatus according to each of the above-mentioned embodiments, it is favorable that the irradiation spot R is circular on the surface of the living body  600  as viewed from the measurement light receiving element in terms of the utilization efficiency of light. Therefore, in a case of making the laser light L obliquely incident on the living body  600  as described above, the spot R on the surface of the living body  600  can be made to have a shape close to a circular shape by making the beam shape of the laser light L elliptical. 
       FIG. 88  is a schematic diagram showing a configuration of a light emitting unit  1900  for making the spot R circular on the surface of the living body  600 .  FIG. 88( a )  is a diagram viewed in the Y direction.  FIG. 88( b )  is a diagram viewed in the X direction. It should be noted that X-Z plane coincides with the plane E (see  FIG. 72 ). 
     As shown in the figure, the light emitting unit  1900  includes a prism  1962 . The prism  1962  has a cylindrical lens  1962   a  and an aspherical lens  1962   b . Further, the starting point on the optical axis of the laser light L emitted from the prism  1962  as a point G. 
     As shown in  FIG. 88( a ) , as viewed in the Y direction, only by the amount of reduction in the emission angle of the laser light L in the cylindrical lens  1962   a , a beam diameter ϕ 1  of the laser light L in the aspherical lens  1962   b  is reduced. 
     Further, as shown in  FIG. 88( b ) , as viewed in the X direction, the beam diameter ϕ 2  of the laser light L increases with respect to the X direction because the cylindrical lens  1962   a  has no lens effects. Accordingly, it is possible to shape the laser light L into a large elliptical beam shape having a smaller diameter in the X direction at the point G and a larger diameter in the Y direction and form a circular irradiation spot shape in front of the measurement light receiving element. 
     (Application Examples of Present Technology) 
     In each of the above-mentioned embodiments, in the present embodiment, the blood flow measurement apparatus for detecting a motion of a red blood cell flowing in blood flow has been described. In addition to the blood flow, the present technology can be used for measurement of moving particles as long as those causes the laser Doppler effect. Accordingly, it is possible to realize a scattered light signal measurement apparatus capable of eliminating noise caused by the relative motion between the sensor and the measurement target. 
     In this case, the blood flow-related information calculation unit described in each of the above-mentioned embodiments functions as a particle velocity information calculation unit that calculates particle velocity information which is information regarding a particle velocity. The noise cancellation unit is capable of cancelling noise caused by the relative motion between the sensor and the measurement target on the basis of the particle velocity information. 
     It should be noted that the present technology may also take the following configurations. 
     (1) A scattered light signal measurement apparatus, including: 
     a light receiving element that receives scattered light obtained when coherent light emitted to a measurement target is scattered by the measurement target; and 
     an incident angle limiting unit that limits an incident angle of the scattered light that is incident on a single point of the light receiving element to be equal to or smaller than a predetermined angle and controls the single light receiving element to increase an aperture ratio. 
     (2) The scattered light signal measurement apparatus according to (1), in which 
     the incident angle limiting unit limits a light beam diameter of the scattered light that is incident on the single point of the light receiving element to be smaller than an entire light beam diameter that is incident on the light receiving element. 
     (3) The scattered light signal measurement apparatus according to (2), in which 
     the incident angle limiting unit concentrates the scattered light at a different point of the light receiving element in accordance with an incident angle. 
     (4) The scattered light signal measurement apparatus according to (3), in which 
     the incident angle limiting unit is a lens array in which a plurality of lenses is arranged. 
     (5) The scattered light signal measurement apparatus according to (4), in which 
     the light receiving element is at least one, and 
     the plurality of lenses concentrates scattered light at an incident angle limited by each of a plurality of lenses to the single light receiving element. 
     (6) The scattered light signal measurement apparatus according to (3) or (4), in which 
     the incident angle limiting unit is a lens, further including 
     a light shield that prevents mixing of light incident on the lens and light incident on a portion other than the lens. 
     (7) The scattered light signal measurement apparatus according to (2), in which 
     the light receiving element is at least one, and 
     the incident angle limiting unit that
         sets a plurality of pin holes with respect to the single light receiving element,   limits the incident angle of the scattered light that is incident on the single point of the light receiving element to be equal to or smaller than the predetermined angle, and   concentrates scattered light having an amount larger than an amount of scattered light, which is concentrated by a single pin hole, to the light receiving element.       

     (8) The scattered light signal measurement apparatus according to any one of (1) to (9), further including: 
     a coherent light source that emits the coherent light; and 
     an irradiation light control unit that controls an irradiation diameter of the coherent light emitted from the coherent light source. 
     (9) The scattered light signal measurement apparatus according to (8), in which 
     the irradiation light control unit makes the coherent light collimated, the coherent light being emitted from the light source. 
     (10) The scattered light signal measurement apparatus according to (9), in which 
     the irradiation diameter is 0.5 mm or more and 2 mm or less. 
     (11) The scattered light signal measurement apparatus according to any one of (1) to (10), further including: 
     a first coherent light source that emits first coherent light having a first wavelength range to the measurement target; and 
     a second coherent light source that emits second coherent light having a second wavelength range different from the first wavelength range to the measurement target. 
     (12) The scattered light signal measurement apparatus according to any one of (1) to (10), further including: 
     a coherent light source that emits the coherent light to the measurement target; 
     a reflector that reflects part of the coherent light emitted from the coherent light source; and 
     a light receiving element that receives the coherent light reflected by the reflector. 
     (13) The scattered light signal measurement apparatus according to any one of (1) to (10), further including 
     a light receiving element array in which a plurality of light receiving elements is arranged. 
     (14) The scattered light signal measurement apparatus according to any one of (1) to (13), in which 
     the incident angle limiting unit includes one louver layer or a plurality of louver layers, 
     the louver layer includes a plurality of louvers extending in parallel to each other, and 
     in a case of including the plurality of louver layers, the incident angle limiting unit is an optical path limiting filter in which directions in which the respective louvers of the layers extend cross each other. 
     (15) The scattered light signal measurement apparatus according to (14), in which 
     the incident angle limiting unit is a lens array in which a plurality of lenses is arranged, and 
     the louver has a height that enables mixing of light passing through the filter between adjacent lenses of the plurality of lenses to be prevented. 
     (16) The scattered light signal measurement apparatus according to any one of (1) to (15), further including 
     a polarization filter disposed in any optical path between the measurement target and the light receiving element. 
     (17) The scattered light signal measurement apparatus according to (16), in which 
     the polarization filter permits light having a polarization direction in a direction perpendicular to a plane formed by an optical axis of the coherent light and an optical axis of the scattered light to pass therethrough. 
     (18) The scattered light signal measurement apparatus according to any one of (1) to (17), in which 
     the coherent light source and the irradiation light control unit make the coherent light incident on the measurement target in a direction inclined with respect to an optical path of scattered light travelling toward the light receiving surface from the measurement target. 
     (19) The scattered light signal measurement apparatus according to (18), in which 
     the irradiation light control unit shapes the coherent light into an elliptical shape and emits the shaped coherent light such that an irradiation spot of the coherent light at the measurement target has a circular shape. 
     (20) The scattered light signal measurement apparatus according to any one of (1) to (19), in which 
     the coherent light source emits coherent light having a polarization direction in a direction perpendicular to a plane formed by an optical axis of the coherent light and an optical axis of the scattered light. 
     (21) The scattered light signal measurement apparatus according to any one of (1) to (19), in which 
     the coherent light source emits coherent light having a polarization direction in a direction parallel to a plane formed by an optical axis of the coherent light and an optical axis of the scattered light. 
     (22) An information processing apparatus, including: 
     a light receiving element that receives scattered light obtained when coherent light emitted to a measurement target is scattered by the measurement target; 
     an incident angle limiting unit that limits an incident angle of the scattered light that is incident on a single point of the light receiving element to be equal to or smaller than a predetermined angle and controls the single light receiving element to increase an aperture ratio; and 
     a particle velocity information calculation unit that calculates particle velocity information that is information regarding velocity of a particle inside the measurement target on the basis of a signal output from the light receiving element. 
     (23) The information processing apparatus according to (22), further including 
     a noise cancellation unit that cancels noise caused by a relative motion between the sensor and the measurement target from the particle velocity information on the basis of two signals respectively generated from two types of light different in optical characteristic. 
     (24) The information processing apparatus according to (23), in which 
     the noise cancellation unit cancels the noise on the basis of a first signal generated by first coherent light being scattered by the measurement target and received by the light receiving element and a second signal generated by second coherent light different in wavelength range from the first coherent light being scattered by the measurement target and received by the light receiving element. 
     (25) The information processing apparatus according to (23), in which 
     the noise cancellation unit cancels the noise on the basis of a first signal generated by the coherent light being scattered by the measurement target and received by the light receiving element and a second signal generated by light being received by the light receiving element, the light including scattered light obtained when the coherent light is scattered by the measurement target and the coherent light. 
     (26) The information processing apparatus according to (22), further including 
     a noise cancellation unit that detects a relative motion between the sensor and the measurement target on the basis of an output of a light receiving element array in which a plurality of light receiving elements provided in the sensor is arranged, and cancels noise caused by the relative motion between the sensor and the measurement target on the basis of the particle velocity information. 
     (27) The information processing apparatus according to any one of (22) to (26), in which 
     the measurement target is blood, and 
     the particle velocity information calculation unit calculates blood flow velocity information that is information regarding velocity of a red blood cell inside the measurement target. 
     (28) The information processing apparatus according to any one of (22) to (27), in which 
     the information processing apparatus includes a wearable device. 
     REFERENCE SIGNS LIST 
     
         
           100 ,  200 ,  300 ,  400  blood flow measurement apparatus 
           110 ,  210 ,  310 ,  410  sensor head 
           111 ,  211 ,  311 ,  411 ,  1800 ,  1900  light emitting unit 
           112112 ,  312 ,  313 ,  412 ,  1100 ,  1200 ,  1300 ,  1400 ,  1500 ,  1600 ,  1700  light receiving unit 
           151 ,  251 ,  252 ,  351 ,  451 ,  1810 ,  1910  laser light source 
           161 ,  261 ,  361 ,  461 ,  1210 ,  1310 ,  1410 ,  1510 ,  1610 ,  1710  measurement light receiving element 
           162 ,  165 ,  168 ,  171 ,  262 ,  362 ,  364 ,  462  incident angle limiting unit 
           163 ,  164  light shield 
           352  reflector 
           363  reflected light receiving element 
           463  light receiving element array 
           1120 ,  1230  louver filter 
           1320  louver align lens array 
           1630 ,  1730  polarization filter