the gaseous constituents of blood have been found to have a unique absorption spectrum in the infrared ( ir ) region . in fig1 there is shown by graph the infrared absorption spectra of blood containing co and blood virtually free from co from around 3 to 14 microns . it has been discovered that there is one distinct and strong absorption peak in this wavelength range for blood due to carbon monoxide in the bloodstream . an inspection of fig1 reveals absorption spectra which are nearly identical over a substantial portion of the wavelength region shown here . note , however , the distinct co absorption peak at λ ˜ 5 . 13 μm . at this wavelength , the presence of co in the blood sample reduces the sample &# 39 ; s transmission . it is , therefore , at this wavelength that the absorption tests for co will be taken . since absorption is directly related to concentration , by simply measuring absorption at that particular wavelength in the ir and comparing the absorption with that at a second &# 34 ; reference &# 34 ; wavelength , the concentration of carbon monoxide at a specific time and specific portion of the circulatory system can be determined . although not shown in fig1 similar considerations hold for other blood gases such as o 2 and co 2 . for blood with a large co 2 concentration , an absorption at 4 . 3 μm is observed , and for o 2 , 9 . 0 μm . referring now to fig2 a there is shown an embodiment designed to detect the relative concentration of co in blood by employing a single wavelength detector . an ir laser ( emitting a single wavelength ) or lamp 10 ( emitting a broad band of light ) emits ir radiation which then passes through a body tissue specimen 12 , such as the flesh of a relatively thin body portion , for example , an ear lobe or the skin between the thumb and index finger . a band pass filter 14 ( λ ˜ 5 . 13 μm ) is situated behind the tissue for passing only radiation that lies in a narrow spectral region centered at 5 . 13 μm to detector 16 . as noted above , light source 10 may be either a broad band lamp or a laser ( usually a gas laser ). when a laser is used for the light source , filter 14 may be omitted , and the required wavelength light transmitted through the tissue specimen and measured directly by detector 16 . the arrangement in fig2 a is capable of providing a relative reading of the absorption of light in blood at around 5 . 13 μm through a blood - containing body tissue . while the co content reading , obtained by the arrangement illustrated in fig2 a , will be accurate for that particular tissue specimen , it has no frame of reference , and as such cannot distinguish between tissue specimen from person to person whose body portions have different absorptive characteristics or even thicknesses . accordingly , a further embodiment of the invention is disclosed in fig2 b which overcomes this shortcoming . it measures the absolute co content by using double wavelength detection . there is provided a source of light , filters and detectors , much in the same manner as in fig2 a . an ir lamp 20 ( or laser ) passes light through body tissue specimen 22 . there is provided , however , in this arrangement a beam splitter 24 whereby light emerging from tissue 22 and falling on beam splitter 24 is divided into two equal beams . one beam passes directly through the specimen to bandpass filter 26 ( approximately 5 . 13 μm ) and on to detector 34 . the other half is reflected to mirror 28 to pass through bandpass filter 30 ( approximately 5 . 05 μm ) and on to detector 32 . the filters admit light in their respective bands to a pair of light detectors 32 and 34 similar to the one provided in fig2 a . note in fig1 that light in the wavelength of approximately 5 . 05 μm is transmitted freely through tissue sample 22 independent of the carbon monoxide content of the blood in the sample . the intent is to use the total absorption and scattering at 5 . 05 μm as a reference against which the carbon monoxide concentration can be measured exactly . for the carbon monoxide rich blood sample shown in fig1 the transmission at the 5 . 13 μm co peak is decreased whereas no co - related absorption peak is evident at 5 . 05 μm . note also that the &# 34 ; background &# 34 ; absorption ( i . e . the blood free of carbon monoxide - dashed line ) which is due to h 2 o and other blood constituents is almost the same at 5 . 13 μm as it is at 5 . 05 μm . consequently , the ratio of the transmitted intensities at 5 . 13 and 5 . 05 μm can be used to calculate the absolute carbon monoxide blood concentration . specifically , if we let i t ( 5 . 13 μm ) and i t ( 5 . 05 μm ) be the transmitted intensities at 5 . 13 and 5 . 05 μm , respectively , and if i i is the infrared intensity incident on the sample , then ( 1 ) i i is taken to be constant over the 5 . 05 & lt ; λ & lt ; 5 . 13 μm wavelength interval . and ( 3 ) α is the absorption coefficient of our co rich blood sample which , of course , is wavelength dependent . for the arrangements of fig2 a and 2b , due to the presence of skin and perhaps other tissue , the absorption coefficient will be composed of several parts . for example : α ( 5 . 13 μm )= α ( due to co )+ α ( background absorption , due to h . sub . 2 o & amp ; other constituents )+ α ( skin & amp ; tissue ) once it is realized that the absorption of skin and the blood &# 39 ; s &# 34 ; background &# 34 ; absorption vary only slightly around 5 . 13 μm ( 5 . 03 - 5 . 13 μm ), then the ratio of the transmitted light at 5 . 13 μm to that at 5 . 05 μm is written : ## equ1 ## but α co is related to the co contentration in the blood ( denoted as [ co ]) by the where α co is the co absorption cross - section in blood at 5 . 13 μm . remember , l is known and σ co is a constant independent of co concentration and can readily be determined . therefore , by simply taking the outputs of detectors 32 and 34 in fig2 b , electronically computing their ratio and solving equations ( 4 ) and ( 5 ), then the absolute co concentration can be rapidly determined . this invention rests on the rapid variation of blood &# 39 ; s co absorption around 5 . 13 μm as opposed to only a weak variation for skin , tissue and h 2 o in that same wavelength region . consequently , these extraneous contributions to the absorption coefficient cancel out when the ratio in equation ( 4 ) is taken . similar considerations hold when probing for co 2 at 4 . 3 μm or for o 2 at 9 . 0 μm . the same technique can be easily applied to measuring oxygen in blood concentrations at critical internal points of the human body circulatory system , such as in the aorta . presently , catheters are commonly used to allow access of various probes to the circulatory system , and they may take the form of a needle having an internal diameter sufficiently large to carry very small diameter optical fibers . the needle may actually provide for two fibers 52 and 54 , as illustrated in fig3 a , one for the transmission of the light signal to the test area and the other for returning the signal to the detector . the probe can be maneuvered to the desired area ( usually a cavity ) and measurements made of light transmitted from one of the fibers across a thickness of blood and received by the other optical fiber . there is illustrated in fig3 a an enlarged cross - sectional view of a small diameter probe ( needle ) 50 carrying two ( coextensive ) optical fibers which terminate near its sharp point 56 . since the outside diameter of the optical fibers , including cladding , is only around 100 μm , it will be appreciated that the overall probe diameter which enters the human body may be relatively small . the optical fibers are spaced apart a prescribed distance at the ends to allow blood to flow between them , and their ends are chopped , or otherwise terminated , at near 45 ° so that a beam of infrared radiation from one fiber is transmitted laterally across the blood sample therebetween to the other optical fiber and returned to the detection system . in the alternative ( not illustrated in the drawings ), the probe , catheter , or needle may carry only a single optical fiber through which the beam is both sent and returned . this single optical fiber may terminate at a fixed distance from a mirror or reflective surface carried by the probe or needle body . blood is allowed to enter through an opening at the end of the probe or needle body , as with the fig3 a embodiments , and an infrared beam emitted from the end of the fiber passes through a blood dimension and is reflected by the mirror or backscattered by the blood back into the same fiber and separated from the sent beam at the detection system . since the beam is reflected , it traverses the blood dimension twice , and this must be taken into account when measuring its relative absorption . ( i . e ., l becomes 2 l in equation 4 ). in the arrangements , whether a probe , catheter or needle employing a single or double optical fiber , optical absorption data is transmitted by the return fiber or fibers to a detector or detector system as illustrated by the block diagrams of fig3 and 4 . the significance of the catheter arrangement and the technique involved is that a small probe formed of inert material such as glass or steel may be substantially all that enters the patients body . also , relatively flexible infrared optical fibers made of krs - 5 have been recently demonstrated and would be ideal for this application . if simultaneous monitoring of a large number of patients is desired , then the arrangement depicted in fig4 could be used . for example , the ir laser and the detection system could be situated wherever convenient , either in a patient &# 39 ; s hospital room or at some other central control location for monitoring plural units through multiplexing . such a system has the advantage of electrical isolation of the individual patients . one of the advantages of the invention is that blood - co ( or any other blood gas , for that matter ) concentrations can be read instantaneously by simply comparing the absorption of the sample at a particular wavelength ( e . g ., 5 . 13 μm ) to that for a reference sample , or comparing absorption of two different wavelengths . thus , continuous monitoring of co concentration during oxygen therapy while hospitalized or on the scene of the accident , for example , may be realized . when arranged in a monitor , less expense is involved than with individual units , and it provides a convenient means for centrally monitoring each patient without the necessity for electrical connections . also , the use of optical fibers would minimize interference to the system from , for example , a high voltage x - ray source in the hospital . at the patient &# 39 ; s end of the system , the optical terminal is simply a low cost disposable needle catheter or an external clip - on unit . while the invention has been particularly shown and described with reference to specific embodiments , it will be understood by those skilled in the art that various changes in form and in details may be made therein to without departing from the spirit and scope of the invention . the invention is meant to be limited only by the scope of the claims annexed hereto .
a simple compact optical device and method for quickly measuring the concentration of co and co 2 , bound to hemoglobin or dissolved in a person &# 39 ; s blood using optical techniques which do not require removing a blood sample from the body . it also provides a simple fiber optic device for measuring blood - gas concentrations of critical internal points of the circulatory system such as the aorta .