Surgery – Diagnostic testing – Measuring or detecting nonradioactive constituent of body...
Reexamination Certificate
2001-08-09
2003-10-28
Winakur, Eric F. (Department: 3736)
Surgery
Diagnostic testing
Measuring or detecting nonradioactive constituent of body...
C356S367000
Reexamination Certificate
active
06640116
ABSTRACT:
BACKGROUND OF THE INVENTION
Various optical spectroscopic measurement systems have been developed for the noninvasive monitoring of blood constituent concentrations. In such systems, light of multiple wavelengths is used to illuminate a thin tissue portion of a person, such as a fingertip or earlobe, to obtain a spectrum analysis of the light absorbed by blood flowing through the tissue site. Pulse oximetry systems, which perform such measurements to monitor blood oxygenation of hemoglobin constituents, have been particularly successful in becoming the standard of care. Extending this technology to the noninvasive monitoring of other blood constituents, such as blood glucose, is highly desirable. For example, current methods for accurately measuring blood glucose involve drawing blood from the subject, which can be onerous for diabetics who must take frequent samples to closely monitor blood glucose levels.
FIG. 1
illustrates an optical spectroscopic measurement system
100
. A multiple wavelength light source
110
produces incident light
112
of intensity I
0
and wavelength &lgr;, I
0,&lgr;
, which illuminates a sample
120
having multiple constituents, each of concentration c
i
. The incident light
112
is partially absorbed by the sample
120
, and transmitted light
130
of intensity I emerges from the sample
120
. A detector
140
provides an output signal
142
that is proportional to the transmitted light
130
. A signal processor
150
operates on the detector output signal
142
to provide a measurement
152
that is indicative of one or more of the constituent concentrations c
i
in the sample
120
, based upon the known extinction coefficients &egr;
i,&lgr;
of the sample constituents.
SUMMARY OF THE INVENTION
The attenuation of light through a homogenous, non-scattering medium of thickness d having n dissolved, absorbing constituents is described by the Beer-Lambert Law
I
=
I
0
,
λ
ⅇ
-
∑
i
=
1
n
ϵ
i
,
λ
·
c
i
·
d
(
1
)
Dividing both sides by I
0,&lgr;
and taking the logarithm yields
ln(
I/I
0
)=−&mgr;
a
·d
(2a)
μ
a
=
∑
i
=
1
n
ϵ
i
,
λ
·
c
i
(2b)
where &mgr;
a
is the bulk absorption coefficient and represents the probability of absorption per unit length. Measurements are taken at n wavelengths to yield n equations in n unknowns
[
ln
(
I
λ
1
I
0
,
λ
1
)
ln
(
I
λ
2
I
0
,
λ
2
)
⋮
ln
(
I
λ
n
I
0
,
λ
n
)
]
=
-
[
ϵ
1
,
λ
1
ϵ
2
,
λ
1
⋯
ϵ
n
,
λ
1
ϵ
1
,
λ
2
ϵ
2
,
λ
2
⋯
ϵ
n
,
λ
2
⋮
⋮
⋰
⋮
ϵ
1
,
λ
n
ϵ
2
,
λ
n
⋯
ϵ
n
,
λ
n
]
[
c
1
c
2
⋮
c
n
]
d
(
3
)
which can be written in matrix notation as
I=−A
(&lgr;)
Cd
(4)
Solving for the constituent concentrations yields
C
=
-
1
d
A
(
λ
)
-
1
I
(
5
)
If the medium is a tissue portion of a person, such as a fingertip, it includes a number of constituents that absorb light. Some of the principal absorbing constituents in tissue include water, oxyhemoglobin, reduced hemoglobin, lipids, melanin and bilirubin. A drawback to applying the Beer-Lambert Law to determine the concentrations of absorbing constituents, however, is that tissue is a turbid media, i.e. strongly scatters light, which violates an underlying assumption of equation (1). Scattering in tissue is due, in part, to the variations in refractive index at the boundaries of cells or other enclosed particles, such as collagen fibers, mitochondria, ribosomes, fat globules, glycogen and secretory globules.
FIG. 2
illustrates a particular photon path
200
as it travels through a turbid medium
202
. The photon path
200
is shown as a series of connected vectors {right arrow over (p)}l
i
each representing the direction and pathlength of a particular photon between collisions. The total pathlength traveled by the photon is
pl
=
∑
i
=
1
n
pl
ix
2
+
pl
iy
2
+
pl
iz
2
(
6
)
As shown in
FIG. 2
, the effect of scattering is to substantially increase the photon pathlength and, hence, the probability of absorption. Thus, when a turbid media is considered, the Beer-Lambert Law is modified to include the effective pathlength, pl, which is a function of wavelength. The Beer-Lambert Law is also written in terms of transmission, T, to differentiate reflected light due to back-scattering of the incident light.
T
=
T
max
ⅇ
-
∑
i
=
1
n
ϵ
i
,
λ
·
c
i
·
pl
λ
(
7
)
where T
max
is the maximum transmitted light without absorption.
[
ln
(
T
λ
1
T
max
,
λ
1
)
ln
(
T
λ
2
T
max
,
λ
2
)
⋮
ln
(
T
λ
n
T
max
,
λ
n
)
]
=
-
[
pl
λ
1
0
0
0
0
pl
λ
2
0
0
0
0
⋰
0
0
0
0
pl
λ
n
]
[
ϵ
1
,
λ
1
ϵ
2
,
λ
1
⋯
ϵ
n
,
λ
1
ϵ
1
,
λ
2
ϵ
2
,
λ
2
⋯
ϵ
n
,
λ
2
⋮
⋮
⋰
⋮
ϵ
1
,
λ
n
ϵ
2
,
λ
n
⋯
ϵ
n
,
λ
n
]
[
c
1
c
2
⋮
c
n
]
(
8
)
T=−X
(&lgr;)
A
(&lgr;)
C
(9)
C=−A
(&lgr;)
−1
X
(&lgr;)
−1
T
(10)
FIG. 3
illustrates one method of measuring the effective pathlength through a sample. A picosecond pulse laser
310
and an ultra-fast detector
340
directly measure the photon “time of flight” through a sample
320
. A single pulse
360
with a duration on the order of a picosecond is directed through the sample
320
. The detector
340
measures the corresponding impulse response
370
. The time difference between the light entering the sample
312
and the mean time of flight, {overscore (t)}
380
, of light having traversed the sample
330
yields the mean optical pathlength, i.e. the effective pathlength
mpl=c
v
{overscore (t)}
s
(11a)
t
_
=
∫
0
∞
T
(
t
)
t
ⅆ
t
/
∫
0
∞
T
(
t
)
ⅆ
t
(11b)
where c
v
is the speed of light in a vacuum and n
s
is the refractive index of the sample.
An analytic expression for the shape of the impulse response of a narrow collimated pulsed light beam normally incident on the surface of a semi-infinite homogeneous tissue slab of thickness d, derived from the diffusion approximation to radiative transfer theory, is
T
(
d
,
t
)
=
(
4
π
Dc
)
-
1
2
t
-
3
2
e
-
μ
a
ct
f
(
t
)
(
12a
)
f
(
t
)
=
{
(
d
-
z
0
)
e
-
[
(
d
-
z
0
)
2
4
Dct
]
-
(
d
+
z
0
)
e
-
[
(
d
+
z
0
)
2
4
Dct
]
+
(
3
d
-
z
0
)
e
-
[
(
3
d
-
z
0
)
2
4
Dct
]
-
(
3
d
+
z
0
)
e
-
[
(
3
d
+
z
0
)
2
4
Dct
]
}
(
12b
)
D=
{3[&mgr;
a
+(1−
g
)&mgr;
s
]}
−1
(12c)
z
0
=[(1−
g
)&mgr;
s
]
−1
(12d)
where T(d, t) is the spatially integrated transmittance, D is the diffusion coefficient, c is the speed of light in the tissue, &mgr;
a
is the bulk absorption coefficient, &mgr;
s
is the bulk scattering coefficient and g is the anisotropy, which is the mean cosine of the scattering angle. Equations (12a)-(12d), therefore, are an approximation of the impulse response
370
shown in FIG.
3
. The derivation of equations (12a)-(12d) and a description of the model upon which that derivation is based, is given in
Time Resolved Reflectance and Transmittance for the Noninvasive Measurement of Tissue Optical Properties,
Patterson et al., Applied Optics, Vol. 28, No. 12, Jun. 15, 1989, Optical Society of America, inc
Knobbe Martens Olson & Bear LLP
Masimo Corporation
McCrosky David J.
Winakur Eric F.
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