optical performance monitoring a necessity in transparent optical n/w

7/31/2019 optical performance monitoring a necessity in transparent optical n/w

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Model Engineering College Micro-optical Medical Monitoring

1

Department of Electronics Engineering

Chapter 1

INTRODUCTION

Biophotonics is a dynamic discipline found at the interfaces between

optics, photonics, biology and medicine. Considerable research and development work in

the areas of advanced microscopy, biological imaging, optical diagnostics, laser surgery

and optical sensing has been undertaken and many of the associated technologies are

becoming industrial standards in the biological and medical areas.

Developments in micro-optics and microsystems have supported the

strong trend to miniaturization of biophotonic systems. Whereas microscopes, external

imagers, lasers for surgery and much diagnostic equipment is decidedly macroscopic,

advances in miniaturization have led to new types of endoscopic systems, implantable

monitors and highly-parallel optical diagnostic systems.A relatively new but very promising area is that of tunable micro-

optics. A controlled change in the optical properties of lenses, mirrors or filters to allow a

tuning of focal length, position, lens curvature or transmission wavelength will provide a

wide range of functionality for optical microsystems and enable a broad spectrum of new

applications. In macroscopic optical systems, tunability is almost exclusively realized by

mechanical motion of lenses or other components with respect to each other. Micro-

optical systems, however, benefit from physical and chemical effects usually not

available at the macro-scale to achieve the same results. Surface tension, membrane

distention or polymer swelling are examples of mechanisms which are employed to

change the optical properties of micro-optical devices and these give result in new forms

of tunability. The medical field stands to benefit strongly from these developments.


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Chapter 2

MICRO-OPTICS

Advances in microsystems engineering and semiconductor processing

have led to the miniaturization of optical components and systems; the field of micro-

optics has thus grown rapidly to include a wide variety of optical devices.

Optical microsystems can take advantage of a wide variety of

miniaturized components. Micro-lenses, both refractive and diffractive, exist in many

variations, as do a broad spectrum of diffractive optical elements ranging from gratings to

complex holograms. These may be fabricated as single components or, more often, in

arrays, as shown in fig 2.1. Micro-mirrors, filters, and waveguides are widely available,

and the resultant systems rely on numerous types of active micro-optical components,

such as laser diodes, LEDs, modulators and detectors, for generating and detectingoptical signals.

Fig 2.1: An array of micro-lenses [1]


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Chapter 3

TUNABLE MICRO-OPTICS

A relatively new but very promising area is that of tunable micro-

optics. Tunability is one of the most desirable feature in an optical component or system.

Controlled variation of focal length, aperture, transmission wavelength, or propagation

direction provides significant design flexibility and broadens the spectrum of possible

applications significantly. Whereas in macroscopic optical systems tunability is

frequently achieved through mechanical movement of components, this approach

becomes less and less suitable when component sizes shrink. Advances in microsystems

engineering, however, have led to the development of a host of new effects, many

without a macroscopic counterpart, which may be used to tune the characteristics of

micro-optical components.

Some of the technologies and devices being employed for the

fabrication of tunable micro-optics and photonics include tunable liquid lenses and lens

arrays, polymer membrane-based micro-lenses (tunable microfluidic microlenses),

tunable micromirrors, tunable optical filters, Bragg gratings etc. Liquid lenses have seen

extensive development and rely on controlled surface tension for tunability. Using an

applied electric field, the contact angle of a liquid droplet, whose hemispherical shape

forms an excellent lens, may be tuned over a wide range using electrowetting. Through

the use of extended electrode arrays, the liquid lenses may not only be tuned in focal

length but also repositioned two-dimensionally. Alternative tunable micro-lens

approaches use liquid filled micro-cavities bounded by highly-distensible polymer

membranes. By applying pressure to the liquid in the micro-cavity, the curvature of the

membrane may be changed, resulting in a variation of the focal length. Research is

active on the development of new types of tunable micro-optics, including tunable

variable apertures and photonic crystals.


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Chapter 4

TUNABLE MICRO-OPTIC COMPONENTS

4.1 Liquid Lenses

One representative technology for tunable micro-optics is that of

liquid microlenses. For sizes smaller than about a millimeter, liquid droplets on a

transparent surface form excellent lenses, to which anyone who has carefully examined a

window after a rainstorm will attest. Surface tension leads to spherical planoconvex

lenses whose curvature is defined uniquely by surface tension and volume; for

sufficiently small sizes, gravitational effects are negligible. This effect has been known

for two centuries and forms the basis for the well-established technology of microlens

fabrication.

Less well known is that such liquid droplets may also be electrically

manipulated. Using electrowetting, the contact angle of a liquid droplet may be changed

by application of a voltage between it and the substrate. First experiments on

electrocapillarity, i.e., the influence o f an electric potential on the surface tension of

liquids, were conducted by the Nobel laureate Gabriel Lippmann in 1875. A closely

related effect termed electrowetting was reported in 1993 by Berge. In this, a thin

dielectric layer is artificially introduced between a conducting liquid droplet and a

conductive substrate and observed a decrease of the contact angle of the liquid upon

application of a voltage. This advance, now known as electrowetting-on-dielectrics

(EWOD), allowed the use of electrowetting with nearly any liquid, provided it is

conductive or, at least, polar.

The physical reason for using liquids as optical elements results

from capillary and interface thermodynamics. For small liquid objects typically

occurring in microsystems engineering gravity effects will be neglected. Small liquid


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droplet placed on a planar substrate forms a spherical cap due to the influence of surface

tension as in fig 4.1.

Fig 4.1: Liquid droplet placed on a substrate[2]

Here the system consists of three phases: a substrate S, a liquid object

L, and the ambient V, which may be a gas or a second liquid. Its contact angle () is

uniquely given by the interfacial energies between the substrate and the liquid; the liquid

and the ambient; and the substrate and the ambient by (4.1)

cos =

( 4.1)

The focal length fof a liquid microlens is given as a function of the

diameterD of the droplet on the surface, contact angle, and the refractive indices n1,

n2 of the liquid and the ambient ( Fig 4. 2)

Fig 4.2 Focal length of a liquid lens [2]


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Focal length f is given by (4.2)

=

(4.2)

Replacingin (4.2) with the value from (4.3), the focal length of a liquid lens is seen to

be well defined by several parameters: the refractive indices of the liquid and the

ambient, the volume of the liquid, and the involved surface energies. By controlling these

parameters, the focal length of microlenses may be tuned. In electrowetting-based

microoptics, the influence of electric fields on the interfacial energy balance is used

deliberately to control the focal length precisely and reversibly.

Electrowetting

A typical EWOD setup is shown in fig 4.3.A droplet is placed on

an electrode covered by a dielectric layer (thickness d and dielectric constant ).

Application of the voltage between the electrode and the droplet leads to an adsorption of

charges at the solidliquid interface, thereby effectively lowering the corresponding

interfacial energy. Consequently, the contact angle is reduced by the voltage.

Fig 4.3: Schematic view of a typical electrowetting experiment[2]


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the effective interfacial energy of the corresponding interface is reduced with respect to

the unbiased interface by an electrostatic term given by (4.3)

=

(4.3)

Here is the electrostatic energy stored in the capacitance consisting of the conductive

droplet plus the substrate with the dielectric in between.Using a parallel plate capacitor

model for the calculation ofES, one obtains

=

(4.4)

Replacing by

from (4.3) in (4.1) and considering (4.4), we finally obtain

cos = +

(4.5)

which is the fundamental equation for EWOD. From (4.5), it can be seen that the

resulting contact angle is reduced upon application of a bias voltage to the system shown

in fig 4.3. This reduction of the contact angle is a reversible effect; as far as the free

motion of the contact line on the substrate is guaranteed, the droplet readily contracts

again if the voltage is turned off.Equation (4.5), together with (4.2), is the basis for focal

length tuning of liquid lenses by means of EWOD.

Due to the reduction of the contact angle upon application of a

voltage, movement of liquid droplets may also be achieved by EWOD. For achieving

movement, the electrode is separated into multiple individually addressable electrodes.

Upon biasing one of these electrodes, the contact angle is reduced only in those parts of

the droplets contact line located above the activated electrode as in fig 4.4. The other

parts of the contact line remain unaffected. Since a lower contact angle on a substrate

implies an energetically favorable situation for a liquid, the droplet tends to wet those

parts of the substrate located over the activated electrode, i.e., it experiences a net force

directed toward the activated electrode.


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Fig 4.4: Schematic view of droplet movement by EWOD [2]

This mechanism allows fluidic operations such as splitting and

merging of droplets by means of EWOD. For example, a droplet may be split by first

spreading it over three or more electrodes and subsequently switching off the bias voltage

from the middle electrode, leading to the formation of a liquid neck and subsequent

rupture of the droplet over the middle electrode, as shown in fig 4.5.

Fig 4.5: Illustration of the process steps for splitting a liquid droplet by EWOD.[2]


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Similarly, merging of two droplets may be achieved by simply moving one droplet

toward another one. Thus, microlens arrays capable of changing the volume or/and the

number of lenses can also be developed.

4.2 Tunable Microfluidic Microlenses

Microlenses with intrinsic tunability are of interest for a wide variety

of applications, particularly beam shaping, optical interconnection, and imaging. The last-

named application is of particular relevance for endoscopy, in which a variable imaging

system provides an advantage for tissue imaging; no dynamic micro-optical systems

suitable for use in extremely small fiber endoscopes, with diameters of as little as 0.5 mm

at the distal end, have been demonstrated. A limited number of concepts for tunable

microlenses have been demonstrated by diverse approaches and technologies, such as by

electro-wetting, through the use of liquid-crystal microlenses etc. The limitations on

liquid lenses tuned by the electrowetting approach include requirement for high driving

voltages and the requirement that only polar liquids be used. The liquid-crystal technique

is hampered by the possibility of nonuniformities in the electric field, which could lead to

optical aberrations. Moreover, both methods require the presence of electrodes, which

can disturb the optical performance. In order to overcome a number of these limitation,

microfluidic microlens is used.

Fig 4.6 Cross-sectional diagram of a membrane-based microfluidic

microlens[4]


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As shown in fig 4.6, the membrane based, liquid-filled tunable

microlenses employ a 50 m thick, pressure-actuated PDMS (poly dimethyl siloxane)

membrane structured upon a silicon fluidic chip. This chip is mounted onto a Pyrex

substrate, thereby forming a fluidic cavity filled with a liquid of suitable refractive index.

The body of the lens is then formed by the liquid, whereby the membrane curvature

provides the lens shape. Varying the pressure of the liquid by using the fluidic system

changes the shape of the membrane and thus the optical power of the lens. As shown in

the fig 4.6, both planoconvex and planoconcave lenses may be obtained. A

microphotograph of two distended, 400m-diameter convex lenses is shown in fig 4.7

Fig 4.7: Microphotograph of two distended planoconvex lenses

achieved by application of positive fluidic pressure[4]

Design and Fabrication

The membrane-based microlens is fabricated on a silicon substrate

bonded to a Pyrex substrate; the membrane itself is a PDMS layer on the front side of the

silicon wafer. The use of these materials means that no cavity deformation takes place

when pressure is applied and no critical alignment or adhesion problems occur. The use

of standard micro-electromechanical systems processing techniques, under clean-room

conditions to ensure particle-free optical surfaces, implies that low-cost mass fabrication

of these lenses is possible.


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Microlens fabrication is thus based on silicon process technology

combined with a number of non silicon materials to yield the complete structure; a

summary of the process is shown schematically in Fig. 4.8

Fig 4.8: Process summary[4]

(a) silicon wafer with backside SiO2 layer, opened by RIE

(b) structured photoresist

(c) first ICP RIEetch

(d) front-side spin-coating of primer and PDMS

(e) second ICP RIE etch

(f) oxide removal

(g) bonding of the pyrex glass wafer and patterning of the backside chromium layer

(h) filling of the devices with the optical liquid.

Processing begins with photolithography on a silicon wafer

polished on both sides in which the lens openings and the microfluidic channels are


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defined. In wet oxidation and subsequent plasma-enhanced chemical-vapor deposition

oxidation steps, a masking layer is deposited. This layer is structured by photolithography

and subsequently opened by reactive ion etching (RIE). The first oxide mask defines the

contours of the lens chamber, the fluidic channels, and the reservoir with a typical

structure as in fig 4.9 . Circular lens openings with diameters of300600m were defined

along with the fluidic channels and the reservoirs. In a second photolithography step, a

12m-thick photoresist layer is patterned. This second mask step defines the contours of

the lens chambers and the reservoirs.

Defined by the second photoresist mask, the lens chambers and the

reservoirs are then etched through half of the silicon wafer by an inductively-coupled

plasma (ICP) RIE process. Afterward, the resist layer is removed in a plasma asher. This

step is followed by front-side coating of a thin PDMS film. To improve the adhesion of

the PDMS to the silicon or silicon oxide, the substrate is coated with a primer ( 3 m).

The thickness of the PDMS film is 50 m. This film is cured for 15 min at 150 C. This

PDMS or silicone elastomer film forms the pressure-actuated membrane, which encloses

the working liquid and ultimately defines the lenss curvature.

Fig 4.9: Layout for a lens array with 400m lenses[4]

As the lenses are arranged as arrays with different dimensions and

contours, one fluidic reservoir will provide nine lenses with the required working liquid

and pressure; in the future each lens will be controlled by an independent microfluidic

system. In a second ICP RIE etch step the lens chambers, the reservoirs, and the fluidic


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channels are etched simultaneously, passivated by the structured SiO2 layer. The etching

stops abruptly when the SiO2 layer at the front side is reached because the etch

selectivity to SiO2 is approximately 200 to 1.

Finally, the front-side SiO2 layer is removed in a last RIE step,

resulting in the PDMS membranes now being freely stretched over the silicon lens

chamber. To enclose the liquid cavity, a thin Pyrex wafer is subsequently bonded to the

back side of the silicon substrate to cap the fluidic channels and chambers. After

alignment, a UV flood exposure followed by an anneal at a temperature of50 Cfor 24 h

results in covalent bonds between the Pyrex and the silicon substrates, and this ensures

that the structure remains mechanically stable.

Because the Pyrex wafer is fully transparent, not only the lens

chambers but also the fluidic channels and the reservoirs are visible from the back of the

bonded structure. To minimize scattering and cross talk between the single lenses of the

array, an aperture stop on the unpatterned Pyrex wafer is required. Such an aperture is

made by use of a 100nm-thick vapor-deposited chrome layer, subsequently structured by

a wet etch.

The wafer stack is then separated into individual chips, each with

an array of nine lenses and one reservoir. During this last step, the delicate PDMS

membrane must be protected by a photoresist layer to prevent scratching during chip

separation.

To prepare the fluidic lenses for operation, one has to introduce

the liquid optical medium into the microfluidic chambers; one accomplishes this filling

by taking advantage of capillary forces that essentially result in a self-filling of the

cavities. Two liquid media were employed: a mixture of 50wt. % water and 50wt. %

ethanol and a proprietary liquid from Cargille with a high refractive index (n = 1.600). In

the former case the addition of ethanol is important to reduce surface tension and thus

provide better filling, caused by capillary forces. The H2O:ethanol mixture working

liquid provides a refractive index of n =1.3534. Cavity filling is accomplished in a

vacuum chamber to deaerate the liquid enclosed inside the small fluidic channels. Any


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remaining bubbles quickly diffuse out of the system because the PDMS membrane is gas

permeable.

The devices are operated by use of a silicone tube glued to the fluidic

reservoir entrance with UV curable adhesive. A manually operated syringe and a

calibrated pressure sensor were then used to define the pressure exerted on the liquid,

thereby leading to the desired lens distension.

Characterization of the microlenses

The pressure of the liquid optical medium was measured with a pressure

sensor with a resolution of 90 mV/kPa. Determination of the maximum lens height

allowed the radius of curvature to be calculated through

=

+

(4.6)

whereR is the radius of the lens, dis the lens diameter, and h is the maximum height. The

focal length could then be determined simply by

=

(4.7)

wherefis the focal length and n1 is the refractive index of the liquid optical medium

The pressure range in the measurement series varied from 0 to 54 kPa,

giving the results shown in fig 4.10. For the highest pressures, the curvature radius

saturates at 0.2 mm, with controllable values ranging from 0.2 to 6 mm

Fig 4.10: Lens radius of curvature as a function of pressure applied

to the liquid optical medium (ethanol and water)[4]


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As can be seen from fig 4.11, the lens profile was measured as a

function of applied pressure in the range 454 kPa; the corresponding radii of curvature

are given in fig 4.10. The lens profile is not ideally spherical for two reasons: the

nonlinear deflection of the PDMS membrane and the fixed edge of the membrane, which

leads to an inflection point at the lens edge.

Fig 4.11: Complete lens profile for a 500m-diameter lens as a function of applied

pressure in the range 454 kPa.[4]

The focal length is shown in Fig 4.12 as a function of pressure. Themeasurements were performed with the two liquid optical media available, with n1=1.35

and n2=1.6. It can be seen that the focal length varies from 1 to 18 mm, saturating at

approximately 1 mm for high pressures.

Fig 4.12: Focal length of the membrane lenses as a function of

applied pressure[4]


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Further design considerations for this lens are the means for

application of pressure, pressure variations, repeatability, and aging. The application of

constant pressure over an extended period of time can be achieved only by use of a

pressure controller with a precise microvalve. Maintaining constant pressure is important

to prevent variations in the focal length during the imaging period. Extremely high

pressures may overstretch the PDMS membrane, causing plastic deformation and thus

leading to non-reproducibility effects. The PDMS membrane is stable and flexible

temperature range from 50 C to 200 C, but the stability of the refractive-index liquid

must also be considered. Volatile liquid combinations, such as H2O: ethanol mixtures,are therefore not suitable for long-term use in such a lens system.

4.3 Deformable Micro-mirrors

Deformable micro-mirrors may be used in a variety of optical

applications, including laser beam steering, beam shaping, and aberration correction. In

particular, beam shaping is often required for aberration correction and is used to improve

imaging quality, predominantly in ophthalmology and astronomy. Wavefront correction

is accomplished by applying the inverse of the undesired phase deformation using a

deformable mirror, thereby improving the optical quality of the image.

Commercially available deformable membrane mirror devices are

usually made of silicon. The disadvantage of this material is its high stiffness: for

electrostatic actuation, a thick silicon membrane requires a high driving voltage. We may

overcome this restriction by using alternative materials for the fabrication of deformable

mirrors. The selected material should feature a lower stiffness but remain robust. The

photo-structurable polymer SU-8 has proven to be suitable as a structural material for this

application: it has excellent mechanical properties, which are coupled with ease of

fabrication. For electrostatically actuated mirrors, two electrodes forming a capacitor are

required. In the case of a membrane mirror, the first electrode is fixed on a bottom

substrate and the second electrode is the membrane itself. If the deformable mirror is

designed to be used in an adaptive optical system, the membrane has to be deformed into

any desired shape. Therefore, numerous electrodes with varying applied voltages on


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either side of the capacitor are required so that a free-standing membrane is mounted over

an array of electrodes, as seen in Fig4.13. The membrane itself is clamped inside a silicon

Fig 4.13:Schematic sketch of the free-standing deformable membrane mirror

[3]

frame to increase the mechanical stability of the device. Between the frame and the

underlying substrate is a spacer layer: this layer defines the maximum deflection of the

membrane and the required voltage. With this setup, different positive voltages can be

applied to the individual electrodes. The varying voltages lead to corresponding

electrostatic forces and these forces are used to shape to the membrane.

4.4 Tiltable Micromirrors

Actuated optical micromirrors are essential components in high-

density optical interconnects, high-accuracy scanning systems and in adaptive optical

systems for applications ranging from astronomy to medicine. Increasing complexity and

density of such systems place increasingly large demands on positional accuracy in

angular mirror attitude, and actuation employing pre-calibrated lookup tables and similar

mechanisms is often insufficiently precise. Future generations of micromirror arrays will

require integrated feedback mechanisms for determining and maintaining the desired

angular position of each individual mirror.

To address these issues, a monocrystalline micromirror with a

monolithically integrated diffraction grating is introduced. Light incident on the mirror is

predominantly reflected into the target fiber but a small fraction is transmitted through

the grating. The diffracted transmitted light gives rise to two diffraction maxima whose

position is measured with a position sensitive detector (PSD) mounted below the mirror.


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Since these diffraction maxima move as the mirror is tilted, assuming that the incident

light wavelength, the grating period and the grating-to-detector distance are known, the

mirror tilt angle may then easily be calculated from the PSD intensity signal. This system

provides the basis for highly accurate mirror position control, as well as for individual in-

situ calibration of every switching mirror, even after packaging.

4.4.1 Design

As illustrated in fig 4.14, the device is composed of three layers: on

top, a silicon mirror layer; then a Pyrex contact layer using indium tin oxide (ITO)

electrodes; and, at the bottom, a position sensitive detector.

The tilt sensing principle is based upon the diffraction of laser light

at a transmission grating, etched into a portion of the mirror surface. Laser light strikes

the mirror on the front side and is reflected into the desired direction, as is the zeroth-

order maximum of light being diffracted at the grating in forward direction. As can be

seen in fig 4.14, a small part of the incident light is transmitted through the grating and

the diffracted light is emitted from the mirror backside.

Fig 4.14: Schematic of the tilt sensing mechanism[5]


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4.4.2 Diffraction orders

Diffraction occurs for both the reflected and transmitted light; whereas on

the front side both diffraction and reflection occur, on the backside, only diffraction can

be observed. The angle under which the light is diffracted in the m-th order is

determined by the tilt angle of the mirror , the wavelength and the grating period .

The angle is measured relative to the surface normal of the grating. If the mirror, and

thus the grating, is orthogonal to the incident laser beam, the grating equation predicts

that constructive interference from the grating backside will occur when the optical path

difference x1 of two light waves originating from two neighboring slits is

1 = sin = , m=0,1,2,3,.. (4.8)

This diffraction results in light intensity maxima of the order 0,1,2 and

higher. For an untilted grating, the diffraction angles 1, 2, . . . are symmetrically

distributed to the left and right of the normal to the grating surface.

As can be seen in figures 4.15(a) and (b), when the mirror plane is tilted

by the angle from its horizontal position there is an additional optical path difference

2 = sin (4.9)

between the paths 1 and 2 of the diffracted light.

(a) (b)Fig 4.15: The optical path differences of two light waves on a tilted grating[5]In (a) x1

andx2 add up whereas in (b) they are subtracted to give the total optical path

difference.


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The total optical path difference and therefore the diffraction angle on the

grating backside are different for the positive and negative maxima, depending on

whether these are diffracted in the direction or against the direction of tilt motion. This

effect occurs because the two path differencesx1and x2 contribute differently to the

total path difference x. As can be seen in figure 4.15(a), for the case of diffraction in the

direction of tilt motion, the total path difference between path 1 and path 2 is

x=x1 +x2 (4.10)

whereas in figure 4.15(b) , in the case of diffraction in the direction opposite to the tilt

motion, the total path difference x is given by

x=x1 - x2 (4.11)

We can thus relate the diffraction angle , referenced to the surface normal

of the tilted grating, to the tilt angle of the grating as

= 1 2 = = (4.12)

As can also be seen in figure 2, to obtain the diffraction angle of maximal

constructive interference with reference to the surface normal of the untilted grating, we

have to add (fig 4.15(a)) or subtract (fig 4.15(b)) the tilt angle from the angle , leading

to

= = arcsin

sin (4.13)

Finally, the position of the interference maxima on the PSD, xPSD, placed

at a distance dfrom the mirror, for a mirror tilt angle , is thus given by

PSD = d tan tan0 (4.14)

Based on the above equations, the diffraction angle is plotted as a function

of the mirror tilt angle in fig 4.16. It can clearly be seen that the positive and negative

diffraction maxima do not shift symmetrically upon the tilting of the mirror grating and

that the diffraction angle is larger for a given order in the direction opposite to mirror tilt.


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Fig 4.16: Relationship between the diffraction angle and the

mirror tilt angleof the 1st and 2nd diffraction orders[5]

It isthus reasonable to use those diffraction orders (as depictedin figs 4.15(a) and (b))

for tilt sensing, since a larger shift allowshigher measurement sensitivity. Moreover, this

asymmetry also allows the determination of the tilt direction. To further increase the

angular diffraction shift, and with it the tilt sensing accuracy, one can also use the

second-order diffractionmaxima.

Fig 4.17: Silicon-based two-dimensionally tiltable micro-mirrors designed for use in

endoscopic imaging systems. [1]


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Chapter 5

APPLICATION OF MICRO-OPTICS IN MEDICAL

DIAGNOSTICS

5.1 Implantable Blood-Oxygen Monitors

There is considerable interest in continuous in vivo monitoring of

certain parameters in the cardiovascular system, particularly blood oxygenation, pulse,

blood pressure and concentration of various dissolved gasses. As seen in the absorption

spectrum of oxygenated and deoxygenated hemoglobin in fig 5.1, the oxygenation of

blood can be determined by a differential absorption spectroscopy measurement at two or

more wavelengths.

Fig 5.1: Absorption of spectra of oxygenated and deoxygenated hemoglobin. [1]

Pulse oximetry is established as one of the standard monitoring

techniques allowing the monitoring of the oxygenation of a patients blood and is

nowadays ubiquitous in hospital critical care and surgery, since it provides an early

indication of problems in the oxygen transport to the tissue.

Common pulse oximeter designs use time-multiplexed red and infrared

LEDs and a photodetector which detects the light attenuated by the perfused tissue . The


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probe is usually attached to the patients fingertip, earlobe, forehead, or foot, i.e., to a

region of tissue which is well perfused and translucent. The embodiment of the probe can

be of transmissive or reflective type, which means that either the intensity of the

transmitted light through the tissue or that of the backscattered light from the tissue is

measured. The main advantage of those systems is their noninvasive character which

makes the method quick, safe, and easy to apply.

5.1.1 Optical Transmission Oximetry

For transmission pulse oximetry, the AC signal is generally low: it

ranges from 1% to 10% of the total transmitted light. Also due to the extracorporeal

sensor position, the measurement period is limited: the time required is usually between a

couple of seconds for short-term diagnostics up to several days during inpatient health

care.

A new type of implantable pulse oximeter offers some advantages

over noninvasive standard pulse oximeters, especially for high-risk cardiovascular

patients in intensive care or for surveillance of people in critical condition. Once

implanted during a surgical procedure which had to be performed anyway, the sensor

implant offers improved patient comfort and signal quality. The key benefits of an

implantable oxygen sensor are the capability for longterm monitoring and assessment,

emergency detection, and, if necessary, for efficient treatment or medication. Due to its

location directly on an arterial vessel, an enhanced AC signal is expected, one less

influenced by the limitations associated with standard pulse oximetry such as ambient

light, low perfusion state, skin pigmentation, nail polish, intravenous dyes, etc.

Schematic of an implantable micro-optical blood oxygen monitor is

shown in fig 5.2. The design uses a sensor stripe fitted with optoelectronic components

(photodetectors and LEDs) which is wrapped directly around an artery. Transmissionintensities through the blood are compared and allow a determination of oxygenation

continuously and in real time. Several design constraints are given due to the positioning

of the implant. As a first practical aspect the sensor has to be easily implantable by the

surgeon, preferably with standard methods; the surgery must not be made more

complicated due to implantation of the sensor. To fulfill this requirement, the sensor


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stripe is fixed by a titanium ligature clip after wrapping around the vessel.

Fig 5.2: Model of the implantable pulse oximeter sensor. [6]

The stripe material itself has to be

biocompatible, optically transparent, and at least as elastic as the vessel in order to

minimize vessel restriction. Due to the strong attenuation of light in blood, the maximum

diameter of the arterial vessel and thus the sensor rings inner diameter is limited to about

5 mm. As is common for sensors relying on optical transmission, the LEDs and

photodetector face each other when placed around the vessel. The optical paths from the

LEDs to the photodetector have to be equal and not influenced by slightly differing vessel

diameters; a side by side arrangement of the two LEDs meets this condition. The

fabricated system is shown in fig 5.3

Fig 5.3: Photograph of the implantable blood oxygenation sensor. [6]


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Main criteria for the selection of the LEDs and the photodetector

are their size, spectral characteristics, and power. A red LED and an infrared LED, with

peak wavelengths 659 and 942 nm, were chosen as light sources.This approach provides

an implantable solution as close as possible to the quantity to be measured but without

direct blood contact. The AC signal of the arterial pulse oximeter sensor ring ranges from

20% to 40%, which is an excellent value when compared to standard finger tip pulse

oximeters.

5.1.2 Optical Reflection Oximetry

A further micro-optical approach providing great flexibility in

monitoring physiological parameters is by the use of reflection from perfused tissue, as

shown schematically in Fig 5.4.

Fig 5.4: The principle behind reflectance oximetry.[7]

For measurements of arterial oxygen saturation and the concentration of

other hemoglobin derivatives, the sensor utilizes the method of reflectance pulse

oximetry. Light of multiple wavelengths is emitted into tissue which is well supplied with

blood, and the backscattered light is detected . With the different absorption spectra

shown in Fig 5.5, the different kinds of hemoglobin, such as oxyhemoglobin (O2Hb),

deoxygenated hemoglobin (HHb), carboxyhemoglobin (COHb) and methemoglobin

(MetHb), cause a specific attenuation of light, dependent on their concentration.


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Fig 5.5: Absorption spectra of different hemoglobin derivatives.[7]

The sensor consists of a flexible polymer substrate with LED-chips

and a photodiode as well as an amplification circuit placed close to the photodetector toprovide a good signal-to-noise ratio. Direct illumination of the photodetector by stray

light from the LEDs is prevented by an interstitial optical barrier (Fig. 5.6). The sensor is

housed in biocompatible silicone and provides mounting holes to suture the sensor onto

the underlying tissue.

Fig 5.6: Sketch and photograph of the reflectance oximeter, showing the 8 LEDs and

detector.[7]


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5.2 Endoscopic OCT

Fig5.7: Schematic cross-section of the OCT measurement head.[8]

In addition to implantable optical systems, micro-optics is enabling the

development of advanced endoscopic diagnostics, using miniaturized optical

components, thereby also supporting the movement toward minimally-invasive surgery.

Optical coherence tomography (OCT) is a non-destructive imaging technique

conceptually similar to white-light interferometry. Using a scanning mirror as part of a

Michelson interferometer, an OCT system can image up to several millimeters below the

surface of tissue with a resolution below 10 m. Tunable micro-optics has enabled the

conception of an OCT system which may be integrated into an endoscope.

Fig 5.8: OCT scan through a biological sample. [8]


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The OCT measurement head, which is found at the endoscope tip,

is shown schematically in fig 5.7. A two-lens system is used to dynamically focus the

beam and a two-axis mirror (tiltable mirror) serves to scan the beam over a given surface

area of the tissue. OCT imaging resolution can be significantly improved if the beam

focus follows with depth scan. All of these components have been integrated into a single

measurement head, with a diameter less than 3 mm, and thus compatible with endoscopic

dimensions. The assembled system, based on silicon micro-machined components, an

aluminum sub-mount, a pneumatic tunable micro-lens and the two-axis scanning mirror,

is shown in Fig 5.9.

Fig 5.9: Complete OCT measurement head on an aluminum sub-mount.[8]

Bo

A complete scan, shown in Fig 5.8, demonstrates the benefit of the

tunable lens, leading to dynamic focussing. For the scans from left to right, the focus

moves deeper into the material, revealing more detailed features. As a result, the

incorporation of the tunable photonic components results in depth-independent resolution

for this endoscopic system.


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Chapter 6

FUTURE DEVELOPMENTS

Since micro-optics is a newly developing area, new technologies

may be expected in various micro-optic components which will extend their use in a wide

variety of applications. Future developments in liquid lenses will extend the concept of

electrowetting to allow for the independent movement of multiple lenses on a single

substrate, as well as to provide the means for generating a variable number of lenses

based on the immediate demands of a particular application. The performance of tiltable

micromirrors can be improved if it uses two tilt axes. Future developments of implantable

pulse oximeter will include the development of implant controller electronics,

rechargeable power supply, processing and communication capabilities all associated

with low power instrumentation and system miniaturization. Developments in

microlenses and mirrors will also improve the performance of OCT system, which will

provide good quality images.


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Chapter 6

CONCLUSION

Developments in micro-optics and microsystems have supported the

strong trend to miniaturization of biophotonic systems. These developments include a

number of tunable micro-optical components which employ physical phenomena on the

micro-scale for achieving optical variability. Application areas for these devices arewidespread: tunable micro-lenses and micromirrors are of considerable value in compact

optical systems with enhanced functionality and are being developed, for example, for

use in endoscopic optical coherence tomography systems. Numerous other technologies

have become relevant for tunable micro-optics in, particularly silicon-based opto-thermal

filters and polymer structures with controlled swelling, and further developments in

these areas will also be expected.


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optical performance monitoring a necessity in transparent optical n/w

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