3.1.

Light Guiding Principles in Conventional Optical Fibers

An optical fiber is a dielectric waveguide that operates at optical frequencies. This fiber waveguide is normally cylindrical in form. It confines electromagnetic energy in the form of light within its surfaces and guides the light in a direction parallel to its axis. The propagation of light along a waveguide can be described in terms of a set of guided electromagnetic waves called the modes of the waveguide. Each guided mode is a pattern of electric and magnetic field distributions that is repeated along the fiber at equal intervals. Only a certain discrete number of modes are capable of propagating along the waveguide. These modes are those electromagnetic waves that satisfy the homogeneous wave equation in the fiber and the boundary condition at the waveguide surfaces.

Figure 2 shows a schematic of a conventional optical fiber, which consists of a cylindrical silica-based glass core surrounded by a glass cladding that has a slightly different composition.^17–19 The core of diameter 2a has a refractive index $n_1$ and the cladding has a slightly lower refractive index $n_2$. Surrounding these two layers is a polymer buffer coating that protects the fiber from mechanical and environmental effects. The refractive index of pure silica varies with wavelengths ranging from 1.453 at 850 nm to 1.445 at 1550 nm. By adding certain impurities, such as germanium dioxide (${\mathrm{GeO}}_2$), to the silica during the fiber manufacturing process, the index can be slightly changed. This is done so that the refractive index $n_2$ of the cladding is slightly smaller than the index of the core (i.e., $n_2<n_1$). This condition is required so that light traveling in the core is totally internally reflected at the boundary with the cladding, which is the physical mechanism that guides light signals along a fiber.

Fig. 2

Schematic of a conventional silica fiber structure.

The variations in material and size of the conventional solid-core fiber structure dictate how a light signal is transmitted along a fiber and also influence how the fiber performance responds to environmental perturbations, such as stress, bending, and temperature variations. Variations in the material composition of the core give rise to two commonly used fiber types, as shown in Fig. 3. In the first case, the refractive index of the core is uniform throughout and undergoes an abrupt change (or step) at the cladding boundary. This is called a step-index fiber. In the second case, the core refractive index varies as a function of the radial distance from the center of the fiber. This type is a graded-index fiber.

Fig. 3

Comparison of conventional single-mode and multimode step-index and graded-index optical fibers.

Both the step- and the graded-index fibers can be further divided into single-mode and multimode classes. As the name implies, a single-mode fiber (SMF) sustains only one mode of propagation, whereas a multimode fiber (MMF) contains many hundreds of modes. A few typical sizes of SMF and MMF are given in Fig. 3 to provide an idea of the dimensional scale. MMFs offer several advantages compared with SMFs. The larger core radii of MMFs make it easier to launch optical power into the fiber and to collect light emitted or reflected from a biological sample. SMFs are more advantageous when delivering a narrow light beam to a specific tissue area and also are needed for applications that deal with coherence effects between propagating light beams.

The remainder of Sec. 3.1 describes the operational characteristics of step-index fibers, and Sec. 3.2 describes graded-index fiber structures.

3.1.1.

Ray optics concepts

To get an understanding of how light travels along a fiber, first consider the case when the core diameter is much larger than the wavelength of the light. For such a case, a simple geometric optics approach based on the concept of light rays can be used. Figure 4 shows a light ray entering the fiber core from a medium of refractive index $n$, which is less than the index $n_1$ of the core. When the ray meets the fiber end face, it is refracted into the core and propagates at an angle $\theta$, which is smaller than the entrance angle ${\theta }_0$ of that ray. Inside the core, the ray strikes the core-cladding interface at an angle $\varphi$ relative to the normal to the surface. If the light ray strikes this interface at such an angle that it is totally internally reflected, the ray becomes confined to the core region and follows a zigzag path as it travels along the fiber.

Fig. 4

Ray optics representation of the propagation mechanism in an optical fiber.

From Snell’s law, the minimum or critical angle ${\varphi }_c$ that supports total internal reflection is given by

Eq. (1)

$$\sin {\varphi }_c=\frac{n_2}{n_1}.$$

Rays striking the core-cladding interface at angles less than ${\varphi }_c$ will refract out of the core and be lost in the cladding as the dashed line shows. The condition of Eq. (1) can be related to the maximum entrance angle ${\theta }_{0,\max }$, which is called the acceptance angle ${\theta }_A$, through the relationship

Eq. (2)

$$n\sin {\theta }_{0,\max }=n\sin {\theta }_A=n_1\sin {\theta }_c={(n_1^2-n_2^2)}^{1/2},$$

where ${\theta }_c=\pi /2-{\varphi }_c$. Thus, those rays having entrance angles ${\theta }_0$ less than ${\theta }_A$ will be totally internally reflected at the core-cladding interface. Thus, ${\theta }_A$ defines an acceptance cone for an optical fiber. Rays outside of the acceptance cone, such as the ray shown by the dashed line in Fig. 4, will refract out of the core and be lost in the cladding.

The critical angle also defines a parameter called the numerical aperture (NA), which is used to describe the light acceptance or gathering capability of fibers that have a core size much larger than a wavelength.^5,17,18 This parameter defines the size of the acceptance cone shown in Fig. 4. NA is a dimensionless quantity that is less than unity, with values nominally ranging from 0.14 to 0.50. NA is given by

Eq. (3)

$$\mathrm{NA}=n\sin {\theta }_A=n_1\sin {\theta }_c={(n_1^2-n_2^2)}^{1/2}\approx n_1\sqrt{2\mathrm{\Delta }}.$$

The parameter $\mathrm{\Delta }$ is called the core-cladding index difference or simply the index difference. It is defined through the equation $n_2=n_1(1-\mathrm{\Delta })$. Typical values of $\mathrm{\Delta }$ range from 1 to 3 percent for MMF and from 0.2 to 1.0 percent for SMF. Thus, since $\mathrm{\Delta }$ is much less than 1, the approximation on the right-hand side of Eq. (3) is valid. Because NA is related to the maximum acceptance angle, it is commonly used to describe the light acceptance or gathering capability of an MMF and to calculate the source-to-fiber optical power coupling efficiencies. The NA value is listed on vendor data sheets for fibers.

3.1.2.

Modal concepts

Although the ray representation gives a general picture of light propagation along a fiber, mode theory is needed for a more detailed understanding of concepts such as mode coupling, dispersion, coherence or interference phenomena, and light propagation in single-mode and few-mode fibers. Figure 5 is a longitudinal cross-sectional view of an optical fiber that shows the field patterns of some of the lower-order transverse electric. The order of a mode is equal to the number of field zeroes across the guiding core. The plots show that the electric fields of the guided modes are not completely confined to the core but extend partially into the cladding. The fields vary harmonically in the core region of refractive index $n_1$ and decay exponentially in the cladding of refractive index $n_2$. For low-order modes, the fields are tightly concentrated near the axis of an optical fiber with little penetration into the cladding region. Higher-order mode fields are distributed more toward the edges of the core and penetrate farther into the cladding.

Fig. 5

Electric field distributions of lower-order guided modes in an optical fiber (longitudinal cross-sectional view).

As the core radius $a$ shown in Fig. 2 is made progressively smaller, all modes except the fundamental mode shown in Fig. 5 will start getting cut off. A fiber in which only the fundamental mode can propagate is an SMF. An important parameter related to the cutoff condition is the $V$ number defined by

Eq. (4)

$$V=\frac{2\pi a}{\lambda }{(n_1^2-n_2^2)}^{1/2}=\frac{2\pi a}{\lambda }\mathrm{NA}\approx \frac{2\pi a}{\lambda }{n}_1\sqrt{2\mathrm{\Delta }},$$

where the approximation on the right-hand side comes from Eq. (3). This parameter is a dimensionless number that determines how many modes a fiber can support. Except for the lowest-order fundamental mode, each mode can exist only for values of $V$ that exceed the limiting value $V=2.405$ (with each mode having a different $V$ limit). The wavelength at which all higher-order modes are cut off is called the cutoff wavelength ${\lambda }_c$. For example, if $a=8.0\mu \mathrm{m}$ and $\mathrm{\Delta }=0.01$, then from Eq. (4), with $V=2.40$, the cutoff wavelength is ${\lambda }_c=1481\mathrm{nm}$. That is, only the fundamental mode will propagate in the fiber for wavelengths $>1481\mathrm{nm}$. The fundamental mode has no cutoff and ceases to exist only when the core diameter is zero. This is the principle on which SMFs are based.

The $V$ number can be also used to express the number of modes $M$ in a multimode step-index fiber when $V$ is large. For this case, an estimate of the total number of modes supported in such a fiber is

Eq. (5)

$$M=\frac{1}{2}{\left(\frac{2\pi a}{\lambda }\right)}^2(n_1^2-n_2^2)=\frac{V^2}{2}.$$

Because the field of a guided mode extends partly into the cladding, as shown in Fig. 5, another quantity of interest for a step-index fiber is the fractional power flow in the core and cladding for a given mode. As the $V$ number approaches cutoff for any particular mode, more of the power of that mode is in the cladding. At the cutoff point, all the optical power of the mode resides in the cladding. For large values of $V$ far from cutoff, the fraction of the average optical power residing in the cladding can be estimated by

Eq. (6)

$$\frac{P_{\mathrm{clad}}}{P}\approx \frac{4}{3\sqrt{M}},$$

where $P$ is the total optical power in the fiber.

In an SMF, the geometric distribution of light in the propagating mode is needed when predicting the performance characteristics of these fibers, such as splice loss, bending loss, cutoff wavelength, and waveguide dispersion. Thus, a fundamental parameter of an SMF is the mode-field diameter (MFD), which can be determined from the mode-field distribution of the fundamental fiber mode.^17,18,21 MFD is a function of the optical source wavelength, the core radius, and the refractive index profile of the fiber. MFD is analogous to the core diameter in an MMF, except that in an SMF, not all the light that propagates through the fiber is carried in the core.

A standard technique to find MFD is to measure the far-field intensity distribution $E^2(r)$ and then calculate the MFD using the Petermann II equation.^17,18

Eq. (7)

$$\mathrm{MFD}=2w_0=2{\left[\frac{2{\int }_0^{\infty }{E}^2(r)r^3\mathrm{d}r}{{\int }_0^{\infty }{E}^2(r)r\mathrm{d}r}\right]}^{1/2},$$

where the parameter $2w_0$ (with $w_0$ being called the spot size or the mode field radius) is the full width of the far-field distribution. For calculation simplicity, the exact field distribution can be fitted to a Gaussian function.

Eq. (8)

$$E(r)=E_0\exp (-r^2/w_0^2),$$

where $r$ is the radius and $E_0$ is the field at zero radius, as shown in Fig. 6. Then MFD is given by the $1/e^2$ width of the optical power. The Gaussian pattern given in Eq. (8) is a good approximation for values of $V$ lying between 1.8 and 2.4, which designates the operational region of practical SMFs. An approximation of the relative spot size $w_0/a$, which, for a step-index fiber, has an accuracy better than 1% in the range $1.2<V<2.4$, is given by

Eq. (9)

$$\frac{w_0}{a}=0.65+1.619V^{-3/2}+2.879V^{-6}.$$

Fig. 6

Distribution of light in a single-mode fiber (SMF) above its cutoff wavelength. For a Gaussian distribution, the mode-field diameter is given by the $1/e^2$ width of the optical power.

Manufacturers typically design an SMF to have $V$ values $>2.0$ to prevent high cladding losses, but $<2.4$ to avoid the possibility of having more than one mode in the fiber.

3.2.

Graded-Index Optical Fibers

3.3.

Performance Characteristics of Generic Optical Fibers

When considering the type of fiber to use in a particular biophotonics system application, some performance characteristics that need to be taken into account are optical signal attenuation as a function of wavelength, optical power-handling capability, the degree of signal loss as the fiber is bent, and mechanical properties of the optical fiber.

4.1.

Conventional Optical Fibers

As a result of extensive development work for telecom networks, conventional solid-core silica-based optical fibers are highly reliable and are available in a variety of core sizes. These fibers are used worldwide in telecom networks and in many biophotonics applications. Figure 7 shows the optical signal attenuation per kilometer for a standard ${\mathrm{GeO}}_2$-doped silica fiber as a function of wavelength and photon energy.¹⁷ The shape of the attenuation curve is due to three factors. Intrinsic material absorption due to electronic absorption bands causes high attenuations in the UV region for wavelengths $<\sim 500\mathrm{nm}$. The Rayleigh scattering effect then starts to dominate the attenuation for wavelengths $>500\mathrm{nm}$, but diminishes rapidly because of its $1/{\lambda }^4$ behavior. Intrinsic absorption associated with atomic vibration bands in the basic fiber material increases with wavelength and is the dominant attenuation mechanism in the IR region above $\sim 1500\mathrm{nm}$. As shown in Fig. 7, these attenuation mechanisms produce a low-loss region in silica fibers in the spectral range of 700 to 1600 nm, which matches the low-absorption biophotonics window illustrated in Fig. 1. The attenuation spike around 1400 nm is due to absorption by residual water ions in the silica material. Greatly reducing these ions during the manufacturing process results in a special low-water-content or low-water-peak fiber in which the attenuation spike has been greatly reduced.

Fig. 7

Typical attenuation curve of a ${\mathrm{GeO}}_2$-doped silica fiber as a function of wavelength.

Conventional optical fibers fall into MMF and SMF categories. Because an SMF sustains only one mode of propagation, whereas MMFs contain many hundreds of modes, an MMF can carry more optical power than an SMF. However, an SMF is needed for applications that deal with coherence effects between propagating light beams. Commercially available MMF can be purchased with standard core diameters of 50, 62.5, 100, $200\mu \mathrm{m}$, or larger. As Sec. 5 describes, these MMFs are used in areas such as laser or light-emitting diode (LED) light delivery, biostimulation, optical fiber probes, and phototherapy. SMFs have core diameters around $10\mu \mathrm{m}$, the exact value depending on the wavelength of interest. Applications of SMFs include biomedical fiber sensors used in endoscopes or catheters, in imaging systems, such as optical coherence tomography (OCT) (see Sec. 5.2), and in healthcare sensor systems based on optical fiber sensors.

4.2.

Specialty Solid-Core Fibers

Specialty solid-core fibers are custom-designed for functions such as manipulating lightwave signals to achieve some type of optical signal-processing function, extending the spectral operating range of the fiber, or sensing variations of a physical parameter, such as temperature or pressure. Incorporation of such features into an optical fiber is achieved through either material or structural variations. For biophotonics applications, the main specialty fiber types are photosensitive fibers for creating internal gratings, fibers resistant to darkening from UV light, bend-loss-insensitive fibers for circuitous routes inside bodies, and polarization-preserving optical fibers for imaging and for fluorescence analyses in spectroscopic systems.

4.2.1.

Photosensitive optical fiber

A photosensitive fiber is designed so that its refractive index changes when it is exposed to UV light. For example, doping the fiber core material with germanium and boron ions may provide this sensitivity. The main application for such a fiber is to create a short fiber Bragg grating (FBG) in the fiber core, which is a periodic variation of the refractive index along the fiber axis.^24–26

This index variation is illustrated in Fig. 8, where $n_1$ is the refractive index of the core of the fiber, $n_2$ is the index of the cladding, and $\mathrm{\Lambda }$ is the period of the grating, that is, the spacing between the maxima of the index variations. If an incident optical wave at a wavelength ${\lambda }_B$ (which is known as the Bragg wavelength) encounters a periodic variation in the refractive index along the direction of propagation, ${\lambda }_B$ will be reflected back if the following condition is met: ${\lambda }_B=2n_{\mathrm{eff}}\mathrm{\Lambda }$. Here, $n_{\mathrm{eff}}$ is the effective refractive index, which has a value falling between the refractive indices $n_1$ of the core and $n_2$ of the cladding. When a specific wavelength ${\lambda }_B$ meets this condition, this wavelength will get reflected and all others will pass through. FBGs are available in a selection of Bragg wavelengths with spectral reflection bandwidths at a specific wavelength varying from a few picometers to tens of nanometers.

Fig. 8

A periodic index variation in the core of an SMF creates a fiber Bragg grating (FBG).

As described in Sec. 5.9, an important biophotonics application of an FBG is to sense a variation in a physical parameter, such as temperature or pressure. For example, an external factor, such as strain, will slightly stretch the fiber. This stretching will lengthen the period $\mathrm{\Lambda }$ of the FBG and, thus, will change the value of the specific reflected wavelength. Similarly, rises or drops in temperature will lengthen or shorten the value of $\mathrm{\Lambda }$, respectively, thereby changing the value of ${\lambda }_B$.

4.2.4.

Polarization-maintaining fiber

In a conventional SMF, the fundamental mode consists of two independent orthogonal polarization modes.^17,18 These modes may be arbitrarily chosen as the horizontal and vertical polarizations in the $x$ direction and $y$ direction, respectively, as shown in Fig. 9. In general, the electric field of the light propagating along the fiber is a linear superposition of these two polarization modes and depends on the polarization of the light at the launching point into the fiber.

Fig. 9

Two polarization states of the fundamental mode in an SMF.

In ideal fibers with perfect rotational symmetry, the two modes are degenerate with equal propagation constants (${\beta }_x=n_{x}2\pi /\lambda ={\beta }_y=n_{y}2\pi /\lambda$, where $n_x$ and $n_y$ are the effective refractive indices along the $x$ and $y$ axes, respectively), and any polarization state injected into the fiber will propagate unchanged. In actual fibers, there are imperfections, such as asymmetrical lateral stresses, noncircular cores, and slight variations in refractive-index profiles. These imperfections break the circular symmetry of the ideal fiber and lift the degeneracy of the two modes. The modes then propagate with different phase velocities, and the difference between their effective refractive indices is called the fiber birefringence.

Eq. (16)

$$B_f=\frac{\lambda }{2\pi }({\beta }_x-{\beta }_y).$$

If light is injected into the fiber so that both modes are excited, then one mode will be delayed in phase relative to the other as they propagate. When this phase difference is an integral multiple of $2\pi$, the two modes will beat at this point and the input polarization state will be reproduced. The length over which this beating occurs is the fiber beat length, $L_B=\lambda /B_f$.

In conventional fibers, the small degrees of random imperfections in the core will cause the state of polarization to fluctuate as a light signal propagates through the fiber. In contrast, polarization-maintaining fibers have a special core design that preserves the state of polarization along the fiber with little or no coupling between the two modes. Figure 10 illustrates the cross-sectional geometry of four different polarization-maintaining fibers. The light circles represent the cladding and the dark areas are the core configurations. The goal in each design is to use stress-applying parts to create slow and fast axes in the core. Each of these axes will guide light at a different velocity. Crosstalk between the two axes is suppressed so that polarized light launched into either of the axes will maintain its state of polarization as it travels along the fiber. These fibers are used in special applications, such as fiber optic sensing and interferometry, where polarization preservation is essential.^33,34

Fig. 10

Cross-sectional geometry of four different polarization-maintaining fibers.

A further structural refinement uses the bow-tie geometry shown in Fig. 10 in order to create an extreme birefringence in a fiber. Such a birefringence allows one and only one polarization state of the fundamental mode to propagate with all other polarization modes being greatly suppressed. In these fibers, single-polarization guidance occurs in only a limited wavelength range of $\sim 100\mathrm{nm}$. Outside of that spectral range, either both of the polarization states or no light at all may be guided. These fibers are used in special fiber optic biosensing applications where it is desirable to monitor a single state of polarization.

4.3.

Double-Clad Fibers

A DCF has found widespread use in constructing optical fiber lasers and now is being used in the medical field for imaging systems, such as endoscopy.^35–38 More details on these applications are given in Secs. 5.3 and 5.4. As indicated in Fig. 11, a DCF consists of an inner cladding, an outer cladding, and a core region arranged concentrically. Typical dimensions of a commercially available DCF are a $9\text{-}\mu \mathrm{m}$ core diameter, an inner cladding diameter of $105\mu \mathrm{m}$, and an outer cladding with a $125\text{-}\mu \mathrm{m}$ diameter. Light transmission in the core region is single-mode, whereas in the inner cladding, it is multimode. The typical index of a DCF for biomedical applications is decreasingly cascaded from the core center to the cladding boundary.

Fig. 11

The cross-section of a typical double-clad fiber (DCF) and its index profile.

The integration of both single-mode and multimode transmissions allows using a single optical fiber for the delivery of the illumination light (using the single-mode core) and the collection of the tissue-reflected light (using the multimode inner cladding).

4.4.

Hard-Clad Silica Fibers

A HCS optical fiber is a multimode design that consists of a silica glass core surrounded by a hard thin plastic cladding. The hard cladding increases fiber strength and reduces static fatigue in humid environments. These fibers also feature bend insensitivity, long-term reliability, ease of handling, and resistance to harsh chemicals.^39,40 Core diameters of commercially available HCS fibers range from 200 to $1500\mu \mathrm{m}$. Table 1 lists some performance parameters of three selected HCS fibers. A common HCS fiber for industrial and medical applications has a core diameter of $200\mu \mathrm{m}$ and a cladding diameter of $230\mu \mathrm{m}$, which results in a very strong optical fiber with low attenuation ($<10\mathrm{dB}/\mathrm{km}$ at 820 nm), an NA of 0.39, negligible bending-induced loss for bend diameters $>40\mathrm{mm}$, and an extremely high core-to-clad ratio to enable efficient light coupling into and out of the fiber. The fibers are available in both high ${\mathrm{OH}}^{-}$ and low ${\mathrm{OH}}^{-}$ content for operation in the UV, visible, and NIR regions. The mechanical, optical, and structural properties of HCS fibers are especially useful in applications like laser delivery, endoscopy, photodynamic therapy (PDT), and biosensing systems.

Table 1

General specifications of selected hard-clad silica (HCS) fibers.

4.5.

Coated Hollow-Core Fibers

Because a solid-core silica fiber exhibits tremendously high material absorption above $\sim 2\mu \mathrm{m}$, fibers that have a hollow core with an internal reflection coating provide one alternative solution for the delivery of mid-IR (2 to $10\mu \mathrm{m}$) light to a localized site with a low transmission loss. Section 4.9 describes middle-IR fibers as another alternative for operation above $2\mu \mathrm{m}$. Examples of light sources operating above $2\mu \mathrm{m}$ include ${\mathrm{CO}}_2$ ($10.6\mu \mathrm{m}$) and Er:YAG ($2.94\mu \mathrm{m}$) lasers, which have wide applications in urology, dentistry, otorhinolaryngology, and cosmetic surgery. In addition, hollow-core fibers are useful in the UV region, where silica-based fibers also exhibit high transmission losses.

As shown in Fig. 12, a hollow-core fiber is composed of a glass tube with metal and dielectric layers deposited at the inner surface plus it has a protective jacket on the outside.^41–47 In the fabrication process of these fibers, a layer of silver (Ag) is deposited on the inside of a glass tube, which then is covered with a thin dielectric film, such as silver iodide (AgI). Light is transmitted along the fiber through mirror-type reflections from this inner metallic layer. The thickness of the dielectric film (normally $<1\mu \mathrm{m}$) is selected to give a high reflectivity at a particular IR wavelength or a band of wavelengths. Tubes of other materials, such as plastic and metal, are also employed as hollow-core waveguides, but glass hollow-core fibers provide more flexibility and better performance. The bore sizes (diameters of the fiber hole) can range from 50 to $1200\mu \mathrm{m}$. However, since the loss of all hollow-core fibers varies as $1/r^3$, where $r$ is the bore radius, the more flexible smaller hollow-core fibers have higher losses.

Fig. 12

The cross-section of a typical hollow-core fiber (not to scale).

Compared to solid-core fibers, the optical damage threshold is higher in hollow-core fibers, there are no cladding modes, and there is no requirement for angle cleaving or antireflection coating at the fiber end to minimize laser feedback effects. To direct the light projecting from the hollow-core fiber into a certain direction, sealing caps with different shapes (e.g., cone and slanted-end) at the distal end of the fiber have been proposed and demonstrated by using a fusing and polishing technique.⁴⁵

4.6.

Photonic Crystal Fibers

A PCF differs from a conventional optical fiber in that the PCF has a geometric arrangement of air holes that run along the whole length of the fiber.^48–52 A PCF is also referred to as a microstructured fiber or holey fiber. The core of the PCF can be solid or hollow as shown in Fig. 13 by the cross-sectional images of two typical PCF structures. This structural arrangement creates an internal microstructure, which offers another dimension of light control in the fiber. The arrangement, size, and spacing (known as the pitch) of the holes in the microstructure and the refractive index of its constituent material determine the light-guiding characteristics of a PCF. Depending on the PCF structure, light is guided along the fiber either by total internal reflection or by a bandgap effect.

Fig. 13

Sample structural arrangements of air holes in solid-core (a) and hollow-core (b) photonic crystal fiber (PCF).

The fact that the core can be made of pure silica gives the PCF a number of operational advantages over conventional fibers, which typically have a germanium-doped silica core. These advantages include very low losses, the ability to transmit high optical power levels, and a high resistance to darkening effects from nuclear radiation. The fibers can support single-mode operation over wavelengths ranging from 300 to $>2000\mathrm{nm}$. The mode field area of a PCF can be $>300\mu {\mathrm{m}}^2$ compared to the $80\text{-}\mu {\mathrm{m}}^2$ area of a conventional SMF. This allows the PCF to transmit high optical power levels without encountering the nonlinear effects exhibited by conventional fibers.

For biophotonics applications, some hollow-core PCFs, such as the Kagome fiber, are promising substitutes for the hollow-core fiber mentioned in Sec. 4.5. These PCFs have a broadened transmission window for delivering middle-IR light with reduced transmission or bending loss.^53,54 PCFs are also being used in optical fiber-based biosensors. Section 5.8 discusses a number of these biosensor applications.

4.7.

Plastic Optical Fibers

POFs are an alternative to glass optical fibers for applications such as biomedical sensors.^55–61 POFs were first fabricated in the late 1960s, which makes them about the same age as silica fibers. However, compared to silica fibers, POFs did not become as popular due to factors such as high attenuation, immature fabrication techniques, and low temperature tolerance ($\sim 100°\mathrm{C}$). Most POFs are made of polymethylmethacrylate with a refractive index of $\sim 1.492$, which is slightly higher than the index of silica. The size of a POF is normally larger than typical silica fibers, with diameters ranging up to 0.5 mm.

In addition to POFs with large core diameters, currently both multimode and single-mode POFs with telecom core sizes are commercially available. The standard core sizes of a multimode POF include 50- and $62.5\text{-}\mu \mathrm{m}$ diameters, which are compatible with the core diameters of standard glass telecom MMFs. The development of single-mode POF structures enables the creation of FBGs inside plastic fibers, which, therefore, provides more possibilities for POF-based biosensing. Together with the advantages of low cost, inherent fracture resistance, low Young’s modulus, and biocompatibility, POF has become a viable alternative for silica fibers in biosensing. For example, some special fiber sensor designs, such as the exposed-core technique now being used in etched-core FBG sensors, were first realized by using polymer optical fibers.^62–64 Although most POFs have a higher refractive index than silica, a single-mode perfluorinated POF with a refractive index of 1.34 has been fabricated.⁶⁰ This low-index fiber has the potential for improvement in the performance of biosensors, because it results in a stronger optical coupling between a light signal and the surrounding biosensor analyte. Section 5.8 has more details on biosensors.

4.8.

Side-Emitting and Side-Firing Fibers

In most cases, the function of optical fibers is to transport light to its intended destination with as little optical power loss and signal distortion as possible. However, a different category of optical fiber that emits light along its entire length is finding increased biomedical applications.^65–71 Such a side-emitting fiber or glowing fiber acts as an extended emitting optical source. These side-emitting fibers are widely available and their popular nonmedical applications include decorative lamps, submersible lighting, event lighting, and lighted signs. For medical use, wearable luminous fabric devices consisting of textiles imbedded with glowing fibers are being considered for use as a pulse oximeter and for inflammation and pain reduction, edema treatment, sunburn relief, and wound healing.^67,68

Another embodiment for lateral emission of light is the side-firing fiber. As described in Sec. 5.1, this structure consists of an angled end face at the distal end of the fiber. This enables the light to be emitted almost perpendicular to the fiber axis. Uses of side-firing fibers for biomedical procedures include PDT, therapy for atrial fibrillation, treatment of prostate enlargements, and dental treatments.^69–71

A variety of fiber materials and construction designs have been used to fabricate side-emitting fibers. Both plastic and silica fibers are available. One popular method of achieving the fiber glowing process is to add scattering materials into either the core or cladding of the fiber. This creates a side-emitting effect through the scattering of light from the fiber core into the cladding and then into the medium outside of the fiber. As Fig. 14 shows, one implementation of a side-emitting fiber is to attach a length of this type of fiber (e.g., 10 to 70 mm for short diffusers or 10 to 20 cm for long diffusers) to a longer delivery fiber (nominally 2 to 2.5 m), which runs from the laser to the side-emitting unit.

Fig. 14

The red input light is scattered radially along the length of the side-emitting fiber.

As a result of scattering effects, the light intensity in a side-emitting fiber exponentially decreases with distance along the fiber. Therefore, the intensity of the side-emitted light will decrease as a function of distance along the fiber. If it is assumed that the side-emitting effect is much stronger than light absorption and other losses in the fiber, then the side-glowing radiation intensity, $I_S$, that is emitted in any direction per steradian at a point located a distance $x$ from the input end of the diffuser is^65,66

In many practical biomedical applications where the glowing fiber lengths are of the order of 10 or 20 cm, the effect of an exponential decrease in side-emitted light intensity may not be a major problem. However, there are two methods for creating a more uniform longitudinal distribution of the side-emitted light. One method is to simultaneously inject light into both ends of the fiber if it is possible to do so. In other applications, a reflector element attached to the distal end of the fiber can produce a relatively uniform emitted light distribution of the combined transmitted and reflected light if the fiber diffuser is not too long.

4.9.

Middle-Infrared Fibers

A middle-IR, or simply an IR, optical fiber transmits light efficiently at wavelengths $>2\mu \mathrm{m}$.^72–81 Compared to conventional telecom silica fibers, IR fibers usually are more fragile and have larger refractive indices, lower melting or softening points, and greater thermal expansion. Typically, in biomedical photonics, IR fibers are used in lengths $<2$ to 3 m for fiber optic sensors and optical power delivery applications. Based on the fiber material and structure used for their fabrication, IR fibers can be classified into glass, crystalline, PCFs, and hollow-core waveguide categories. This section describes glass and crystalline fiber classes. Hollow-core waveguides and PCFs are described in Secs. 4.5 and 4.6, respectively.

Glass materials for IR fibers include heavy metal fluorides,^72–74 fluoroindates,⁷⁵ chalcogenides,^76,77 and heavy metal germinates.⁷⁸ Heavy metal fluoride glasses (HMFG) are the only materials that transmit light from UV to mid-IR without any absorption peak. These HMFG fibers are drawn from a preform using the same techniques as are used for silica fibers. Commercial HMFG fibers include ${\mathrm{InF}}_3$ (with attenuations of $<0.25\mathrm{dB}/\mathrm{m}$ between 1.8 and $4.7\mu \mathrm{m}$) and ${\mathrm{ZrF}}_4$ (with attenuations of $<0.25\mathrm{dB}/\mathrm{m}$ between 1.8 and $4.7\mu \mathrm{m}$). Such optical losses allow for practical short- and medium-length applications of a few meters. The reliability of HMFG fibers depends on protecting the fiber from moisture and on pretreatment of the preform to reduce surface crystallization. Thus, HMFG fibers are a good choice for the 1.5- to $4.0\text{-}\mu \mathrm{m}$ spectrum, which is increasingly used in medical applications.

Chalcogenide glasses are based on the chalcogen elements sulfur, selenium, and tellurium together with the addition of other elements, such as germanium, arsenic, and antimony. These glasses are very stable, durable, and insensitive to moisture. Chalcogenide fibers are able to transmit longer wavelengths in the IR than fluoride glass fibers. Longer wavelengths are transmitted through the addition of heavier elements. All chalcogenide fibers have strong extrinsic absorption resulting from contaminants such as hydrogen. Although the losses of chalcogenides are generally higher than that of the fluoride glasses, these losses are still adequate for about 2-m transmissions in the spectral band ranging from 1.5 to $10.0\mu \mathrm{m}$. Some typical loss values are $<0.1\mathrm{dB}/\mathrm{m}$ at the widely used 2.7- and $4.8\text{-}\mu \mathrm{m}$ wavelengths. The maximum losses for the three common chalcogenides are as follows:

Applications of chalcogenide fibers include Er:YAG ($2.94\mu \mathrm{m}$) and ${\mathrm{CO}}_2$ ($10.6\mu \mathrm{m}$) laser power delivery, microscopy and spectroscopy analyses, and chemical sensing.

${\mathrm{GeO}}_2$ glass fibers generally contain heavy metal oxides (for example, PbO, ${\mathrm{Na}}_2\mathrm{O}$,  and  ${\mathrm{La}}_2{\mathrm{O}}_3$) to shift the IR absorption edge to longer wavelengths. The advantage of ${\mathrm{GeO}}_2$ fibers over HMFG fibers is that ${\mathrm{GeO}}_2$ glass has a higher laser damage threshold and a higher glass transition temperature, making ${\mathrm{GeO}}_2$ glass fibers more mechanically and thermally stable than, for example, tellurite glasses. ${\mathrm{GeO}}_2$ glass has an attenuation of $<1\mathrm{dB}/\mathrm{m}$ in the spectrum ranging from 1.0 to $3.5\mu \mathrm{m}$. A key application of ${\mathrm{GeO}}_2$ fibers is for laser power delivery from the laser to a patient from Ho:YAG ($2.12\mu \mathrm{m}$) or Er:YAG ($2.94\mu \mathrm{m}$) lasers, where the fiber losses nominally are 0.25 and $0.75\mathrm{dB}/\mathrm{m}$, respectively. These fibers can handle up to 20 W of power for medical procedures in dermatology, dentistry, ophthalmology, orthopedics, and general surgery.

Crystalline IR fibers can transmit light at longer wavelengths than IR glasses, for example, transmission up to $12\mu \mathrm{m}$ is possible.^79–81 These fibers can be classified into single-crystal and polycrystalline fiber types. Although there are many varieties of halide crystals with excellent IR transmission features, only a few have been fabricated into optical fibers because most of the materials do not meet the required physical property specifications needed to make a durable fiber. Polycrystalline silver-halide fibers with AgBr cores and AgCl claddings have exhibited excellent crystalline IR fiber properties.^78–81 There are several extrinsic absorption bands for Ag-halide fibers, for example, 3 and $6.3\mu \mathrm{m}$ due to residual water ions. The attenuation of Ag-halide fibers is normally $<1\mathrm{dB}/\mathrm{m}$ in the 5- to $12\text{-}\mu \mathrm{m}$ spectral band and 0.3 to $0.5\mathrm{dB}/\mathrm{m}$ at $10.6\mu \mathrm{m}$ and can be as low as $0.2\mathrm{dB}/\mathrm{m}$ at this wavelength.

4.10.

Optical Fiber Bundles

To achieve high throughput with flexible glass or plastic fibers, multiple fibers are often arranged in a bundle.^81–83 Each fiber acts as an independent waveguide that enables light to be carried over long distances with minimal attenuation. A typical large fiber bundle array can consist of a few thousand to 100,000 individual fibers, with an overall bundle diameter of $<1\mathrm{mm}$ and an individual fiber diameter between 2 and $20\mu \mathrm{m}$. Such fiber bundles are mainly used for illumination purposes. In addition to flexibility, fiber bundles have other potential advantages in illumination systems: (1) illuminate multiple locations with one source by splitting the bundle into two or more branches, (2) merge light from several sources into a single output, and (3) integrate different specialty fibers into one bundle.

For bundles with a large number of fibers, the arrangement of individual fibers inside the bundle normally randomly occurs during the manufacturing process. However, for certain biomedical applications, it is required to arrange a smaller number of fibers in specific patterns. Optical fibers can be bundled together in an aligned fashion such that the orientations of the fibers at both ends are identical. Such a fiber bundle is called a coherent fiber bundle or an ordered fiber bundle. Here the term coherent refers to the correlation between the spatial arrangement of the fibers at both ends and does not refer to the correlation of light signals. As a result of the matched fiber arrangements on both ends of the bundle, any pattern of illumination incident at the input end of the bundle is maintained and ultimately emerges from the output end.

Figure 15 illustrates illumination and coherent fiber bundle configurations. As shown in Fig. 15(a), the optical fibers are randomly oriented on both ends in a bundle configuration that contains hundreds or thousands of fibers. Figure 15(b) shows one possibility of an ordered fiber arrangement in a bundle. In this hexagonal packaging scheme, the central fiber is used for illuminating a target tissue area and the surrounding fibers are used for fluorescent, reflected, or scattered light returning from the tissue. The discussions in Sec. 5 give further examples.

Fig. 15

(a) Randomly arranged fibers in a bundle containing many optical fibers. (b) Coherent bundle cable with identical arrangements of optical fibers on both ends.

When using optical fiber bundles, there is a trade-off between high light collection efficiency and good image resolution. A large core diameter enables high light transmission but poor image resolution. A thicker cladding avoids crosstalk among individual fibers but limits light collection and image resolution because more of the light-emitting area is blocked when a thicker cladding is used. In practice, the core diameter of individual fibers is 10 to $20\mu \mathrm{m}$ and the cladding thickness is $\sim 1.5$ to $2.5\mu \mathrm{m}$. Coherent fiber bundles are the key components in modern fiber optic endoscopes. Current sophisticated ear, nose, throat, and urological procedures utilize high-resolution flexible image bundles for image transfer.

5.1.

Optical Fiber Biomedical Probes

Optical fibers are widely used as biomedical probes. An important factor to consider when using optical fiber probes for light delivery and collection functions is the tip geometry of the fiber at its distal end. The shape of the tip determines how the light emerging from the fiber is distributed on a tissue sample and also determines the light collection efficiency for viewing the irradiated tissue sample or examining scattered or fluorescing light emitted from the sample.^84–96

When deciding on the geometry of an optical fiber tip, parameters that need to be evaluated include the sizes of the illumination and light-collection areas, the collection angle (related to the fiber NA), and the fiber diameter. Another important factor to keep in mind is that biological tissue has a multilayered characteristic from both compositional and functional viewpoints. Specific biological processes and diseases occur at different levels within this multilayered structure. Thus, when designing an optical fiber probe, it is necessary to ensure that the light that is aimed at the targeted biological process or disease penetrates the tissue down to the desired evaluation layer.

A generic schematic of an optical fiber-based probe system is shown in Fig. 16. The excitation light can come from optical sources, such as a white light-emitting diode, a deuterium lamp, or a laser diode. Devices such as a dielectric spectral bandpass filter or a monochromator can be used to select a very narrow spectral width for delivery to a tissue sample. The transport fiber to the sample can be any of the fibers described in Sec. 4. If it is used, a return fiber can be selected to be an optical fiber bundle, an SMF, an MMF, a polarization-preserving fiber, a PCF, an IR fiber, or a DCF. In the case of a DCF, the light is typically delivered to the sample using the single-mode core of the fiber and the return path uses the large inner multimode cladding of the fiber for increased light collection efficiency. The photodetector that interprets the returned light can be an instrument such as a viewing scope, a photomultiplier tube (PMT), a charge-coupled device (CCD) camera, or an optical spectrum analyzer.

Fig. 16

Schematic of a generic optical fiber probe system.

A great number of fiber probe configurations have been analyzed, designed, and implemented. As Figs. 17(a) and 17(b) show, to distribute the light emerging from the fiber, the end face can be cleaved or polished flat or the exit surface can be polished at an angle relative to the fiber axis to deflect the light to the selected area. In other probe embodiments, a miniature lens can be formed or attached to the fiber end or a light-diffusing fiber segment can be attached (see Fig. 14). If the end face angle is equal to the critical angle for total internal reflection [Fig. 17(c)], then the light will leave the fiber through its side. This is the basis of the side-firing fiber. To increase the light illumination or collection efficiency, probe designs with two or more fibers are used. Figures 17(d) and 17(e) show two probe designs based on using either two flat-polished or beveled end faces.

Fig. 17

Examples of fiber distal end tips. (a) Flat-polished fiber. (b) Angle-polished fiber. (c) Side-emitting fiber. (d) Dual flat-polished fibers. (e) Dual angle-polished fibers. (Reproduced with permission from Ref. 84.)

At the expense of having a larger probe, the efficiency can be increased by using more than two fibers for either or both of the illumination and collection functions. Figure 18 shows a classic example of hexagonal packing for a fiber probe. This arrangement enables the use of one or more fibers for illumination or fluorescence excitation with the remaining fibers used as the collection channels. Optionally, optical filters with different spectral pass bands can be deposited on the end faces of individual collection fibers.

Fig. 18

End-face views of two different arrangements for excitation and collection functions of a hexagonal packing of multiple-fiber probes.

5.2.

Optical Coherence Tomography

Optical imaging is playing an increasingly important role for the diagnosis of living cells and biological tissue. Among the successful medical imaging techniques, OCT has become an important and established noninvasive optical-based tissue diagnosis procedure.^97–103 This section addresses three diverse areas of optical fiber uses in OCT. The first area is the use of single-mode silica fibers for transmitting the incident and reflected light in the OCT imaging process. The second application is the use of erbium-doped or ytterbium-doped optical fibers to construct wideband low-coherence fiber laser sources that enable high-resolution imaging. A third application is the use of various types of optical fibers to broaden the emission spectrum from commercially available femtosecond solid-state lasers for image-resolution enhancement.

5.2.1.

OCT imaging process

From the time it was developed for noninvasive imaging of biological tissue in the early 1990s, OCT has been successfully applied to many of the major human biomedical procedures for in vitro and in vivo imaging. As noted in Sec. 2, a biological window exists between $\sim 600$ and 1500 nm, where light can penetrate several millimeters into tissue. This is the operating region used for OCT, because light attenuation here is more dependent on scattering processes than on absorption processes. OCT performs ranging-based optical imaging in biological tissue due to the fact that a portion of the incident light is reflected by each interface whenever there is a change in refractive index in the biological tissue. Because the velocity of light is extremely high, low-coherence interferometry is utilized to measure the time delay of the reflected light. Low-coherence interferometry is performed using a Michelson-type interferometer. The use of optical fibers enables OCT systems to be compact and portable.

Figure 19 shows the schematic of a basic optical fiber-based OCT system. The light from the broadband optical source is divided in half by an optical splitter or coupler and the two halves are sent to the reference arm and the sample arm separately through SMFs. Reflections from the reference arm mirror and from the sample are recombined by the optical coupler and sent to the detector. Because low-coherence light from a broadband source is used, constructive interference of the light only occurs when the optical path lengths are matched between the two arms of the interferometer. In order to achieve a high image resolution, it is necessary to use SMFs. By scanning across the face of the reference arm mirror, a single axial depth scan slice of reflection data is acquired from the sample. By slightly moving the position of the reference mirror, another scan slice is obtained at a different axial depth.

Fig. 19

A typical optical coherence tomography (OCT) imaging setup (reproduced with permission from Ref. 103).

One important parameter for characterizing an OCT system is the imaging resolution. The axial resolution $\mathrm{\Delta }x$ is determined by the coherence length of the source. Given that an OCT system uses a broadband probing light source with a center wavelength $\lambda$, the standard equation for the axial resolution is

Eq. (18)

$$\mathrm{\Delta }x=\frac{2\ln 2}{\pi n\mathrm{\Delta }\lambda }{\lambda }^2,$$

where $n$ is the optical fiber refractive index and $\mathrm{\Delta }\lambda$ is the spectral width of the probing source. Therefore, the axial resolution is directly proportional to the square of the source central wavelength and inversely proportional to the source spectral bandwidth. Thus, to achieve high axial resolution, optical sources with broad spectral bandwidths are desired. It is in the area of spectral broadening that various types of optical fibers are making a significant impact on OCT. Higher resolution could also be achieved by decreasing the center wavelength from the standard 1300-nm value. However, shorter wavelengths are more highly scattered in biological tissue, resulting in less imaging penetration.

5.3.

Optical Spectroscopy

A number of optical spectroscopic methodologies that make use of optical fiber technology are being used in research laboratories and medical clinics for rapid, accurate, and noninvasive detection and diagnosis of premalignant and malignant changes in tissue. In contrast to a standard biopsy that involves the removal of one or more tissue samples for later laboratory evaluation, optical spectroscopic methods eliminate this need for surgical removal of tissue. Instead, by using an optical probe placed on or near the surface of the tissue to be evaluated, an imaging system records some in vivo form of spectral analysis of the tissue. Depending on the method used, the diagnostic information obtained from the biological tissues can be at the biochemical, molecular structural, or physiological levels.^113,114 Selections of optical fiber probe applications illustrated here from among the numerous optical spectroscopic methodologies are fluorescence spectroscopy,^115,116 elastic scattering spectroscopy,^117–119 diffuse correlation spectroscopy (DCS),¹²⁰ and coherent anti-Stokes Raman scattering (CARS) spectroscopy.¹²¹ Each spectroscopic discipline is continuously adopting more sophisticated optical fiber-based systems for delivering probing light to an investigation site, for collecting fluorescence products or scattered imaging light, and for returning this light to photodetection and analysis instruments.

5.3.1.

Fluorescence spectroscopy

Fluorescence spectroscopy instruments generally employ some variation of the generic probe system shown in Fig. 16. In such a setup, to be of analytical use, the wavelength of incident radiation needs to be selectable and the detector signal should be capable of precise spectral manipulation and presentation. Therefore, monochromators or spectrometers are provided for both the selection of the excitation light and the analysis of the emission from the sample.

In a fluorescence process, photons from a UV or visible external source are sent by means of an optical fiber to a target area. There the photons can give up their energy to electrons in a molecule, thereby exciting these electrons from their ground state to higher electronic states. Fluorescence occurs when the excited electrons return to their initial state, thereby emitting a photon. The photon energy is equal to the difference in the excited and ground state electron energy. For a given instrument and environmental condition of the tissue being analyzed, the excitation and emission properties of a compound are fixed and can be used for identification and quantification of molecules and molecular processes. The key advantage of fluorescence spectroscopy in the diagnosis of tissue pathology is that the excitation or emission spectra are sensitive to the biochemical composition of the tissue. This feature is helpful in assessing whether the biochemical state of the tissue is normal or malignant.

A proper fluorescence probe design is important for efficient transport and collection of photons, which ultimately helps in quantifying the resultant emissions and in understanding light-matter interaction. The fiber parameters that need to be taken into consideration include fiber core and cladding diameters, NA, and core-clad ratio. Numerous embodiments of multiple-fiber and single-fiber probes have been implemented for fiber-based fluorescence spectroscopy systems.^86–90,122

One method for using a multiple-fiber probe is to couple the output of the light from the excitation monochromator into one arm of a bifurcated fiber bundle, which delivers the excitation light to the sample. The other arm of the bifurcated fiber bundle collects the emission light and couples the light to the emission monochromator. In another multiple-fiber embodiment, the excitation light can be guided to the tissue target through a large-core fiber (e.g., 400- or $600\text{-}\mu \mathrm{m}$ core diameter), while the collection and transmission of fluorescence emission can be realized by two or more collection fibers that are arranged circularly around the excitation fiber.

An example schematic of a multifiber fluorescence spectroscopy probe⁸⁷ is shown in Fig. 20. As the cross-section of the probe end shows, four collection fibers are symmetrically arranged around an excitation fiber. The core and cladding diameters of the excitation fiber are 200 and $220\mu \mathrm{m}$, respectively, and are 400 and $440\mu \mathrm{m}$, respectively, for the collection fibers. This particular probe was designed for immersion in a liquid.

Fig. 20

Schematic of an example fluorescence spectroscopy probe that uses one excitation fiber and four collection fibers (reproduced with permission from Ref. 87).

Currently, single-fiber probes are being used extensively for both in vitro and in vivo fluorescence research. The main configurations use MMFs, DCFs, or PCFs. In these applications, MMFs typically have core diameters of 200 to $600\mu \mathrm{m}$. NA of these MMFs ranges from 0.2 to 0.48, which corresponds to delivery and acceptance cone half-angles of 8 to 21 deg. A single-fiber probe consisting of a single optical fiber with a core diameter of $600\mu \mathrm{m}$ allows for common illumination and light collection channels by means of a dichroic beam splitter arrangement.

Figure 21 shows an example of a fluorescence spectroscopy application that uses a combination of SMF, DCF, and MMF.⁹⁰ DCF had an inner core diameter of $7.4\mu \mathrm{m}$ with an NA of 0.12. The inner cladding diameter was $125\mu \mathrm{m}$ with an NA of 0.44 and the outer cladding diameter was $180\mu \mathrm{m}$. The function of the DCF is to enable simultaneous illumination of the sample and collection of the excitation emissions. A high level of excitation efficiency into the single-mode core of the DCF was achieved by sending 488-nm excitation light from an argon laser into a short SMF that was fusion-spliced to the DCF. The inner cladding layer of the DCF was used to collect the fluorescence signals emitted by the sample. At the receiving spectrometer, the fluorescent light collected by the DCF was coupled into a $50\text{-}\mu \mathrm{m}$ core diameter MMF by means of collimator lenses $C_1$ and $C_2$. The neutral-density filter (ND) adjusted the intensity of the input beam, a bandpass filter (BP) removed the background noise of the laser, and a longpass filter (LP) blocked the fundamental excitation light that was elastically scattered back from the sample.

Fig. 21

Setup of a fluorescence spectroscopy system that uses SMF, DCF, and multimode fiber (reproduced with permission from Ref. 90).

PCFs are also exploited to enhance the fluorescence emission. For example, in one experiment, a gold-coated side-polished D-shape microstructure optical fiber was employed to excite a surface plasmon, which could confine and amplify the leaky evanescent field from the fiber core. The plasmonic effect alone was found to provide an immediate fluorescence enhancement factor of two.¹²²

5.3.2.

Elastic scattering spectroscopy

Elastic scattering spectroscopy (ESS), which is also known as diffuse reflectance spectroscopy, is based on the elastic scattering of photons.^117–119 In an elastic scattering process, photons hit tissue components and are scattered without experiencing a change in their energy, which means without a change in wavelength. The relative intensity of the backscattered light that can be collected by an optical fiber probe depends on the sizes and concentrations in the tissue of scattering components (e.g., nuclei, mitochondria, and connective tissue) and absorbing components (e.g., hemoglobin). Because the sizes and densities of these organelles change when the tissue becomes premalignant or malignant, the ESS process assists pathologists in diagnosing dysplasia.

The scattered light may be detected by either the same fiber that is used for illumination or by a separate collection fiber. An example of an ESS optical fiber probe that uses separate illuminating and collecting fibers is shown in Fig. 22. A white light source is used in ESS in order to examine both the scattering and absorption properties of the tissue over a wide range of wavelengths. The optical probe passes through an endoscope channel and is placed in direct contact with tissue. A flash of white light illuminates a cylinder of tissue $\sim 0.5\mathrm{mm}$ in diameter and nominally 1 mm deep. The backscattered signal is available for diagnosis within milliseconds. Note that only light that has undergone multiple scatterings can enter the collection fiber.

Fig. 22

Example of an elastic scattering spectroscopy probe with separate illumination and collection fibers.

5.3.3.

Diffuse correlation spectroscopy

DCS is a continuous noninvasive technique that uses the time-averaged intensity autocorrelation function of a fluctuating diffusely reflected light signal to probe the flow of blood in thick or deep tissue vasculature, such as brain, muscle, and breast.^120,123 This technique is also known as diffusing-wave spectroscopy. DCS operates in the 600- to 900-nm spectral range, where the relatively low tissue absorption enables deep penetration of light (see Fig. 1). This technique accounts for multiple scatterings of photons in thick tissues and quantifies blood flow in these tissues. The flow measurements are achieved by monitoring the speckle fluctuations of photons that are diffusely scattered by the movements of blood cells in the tissue.

A typical DCS setup uses a laser that has a long coherence length, a single-photon-counting avalanche photodiode (APD) or a PMT as the photodetector, and a hardware autocorrelator. The schematic of the experimental setup that is shown in Fig. 23 is representative of an actual DCS measurement system. The purpose of the autocorrelator is to compute the autocorrelation function of the temporal fluctuation of the measured light intensity. To measure the blood flow, NIR light from the long-coherence-length laser is launched into the tissue through an MMF that has its distal end placed on the tissue surface. The light that is scattered through the tissue is collected by an SMF or a few-mode fiber that is placed a few millimeters or centimeters away from the source fiber. The observed temporal fluctuation of the light intensity in a single speckle area can be associated with the movements of the scattering components, which mainly are red blood cells in microvasculature. The blood flow can be quantified by calculating the decay of the light intensity, which is derived from the autocorrelation function results.

Fig. 23

Schematic of a representative experimental diffuse correlation spectroscopy setup (reproduced from Ref. 120).

5.3.4.

Raman spectroscopy

Raman spectroscopy is a noninvasive, label-free biomedical optics tool that is used for evaluating the chemical composition of biological tissue samples. The compositional information is obtained by analyzing photons that have been inelastically scattered from the vibrational and rotational transitions in chemical bonds of the tissue sample, thereby providing detailed information about the biochemical composition of a sample. A number of variations have been investigated within this discipline, such as CARS, time-resolved spectroscopy, polarization modulation, and wavelength-modulated Raman spectroscopy.^121,124,125

Miniaturized fiber-optic Raman probe designs have included a simple single fiber, two fibers with lenses for better sensitivity, and bifurcated coherent fiber bundles with integrated optical filtering modules. The applications include in vivo biomedical Raman spectroscopy for examining the bladder, the oral cavity, the larynx, the cervix, the lung, the breast, the colon, the esophagus, and the stomach. Figure 24 gives a general illustration of three Raman probe designs.¹²¹ In such probes, laser light emerging from an excitation fiber is scattered by the tissue and is then coupled into a collection fiber. Because only inelastically scattered photons are of interest in Raman spectroscopy, the elastically scattered photons are suppressed by longpass or notch filters. After emerging from the collection fibers, the inelastically scattered photons are dispersed in a spectrograph and registered by a CCD detector array.

Fig. 24

Three generic probe configurations used in Raman spectroscopy: (a) unfiltered single fiber, (b) two fibers with optical filters and a lens, and (c) fiber bundles with on-the-tip filters. BS, beamsplitter; LP, longpass; M, mirror; L, fiber coupling lens; EF, excitation fiber; CF, collection fiber; and BP, bandpass. (Reproduced with permission from Ref. 121.)

The simple probe configuration shown in Fig. 24(a) uses the same fiber for photon excitation and collection. Although this geometry offers a small, low-cost probe, its disadvantages include divergence of the beam at the distal end of the fiber and the generation of intense Raman signals within the core of the fiber, which interfere with the detection of photons below $1500{\mathrm{cm}}^{-1}$. The lens-based probe design shown in Fig. 24(b) offers better sensitivity owing to improved focusing of the light onto the tissue sample and its higher collection efficiency. This design suppresses the Raman signals that are generated in the excitation fiber and it enables detection of the entire desired 300- to $4000\text{-}{\mathrm{cm}}^{-1}$ wave-number range. The probe design illustrated in Fig. 24(c) offers further miniaturization because the optical longpass and bandpass filters are placed directly on the fiber ends and a fiber bundle that surrounds a central excitation fiber improves the Raman signal collection efficiency.

5.4.

Endoscopy

The field of endoscopy deals with using a photonics-based instrument, called an endoscope, to look directly inside a hollow organ or body cavity for medical reasons. An endoscopic procedure is often called by names that depend on the body region being examined or treated, for example, laparoscopy for operations within the abdominal or pelvic cavities (also used for referring to minimally invasive surgery), rhinoscopy for examinations of the nasal cavity, arthroscopy for procedures involving joints, and bronchoscopy for visualizing the inside of the airways for diagnostic and therapeutic purposes. A diverse collection of endoscope designs have been progressively implemented to obtain finer details from biological tissues, such as acquiring images in various colors, obtaining greater depth resolution, or observing molecular or metabolic functions. Various advanced technologies are enabling these capabilities, for example, the development of new optical fiber types and miniaturized cameras. This section first discusses some endoscopic applications using different optical fiber designs, then gives examples of endoscopic or minimally invasive surgery, and concludes with a description of a recent technique called tethered capsule endomicroscopy.

5.4.1.

Basic endoscopy

A basic endoscope system consists of the following units:

• • A flexible outer tube containing one or more of the optical fibers described in Sec. 4 for illumination and viewing functions. As shown in Fig. 25, the encapsulating endoscope tube can also contain air, water, and suction or biopsy tubes, plus tip control wires.

• • An external light source used to illuminate (via the optical fibers) the organ, tissue area, or object being diagnosed

• • A lens system that collects reflected or fluorescing light from the diagnostic site for transmission via optical fibers to a viewer

• • A viewing mechanism (for example, a simple eyepiece, a video screen, or a camera).

Fig. 25

Cross-sectional view of an endoscope channel showing various fibers and tubes.

An endoscope can also contain one or more additional channels to allow for the insertion of capillaries for specimen collection, manipulators for microsurgery, or additional optical fibers for therapy purposes. Traditionally, endoscopes have used a small flexible optical fiber bundle to transmit light to a diagnostic area and a coherent fiber bundle for image retrieval.

Although the use of fiber bundles results in a noninvasive exploratory instrument, the number of individual fibers in the bundle and their characteristics limit the image resolution. In addition, the size and limited flexibility of the endoscope tube require patient sedation during use of the endoscope. Consequently, during the evolution to modern endoscopes, a replacement for the fiber bundle was sought using single optical fibers, such as SMF, MMF, DCF, or hollow-core PCF, to reduce the size and to increase the flexibility of endoscopes, thereby enabling safer, faster, and less expensive clinical procedures. In some endoscope designs, a single fiber is used for both delivery of the incident light and image transport. Other endoscope designs use a central fiber for illuminations and either another individual fiber or a circular array of fibers for image retrieval.^126–131

A major limitation when using MMFs is the temporal variation of the modal distributions in the fiber, which can lead to image flickering. Actually this situation occurs in all imaging systems that employ MMFs both for the delivery and the collection of light. SMFs offer better resolution and a decrease in speckle noise compared with MMFs. Compared to conventional SMF and MMF, the DCF types described in Sec. 4.3 can support single-mode light transmission through the core for illumination of a target area and multimode image transfer (consisting of partially incoherent light reflected from the sample) through the inner cladding back to an imaging instrument. Thus, a DCF achieves high-resolution imaging, reduced speckle effects, and a large depth of field resulting from the larger collection diameter of the cladding.

In conjunction with evaluating the potentials of using single fibers are innovative optical designs, miniature image scanning mechanisms, and specific molecular probes. The aims of the combined fiber plus technology advances are to provide images with various colors, more depth, and multiple dimensions in order to acquire greater biological details from tissue.^132–134

5.5.

Fiber-Based Confocal Microscopy

Confocal microscopy is a well-established and widely used tool in cell biology because of its optical sectioning capability. This feature enables the analysis of morphologic changes in thick biologic tissue specimens with subcellular resolution. Traditionally, confocal microscopy was limited to in vitro studies of biopsy specimens and to in vivo analyses of easily accessible sites, such as the cornea, the skin, and lip and tongue areas, because it required large microscope objectives and relatively long image acquisition times.

These limitations were overcome through the use of optical fiber technology.^139,140 As one example, Fig. 26 shows a concept based on using a single optical fiber and a miniaturized microelectromechanical system (MEMS) scanning mirror.¹³⁹ In this setup, 25 mW of linearly polarized 635-nm light from a laser diode (LD) was coupled through a bulk-optic polarizing beam splitter (PBS1) to a single-mode polarization-maintaining fiber for delivery to the tissue sample. At the distal end of this fiber, the beam was raster scanned on the sample by means of the MEMS scanner. The functions of the two PBS elements and the quarter-wave plate were to minimize the amplitudes of unwanted reflections in the optical path. The photodetector was an APD, which was connected to an image reconstruction and display unit.

Fig. 26

Example of a single-fiber confocal microscope system (reproduced with permission from Ref. 139).

5.6.

Optical Fibers in Dentistry

The type of optical fiber used in dentistry depends on the specific application, such as tooth erosion studies, gum disease treatments, early dental caries detection, or ablation of hard dental tissues. An important reason for the use of optical fiber-based diagnostic and therapeutic dental tools is that the fibers allow noninvasive and nondestructive evaluation of dental tissues. The use of optical fibers also enables a wide range of optical power levels at different wavelengths to be delivered to specific viewing or treatment areas in an oral cavity.^141–146

Conventional optical fibers can be used for procedures such as early dental caries detection, oral cavity and caries imaging, and discrimination of periodontitis and gingivitis from healthy gingiva. PCFs and conventional hollow-core fibers are commonly used for delivery of optical power from $1.06\text{-}\mu \mathrm{m}$ Nd:YAG, $2.94\text{-}\mu \mathrm{m}$ Er:YAG, and $10.6\text{-}\mu \mathrm{m}$ ${\mathrm{CO}}_2$ lasers for ablation of hard dental tissues.

Figure 27 gives an example of a dental application of the fiber-based confocal microscope techniques discussed in Sec. 5.5. In this setup, 488-nm light is focused into an SMF for delivery to the probe. The three mirrors, the beam splitter, and an optical filter establish a path from the probe to a PMT for the 508-nm fluorescent signal emitted by the target tissue. Within the probe, the SMF is attached to a miniaturized tuning fork that is made to vibrate and rotate, thereby providing a $390\mu \mathrm{m}\times 390\mu \mathrm{m}$ two-dimensional scan of the laser beam on the tissue. The objective lens in the probe is used both to focus the excitation laser into the tissue and to collect the fluorescence signal for transmission to the PMT.

Fig. 27

Example of a confocal laser endomicroscope system for fluorescent imaging of the oral cavity (reproduced with permission from Ref. 145).

5.7.

Photodynamic Therapy

PDT is a treatment modality for diseases and cosmetic conditions ranging from acne to cancer that involve fast-growing or pathological cells or bacteria.^147–151 Key uses of PDT include age-related macular degeneration, arteriosclerosis, Barrett’s syndrome, and cancer therapy. The PDT process uses laser light in conjunction with a light-sensitive drug that is administered before exposure to the light. The selected laser emits at a wavelength that corresponds to the absorption peaks of the photosensitive agent. The photosensitive agent accumulates in malignant cells and its role is to absorb the light energy, which then is transferred to these cells. The process of absorbing energy from the light results in the formation of singlet oxygen, which then causes chemical destruction of the tumor cells.

Currently, the most commonly used light source is a laser diode that is selected to emit at the absorption wavelength of the photosensitive agent. Depending on the drug used, this wavelength can range from 620 to 760 nm, which is a region of high penetration depth into tissue (see Fig. 1). An advantage of using laser diode sources is that the light can be transmitted directly from the laser to an internal treatment site (such as the esophagus, the stomach or intestines, the urinary bladder, and the lungs) by means of optical fibers. The exit port of the fiber can be a flat or an angle-polished bare end or some type of diffuser that can be attached to the fiber tips for homogeneous spot or cylindrical light distribution. The diffuser could be a microlens or a short length of side-emitting fiber. In addition to using fibers to deliver the therapeutic light, optical fibers are also used to monitor the therapeutic light fluence, the concentration of the tumor sensitizer, and the oxygen levels produced by the PDT process.

A wide range of optical fiber types can be used in PDT depending on the medical application and the desired light distribution pattern. Bare-ended fibers are typically silica glass fibers with large NAs (normally $<0.60$) and core diameters ranging from 200 to $1000\mu \mathrm{m}$. These fiber types allow the setting of a well-defined distance between the fiber end and the tissue being treated or examined and, thus, can enable precise irradiations to be administered or accurate values of tissue properties to be collected.

As shown in Fig. 14, a cylindrical light diffuser is typically a plastic or a specially constructed glass side-emitting fiber that produces a radial light pattern, which is homogeneous along its entire length. Standard cylindrical diffuser fiber sections range from 10 to 70 mm in length, have a nominal core diameter of $500\mu \mathrm{m}$, an NA of 0.48, and a minimum bend radius of 10 mm. Their main uses are for intraluminal PDT, such as in arteries or veins, in the esophagus, in gastrointestinal tracts, in urinary tracts, or in bronchi in the lungs. The side-firing (or sideward irradiating) tip shown in Fig. 17(c) can be used with a needle (e.g., 0.5 to 0.9 mm diameter) for treating interstitial tumors. Intracavity or superficial tumors can be treated directly with light emerging from the end of a fiber with a core diameter ranging from 200 to $1000\mu \mathrm{m}$.

5.8.

Optical Fibers Used in Biomedical Sensing

Numerous embodiments of optical fiber structures using different materials and configurations have been investigated to form biosensors.^5,152–159 A biophotonic detection process can involve the sensing of a change in a physical parameter, such as the refractive index of a material or optical power loss due to fiber movement or microbending. For example, an optical fiber-based device can be utilized to sense perturbations in the evanescent field of the propagating modes in an optical fiber. Technology platforms that have been examined for waveguide biosensors include various surface plasmon resonance (SPR) configurations, Mach-Zehnder and other interferometers, and photonic crystal-based devices.

The attractions of such sensors are their small size, excellent integration capability, and relatively high sensitivity to detect analytes (substances being identified and measured). Figure 28 shows the basic operation of one type of fiber biosensor. This class of sensors uses absorbance measurements to detect any variations in the concentrations of substances of interest that absorb a specific wavelength of light. First, the fiber is coated with a biorecognition material that is referred to as an antibody. This antibody has the characteristic of being able to absorb or capture a specific analyte or antigen. The absorption process changes the effective refractive index of the fiber coating at a given wavelength of light. As noted in Fig. 5, because part of the energy of a propagating mode is in an evanescent field that travels in the cladding or coating, a change in the refractive index of the coating will cause a slight perturbation in the mode near the fiber-coating interface, which induces a variation in the optical power level in the fiber core. Now, the variation in the amount of light transmitted through the fiber at that wavelength can be related to the concentration of the absorbed analyte. This methodology has been developed for sensing conditions such as glucose levels, pH levels, oxygen levels, antibodies, and pressure.

Fig. 28

Analyte sensing via an optical fiber waveguide evanescent field perturbation.

A variation on this class of antibody/antigen fiber sensor is shown in Fig. 29. The operational method is based on detecting fluorescence emitted from the addition of a detection antibody that can be made to bind to the absorbed antigen. The key characteristic of the detection antibody is that it will fluoresce at a specific excitation wavelength. In this sensor, first a fiber is coated with an antibody and then immersed in a medium containing antigens. Next a layer of the detection antibodies is attached to the captured antigens. Thereby, a sandwich structure is formed consisting of the basic antibody layer, the antigen layer, and the detecting antibody layer. After the detection antibody is attached to the antigen, a laser light is sent through the fiber, causing the detection antibody to fluoresce. The degree of fluorescence produced in this process is then a measure of the concentration of the antigen.¹⁶⁰

Fig. 29

Fiber-based biosensor that uses a fluorescing antibody attached to the antigens.

As another example, Fig. 30 shows the operational scheme of an optical fiber-based glucose sensor that uses a fiber loop ringdown technique.¹⁶¹ First, glucose oxidase is deposited on the surface of a section of SMF ($\sim 16\mathrm{cm}$ long), which functions as the sensor head. When the coated sensor head is exposed to a glucose solution, the glucose oxidase reacts with the glucose and generates gluconic acid. This interaction results in a change in the refractive index around the surface of the sensor head. A measure of the glucose concentration is obtained by comparing the differences in the ringdown times before and during the exposure event.

Fig. 30

Concept of an evanescent-field fiber loop ringdown glucose sensor system (reproduced with permission from Ref. 161).

Some common optical fiber-based pressure sensors function by measuring variations in backreflected light, changes in light levels coupled between two fibers, or losses due to changes in microscopic bends in the fiber axis.^162,163 The variations in the light intensities can then be directly correlated to changes in pressure. Two methods among many are shown in Fig. 31. In Fig. 31(a), light emerges from a fiber and is reflected back toward the fiber by a mirror or a diaphragm located at a distance $d$ from the end of the fiber. Part of the reflected light is partially captured by the same fiber core from which the light emerges and is transmitted back to a photodetector that senses the light intensity. If the pressure on the mirror or diaphragm increases or decreases, the distance $d$ will decrease or increase correspondingly. Thus, the reflected light captured by the fiber will increase or decrease, respectively. The variation of the light intensity at the photodetector, thus, is a measure of the change in pressure. In Fig. 31(b), light emerging from a fixed fiber is coupled into a second movable fiber separated a distance $d$ from the fixed fiber. The second fiber can move when there is a change in pressure, thus changing the distance $d$ between the two fiber ends. If the distance $d$ increases, then less light will be coupled into the second fiber and vice versa. Again, the variation of the light intensity at the photodetector is a measure of the change in pressure.

Fig. 31

Two examples of simple optical fiber-based pressure sensors: (a) reflecting mirror or diaphragm and (b) moving acceptor fiber.

PCFs are being applied in several categories of biosensors. One reason is that the air holes in a PCF serve as natural chambers for a liquid sample to enter. This feature significantly reduces the volume of a sample that is typically required. The infiltration of liquid into the PCF can be realized by simply immersing the fiber end in the liquid.^164–166 However, afterward if the liquid needs to be removed from the PCF, which often happens in multistep biochemical processes, vacuum or pressure needs to be applied to the fiber end. Different selective filling techniques have been reported using a UV polymer,¹⁶⁷ a femtosecond laser,¹⁶⁸ a focused ion beam,¹⁶⁹ or even normal glue.¹⁷⁰ The sensing principle for the analyte solution inside the PCF can be based on an FBG, interferometry, mode coupling, evanescent field monitoring, or observation of a bandgap shift.

Another fiber sensor type that uses a PCF as the sensing element is based on intermodal interference between the forward-propagating core and the cladding modes in a PCF.¹⁷¹ The basic configuration of a two-beam interference method uses the core mode as the reference beam and the cladding modes as the sensing beam. An example of such a PCF-based interferometer is shown in Fig. 32. Here, a short length of PCF is sandwiched between input and output SMFs (such as the conventional industry standard G.652 fibers) to form an inline core-cladding intermodal interferometer. Different higher-order cladding modes can be excited in the PCF from the fundamental mode (called the ${\mathrm{LP}}_{01}$ mode) in the SMF by having abrupt fiber tapers formed at the interfaces between the PCF segment and the single-mode input and output fibers.

Fig. 32

Concept of an inline core-cladding intermodal interferometer based on the mode-field mismatch method (reproduced with permission from Ref. 171).

When light from a broadband source is sent through the device, the interference between the core and cladding modes in the PCF sensing section will result in a transmission-versus-wavelength interference pattern. Since the higher-order mode is a cladding mode of a glass-air waveguide, any change of the outer refractive index caused by immersion of the PCF sensing segment in a liquid will change the propagation constant of the higher-order mode and, consequently, will cause a shift in the interference fringes.

Figure 33 illustrates an application of this method for the design of a compact and highly sensitive biosensor.¹⁷² In this design, conventional SMFs were arc-fusion spliced to both ends of an $\sim 3\text{-}\mathrm{mm}$-long segment of PCF. The air-hole arrangement of the PCF is shown in Fig. 33(a). As shown in Fig. 33(b), the air holes were completely collapsed in the splice region and an air bubble was formed as a result of air being pushed out during the splicing process. As shown in Fig. 33(c), this air bubble acts as a diverging lens, which causes a substantial amount of light emerging from the SMF to be refracted toward the outer cladding of the PCF, thereby exciting the PCF cladding modes.

Fig. 33

Design of a compact and highly sensitive PCF-based biosensor with an internal air bubble lens (reproduced with permission from Ref. 172).

Sensitivities of up to 320 nm per relative index of refraction unit were demonstrated for this sensor for evaluating bulk liquids in the biosensing refractive index range of 1.33 to 1.34. The external bioaffinity sensing of biomolecules has been demonstrated with this device using biotin as the receptor and streptavidin as the analyte.

5.9.

Health Status Diagnostics and Monitoring

A highly successful precise wavelength selection component developed in the telecom industry is an FBG. As Sec. 4.2.1 describes, an FBG is a narrowband reflection-grating filter that is constructed within an optical fiber core. For biophotonics applications, an external force, such as a strain induced on the device by the weight of a person, will slightly stretch the fiber, thereby changing the length of the FBG and, thus, changing the reflected wavelength. To measure the induced strain, the FBG sensor is typically glued to or embedded in a specimen that responds to the external strain. The one precaution that needs to be taken when using such a sensor is to realize that the FBG is temperature sensitive. Thus, either the substrate on which the FBG is glued has to be temperature insensitive or some type of temperature compensating method has to be deployed along with the strain sensor.

A wide selection of biosensor applications to diagnose or monitor the health status in a person has been realized using FBGs. Among these are the following:

More details for two of these examples are given below. Both applications used wavelength division multiplexing techniques for simultaneous querying of an array of FBG sensors along a single fiber line.

5.9.1.

Smart-bed FBG system

One implementation of an array of FBG sensors is a smart-bed system for healthcare applications. In this particular case, the smart-bed setup consisted of 12 FBG sensors with different Bragg wavelengths cascaded along a single fiber.^175,176 The line of sensors was mounted on the surface of a bed to form a 3 by 4 matrix array, which then was covered by a standard mattress, as shown in Fig. 34. Each FBG sensor was specially packaged into an arc-shaped elastic bending beam using a carbon fiber reinforced plastic material, which ensures excellent sensitivity and good linear translation from a lateral force exerted on the apex of the sensor into axial strain of the FBG when a subject is on the bed. Thereby, patient movements and respiratory rates cause different pressures on individual FBG sensors, so that changes in the Bragg wavelength of specific FBG sensors allow the monitoring of both healthy and abnormal conditions of a patient.

Fig. 34

FBG bed sensor array for patient condition monitoring (reproduced with permission from Ref. 176).

5.9.2.

Distributed FBG-based catheter sensor

A second example of an FBG-based biosensor is a catheter that does distributed in vivo sensing of pressure in the gastrointestinal tract.¹⁸¹ The catheter was formed from a serial array of 12 FBG sensors that were fabricated into a continuous length of SMF. As shown in Fig. 35, each FBG was attached to a localized pressure-sensitive structure consisting of a rigid metallic substrate and a flexible diaphragm. Each FBG element was 3 mm long and had a full-width half-maximum spectral response of 0.6 nm for Bragg wavelengths spaced 1.3 nm apart in the 815- to 850-nm sensing range. The FBG sensor elements were spaced 10 mm apart, resulting in a catheter with a 12-cm sensing length. The device was designed to measure pressure changes between $-50$ and $+300\mathrm{mmHg}$, which adequately covers the range of pressures normally encountered in esophageal tracts. As shown in Fig. 35, a circulator was used to insert light from a broadband optical source into the sensor array. Optical signal variations returning from the sensor array reentered the circulator and were sent to an optical detector from the third port of the circulator.

Fig. 35

Data acquisition equipment to record pressure changes along the catheter. Inset: schematic of an FBG sensor (modified with permission from Ref. 181).

5.10.

Photorejuvenation

In addition to healthcare monitoring applications, optical fibers are also used in laser resurfacing or photorejuvenation for the treatment of certain skin conditions and for wrinkle removal. The dermatological conditions include skin atrophy, skin thickening, varicose veins, vascular lesions, unwanted hair, age spots, surgical and acne scars, and pigmented lesions. The photorejuvenation process uses intense pulses of light to ablate a skin area by inducing controlled wounds on the skin stimulating the skin to heal itself through the creation of new cells. Two lasers that have been historically used in free-space or direct-beam setups are the Er:YAG laser (emitting at 2940 nm) and the ${\mathrm{CO}}_2$ laser (emitting at $10.6\mu \mathrm{m}$).

To reduce the side-effects associated with large-area irradiation, techniques such as fractional lasers and fiber-coupled lasers have been introduced. In the fractional laser technique, the skin is irradiated with a dense array of pinpoints of light.^183,184 This process leaves healthy skin between the ablated areas so that more rapid healing can take place.

Transmitting laser light via an optical fiber allows a more precise delivery of this light to a localized skin area. The following are examples of fiber-coupled lasers, which are available for research on the basic cell photorejuvenation mechanisms^185–188 and for clinical use.^189–192 As discussed in Sec. 4, the particular optical fibers that are coupled to these light sources depend on the efficiency with which the fibers transmit light in the spectral emission region of the source.

5.11.

Microscope-in-a-Needle Concept

A number of minimally invasive needle-based optical imaging techniques that use optical fibers are being investigated. One method is to perform confocal fluorescence microendoscopy or OCT by means of an optical fiber placed inside small-gauge needles for optical imaging within soft tissues or lesions.^193–195 The basic technology goal is to put a microscope inside a standard hypodermic needle to enable three-dimensional (3-D) scanning. This technique enables surgeons to locate and diagnose very small tumor elements, thereby avoiding the need for more invasive surgery and allowing for a speedy patient recovery from the examination process.

Figure 36 shows a prototype of an ultrathin 30-gauge needle ($310\mu \mathrm{m}$ in diameter) with an enclosed SMF. The distal elements consist of a $260\text{-}\mu \mathrm{m}$ section of a no-core fiber (NCF) and a $120\text{-}\mu \mathrm{m}$ section of a graded-index (GRIN) fiber that is spliced to the end of an SMF. Following the GRIN section is another section of NCF, which is polished at an angle of 45 deg using a fiber connector-polishing machine. This angled NCF allows side illumination and viewing. A laser-drilling method is used to create a window in the side of the needle. This structure creates an elliptical output beam that has a full-width half-maximum transverse resolution $<20\mu \mathrm{m}$ over a distance of $\sim 740\mu \mathrm{m}$ in tissue of refractive index $\sim 1.4$. The peak resolution was at 8.8 and $16.2\mu \mathrm{m}$ in the $x$- and $y$-directions, respectively.

Fig. 36

Design and images of an ultrathin OCT probe. (a) Layout of the components. (b) Microscope image of the component assembly. (c) SEM image of the drilled hole. (d) Position of the hole in the needle. (Reproduced with permission from Ref. 194.)

5.12.

Lab-on-Fiber Concepts

The recent joining of nanotechnology and optical fiber technology has led to the lab-on-fiber concept.^196–199 In this currently evolving technology, major efforts include attaching patterned layers of nanoparticles either on the tip of a fiber or as a nanocoating over a Bragg grating written inside a microstructured optical fiber. When a patterned nanostructure is deposited on the tip of a standard optical fiber, localized surface plasmon resonances (LSPRs) can be excited by an illuminating optical wave because of a phase matching condition between the scattered waves and the modes supported by the nanostructure. Figure 37 shows an example in blue of the SPR wave associated with a nanostructure pattern for the condition when no analyte is covering the pattern on the end of a fiber. These LSPRs are very sensitive to the surrounding refractive index at the fiber tip. Thus, inserting the fiber tip into a fluid will cause the analyte liquid to cover the nanoparticle pattern. This action changes the refractive index of the nanolayer-fluid interface, thereby causing a wavelength shift in the LSPR peak due to a change in the phase matching condition, as shown by the red curve in Fig. 37. Thereby, highly sensitive physical, chemical, and biological sensing functions can be enabled.

Fig. 37

Example of the shift in the surface plasmon resonant peak when there is a relative index change at the tip of a fiber covered with a nonoarray pattern.