The majority of commercial sensors widely used in biomedicine are electrically active and hence not appropriate for use in biomedical applications, in particular, in high microwave/radiofrequency fields, ultrasound fields, or laser radiation associated with hyperthermia treatment, due to local heating of the sensor head and the surrounding tumor due to the presence of metallic conductors and corresponding electromagnetic interference, resulting in erroneous readings. Fiber-optic sensors can overcome these problems as they are virtually dielectric [19].

Today, most biomedicine or biochemical sensing takes place in laboratories. Tissue is invasively removed by biopsies and then analyzed in vitro. This is an invasive, time-consuming, and expensive process. Fiber-optic sensors have become a more attractive means because they can be integrated into hypodermic needles and catheters owing to their small size. Through the use of appropriate sensing method, they offer the capability to provide immediate in vivo monitoring of biochemical species and concentrations of interest in a minimally invasive manner.

Extrinsic fiber-optic sensor is an effective way for measuring biochemical and biomedicine parameters, in which some active material’s properties are modulated by molecular concentrations of interest. An FP cavity will modulate the light spectrum incident upon it as a function of the reflectivity of its two parallel surfaces, their separation, and the indices of refraction inside and outside of the cavity.

A miniature fiber pressure sensor based on FP WLI principle with a desired diameter of 550 p,m have been successfully developed by Pinet, Pham, and Rioux as shown in Figure 6.8 [20]. The sensor comprises a micromachined silicon diaphragm membrane adhered to a cup-shaped glass base as the pressure sensitive element. The FP cavity length varies as the environmental pressure changes. White light from a lamp is guided directly into the FP cavity, which modulates the signal with a low-coherence interference, thus coding the sensor cavity length.

Intra-aortic balloon pumping (IABP) therapy, which usually consists of inserting a catheter terminated by an inflatable balloon through the femoral artery, has been developed for over three decades. It has been one of the most efficient and popular technologies for life support especially when pharmacologic therapy fails or presents a high risk of mortality [21].

The experimental setup is presented in Figure 6.9 [20]. To test the sensor performance, a Bio-Tek Instruments’ blood pressure system

Miniature FP pressure sensor. (From Pinet, E., Pham, A., and Rioux, S. 2005. Miniature fiber optic pressure sensor for medical applications

Figure 6.8 Miniature FP pressure sensor. (From Pinet, E., Pham, A., and Rioux, S. 2005. Miniature fiber optic pressure sensor for medical applications: An opportunity for intra-aortic balloon pumping (IABP) therapy. In Voet, M., Willsch, R., Ecke, W., Jones, J., and Culshaw, B., eds. 17th International Conference on Optical Fibre Sensors, Bruges, Belgium. International Society for Optics and Photonics. vol. 5855, 234-237.

Experimental setup for pressure recording

Figure 6.9 Experimental setup for pressure recording.

calibrator Model 601A was used as a calibrated waveform generator. This system simulates different waveforms including typical aortic pressure waves in a sealed liquid environment (water). The pressure was measured with the commercial FOP-MIV miniature pressure sensor connected to a commercial PM-250 signal conditioner operating at 250 Hz. A real-time signal was visualized on a digital oscilloscope. Data were transferred to a computer for further analysis.

The sensor is immune to electromagnetic interference that can occur in a hospital environment or during a surgery. Due to the white light cross-correlation technology used in the PM-250 signal conditioner, the pressure readings are not affected by the binding or displacement of the optical fiber, because the system measures the interference position directly related to the EFFPI cavity length of the pressure sensor and not a light intensity level, which would have been affected by optical fiber binding. The performance of the pressure sensor is shown in Figure 6.10.

Figure 6.10b shows that the FOP-MIV sensor is responding to square pressure waveform with high fidelity and without damping effects such as commonly observed in fluidic pressure transduction of commercial IABP systems.

EFFPI sensing in medical or biotechnology is also a useful tool. In 2006, an FP pressure chip for medical application was invented by Belleville and coworkers [22]. The chip has a body including a surface. A diaphragm covers the surface and is affixed to the body,

Pressure fluctuations simulated by Bio-Tek 601A pressure waveform generator as recorded using an FOP-MIV pressure sensor with a PM-250 conditioner

Figure 6.10 Pressure fluctuations simulated by Bio-Tek 601A pressure waveform generator as recorded using an FOP-MIV pressure sensor with a PM-250 conditioner. (a) Aorta tachycardia 120 bpm (80/140 mmHg). (b) Square pressure 30 bpm (80/140 mmHg).

where the diaphragm defines another surface. The surfaces are separated by a distance and form an FP cavity. The body has another cavity for receiving an extremity of an optical fiber. A small quantity of adhesive is required to secure the extremity of the optical fiber, where the body comprises a Pyrex portion. In 2007, Donlagic et al. developed a new manufacturing method of the FP optical sensor for medical application [23]. A cavity is formed by the shaped fiber end of a multimode step index optical fiber. A diaphragm arranged by a shaped optical fiber end forms a reflecting surface of a FP resonator. With the special manufacturing method, the sensitivity of the sensor is increased, and the production cost is reduced. A multi-cavity FPI sensor was invented in 2007 by Zhang and coworkers [24]. An FP cavity is formed at the end of the sensor to measure and provide temperature information. This device provides highly sensitive detection of the selected material, for example, deoxyribonucleic acid (DNA) hybridization, and bacteria, for the study of the thin-film and for chemical and biological material detection, even at low concentrations and without a need for labeling materials of interest, and prevents interference of the label with a process, for example, chemical reaction.

A simple and efficient approach for H2-breath-test analysis based on FP microcavity has been reported in 2007 [25]. As shown in Figure 6.11, the microcavity is composed of a thin sensitive palladium layer exposed to hydrogen emission, and a GaP dielectric layer joined to a thick metallic mirror. The sensitive area consists only of the Pd film c. Insulating the substrate and the two films as indicated by the labels a and b, the optical properties of the sensor depend on the interaction between the exhaled hydrogen and layer c. In this configuration, layer a acts as a mirror. Its overall thickness is not significant above a certain value that provides maximum reflectivity and an essentially zero transmittivity. Reflectivity and transmittivity through this layer remain constant by increasing the thickness a above the

aforementioned value. One might be tempted to think that the sensitivity of the device might increase by simply increasing the active layer c. However, the sensitivity is also related to the spectral position of the FP cavity resonance, which is uniquely determined by the thicknesses of the alternating layers b and c. The sensitivity strictly depends on the spectral position of the FP resonance, which in turn depends on the permittivity values resulting for different hydrogen concentrations. An extensive examination of the available geometrical parameters allows one to consider a wide number of resonances located at different spectral positions.

In 2011, Fan and coworkers demonstrated an FP cavity label-free biosensor with integrated flow-through micro-/nano-channel, which takes advantages of the large surface-to-volume ratio for analyte concentration and high detection sensitivity and built-in fluidic channels for rapid analyte delivery [26-28]. Recently, a novel EFFPI pressure sensor fabricated with a lensed fiber and a polymeric diaphragm was proposed for application in the medical field such as bladder or intracranial pressure measurement with reliable low-pressure measurement [29]. The structure is shown in Figure 6.12. A lensed SMF and a polymeric diaphragm form the cavity of the EFFPI. Two beams, which are reflected from the end face of the lensed fiber and the gold- coated surface of the polymeric diaphragm, respectively, are interfered through the SMF. Interference spectra of the pressure sensor were obtained while changing external pressure, and the cavity length was calculated by taking inverse fast fourier transform (IFFT) to the obtained interference spectra.

So far, there have been commercial biomedical sensors based on the FFPI technology. FISO Technologies Inc. produces EFFPI sensors for clinical applications [20,30], which are really meaningful and useful for human health monitoring.

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