Skip to main content
NCBI home page
NIHPA Author Manuscripts logoLink to NIHPA Author Manuscripts
. Author manuscript; available in PMC: 2009 Oct 28.
Published in final edited form as: Ann Biomed Eng. 2009 Jun 3;37(8):1638–1645. doi: 10.1007/s10439-009-9718-x

Performance Characteristics of a New Generation Pressure Microsensor for Physiologic Applications

Patrick S Cottler 1, Whitney R Karpen 2, Duane A Morrow 2, Kenton R Kaufman 2
PMCID: PMC2768899  NIHMSID: NIHMS132279  PMID: 19495983

Abstract

A next generation fiber-optic microsensor based on the extrinsic Fabry–Perot interferometric (EFPI) technique has been developed for pressure measurements. The basic physics governing the operation of these sensors makes them relatively tolerant or immune to the effects of high-temperature, high-EMI, and highly-corrosive environments. This pressure microsensor represents a significant improvement in size and performance over previous generation sensors. To achieve the desired overall size and sensitivity, numerical modeling of diaphragm deflection was incorporated in the design, with the desired dimensions and calculated material properties. With an outer diameter of approximately 250 µm, a dynamic operating range of over 250 mmHg, and a sampling frequency of 960 Hz, this sensor is ideal for the minimally invasive measurement of physiologic pressures and incorporation in catheter-based instrumentation. Nine individual sensors were calibrated and characterized by comparing the output to a U.S. National Institute of Standards and Technology (NIST) Traceable reference pressure over the range of 0–250 mmHg. The microsensor performance demonstrated accuracy of better than 2% full-scale output, and repeatability, and hysteresis of better than 1% full-scale output. Additionally, fatigue effects on five additional sensors were 0.25% full-scale output after over 10,000 pressure cycles.

Keywords: Pressure, Fiber-optic sensor, Extrinsic Fabry–Perot interferometric (EFPI)

INTRODUCTION

The use of fiber-optic sensors for physiologic applications has great appeal based on fundamental characteristics, such as miniature size, immunity to electromagnetic interference, and patient isolation from electric current. Based on these characteristics, fiber-optic sensors can be used to monitor various physiological parameters in anatomical structures and in ways that other solid state sensors cannot be used. Fiber-optic sensors are based on the transmission of light through a wave guide and subsequent analysis of specific properties of the returned light, such as intensity, wavelength, phase shift, and/or spectral composition that have been altered directly or indirectly by the parameter to be measured. This flexible platform has been exploited for sensing numerous biomedical parameters, including pressure,4,9 temperature,11 shape and position,7 blood perfusion,3 gases and vapors, pH, presence of organics, specific enzymes, antibodies, DNA, bacterial detection,24 pathogens, biochemical reactions, and toxins.12

There are several fiber-optic-based pressure sensors, with varying performance characteristics that are currently available and have been used in a variety of physiologic applications (Table 1). This brief communication describes a next generation ultraminiature pressure sensor system that uses the extrinsic Fabry–Perot interferometric (EFPI) measurement technique16 and has improved potential for a wide range of medical applications. The first generation miniature fiber-optic physical sensor has been described and characterized previously.1,10,25 Size, accuracy, repeatability, sensitivity, resolution, and hysteresis and fatigue effects are critical performance criteria that were tested in this new sensor.

TABLE 1.

Characteristics of available fiber-optic pressure sensors.

Sensor Application Diameter Dynamic range Frequency
response/Sampling
rate		 Functional specifications
Camino® ICP monitor, Integra™ Lifesciences Corporation (Plainsboro, NJ) • Intracranial pressure8,9,13–15,17–20 1.35 mm −10 to +250 mmHg 100 Hz • Resolution of 1 mmHg
• Linearity and hysteresis of ±7% at 250 mmHg
• Drift of output of ±1–2 mmHg/day13 during indwelling
Fiber-optic-linked pressure sensor (FOLPS), BioTechPlex® (San Marcos, CA) • Airway pressure 1 mm ±7.35 mmHg 100 Hz • Resolution of 0.007 mmHg
FOH200-P, Omega® (Stamford, CT) • Various medical applications 2.5 mm 0 to 760 mmHg 20 Hz • Resolution of ±0.003% full scale
• Precision of ±0.01% full scale
FOP-F125, FOP-M260 FOP-MIV, Fiso Technologies, Inc. (Quebec, Canada) • Various medical applications 125 µm ±300 mmHg 250 Hz (125 Hz dual-channel) • Accuracy of 2–5 mmHg
260 µm Resolution of 0.3–0.4 mmHg
515 µm
Model 420, Samba Sensors (Sweden) • Blood pressure 420 µm −37.5 to +262 mmHg 1–40 kHz • Accuracy of 0.37 mmHg
• Intracranial pressure • Resolution of 0.08 mmHg
• Intervertebral disc pressure
• Bladder and urinary tract pressure
• Tracheal pressure, Gastrointestinal pressures
• Pressure in the uterus, oviducts, and ovarian follicles
First Generation EFPI sensor.1,10,25 Luna Innovations (Roanoke, VA) • Intramuscular pressure5 360 µm 0 to +250 mmHg 66 Hz (~10 Hz with 8 channels) • Accuracy: 1.5% full scale
• Repeatability: 1.5% full scale
• Linearity: 1.6% full scale
• Hysteresis: 4.5% full scale10
• Decreased sensitivity a higher signal frequencies and lower temperatures1
• Silica glass and polyimide ensured biocompatibility in muscle tissue25
Second-generation EFPI sensor. Luna Innovations (Roanoke, VA)—as described herein. • Intramuscular pressure23 280 µm 0 to +250 mmHg (functions to greater than 1500 mmHg) 960 Hz (240 Hz with 4 channels) • Accuracy: 1.8% full scale
• Repeatability: 0.48% full scale
• Hysteresis: 0.60% full scale
• Fatigue resistant over 10,000 cycles
• Biocompatible silicon-based materials
Open in a new tab

METHODS

Sensor Design

These diaphragm-based pressure microsensors function with the formation of a low-finesse Fabry–Perot cavity between the polished end-face of a fiber and a reflective surface that deflects with pressure (Fig. 1). Light emitted from a broadband source is passed through a single fiber, where a portion of the light is reflected off the fiber/air interface (R1). The remaining light propagates through the air gap between the fiber and the diaphragm and is reflected back into the fiber (R2). R1 is the reference (fixed) reflection while R2 is the sensing (variable) reflection. These two light waves interfere constructively or destructively based on the path length difference between the sensing reflection and the reference reflection. The resulting interferogram travels back through the optical fiber to the demodulation unit.10 Absolute gap information is contained in the frequency content of the returned signal (Fig. 1).22 Therefore, as the diaphragm deflects due to pressure changes, the interference pattern of the returned light can be demodulated to calculate the path length difference, which is then correlated with an absolute pressure level.

FIGURE 1.

FIGURE 1

Open in a new tab

Schematic of the EFPI measuring technique. (a) Light propagates through an optical fiber, a portion of the light is reflected by the polished end face of the fiber (R1) and the remaining light travels through the Fabry–Perot cavity and is reflected back by a diaphragm (R2). The optical path length changes as pressure deflects the diaphragm and can be determined through interferometric measurements of R1 and R2. (b) Resultant interferogram based on the reflections of R1 and R2 show the characteristic fringe pattern, with a unique frequency component that increases as the optical path length increases.

The FiberPro 2™ optical demodulation system is capable of accurately measuring a path length change of less than 1 nm. Therefore, to obtain the desired resolution of 0.5 mmHg, the diaphragm deflection profile must be at least 2 nm/mmHg in the desired dynamic pressure range (0–250 mmHg). Diaphragm construction, including size, thickness, and material properties, was designed to achieve the necessary sensitivity based on the DiGiovanni elasticity equation6 for deflection of a rigidly supported diaphragm (1). The model for the rigidly supported diaphragm was chosen due to the fact that it would provide the least sensitive estimate of sensor function.

P=[163(1−μ2)(Eh3a4)]y+[(7−μ)3(1−μ)(Eha4)]y3 (1)

where P represents pressure, y the deflection, a the radius, h the thickness, µ the Poisson’s ratio, and E the Young’s modulus.

Photoetching of silicon-based materials was used to fabricate the sensor heads to these desired specifications, though the construction was not truly a rigidly supported diaphragm. In addition, elasticity analysis was performed on the sensor body to guide the dimensions to ensure deformation of the body would be negligible compared to the diaphragm deflection.

Sensor Calibration

The pressure microsensors were calibrated over the defined dynamic range of 0–250 mmHg by utilizing a digitally controlled hydraulic pressure regulation system (Ruska Model # 7250, GE Sensing, Houston, TX) to systematically step through the pressure profile. Pressure microsensors were placed in a sealed and water-filled pressure chamber (approximate volume: 4 mL) that was in series with pressurized air and the pressure regulation system and maintained in an isothermal ice water bath. The pressure regulation system was programmed to step the pressure through two successive cycles from 0 to 250 mmHg and back to 0 mmHg in steps of 25 mmHg, with a resolution of 0.05 mmHg. Each discrete pressure level was held steady for 20-s prior to advancing to the next pressure level. The 20-s dwell time for each pressure would be reset if the value at any time fell outside of the 0.05 mmHg tolerance that was specified. Data was logged with the custom software for the FiberPro 2™ (Luna Innovations) demodulation system, and used to generate calibration curves for each sensor (pressure vs. sensor response).

Sensor Performance

A custom-designed pressure chamber that was mounted on an ElectroForce 3200 Test Device (Bose Corp., Eden Prairie, MN) was used to test the microsensor performance. The test device was controlled using WinTest Control software (Bose Corp., Eden Prairie, MN). The pressure chamber was filled with degassed water to prevent errant readings caused by air bubbles on the sensor. Actuation of the test device crosshead provided pressure changes of 0–250 mmHg (gage). The microsensor was placed in the pressure chamber and subjected to a 0.02 Hz sawtooth signal that cycled in a linear fashion at the rate of 8.8 mmHg/s from ambient to 250 mmHg and then decreased to ambient pressure for five cycles. The chamber pressure was measured with a NIST Traceable 500 mmHg Sensotec FPG gage pressure transducer (Sensotec, Columbus, OH) that was accurate to 0.25% full scale. The chamber pressure and microsensor pressure were collected simultaneously at 10 Hz.

Testing was performed to determine the transducer response performance characteristics of accuracy, repeatability, and hysteresis. The output of the pressure microsensor (EFPI gap deflection) was plotted as a function of the NIST Traceable reference pressure. The algebraic difference between the “measured” pressure as indicated by the pressure microsensor and the “true” NIST Traceable reference pressure was defined as the transducer’s error. Accuracy was defined as the ratio of the maximum error to the full-scale output (FSO). Repeatability of the transducer was defined as the ability of the transducer to reproduce an output reading when the same measurand value was applied to it consecutively, under the same conditions while the pressure was increased from ambient to 250 mmHg. Hysteresis was defined as the maximum difference between any pair of output readings obtained during a pressure cycle of the transducer when the measurand level was first increased and then decreased throughout the entire pressure range of the transducer. All performance characteristics were specified as a percentage of FSO. Nine sensors were tested.

Fatigue Testing

To investigate fatigue, sensor response was evaluated as a function of cycling pressures. Three standard pressure response curves were generated with the Ruska pressure regulation system, as described (0–250 mmHg) on five individual pressure microsensors before and after being exposed to 1,000 and 10,000 cycles of a 0–300 mmHg, 1 Hz square pressure wave (pre-cycle, post 1,000 cycles, and post 10,000 cycles). Composite response curves for each sensor were obtained by combining and fitting the three pre- and post-cycle data sets. The residuals (difference between each measured and the composite fitted value) were calculated for each individual sensor at the pre-cycle and each of the two post-cycle conditions and analyzed to determine the influence of fatigue. The residuals were analyzed statistically for normality, trends or patterns, and equality of variance. The standard error of the measurement was defined by calculating the pooled standard deviation of the residuals for all the data (3 curves each for 5 sensors) and representing that variability as a percentage of full scale.

RESULTS

Sensor Design

The silicon-based sensor head has the dimensions of 280 µm in diameter and a length of 750 µm, with a wall thickness of approximately 55 µm (Fig. 2). Functionality of the sensor is based on the deflection of a reflective diaphragm. The diaphragm geometry (diameter and thickness) was defined by fabrication specifications and verified with SEM images of the sensor head. By using the rule of mixtures and the individual properties of the silicon-based materials utilized in fabrication, an effective Young’s modulus of 111 GPa and an effective Poisson’s ratio of 0.20 were calculated for the sensor diaphragm. Figure 3 shows the theoretical deflection of a rigidly supported diaphragm using these dimensions and properties, as calculated with Eq. (1). For comparison, the theoretical deflection of a simply supported diaphragm has also been calculated and included.

FIGURE 2.

FIGURE 2

Open in a new tab

Magnified image that shows a human hair compared to the pressure sensor mounted at the end of a 135 µm diameter single-mode optical fiber.

FIGURE 3.

FIGURE 3

Open in a new tab

Comparison of theoretical deflections of the microsensor diaphragm compared to a representative calibration curve of the described sensor. The sensor demonstrated significantly greater sensitivity over the dynamic range of 0–250 mmHg compared to the theoretical deflection of a rigidly supported diaphragm. However, the model for a simply supported diaphragm shows significantly greater deflection than the sensor response. Design of the sensor leads to a hybrid of a simply supported and rigidly support diaphragm.

Sensor Calibration

The sensor function data was best fit by a fourth-order polynomial. Pressure microsensors exhibited excellent precision with an average sensitivity of 3.7–4.0 nm of deflection per mmHg of pressure, over the range of 0–250 mmHg. Therefore, due to the fidelity of the sensor and measurement system, the theoretical resolution is approximately 0.25 mmHg. In addition, mechanical analysis of the sensor body loaded with pressure on the diaphragm face resulted in a theoretical deformation of 4 × 10−4 nm per mmHg of pressure.

Sensor Performance

Characterization results showed that the microsensor has excellent accuracy, repeatability, and minimal hysteresis. Based on sensor calibrations, pressure microsensor responses correlated very closely with the NIST Traceable pressure standard. The average accuracy was better than 2% FSO, and repeatability and hysteresis were better than 1% FSO (Table 2). Individual sensors were shown to perform with similar functional characteristics, as demonstrated by the low standard deviations.

TABLE 2.

Pressure microsensor performance (n = 9).

Mean* Standard deviation
Accuracy 1.79 0.90
Repeatability 0.48 0.29
Hysteresis 0.60 0.18
Open in a new tab
*

Values reported as a percentage (%) of full-scale output (FSO).

Fatigue Testing

Analysis of the response curves generated before and after 1,000 and 10,000 cycles of a 0–300 mmHg square wave revealed that there was no correlation of the residual mean and variance to fatigue cycles (Fig. 4). Additionally, the standard error of these sensors based on the pooled standard deviation from all the sensor fatigue tests was calculated to be 0.25 ± 0.10%, FSO.

FIGURE 4.

FIGURE 4

Open in a new tab

Typical calibration curves for the pressure microsensor showing the deflection of the diaphragm over 250 mmHg before and after 1000, and 10,000 cycles of a 0 and 300 mmHg square wave. Each curve represents the average of two cycles of pressure stepped up from 0 to 250 mmHg and then back to 0 mmHg. Error bars of ±1 standard deviation are shown. A fourth-order polynomial best fits the calibration data. Additional residual analysis demonstrated that there were not any effects in sensor response over the repeated cycling.

DISCUSSION

Using a broadband light source for Fabry–Perot interferometric, analysis of an optical path length has been shown to be very useful in the design and function of this new pressure microsensor. The interferogram of the two reflections (R1 and R2) results in a characteristic fringe pattern based around the center wavelength of the broadband light source with a frequency component proportional to the optical path length (Fig. 1). By performing a frequency analysis on the resultant interferogram, an absolute gap measurement between the end of the fiber and the diaphragm can be made. Utilizing a broadband light source and not a single wavelength, calculations are simplified and results in a unique solution.

This new microsensor, with its size and functionality, represents a significant improvement over previous fiber-optic pressure sensors and opens the door to numerous potential physiologic applications. The Di-Giovanni equation (Eq. 1) for deflection of a rigidly supported diaphragm was used as a conservative guide to determine the necessary sensor dimensions and properties needed to achieve the desired minimum sensitivity. Through an iterative design process, these sensors have been able to achieve sensitivities that exceed the desired minimal response of 2 nm/mmHg. Not surprisingly, each of the analyzed sensors demonstrated a sensitivity that is greater than the theoretical response for a rigidly supported diaphragm (Fig. 3). The variation in sensitivity from the theoretical sensor response could be attributed to many factors. Based on the fabrication process of the sensor, the diaphragm is not anchored to the body in a perfectly defined rigid support. As also seen in Fig. 3, the theoretical deflection of a simply supported diaphragm under load is significantly greater than a rigidly support diaphragm. The DiGiovanni elasticity equation6 for deflection of a simply supported diaphragm is described in Eq. (2), variables are consistent with Eq. (1). Design of this new sensor is closer to a rigidly supported system; however, it is more accurate to describe the support as a hybrid of both. Therefore, it is reasonable for the calculated sensitivity to between a rigidly and simply supported system.

P=[163(1−μ2)(Eh3a4)]*[1+μ5+μ]y (2)

Additional variability can also be caused by differences between the measured and actual thickness (h) and radius (a) of the diaphragm, which have great influence on the deflection in response to pressure. The photoetching process that was utilized to fabricate these sensors has stated tolerances of ±7%, which can account for the majority of this deviation. Other etching issues, such as imperfect etch angle could account for the remaining variability. Lastly, the bulk material properties (Young’s modulus and Poisson’s ratio) were calculated by a weighted average of the individual material properties and the actual material properties of each sensor head may vary from the calculated average. Despite the variation between measured and theoretical function, each sensor was shown to be very robust in its accuracy, repeatability, fatigue, and hysteresis characteristics.

Even though the entire transducer is exposed to the varying pressures during the calibration procedure, mechanical analysis of the sensor body was performed to account for any deformation of the sensor body under pressure. If the sensor is loaded with pressure on the diaphragm face, the resulting deformation was calculated to be approximately 4 × 10−4 nm per mmHg of pressure. Thus, if the analysis was expanded to include buttressing pressure on the transducer side, the deformation would be even less. However, this measurement is multiple orders of magnitude smaller than the noise floor of the FiberPro 2™ demodulation system and not detectable. Therefore, this sensor has achieved its desired design specifications and is well suited for physiologic pressure measurements.

The transducer heads for this new microsensor are fabricated with a photoetching process. Therefore, the production variance has been significantly decreased with the improved performance characteristics compared to first generation sensor that were fabricated individually in a multiple step process from several materials (fused silica glass, polyimide, and aliphatic amine-cured mineral-filled epoxy adhesive) (Table 1).1 Controlled photoetching also allows for significantly easier modification of parameters and dimensions of the sensor transducer to achieve different sizes, sensitivities, and operating pressure ranges for specific applications, greatly increasing the utility of the sensor technology. Sensor components are made of glass and silicon-based materials; therefore, biocompatibility in muscle and connective tissue is also likely to be maintained.21 Another important characteristic of the sensor for biomedical applications is the ability to sterilize for use in vivo. Gas plasma and ethylene oxide have been shown to be appropriate techniques to sterilize these sensors prior to use.

Along with the improved fabrication process, the performance characteristics of this new generation fiber-optic pressure sensor make it ideal for numerous physiologic applications, where low pressure and small magnitude changes need to be accurately measured. Mechanical design of the sensor has resulted in excellent sensitivity, with theoretical resolution of less than 0.5 mmHg over the range of 0–250 mmHg, and accuracy, repeatability and hysteresis of less than 2% FSO. Closely related to the repeatability of the sensor is any effect of fatigue from standard use. Sensor response curves demonstrated excellent stability over repeated pressure cycling (Fig. 4). Based on the residual analysis of sensor response as a function pressure cycling, fatigue does not lead to a change in sensor function. The observed variability of the sensors in the fatigue tests was very small and standard error of measurement was calculated to be 0.25 ± 0.10% FSO for all sensors analyzed. Therefore, these sensors are stable for continuous measurements of small pressure levels and increments.

This new sensor design also lends applicability for direct physiologic pressure measurements. This ultraminiature pressure sensor is 280 µm in diameter and 750 µm in length and is functional over a large dynamic range (> 1500 mmHg, data not shown) that exceeds potential physiologic pressure levels to be measured. The fundamentals of fiber optics as a measurement technique makes the sensors immune to electromagnetic interference, and therefore they can be used effectively in the presence of magnetic fields, microwaves, and radio frequency radiation.

In conclusion, a newly designed sensor has been developed. The sensor is comprised of a monolithic silicon-based transducer head attached to an optical fiber, allowing for a significantly smaller sensor head (280 µm diameter and less than 750 µm in length), and improved control of fabrication. Additionally, an electronic demodulation system has been improved to be able to interrogate a single sensor at 960 Hz or up to four sensors simultaneously at 240 Hz, greatly increasing their physiologic utility. In some instances, a sheath has been placed over the sensor head with a specific geometry on the distal end. This was done to recess the diaphragm at the distal end (in relation to the measurement system) to ensure hydrostatic pressure measurements and create a barb at the proximal end in order to anchor the sensor and effectively make in vivo intramuscular pressure measurements.2,23 The size and design of the sensor make it ideal for catheter-based pressure applications, with potential to measure a pressure gradient down the length of the catheter at discrete points. Other potential applications for this next generation sensor are measurement of intracranial, intraocular, bladder, and respiratory pressures.

ACKNOWLEDGMENTS

The authors would like to acknowledge Dan Blevins for technical assistance in sensor fabrication and David Dausch from RTI International for assistance with the diaphragm design and photoetching processes. This work was supported by the National Center for Medical Rehabilitation Research (NCMRR) within the National Institute of Child Health and Human Development (NICHD) RO1 HD31476, “Microsensor for Intramuscular Pressure Measurement,” as well as the National Eye Institute (NEI) 5R44EY013902-03 “Miniature Non-Invasive IOP Measurement Device.” Its contents are solely the responsibility of the authors and do not necessarily represent the official views of the NCMRR, NEI, or NIH.

REFERENCES

  • 1.Chen S, Pislaru C, Kinnick RR, Morrow DA, Kaufman KR, Greenleaf JF. Evaluating the dynamic performance of a fibre optic pressure microsensor. Physiol. Meas. 2005;26(4):N13–N19. doi: 10.1088/0967-3334/26/4/N02. [DOI] [PubMed] [Google Scholar]
  • 2.Cottler PS, Blevins DC, Averett JS, Wavering TA, Morrow DA, Shin AY, Kaufman KR. Miniature optical fiber pressure microsensors for in vivo measurement of intramuscular pressure. Proc. SPIE. 2007;6433:643304-1–643304-10. [Google Scholar]
  • 3.Cottler PS, Skalak TC. Development of a clinically useful mechanical leech device that promotes flap survival in an animal model of venous-congested skin flaps. Ann. Plast. Surg. 2001;47(2):138–147. doi: 10.1097/00000637-200108000-00006. [DOI] [PubMed] [Google Scholar]
  • 4.Crenshaw AG, Styf JR, Mubarak SJ, Hargens AR. A new “transducer-tipped“ fiber optic catheter for measuring intramuscular pressures. J. Orthop. Res. 1990;8(3):464–468. doi: 10.1002/jor.1100080318. [DOI] [PubMed] [Google Scholar]
  • 5.Davis J, Kaufman KR, Lieber RL. Correlation between active and passive isometric force and intramuscular pressure in the isolated rabbit tibialis anterior muscle. J. Biomech. 2003;36:505–512. doi: 10.1016/s0021-9290(02)00430-x. [DOI] [PubMed] [Google Scholar]
  • 6.DiGiovanni M. Flat and Corrugated Diaphragm Design Handbook (Mechanical Engineering) New York: Marcel Dekker; 1982. [Google Scholar]
  • 7.Duncan RG, Raum MT. Characterization of a fiber-optic shape and position sensor. Proc. SPIE. 2006;6167:616704-1–616704-11. [Google Scholar]
  • 8.Gelabert-Gonzalez M, Ginesta-Galan V, Sernamito-Garcia R, Allut AG, Bandin-Dieguez J, Rumbo RM. The Camino intracranial pressure device in clinical practice. Assessment in a 1000 cases. Acta Neurochir (Wien) 2006;148(4):435–441. doi: 10.1007/s00701-005-0683-3. [DOI] [PubMed] [Google Scholar]
  • 9.Hollingsworth-Fridlund P, Vos H, Daily EK. Use of fiber-optic pressure transducer for intracranial pressure measurements: a preliminary report. Heart Lung. 1988;17(2):111–120. [PubMed] [Google Scholar]
  • 10.Kaufman KR, Wavering TA, Morrow D, Davis J, Lieber JL. Performance characteristics of a pressure sensor. J. Biomech. 2003;36:283–287. doi: 10.1016/s0021-9290(02)00321-4. [DOI] [PubMed] [Google Scholar]
  • 11.Kocincova AS, Borisov SM, Krause C, Wolfbeis OS. Fiber-optic microsensors for simultaneous sensing of oxygen and pH, and of oxygen and temperature. Anal. Chem. 2007;79(22):8486–8493. doi: 10.1021/ac070514h. [DOI] [PubMed] [Google Scholar]
  • 12.Leung A, Shankar PM, Mutharasan R. A review of fiber-optic biosensors. Sens. Actuators B Chem. 2007;125:688–703. [Google Scholar]
  • 13.Martinez-Manas RM, Santamarta D, de Campos JM, Ferrer E. Camino intracranial pressure monitor: prospective study of accuracy and complications. J. Neurol. Neurosurg. Psychaitry. 2000;69(1):82–86. doi: 10.1136/jnnp.69.1.82. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 14.Morgalla MH, Mettenleiter H, Bitzer M, Fretschner R, Grote EH. ICP measurement control: laboratory test of 7 types of intracranial pressure transducers. J. Med. Eng. Technol. 1999;23(4):144–151. doi: 10.1080/030919099294195. [DOI] [PubMed] [Google Scholar]
  • 15.Munch E, Weigel R, Schmiedek P, Schurer L. The Camino intracranial pressure device in clinical practice: reliability, handling characteristics and complications. Acta Neurochir(Wien) 1998;140(11):1113–1119. doi: 10.1007/s007010050224. discussion 1119–1120. [DOI] [PubMed] [Google Scholar]
  • 16.Murphy KA, Gunther MF, Vengsaskar AM, Claus RO. Quadrature phase-shifted, extrinsic Fabry-Perot fiber sensors. Opt. Lett. 1991;16(4):273. doi: 10.1364/ol.16.000273. [DOI] [PubMed] [Google Scholar]
  • 17.Piper I, Barnes A, Smith D, Dunn L. The Camino intracranial pressure sensor: is it optimal technology? An internal audit with a review of current intracranial pressure monitoring technologies. Neurosurgery. 2001;49(5):1158–1164. doi: 10.1097/00006123-200111000-00026. discussion 1164–1165. [DOI] [PubMed] [Google Scholar]
  • 18.Ravi R, Morgan RJ. Intracranial pressure monitoring. Curr. Anaesth. Crit. Care. 2003;14(5–6):229–235. [Google Scholar]
  • 19.Schurer L, Munch E, Piepgras A, Weigel R, Schilling L, Schmiedek P. Assessment of the CAMINO intracranial pressure device in clinical practice. Acta. Neurochir. Suppl. 1997;70:296–298. doi: 10.1007/978-3-7091-6837-0_92. [DOI] [PubMed] [Google Scholar]
  • 20.van den Brink WA, Haitsma IK, Avezaat CJJ. Evaluation of the Camino fiberoptic ICP transducer. Clin. Neurol. Neurosurg. 1997;99 Suppl 1:68. [Google Scholar]
  • 21.Voskerician G, Shive MS, Shawgo RS, von Recum H, Anderson JM, Cima MJ, Langer R. Biocompatibility and biofouling of MEMS drug delivery devices. Biomaterials. 2003;24(11):1959–1967. doi: 10.1016/s0142-9612(02)00565-3. [DOI] [PubMed] [Google Scholar]
  • 22.Wavering T, Meller S, Evans M, Pennington C, Jones M, Murphy K. Interferometric optical fiber microcantilever beam biosensor. Proc. SPIE. 2000;4200:10–16. [Google Scholar]
  • 23.Winters TM, Sepulveda GS, Cottler PS, Kaufman KR, Lieber RL, Ward SR. Correlation between isometric force and intramuscular pressure in rabbit tibialis anterior muscle with an intact anterior compartment. Muscle Nerve. doi: 10.1002/mus.21298. (in press) [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 24.Wolfbeis OS. Fiber-optic chemical sensors and biosensors. Anal. Chem. 2008;80(12):4269–4283. doi: 10.1021/ac800473b. [DOI] [PubMed] [Google Scholar]
  • 25.Yang C, Chunfeng Z, Wold L, Kaufman KR. Biocompatibility of a physiological pressure sensor. Biosens. Bioelectron. 2003;19:51–58. doi: 10.1016/s0956-5663(03)00131-3. [DOI] [PubMed] [Google Scholar]

RESOURCES