Measurement and monitoring of electrocardiogram belt tension in premature infants for assessment of respiratory function
© Ciaccio et al; licensee BioMed Central Ltd. 2007
Received: 17 September 2006
Accepted: 19 April 2007
Published: 19 April 2007
Monitoring of the electrocardiogram (ECG) in premature infants with conventional adhesive-backed electrodes can harm their sensitive skin. Use of an electrode belt prevents skin irritation, but the effect of belt pressure on respiratory function is unknown. A strain gauge sensor is described which measures applied belt tension.
The device frame was comprised of an aluminum housing and slide to minimize the device weight. Velcro tabs connected housing and slide to opposite tabs located at the electrode belt ends. The slide was connected to a leaf spring, to which were bonded two piezoresistive transducers in a half-bridge circuit configuration. The device was tested for linearity and calibrated. The effect on infant respiratory function of constant belt tension in the normal range (30 g–90 g) was determined.
The mechanical response to a step input was second order (fn = 401 Hz, ζ = 0.08). The relationship between applied tension and output voltage was linear in the range 25–225 gm of applied tension (r2 = 0.99). Measured device sensitivity was 2.18 mV/gm tension using a 5 V bridge excitation voltage. When belt tension was increased in the normal range from 30 gm to 90 gm, there was no significant change in heart rate and most respiratory functions during monitoring. At an intermediate level of tension of 50 gm, pulmonary resistance and work of breathing significantly decreased.
The mechanical and electrical design of a device for monitoring electrocardiogram electrode belt tension is described. Within the typical range of application tension, cardiovascular and respiratory function are not substantially negatively affected by electrode belt force.
Electrode belts with standard leads are commonly used devices for monitoring heart rate [1–4] and its variability , to assess physical activity , for instrumentation of ergometers used in sports medicine studies , aerospace medicine studies , for electrocardiogram (ECG) monitoring and defibrillation , as well as for electrical impedance tomography . Similarly, arrays of many electrodes fitted within electrode vests are used to monitor the spatial distribution of heart signals, and can be either strapped to the chest  or held in place by a pneumatic mechanism . The importance heart monitoring with electrode arrays to indicate regions of bioelectric abnormality where cardiac arrhythmias can originate is firmly established [13–16].
Premature infants confined to the neonatal intensive care unit (ICU) are routinely monitored with electrodes to obtain a record of the electrocardiogram (ECG). Application of standard electrodes to a premature infant's skin in the required locations can be tedious and time consuming . Furthermore, the skin of premature infants is delicate and sensitive, and application and removal of contact tape can injure the patient. To overcome this problem, ECG electrodes can be mounted on a rubberized belt and positioned at the required location for recording by wrapping the device about the subject's torso . Electrode belts can be used for both short-term and long-term ECG monitoring . By use of a belt rather than contact tape to position individual electrodes to the chest, skin irritation is avoided. There is also ease of electrode application and removal. However, premature infants typically suffer from diminished respiratory function, which may be complicated by an inappropriately large level of belt tension.
Very premature infants, presently born as young as 20 weeks gestational age , may be as light as 400–500 grams , and have chests approximately 5 cm in diameter. To monitor electrode belt tension would require construction of a miniaturized device that could be attached between the tabs that are used to fasten the belt. Herein we describe the design and implementation of such a device, which is then used to quantify electrode belt tension in a premature infant. We compare monitored respiratory parameters to the measured tension levels that are applied in the ICU.
To measure tension applied normal to the axis of the electrode belt, the device width should be significantly less than chest diameter. Further, to prevent application of a vertical torque during measurement, the frame thickness must be minimized. A stiff, lightweight frame material is needed to prevent bending of the frame and to minimize the weight on the infant's chest. The device must be electrically isolated from the infant's skin. The mechanical and electrical construction of the device took into account these considerations.
Device housing and transducers
Signal conditioning electronics
± 9 volt battery power supply
5 V, low power voltage regulator (provides stable excitation voltage)
LED; on asserts battery low condition
Wheatstone bridge zero adjust
calibration adjust to relate millivolts to grams of applied tension
The current drawn by the resistive circuit elements and voltage regulator provide an estimate of battery life according to the following equation: Icircuit = Ivoltreg + Iamp + Iactivearm + Izero + Ical + Ibatset
7.9 ma = 2.6 ma +3.8 ma + 0.7 ma + 0.7 ma + 0.1 ma + 0.1 ma
The current drawn by the entire circuit Icircuit in Eq. 1 is a summation of the current Ivolt reg drawn by the voltage regulator (reg), the current Iamp of the amplifier (amp), the current Iactive arm through the active arm of the Wheatstone bridge (though 2Rg), the current Izero of the zeroing element (zero), the current Ical through the calibration component (cal), and the current Ibat set through the battery set circuitry (batt set). The lifetime of the power supply to maintain the ≥ 6.2 volts required by voltage regulator for constant 5.0 V output can be calculated as: 0.450 ma·hours/7.9 ma ≈ 55 hours (2)
where the battery life of 0.450 ma·hours was obtained from the battery specifications. When there is less than 6.2 V at the regulator input (battery low condition), the excitation voltage to the Wheatstone bridge is no longer maintained at 5 volts, thereby introducing measurement error. The light emitting diode (LED) alerts the user to a battery low condition. LED off draws approximately no current, however LED on (battery low) was measured to draw an additional 10.2 ma (the forward current requirement). After balancing, if no tension is exerted on the leaf spring, the output will register at zero millivolts. As tension is exerted, a difference in resistance at each strain gauge is created that is translated into a potential difference across the half-active bridge (Fig. 6). The differential bridge voltage is reduced via the calibration adjustment to enable a 100 mV output to correspond to 100 grams of tension (100 g·980 cm/sec2 = 0.98 Newtons).
The piezoresistive transducing elements (Kulite Semiconductor Products, Basingstoke, England) have a stated strain gauge constant of G = 175 ± 5% at 75°C, and resistance R = 5000Ω ± 1%. The maximum strain for linearity for these elements is 1000 με (units are microstrain, which is strain expressed as parts per million). For a rectangular beam with centrally applied force, we calculated the beam thickness needed to maintain linearity. The maximum allowable strain before irreversible deformation can be calculated as follows. The strain is given by : S = ΔL/L = (3 • F • L)/(4 E b h2) (3)
where F is the applied force in Newtons, L is the beam length, b is the beam width (or height), h is the beam thickness, and E is Young's modulus of elasticity. Rearranging the previous equation, the maximum allowable force is: Fmax = Smax • (4 E b h2)/(3 • L) (4)
where the chosen values of parameters L and b were designed to minimize the size and weight of the device, and h was adjusted according to the required maximum force Fmax. From the initial experience with placing the electrode belt about the patient, normally applied levels of belt tension by clinical personnel in the ICU ranged from approximately 30 g–90 g. A maximum force that would be encountered about the belt due to patient breathing and motion might therefore be about 100 g of tension. To provide for maximum sensitivity with a margin of safety double the maximum estimated force to be encountered, we supposed that the maximum belt tension would be 200 g, and: Fmax = 200 g • 9.8 N/1000 g = 1.96 N (5) (5)
Rearranging Eq. 4: h = [(3 • L • Fmax)/(4 • E • b • Smax)]1/2 = [(3 • 1.56 cm • 1.96 N)/(4 • 20.3E6 N/cm2 • 0.244 cm • 0.001 cm/cm) ]1/2 = 0.021 cm = 0.21 mm (6)
which is the required thickness of the steel leaf spring for maximum sensitivity.
For a half-active bridge, the output is : Vout = Vin • ΔRg/2Rg (7)
where Rg is the strain gauge resistance, and the sensitivity is given in units of differential output voltage Vout per gram of tension per bridge input voltage Vin. The relationship between the gauge factor, strain, and resistive changes for a half-active bridge is : ΔR/R = G • ΔL/L = 175 • 5.25 με = 9.19 × 10-4 (8)
where G and ΔL/L are properties of the particular strain gage that was used, and με are units of microstrain. For 1 gram (gm) of tension, the sensitivity is Vout/1 gm tension/5 V input excitation: sensitivity = (5.0 V • 9.19 E-4/2)/gm tension/5 Vex = 2.30 mV/gm tension/5 Vex (9)
where T = tension and Vex is the excitation voltage of the bridge. Therefore, without adjusting the gain, a 100 gm input tension would be expected to generate a 230 mV output. The unit was tested for electrostatic discharge and electrical isolation prior to connecting it to the EKG belt on actual patients.
Clinical monitoring was done at Saint Peter's Medical Center in New Brunswick NJ. An infant's candidacy for monitoring required that an electrode belt was already being used for ECG recording, the subject was sleeping in the supine position, and that the infant's normal schedule of care was not interrupted. The patient's physician applied the electrode belt with tension monitor in place. A technician and an engineer assisted the physician with the device, data processing equipment, and readout. Tension was applied at 30 gm as measured by the tension gauge, and the infant's position was adjusted to allow approximately even distribution of the applied belt tension. After waiting two minutes for the infant and the device to adjust, respiratory function was sampled using a pneumotach (MAS Inc, Hatfield PA). The functions that were measured were: respiratory frequency, tidal volume, and work of breathing. The procedure was then repeated at 50 gm and 90 gm of tension. The mean and standard deviation of each measurement taken at 20 minute intervals over a 2 hour period were tabulated.
ζ = [(ln(x2/x1))2/(π2 + (ln(x2/x1))2]1/2 (10a)
fn = fd/(1 - ζ2)1/2 (10b)
where x1 and x2 are the step input and first maximum, respectively, measured x2/x1 from Fig. 7 was 0.78, and Td, the time from first to second maximum, was 2.5 ms from Fig. 7, and fd = 1/Td = 400 Hz. Thus from Eq. 10, ζ = .08 and fn = 401 Hz. The device was designed to measure the DC level of tension and low frequency components, which are well below the resonant frequency of this device.
Force was then applied to different portions of the frame with varying amplitudes and directions, while output of the tension gauge was tabulated. A 1–3 mV output was observed upon compression of the sensor with approximately 100 Newtons of force applied, which was beyond any force the housing would be expected to encounter. With light pressure the output remained at 0 mV.
The viscoelastic properties of the electrode belt and Velcro connections were considered by wrapping the belt/tension gauge about a cylinder 22 cm in circumference. One hundred grams of tension was applied by tightening the strap; the response indicates a similar viscoelastic time constant to the previous test (Fig. 8, open circles). Thus the material properties of the electrode belt would not be expected to influence the clinical measurement beyond the initial decrease in the output of ~10%.
The actual measured output was found to be 218 mV per 100 g tension (estimate was 230 mV – see Methods). However, for clinical use the output was adjusted through the calibration circuitry so that a 100 mV output reading corresponded to a 100 gm input.
Relationship between Tension Level and Respiratory Parameters
Resp frequency min-1
71.1 ± 3.7
75.7 ± 4.4
71.0 ± 4.1
Tidal volume ml/kg
6.0 ± 0.3
6.4 ± 0.4
7.8 ± 0.9
Min ventilation ml/min/kg
419 ± 20
480 ± 30
517 ± 41
Dyn compliance ml/cm H2O
.32 ± .02
.34 ± .02
.46 ± .07
Pul resistance cm H2O/L/s
186 ± 20
134 ± 11
185 ± 22
Work breathing g·cm/kg
24.1 ± 0.1
23.4 ± 0.2
36.9 ± 0.2
The clinical data showed some change in most respiratory functions during monitoring with three different levels of belt tension. In Table 1, respiratory frequency is higher than the rate for the normal term infant (60/min). The tidal volume increases slightly with increasing tension level but at all three levels it is in range for a normal term infant (6–8 ml/kg). Minute ventilation increased significantly with increased tension level, in tandem with the increase in tidal volume and the approximately flat respiratory frequency. The dynamic compliance also increase significantly with increased tension level, but was still far below the normal for a term infant (1.00 ml/cm H2O). Pulmonary resistance did not change significantly from 30–90 g of tension, but was still far above that of a normal term infant (40 cm H2O/L/s). There was a significant increase in the work of breathing from 30 g to 90 g tension. At all three levels work of breathing was in the normal range for a term infant (10–40 g • cm/kg).
In this study, the design and construction of an electrode belt tension gauge was described. The device was tested for mechanical and electrical response. It was then clinically tested using a preterm infant who was already connected to an electrode belt for monitoring the electrocardiogram in an ICU setting.
The tension gauge was constructed to minimize weight and encumbrance to the infant. The device was designed to convert the mechanical signal (electrode belt tension) to an electrical signal. Two strain gauges bonded to opposite sides of a leaf spring registered differing levels of resistance depending on the degree of tension imparted to the slide mechanism. During bench tests, the device exhibited a linear response to increasing tension levels within the range that would be expected during clinical measurement. The mechanical response to a step input was a second order function. The computed resonant frequency (400 Hz) was above the mechanical frequency range that would be expected to be encountered in a clinical setting.
For the creep tests, a constant load was applied during the test to allow for stiffness measurement . The bonding epoxy did exhibit some creep (~10%) over the course of a few minutes when a constant level of tension was applied. However, the electrode belt was connected to the device frame, not to the leaf spring, and therefore its tension level would be unaffected by the properties of the bonding epoxy used to anchor the strain gauges. Yet, epoxy was also used to cement the Velcro tabs onto the device frame. Creep in this epoxy over time would be expected to diminish the actual tension in the electrode belt to which it was directly connected. This response would be anticipated to affect the measurement (decrease in tension reading over time). However, during actual clinical measurement, the decrease in tension over time was limited to about 10% at each tension level. Epoxies which exhibit less creep when subject to a constant stress should be used in subsequent manifestations of this measuring device . One other difficulty was the nonlinear elastic response of the belt to differing levels of tension (Fig. 10). Thus with increased stretch, tension in the belt increased disproportionately. In terms of clinical recording, this would mean that there would be less compliance of the belt (greater force applied to the infant chest) for increases occurring at higher starting tension levels.
The clinical table (Table 1) and other results suggest that differing levels of tension used to emplace the electrode belt had some effects on respiratory function in this infant. At intermediate belt tension level (50 g), improvement in respiratory function occurred. The intermediate tension level may have stabilized the infant's chest wall, which underwent paradoxical breathing , in a manner that promoted more efficient breathing. Premature infants are known to breath paradoxically during rapid eye movement (REM) sleep, due to the very high level of compliance of their chest walls. In the case of the intermediate tension level, the chest wall may be supported and stabilized in a manner that promotes more efficient breathing. However, it may also be the case that the infant's long-term breathing response will differ from the short-term results depicted in the table. Additional subjects and a longer monitoring time will be required to know if these results will hold for a representative population.
Even distribution of tension about the electrode belt was probably not achieved during clinical measurement. Areas with greater tension were likely located near the tension measuring device, where the belt was connected above the infant's chest. Lesser tension would be expected at the back, where the weight of the infant would partially prevent even distribution of the applied tension. Therefore, we would expect the degree of tension to be less than the read value at some areas about the circumference of the belt. Segments of the belt with reduced tension would not impart as much force on the torso and therefore would contribute less to the effect of the belt on respiratory function.
Limitations and future directions
The piezoresistive transducers were bonded to a leaf spring with an epoxy that exhibited creep over several minutes time. This limited the accuracy of the measurement to determine whether electrode belt tension remained constant over time. Single recordings at varying levels of tension were made on one subject. Due to constraints in interaction with the patient, these recordings were done during a short period of time at one setting. It is uncertain whether the respiratory measurements at fixed levels of tension are time invariant. Supposing that these measurements are approximately time invariant, or that correction or normalization can be used to account for time-varying differences, a large population of subjects would still be required to determine the statistical significance of the effects of belt tension on respiratory function. Thus the approach that we have described is preliminary; other methodology may improve the accuracy of the measurements. Heart rate remained stable during the measurement interval; however, correlation of tension level and respiratory function to heart signals such as electrocardiogram and blood pressure would be useful to state more definitively whether the mechanical properties of the ECG belt affect the heart.
The instrument was not used to measure chest wall motion, but such information could improve understanding of the mechanism of paradoxical breathing. Although the tension gauge monitor was tested with an electrode belt attached to a premature infant, monitoring of belt tension would be useful in settings such as heart rate monitoring and variability [1–5], for sports and aerospace medical activities [6–9], as well as for electrical impedance tomography . Monitoring of tension would assure that any change in respiratory or cardiovascular function is not due to the mechanical constraint offered by the belt. For very long term monitoring, the power supply of the device would need to be upgraded. Since some electrode belts and vests apply tension across the chest in multiple directions, monitoring of each axis will likely be important for understanding respiratory and cardiovascular effects.
Dr. Ciaccio is a recipient of Established Investigator Award #9940237N from the American Heart Association.
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