Measuring respiratory parameters like the breathing
frequency or the tidal volume is essential in intensive care to ensure an
optimal and lung protecting ventilation. A common practice in artificial
ventilation of sensitive patients like infants or neonates is the use of
uncuffed endotracheal tubes in combination with continuous positive airway
pressure (CPAP). This comes with the disadvantage of an unknown leakage
making it difficult to detect spontaneous breathing or to measure the tidal
volume reliable. A novel non-obstructive method to determine respiratory
parameters as well as dynamic changes of thoracic parameters has recently
been presented and uses a pair of coupled UHF (ultra high frequency)
antennae. In this paper, a respective setup is investigated numerically
using finite difference time domain method and experimentally using an
artificial lung phantom. Both approaches show that the investigated method
seems capable of allowing a contactless triggering to synchronize natural
and artificial breathing. The results are compared to derive a better
understanding of influencing factors and opportunities for an optimisation.
Introduction
Medical respirators have become an integral part of nowadays intensive care.
In order to ensure an optimal ventilation that is lung protecting, essential
respiratory parameters like the breathing frequency or the tidal volume have
to be monitored reliably and continuously. The parameters are of significant
importance for controlling the respirator with respect to a correct pressure
and tidal volumes. Additionally, the detection and support of a spontaneous
inhalation is crucial – especially for sensitive patients like infants or
neonates. Due to their low tidal volumes, their relative high breathing
frequency and their sensitive skin, a precise measurement of essential
respiratory parameters is technically challenging, which leads to a reduced
control quality of the respiration parameters. The use of continuous
positive airway pressure (CPAP) together with uncuffed endotracheal tubes is
a common practice in artificial ventilation of premature infants in order to
prevent the development of respirator induced lung injuries, like volu- or
barotrauma (Mahmoud et al., 2015). The disadvantage of this practice is an
unknown time-varying leakage which makes a detection of spontaneous
breathing or the measurement of tidal volume very difficult. Besides flow
and pressure monitoring, other measurement methods like electrical impedance
tomography (EIT) are also inadequate since they require electrodes connected
to the sensitive skin or are of obstructive nature.
A promising approach to overcome the most important technical burdens of
measuring the respiratory parameters are electromagnetic modalities, because
they are leakage independent and unobtrusive. There are approaches measuring
the dielectrical changes of the thorax by capacitive means (Teichmann et al.,
2013; Kundu and Kumagai, 2013; Oum et al., 2008) or measuring induced currents in
the thorax by inductive means (Teichmann et al., 2013,
2015). Another approach is using radar-based methods were the displacement
of the chest due to respiration is measured either via ultra wide-band radar
or continuous wave radar (Leonhardt et al., 2018). In contrast to radar,
which measures reflection of the chest wall, recently a novel setup to
measure the transmission through the thorax was presented (Ringkamp et al.,
2020). With this setup dynamic changes of thoracic parameters due to
breathing are determined with a pair of coupled UHF (ultra high frequency)
antennae. This setup seems to be well suited for contactless triggering to
synchronize natural and artificial breathing, especially in case of infants
and neonates.
In this contribution, the electromagnetic aspects of this approach are
investigated in numerical and metrological manner. The aim of the numerical
field calculations is to get further insight into the influencing factors of
this setup. The pair of coupled antennae operates in the UHF band, which
covers the frequency range from 300 MHz to 3 GHz
or the wavelength range from 1 m to 100 mm
respectively (Meinke and Gundlach, 1992). While the distance
dant=90 mm between the two antennae seems to be
comparably small regarding the wavelength λ, it can be assumed
that the receiving antenna is not placed in far field region, but in the
radiating nearfield, the so-called Fresnel Region (Balanis, 2016). Thus, a
numerical model based on field calculations is necessary. This model is
underpinned with a defined experimental setup consisting of a ventilated
thorax phantom and a set of antennae. Special attention is paid to the
coupling between the antennae as well as to the resonance behaviour of the
coupled antenna pair with respect to the change of air volume in a lung
phantom.
Material and methodsUHF antenna approach
In this contribution, the physics of measuring respiratory parameters with a
pair of coupled UHF antennae is investigated with numerical methods. This is
necessary to get further insight into the wave propagation and occurring
effects. Typically, the receiving antenna of a wireless transmission system
is placed in the far field region of the transmitting antenna, where the
radiated electromagnetic (EM) field propagates as a transversal
electromagnetic (TEM) wave. This means, that the propagating wave has got
only transverse field components (perpendicularly orientated to the
propagation direction), but no field component in propagation direction.
Furthermore, the electric field E and the magnetic field H
are also orientated perpendicular to each other. The free space wave
impedance
Z0=EH=120πΩ≈377Ω
is a constant value in the far field region (Schelkunoff, 1938) and the
Poyting vector
S=E×H
has got only a component in propagation direction, since both, the electric
and the magnetic field strength have none in propagation direction.
In a first step, we will investigate whether the receiving antenna is placed
in the reactive near field region (Rayleigh region), the radiating near
field region (Fresnel region) or the far field region (Fraunhofer region).
Different definitions for the boundaries of the regions could be found in
literature, which have in common that the largest dimension of the antenna
Dmax and the operating frequency f or the corresponding
wavelength λ have to be taken into consideration, respectively. The
IEEE standard definition of the antenna regions defines the reactive near
field region in a distance smaller than λ/2π and the far
field region in a distance greater than 2Dmax2/λ (IEEE Standard #145, 1969), whereas Balanis mentioned
0.62Dmax2/λ and 2Dmax2/λ for the distances of the antenna regions (Balanis, 2016).
Nevertheless, these definitions are only valid for an antenna that is large
compared to the wavelength (Dmax>λ).
Simulation model of the used UHF antenna: The planar antenna is
realized as a slightly meandered dipole on a flexible PCB with the physical
dimensions 79 mm × 10 mm, specified for ISM applications in the frequency
range of from 863 to 870 and 902 to 928 MHz.
In our approach, we use two identical UHF antennae models based on a
commercially available meandered dipole antenna on printed circuit board
(PCB) with a physical length of 79 mm and a width of 10 mm. The chosen antenna is designed and specified for the
transmission in the frequency range of from 863 to 928 MHz.
This antenna type forms an electrical small radiator, since its largest
dimension Dmax=79 mm is significantly smaller than
the corresponding wavelength λ, which covers a range from 347.4
to 323.7 mm. Balanis defines a distance of dfar>λ/2π as the boundary of the far field region
in case of an electrical small radiator (Balanis, 2016), which leads to a
minimal distance of 51.5 mm for far field region in case of the
highest specified operating frequency of 928 MHz. The
calculated distance to the far field region can be verified due to numerical
field calculations. Therefore, the UHF antenna is placed in free space as
illustrated in Fig. 1 and the radiated electromagnetic field is calculated
in different observation points situated on a line parallel to the positive
x-axis. Since the UHF antenna's dipole axis is orientated parallel to the
z-axis, Ez is expected to be the dominant component in the far field
region. Figure 2 shows the three components of the numerical calculated
electric field vector E for a frequency of 900 MHz.
Hereby, Ex represents the component in propagation direction, which
should be as a rule of thumb 100 times or 40 dB smaller than
the dominant transverse components Ez in far field region (Laybros and Combes, 2004). The simulations point out that Ex is the largest component
for distances from 5 to 85 mm with respect to the antenna,
thus it can be assumed that the minimal distance to the far field region is
definitely underestimated and must be larger than 51.5 mm.
Figure 3 presents the simulated electric field in an observation point
located in a distance of 55 mm. The transverse component
Ez becomes the major component in the frequency range above 1 GHz, which seem to be reasonable due to the decreasing wave length
λ. Nevertheless, the 40 dB-criteria is not fulfilled,
which means that neither the infant's thorax nor the receiving antenna are
placed in the far field region of the transmitting antenna, especially if
positioned in close vicinity to each other. This first investigation leads
to the finding to carry out further numerical field calculations in order to
understanding the physics of this setup better. Furthermore, not only the
magnitude of the transfer function between both UHF antennae might be of
interest for measuring the breathing parameter, but also the phase
information, since both the transmitting and the receiving antennae will be
detuned due to their mutual physical presence as well as the infant's
thorax.
Radiated electric field for 900 MHz: Ex denotes the component
in propagation direction, whereas Ez is the transverse component
orientated parallel to the dipole axis.
Radiated electric field in a distance of 50 mm: Ex denotes the
component in propagation direction, whereas Ez is the transverse
component orientated parallel to the dipole axis.
Numerical Field Calculation
Numerical field calculations were carried out with a 3D electromagnetic
simulation tool that is based on the finite difference time domain (FDTD)
method (Troescher et al., 2004). Figure 4 illustrates the simulation model of the infant's thorax and
the UHF antennae arrangement. As transmitting and receiving antennae serve
the UHF antennae, which are also used in the real measurement setup and have
been described more detailed in Sect. 2.1 already. Simple geometries have
been chosen for modeling the infant's thorax and the lung. The simplified
thorax and lung are modeled with only two compartments each with homogenous
material distribution in accordance with the experimental lung phantom. The
thorax is represented by a cuboid filled with water, which includes a second
cuboid forming an air pocket as lung model. Both cuboids are rounded at
their edges in order the reduce resonance effects. The lung's physical
dimensions are varied during the simulation in order to model to dynamic
changes of thoracic parameters due to breathing. The dimensions are listed
in Table 1 and depend on the volume of the lung VLung given in
the unit mL.
Simulation model of the infant's thorax and the UHF antennae
arrangement: The infant's thorax is modeled as a cuboid filled with water,
which includes an air pocket representing the lung. The lung's physical
dimensions are varied during the simulation in order to model to dynamic
changes of thoracic parameters due to breathing.
The antennae were placed at a distance of 10 mm from the thorax.
Since a symmetrical simulation arrangement carries the risk of unwanted
geometry-related interferences, the antennae and the lungs were shifted out
of the symmetry axis of the thorax according to Table 2. In addition, the
antennae were tilted slightly by 3∘.
Position of antenna and lung.
y-shiftz-shiftAntennae-10 mm-20 mmLung-15 mm-20 mm
In the simulation model, the thorax is modeled as a water filled cuboid with
a relative permittivity of er=78 and an electrical
conductivity of σ=1.59Sm. An air pocket
serves as lung. The simulations were carried out with the transient solver
in the frequency range from 300 to 1500 MHz.
Experimental Setup
We developed a dynamic artificial lung phantom in order to provide
reproducible conditions for an experimental validation of the simulation as
described in a previous publication (Ringkamp et al., 2020). The dimensions
are chosen to model an infant's lung within the thorax. The outer dimensions
of the phantom are 135 mm × 76 mm × 65 mm with a
wall thickness of 2.4 mm. The dynamic respiration process is
modelled with a 3D printed artificial lung that consists of two compartments
separated by an elastic diaphragm. Filling the artificial lung with air, the
volume ratio of the two compartment changes, which influences the
transmission path between the antennae.
An elastic diaphragm separates the two compartments. The lower compartment
contains water to simulate the electrical properties of an infant's thoracic
tissue and a balloon. The antennae are mounted in 10 mm
distance to the lower compartment filled with water. Figure 5 contains a
technical drawing of the setup. The inflatable balloon inside the water
compartment serves as artificial lung. The phantom was mounted in an upright
position to centre the balloon by buoyancy. The balloon was filled from
approximately 2 mL up to approximately 22 mL in 5 mL steps using a plastic syringe.
Technical drawing of the used lung phantom. The phantom consists
of two plastic compartments. The lower compartment is filled with water to
model the electric properties of thoracic tissue. The upper compartment is
filled with air and in this work serves no special purpose. An elastic
diaphragm (I) separates the two compartments. The antennae (II) are mounted
at a distance of 1 cm next to the lower compartment. A balloon (III) inside
the lower compartment serves as artificial lung. By inflating the balloon
the diaphragm dilates and is pushed in the upper compartment.
The antennae system consists of two identical UHF antennae strips (meandered
dipole on flexible PCB with 79 mm × 11 mm). Scattering
parameters S11 and S21 of the antennae system are measured using a
vector network analyser (PicoVNA 106, Pico Technology). Two 1 m long U.FL cables connected the antennae to the vector network
analyser. The setup was calibrated by using two U.FL cable of the same
length and type to compensate the attenuation and phase delay caused by the
cables.
Results
We evaluated the simulation by comparing trends in the magnitude of the
scattering parameters S11 and S21 derived from the simulation and
the experiment, which represent the input reflection coefficient of the
transmitting antenna and the transmission function between both antennae.
The air volume in both modalities was varied beginning from a start volume
V0 (in the case of the experiment the start volume is unknown,
in the simulation V0=20 mL). The simulation for V0+10 mL did not converge and is therefore excluded from the further
examination. Figure 6 illustrates the electric field strength distribution
in the x-z-plane calculated numerical for a 900 MHz
continuous wave (CW) excitation signal. The antennae are located at the two
horizontal lines, whereas the other black lines indicate the boundaries of
the thorax and the lung. In the close vicinity of the transmitting antenna,
large field strength values arise. It points out that the electric field
propagates on different paths from one antenna to the other. The first path
is the direct transmission through the thorax and lungs on the shortest
route. In the second path, the electric field propagates into the thorax and
diffracts around the lungs, illustrated as grey arrows in Fig. 6. On the
third path, the propagation wave does not penetrate, but diffracts around
the thorax like a surface wave. The field distribution shown in Fig. 7
delivers comparable results in the x-y-plane. Here, the three
propagation paths can be recognized as well. A principle diagram of the
different transmission mechanisms within and outside the thorax is shown in
Fig. 8.
Electric field strength distribution x-z-plane: The black lines show
the boundaries of the antennae and the thorax model with the lung. The white
arrows illustrate the wave propagation from the transmitting antenna to the
receiving antenna when excited with a 900 MHz CW signal.
Electric field strength distribution x-y-plane: The black lines show
the boundaries the thorax model with the lung. The dipole axes of the
antennae are orientated perpendicular to the cutting plan, at which the
transmitting antenna is located in the red colored region. The white arrows
illustrate the wave propagation from the transmitting antenna to the
receiving antenna when excited with a 900 MHz CW signal.
Illustration of possible propagation paths between the UHF
antennae.
When comparing the magnitude of the transmission factor between the antennae
S21, it is visible that the passband frequency and the maximum differ
in the experiment and the simulation, see Fig. 9. The simulation tends to
have a lower passband edge roughly 2 MHz higher and a passband
roughly 30 dB lower. Furthermore, the passband in the simulation tends to be
broader. However, simulation and experiment both show an erratic behaviour
concerning the dependence on the air volume. There is no clear trend whether
the overall magnitude increases or decrease by increasing the air volume.
Comparison of simulation (a) and experiment (b), magnitude
S21. Simulation and experiment both show an erratic behavior concerning
the trend of the maximum. Depending on the current volume, a further
increase in volume can lead to a global increase of S21 or a global
decrease.
Figure 10 shows the input reflexion S11 of the transmitting antennae,
which has two volume dependent resonances in simulation and experiment. The
behaviour of the volume dependent resonance is the same in simulation and
experiment, even though the simulated volume dependent resonance frequencies
are approx. 160 MHz higher than the measured ones. The resonance
at the lower frequency decreases in magnitude with an increase in volume,
while the resonance at the higher frequency increases in magnitude with
increase in air volume. S11 in the experiment shows a much
stronger resonance leading to a better load matching of the antenna. This
most likely causes the 30 dB increase of the experiments
passband in S21.
Comparison of simulation (a) and experiment (b), magnitude
S11. Simulation and experiment show two volume dependent resonances,
which show the same trend in simulation and experiment. The resonance at
lower frequencies decreases with an increase in volume, while the resonance
at higher frequencies increases.
The two volume dependent resonances in the experiment tend to shift more in
frequency compared to the simulation and differ also much more in magnitude.
This leads to a second passband at slightly higher frequencies in the
experiment, which is not present in the simulation. The second passband
shifts with an increase in volume to even higher frequency but decreases in
magnitude; see Fig. 11.
Comparison of the magnitude responses of S11(a) and S21(b) in the experiment for varying air volumes. The volume dependent
resonance at the higher frequency in S11 leads to a second volume
dependent passband in S21, which shifts to higher frequencies with
increase in air volume but decreases in magnitude.
Discussion and conclusion
In this contribution, we investigated a measuring principle for contactless
respiratory parameters determination, which uses a pair of coupled UHF
antennae. Therefore, we carried out numerical field calculations with a
simplified thorax model as well as experiments by means of a 3D-printed
thorax phantom. The investigations point out that the transfer function
between the UHF antennae is influenced by the dynamic change of the lung
volume. Both, simulation and experimental results show similar resonances in
S11 and S21 as well as trends in these parameters with a change in
air volume within the lung. Differences in the simulated and measured
S11 and S21 could be caused by the simulation model and the lung
phantom used for the measurements, since they differ in their geometrical
dimensions, the shape of the lung and the antenna orientation. A shift in
S11 can be mainly attributed to a detuning of the antenna caused by a
variation of the effective dielectric in close vicinity. However, S21
is influenced by multiple factors when the air volume is changed. These are
the respective and potentially different detuning of transmitting and
receiving antenna, a change in absorption, a change in reflection and
diffraction, and a change in phase offset of the different fractions of the
electromagnetic waves as they interfere. For these reasons, the S21
magnitude shows an erratic behavior and a prediction of the air volume with
only this parameter is not possible. However, S11 magnitude as well as
S21 phase show a more consistent trend and are potentially useful for a
dynamic prediction of air volume. Future work will focus on a systematic
evaluation of these parameters in a more realistic experimental setup as
well as a respective numerical model.
Data availability
The data are available from the corresponding author upon request.
Author contributions
SF, CS and KJD carried out numerical field calculations, JR performed the
experiments and the data analysis. SZ and JL supervised the research project
and gave scientific and conceptual advice. All authors contributed to
discussions and the manuscript.
Competing interests
The authors declare that they have no conflict of interest.
Special issue statement
This article is part of the special issue “Kleinheubacher Berichte 2019”. It is a result of the Kleinheubacher Berichte 2019, Miltenberg, Germany, 23–25 September 2019.
Financial support
The publication of this article was funded by the open-access fund of Leibniz Universität Hannover.
Review statement
This paper was edited by Lars Ole Fichte and reviewed by two anonymous referees.
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