[This paper may be of interest to
respiratory therapists because it also describes the
historical development of the model 3A+ fine particle
nebulizer ...]
Radioaerosol
Inhalation Lung Scanning: A
New Generation
by
Ross Potter and Russell W. King, Medi/Nuclear Corp., Inc.,
Baldwin Park, CA
Radioaerosol
inhalation lung scanning had its beginnings in the mid
1960's shortly after Drs. George Taplin of UCLA and Henry
Wagner of Johns Hopkins independently developed the
macroaggregated particles that made the perfusion lung scan
possible. It didn't take long to realize that, while the
perfusion lung scan could readily show areas of decreased or
absent blood flow, it could say virtually nothing about the
cause. Because of this shortcoming, Dr. Taplin developed the
technique for radioaerosol lung scanning. At the time, about
the only thing then available to evaluate lung ventilation
was Xenon-133 gas. But because no one had yet developed a
delivery/trapping system, its use was limited. Beyond this,
the Anger Camera was not yet in wide use and rectilinear
scanners could not image a gas that did not remain in the
lungs for an extended period. Initially, many isotopes were
tried for the creation of the aerosol, but it was finally
Tc99m Sulfur Colloid that became the agent of choice1.
This
isotope worked well with the ultrasonic nebulizers then in
use, but its attractiveness disappeared with the advent of
the jet nebulizer. This was for the simple reason that jet
nebulizers do not aerosolize suspensions readily. These
early nebulizers produced a range of particles in the
vicinity of 1 M
to 5M,
but no one at the time knew very much about the particle
size range or distribution. At this time, central deposition
was considered diagnostic of COPD and many papers were
written to explain why it occurred.
Over the years, efforts were made to overcome the
shortcoming of central and tracheal deposition. The results
were generally less than fully effective, and because of
this, the radioaerosol inhalation lung scan languished in
just a few centers. While this was happening, the use of
Xenon-133 gas was increasing. Systems had been developed to
deliver the gas to the patient and charcoal traps had been
developed to trap the expired gas, this despite the
limitations of single views and poor resolution. Krypton-87m
generators were developed to overcome these problems, but
cost and logistics argued against widespread use.
Until
the introduction of commercial aerosol systems by Synaco
(later Mallinckrodt) and Cadema (later CIS-US), there were
only minor improvements in the delivery of the radioaerosol.
Taplin, et al, made most of these. They included the use of
a large bag downstream from the nebulizer2 to try to remove
the larger particles by gravity, the addition of a light
bulb3 to try to evaporate the larger particles to make them
smaller and the addition of a cap over the generator in the
nebulizer to cause the larger particles to rain out. Only
the latter was successful and it has led to the development
of the nebulizers in use today. In the early 1990's, Mishkin
developed a system that included a one-way valve at the
mouthpiece. In 1997, he modified a commercial nebulizer
system to duplicate his original system. It was theorized
that without the one-way valve at the mouthpiece, humidified
breath would cause the particles to grow beyond the size
that would evenly distribute throughout the lung and result
in excessive tracheal and bronchial deposition. Testing
demonstrated that removal of the one-way valve resulted in a
35% increase in the mass of particles > 1,
with a corresponding increase in tracheal and central
deposition.4 The system was a fine design, but the necessity
of having the exit filter external to the shield created
problems with commercial viability. There had to be another
way.
Francken, et al, tested a modified commercial system in 1997
that was designed to generate particles almost exclusively
<1µM.5 The control of particle size and its distribution was
accomplished by increasing the diameter of the baffle in the
nebulizer so that the particles generated would be required
to make two 90 degree turns before leaving the nebulizer.
The first turn would cause larger particles to impact on the
baffle and be returned to solution for regeneration. Those
that successfully negotiated this turn were of a mass of
less than 1.
The system proved quite successful in improving peripheral
penetration and reducing tracheal and bronchial deposition.
The only problem that came up was slower delivery of the
aerosol to the patient. This came about because the
extremely small particles could not carry the same amount of
radioactivity carried previously by the larger particles and
because the generation rate of the nebulizer had been
decreased. This was easily compensated for by simply
increasing the concentration of the isotope added to the
nebulizer reservoir. Still, some thought the breathing time
was a little long, especially for patients with breathing
difficulties. Hyun, et al, tested a second system which
incorporated the same nebulizer (giving the same particle
size distribution) but also utilized a holding compartment
and a flow control diaphragm.6 The holding compartment was
intended to hold and retain aerosol generated during the
time of exhalation in order to make it available during the
next inhalation. This would theoretically double the rate of
accumulation in the lungs. In fact, it more than doubled the
accumulation rate. What previously required five minutes of
breathing now required only two minutes. The flow control
diaphragm was intended to make certain that the generated
aerosol was directed to the holding compartment and also to
eliminate or minimize the mixing of humid exhaled breath
with the newly generated aerosol.
An
incidental finding was that the image quality had been
markedly improved.7 To explain this improvement we must
return to Washington, et al. The flow control diaphragm
functions in a manner somewhat similar to the one-way valve
at the mouthpiece in the Mishkin system in that it tends to
prevent saturated breath from the patient exhalation from
mixing with the newly generated aerosol. This, in turn,
prevents some of the growth of the particle that occurs as
they absorb moistures. The primary difference between the
two systems (Washington/Mishkin and Hyun) is that the
Mishkin device allows the one-way valve to strip the larger
particles from the airstream while the Hyun device does not
generate the larger particle in the first place. The
increased breathing restriction in the Mishkin device caused
by the one-way valve is absent in the Hyun device. This lack
of breathing restriction acts to increase both patient
comfort and compliance.
At
this point it is worth noting that mass relative to particle
diameter is not linear, but is in proportion to volume.
The
volume of a sphere is determined by the formula:
Therefore,
if a sphere of 1
diameter has a mass of 1, then a sphere of 3 diameter has a
mass of 27.15 and a sphere of 5
has a mass of 125.66. From this it can be seen that a 5
particle is capable of carrying 125 times the radioactivity
of a 1
particle. If we accept that through moisture absorption,
aerosol particles do grow in size as they penetrate the
passageways of the lung, and that final particle size must
be no greater than 1 to 3
to reach and deposit in the deep lung, then we must start
with very small particles in order to allow for size growth.
Even so, the deposited particles will have carried no more
drug (radioactivity) than they contained originally. At the
same time, any particles larger than 1
which are generated by the nebulizer will contain the much
greater amount of activity related to their mass. Because
the larger particles are unable to negotiate the upper
respiratory tract, they are most likely to deposit in the
trachea and main stem bronchus, thus adding to the count
rate in a much larger proportion than their numbers alone
would indicate.
In
designing a nebulizer to deliver particles largely less than
1,
it became clear that a very small percentage of particles
greater than 1
could produce a noticeable difference in image quality.
Our
first nebulizer produced 35% of particles in the 1 to 5
range. The first modification reduced the percentage of
particles in the 1 to 5
range to 8.2% and the second modification reduced this
number to less than 3%. It was this second modification that
yielded the best results in terms of overall image quality
in a blind reading of 43 consecutive studies. This unit then
moved into production.
As in
all things, there were compromises, the compromise in this
case being decreased aerosol generation efficiency. It now
required 2-4 times longer breathing to obtain the same count
rate as the original system. This was compensated for by
increasing the concentration of isotope placed in the
nebulizer reservoir. This brought breathing times into a
reasonable range for a pre-perfusion study, but
post-perfusion aerosol studies remained somewhat difficult.
The problem at this point was either patient compliance
because of excessive breathing time or a lack of an adequate
concentration of isotope in the evening and early morning
hours.
To
overcome the lack of efficiency accompanying the smaller
particles, a new unit was developed which returned breathing
times to previous levels while still providing the very
small particles needed for superior image quality. This new
unit provides a holding chamber and flow control diaphragm.
The holding chamber enables the unit to retain the aerosol
generated during the time of patient exhalation, so that
upon inspiration, the patient inhales both the aerosol
generated during inspiration and that generated during the
preceding exhalation. The flow control diaphragm directs the
breath to an exit filter during exhalation and opens during
inhalation to give access to both the currently generated
aerosol and the previously generated and stored aerosol.
Where the flow control diaphragm differs from a one-way
valve is in its ability to open completely without
restricting the stream of the aerosol, and to do so without
increasing the breathing resistance. These two
characteristics together prevent pooling of aerosol at the
valve with a consequent loss of efficiency.
TESTING
Particle
size measurements were performed using a Quartz Crystal
Microbalance, 10 Stage Cascade Impactor, manufactured by
California Measurements, Sierra Madre, CA. This device has
been criticized for its inability to handle high
concentrations of aqueous aerosols.9 We have been using this
unit for 10 years and are very pleased with its
reproducibility over this period. The only qualifier is that
it must be used properly, i.e., short sampling times. The
industry standard unit is the Anderson 8 Stage Cascade.
We
have tested both the PC-2 and the Anderson simultaneously
and found that in our hands, there is no essential
difference. We have also had our aerosol tested by TSI Inc.,
St. Paul, MN, using their Model 3934 Scanning Mobility
Particle Sizer and obtained the same results. Testing was
conducted using oxygen as the driving gas at a flow rate of
10L/min., the aerosol being drawn past an isokinetic intake
orifice at 14L/min. and the aerosol drawn through the
cascade stages at 0.24L/min. Sampling times varied from 1 to
5 seconds and settling time in the cascade was 90 seconds.
Aerosol
generation rate, expressed as mL/min., was obtained by
determining weight loss per unit of time. Weights were
obtained using a Mettler Analytical Balance reading to
0.001g. All tests were conducted using oxygen as the driving
gas at a flow rate of 10L/min. as determined with a
calibrated Gilmont Flow Tube. The aerosol generation rate
for our original nebulizer (Medi/Nuclear Corp., NEB-3A) was
0.28mL/min. and for nebulizer modification 2 was 0.14mL/min.
The latter is now in production as Medi/Nuclear Corp.,
NEB-3A+.
Testing
Summary
| Nebulizer |
Particle
Size (µM) |
%
Distribution |
| NEB-3A |
<0.2 |
3.3 |
| |
0.2
to 1.0 |
58.2 |
| |
1.0
to 5.0 |
37.1 |
| |
>5.0 |
1.4 |
| Modification
#1 |
<0.2 |
16.8 |
| |
0.2
to 1.0 |
75.0 |
| |
1.0
to 5.0 |
8.2 |
| |
>5.0 |
0 |
| Modification
#2 |
<0.2 |
6.9 |
| |
0.2
to 1.0 |
90.2 |
| |
1.0
to 5.0 |
2.9 |
| |
>5.0 |
0 |
| NEB-3A+ |
<0.2 |
40.8 |
| |
0.2
to 1.0 |
57.1 |
| |
1.0
to 5.0 |
2.1 |
| |
>5.0 |
0 |
REFERENCES
1.
Potter,
R., Swanson, L.A., Brotherton, J., Suyenaga, K.
Radioaerosol Inhalation Lung Scanning. JNM 1972; 13: 874
2.
Taplin,
G.V. Atlas for Lung Imaging Using Radioaerosols. Div. of
Nuclear Medicine, Dept. of Radiological Sciences, School
of Medicine & Laboratory of Nuclear Medicine &
Radiation Biology, UCLA 1979
3.
Mullins,
J, Hayes, M. Improved Technique for Aerosol Inhalation
Scanning. JNM 1972; 13: 872
4.
Washington,
M.J., Whiteley, M.E., Mishkin, F.S., Potter, R.E., King,
R.W. Factors Influencing Aerosol Ventilation Imaging. JNM
1997; 38: 37P
5.
Francken,
G.A., Waxman, A.D., Potter, R., King, R. The Effect of
Inertial Changes on Particle Size Distribution in
Nebulizer Systems: The Impact on Clinical Studies. JNM
1998; 39: 118P
6.
Hyun,
J., Wawxman, A., Potter, R., King, R. Improvement in
Aerosol Delivery from a Standard Nebulizer System: the Use
of a Holding Compartment. JNM 1998; 39: 117P
7.
Hyun,
M., Waxman, A., D'Angelo, A., Potter, R., King, R. The
Impact of a Holding Compartment on Aerosol Delivery from
Nebulizer Systems. SNM Western Regional Meeting, October
1998.
8.
Raabe,
O.G. The Ideal Particle Size for Optimal Pulmonary
Deposition, Letter of response by O.G. Raabe, PhD. 1996;
Internet, www.vortra.com/RAABE696.htm
9.
Toporkov,V.S.,
Molokeev, V.A. Thermodynamic Model of Lung for
Investigation of Multi-Component Particles in Respiratory
Tract. Journal of Aerosol Medicine 1995; 8: #1, 70
10. Clark,
A.R., Gonda, I., Newhouse, M.T. Journal of Aerosol
Medicine 1998; 11: #1, S-1
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