412
IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. BME-22, NO. 5, SEPTEMBER 1975
An Electromechanical Demand Regulated Liquid Breathing System THOMAS H. SHAFFER AND GORDON D. MOSKOWITZ
Abstract-The process of breathing liquid requires consideration of several problems not typical in gas respiration. These problems stem from the physical properties of liquid such as density, viscosity, and low diffusivity of gases which increase respiratory work. The objective of this study was to develop and study the feasibility of a demand controlled system for providing mechanical assistance and ameliorating respiratory work during liquid breathing. The description, design, and analysis of a demand controlled system for liquid ventilation of the lungs is presented. Results from simulated tests and in viw animal studies are also reported. In vio animal studies demonstrate that demand control is an effective method for providing stability in the control of animal blood gases P02 and PCO2 during liquid ventiation.
INTRODUCTION EARLIER studies in pulmonary gas exchange with hyperbaric saline solution indicated that diffusion rather than flow rate was the rate-limiting factor [1], [2]. A typical problem due to diffusion limitations is the removal of carbon dioxide. In the past the latter problem has usually limited liquid ventilation to short periods of time. Clark and Gollan [3] demonstrated that mice and cats could survive in an oxygenated flurocarbon liquid (FC-80)' at ATPS. However, carbon dioxide clearance and respiratory work difficulties were still evident. If liquid ventilation is to have applications for man, it is necessary that long term survival after liquid breathing be demonstrated. Modell [41 found that life could be supported by ventilation with liquids for prolonged periods and demonstrated long term survival after ventilation with liquid. Most previous studies of liquid ventilation indicate difficulty in maintaining physiologically safe blood CO2 levels. These earlier liquid breathing experiments which have demonstrated difficulty with CO2 removal have been conducted under one of two conditions, manual valving of liquid to an intubated animal or total immersion of the animal in the breathing liquid. Studies employing manual valving provide mechanical assistance at a fixed breathing rate which tended to preclude adequate carbon dioxide clearance rates. Total immersion, although providing free selection of breathing rate, offers no mechanical assistance. Manuscript received May 29, 1973; revised July 8, 1974, and February 18, 1975. This work was supported in part by the National Heart and Lung Institute under Public Health Service Research Grant HE13847. T. H. Shaffer was with the Biomedical Engineering and Science Program, Drexel University, Philadelphia, Pa. 19104. He is now with the Department of Physiology, School of Medicine, University of Pennsylvania, Philadelphia, Pa. 19174. G. D. Moskowitz is with the Biomedical Engineering and Science Program, Drexel University, Philadelphia, Pa. 19104. 1 FC-80 product of 3M Company.
The purpose of the study reported here, was to design a system which employs the animal's internal respiratory controller and sensory mechanisms to set respiratory variables such as tidal volume and breathing rate. The system responds to incipient respiratory efforts and provides for anticipated demands. Demand regulation [5] provides optimum mechanical assistance by following the spontaneous breathing effort. This study has been motivated by a potentially wide variety of applications in medical assistance and research for liquid ventilation of the lungs. Potential applications would be in the following areas: 1. lung lavage; 2. lung resuscitation; 3. controlled hypothermia; 4. aquaspace and aerospace. SYSTEM DESCRIPrION AND ANALYSIS The liquid breathing system consists of the following elements (Fig. 1): 1. electrically operated actuator (Fig. 2, H. Beck & Sons, Inc. #14-101); 2. bellows pumps (designed and fabricated by the authors); 3. electronic controller (circuit shown in Fig. 3); 4. oxygen source (24 ft3, compressed gas bottle); 5. solenoid valves; 6. check valves; 7. pressure transducer (statham p23 dB); 8. temperature transducer (thermistor); 9. flow control valves; 10. function generator for controlled respiration. Actuator deflection causes simultaneous displacement of both bellows. In the compression stroke the inspiration pump displaces fresh liquid to the lungs while the expiration pump displaces spent liquid to the regenerator. Solenoid and check valves restrict flow to the desired directions. In the expansion stroke the expiration pump removes liquid from the lungs while the inspiration pump draws liquid from the regenerator. Bypass lines from each pump to the regenerator provide variability of the maximum flow rate to and from the animal. Actuator control (Fig. 3) is obtained through an electronic controller. Incipient intraesophageal pressure changes are sensed with an esophageal balloon catheter and transduced with a low compliance pressure transducer connected to a D.C. amplifier and electronic controller. Deviation of input esophageal pressure from a preset operating point in the system controller provides constant actuator motion. Actuator displacement produces constant liquid flow from the pumps. Pump displacements are proportional to intraesophageal pressure signal changes. A spontaneous decrease of esophageal pressure from the preset point (a decrease of intrapleural pressure) results in flow to the animal and regenerator. An increase in esophageal pressure results in flow from the animal and regenerator. Based on pulmonary function control measurements of the test -animal (airway resistance and lung compliance) a system gain relationship (the ratio of volume output to pressure signal input) is set in the controller.
413
SHAFFER AND MOSKOWITZ: ELECTROMECHANICAL LIQUID BREATHING SYSTEM
Fig. 1.
Z t II-'
SET GAIN
,CA
Fig. 2.
IS - BELLOWS FLOW
C87 TRANSVERSE BELLOWS COMPLIANCE C82- AXIAL BELLOWS COMPUANCE CL-LINE COMPLIANCE R- LINE RESISTANCE RA- ANIMNAL RESISTANCE CA - ANIMAL COMPLIANCE
IA
-
ANIMAL FLOW
i- FLOW TO C81
PT -PRESSURE TRANSDUCER ED -ESOPHAGEAL BALLOON
PB-
BELLOWS PRESSURE
Fig. 3.
Fig. 4.
Besides the demand mode of operation the liquid breati system has a controlled breathing mode of operation. electronically simulating input pressure variations, contro ventilation can be provided for animals with apnea. In mode flow rate, tidal volume, and frequency can be manu set by an operator to meet physiological needs. The animal lung dynamics were based on a first order linear network. (The pulmonary resistance and compha considered constant.) For a step input of intrapleural p sure, Pip, the volume output is:
V(t) = CAPiP [1 - exp (- t/r)]
(1) Bellows Equations: PB= B,C
APB2 = CB2
T, idt
(2)
1
(3)
where T1 and T2 are initial and final time, respectively. (2) Continuity Equation:
(1)
where 7 = RA CA is the lung time constant and RA and CA are pulmonary resistance and compliance, respectively. The system dynamics are described in a voltage-press analog [61 schematic shown in Fig. 4. The component e quations for the linear actuated system are as follows:
ie
IB = i + IA (3)
(4)
Line Equation:
A PL
RL
RLCLS+III
(4) Lung Equation:
(5)
IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, SEPTEMBER 1975
414
TABLE I ANIMAL TRIALS WITH LIQUID BREATHING SYSTEM (LBS) Dog
Sex
1
O~
2
3
4
APA
RAIA
+
'A
21
9 0'
('
T'rial
12
Yes
2
60
Yes
3
80
Yes
24
2
60
3
73
1
40
2
30
25
1
23
15
1
23
2
60
1
20
20
22
Recovery Rcvr
1
(6)
GA S where s is a Laplace operator. Combining (2) through (6) yields the following linear differential equation:
RC dt +IA =IB
Total Time on Lbs. (min.)
1
20
5 6
sWeight (Pounds)
(7)
where R and C are equivalent system compliances and
resistances. Assuming that a step into an amplitude Pip is applied to the maximum effort flow source IB, the flow response of the system is:
(8) IA(t)=IB [1- exp(-t/r1) where r1 = RC, the equivalent time constant of the system, IA (0) = °, IB(O) = 0, and IB(t) = const. Integrating the flow rate yields the volume response of the
system: V1 (t) = IBt + IBr' [exp (- t/lT) - 1] . (9) The difference, 5 V, in animal volume response V(t) and system volume response V(t)1 results in changes of intrapleural pressure. By internally sensing its own physiological requirements (indicated by blood gas levels, for example) the animal is free to regulate his own minute ventilation by adjusting frequency and volume. Note that the liquid breathing system was not designed to simulate air breathing response while breathing liquid but to reduce animal respiratory work for all breathing frequencies. Complete simulation of liquid breathing for air breathing would most definitely result in severe lung damage. Due to the dynamic viscosity of liquid and the development of abnormal alveolar pressure, circulatory interference, or mechanical lung damage would occur for excessive flow rates.
Yes Yes Died 8 days after trial.
Yes Died 24 Hrs. after trial. Died 18 Hrs. after trial. Yes Yes
Died 24 Hrs.
after trial.
METHODS Six mongrel dogs (Table I) were prepared for liquid breathing trials by anesthetization, cannulation, intubation of both an endotracheal tube and esophageal balloon, and hyperventilation with 100% oxygen for 10 minutes. Complete description of animal preparation procedure has been previously reported [7]. Preanesthetic medication consists of fentanyl and droperidol (Innovaret), 1 ml/20 lig. body at 0.5 h prior to trials and anesthesia is accomplished with pentobarbital sodium intravenously. Additional increments of pentobarbital are administered as needed. Before connecting the animal directly to the system, he was filled from a separate liquid reservoir to help remove gas from his lungs and adjust system operating points. Throughout the trials various physiological parameters were closely monitored and recorded, e.g., breathing frequency, blood pressure, esophageal pressure, Po2, PCO2, pH, ECG, and temperature. Blood gas samples were taken at 8 minute intervals and analyzed with Clark and Severinghaus type electrodes. The maximum instantaneous flow output of the system is 5 1/min at 12 BPM, however, none of the animals received the maximum rate. The average ventilation flow rate ranged between 1-3 1/min. Tidal volume is a function of breathing frequency and varied between 100 and 150 ml. Positive instantaneous esophageal pressures were as high as 50 mmHg in one animal but generally averaged around 10 mmHg. RESULTS
Simulated Animal Testing Simulated animal tests were conducted with a dynamic load (estimated) similar to that of a liquid filled lung. The object of the simulated test was to determine the system volume output over the range of physiological breathing
SHAFFER AND MOSKOWITZ: ELECTROMECHANICAL LIQUID BREATHING SYSTEM
415
400
-\g- SIMULATED -
DESIGN MOEL
300
1
C1Z-J
1
-
0
200
o
a I
00
0
.2
4 .6 .8 BREATHING FREQUENCY (HZ)
1.0
Fig. 5.
frequencies. As long as the simulated volume response exceeds minutes. The first trials of dog I and 2 were performed to the design model response the system is considered adequate. determine optimum system operating procedures and thereSince no values of pulmonary resistance and compliance fore are not included in time or blood gas statistics. were available for the liquid (FC-80) filled lung extrapolated Arterial blood samples were taken during spontaneous air values of air breathing and saline breathing (static compliance) breathing, hyperventilation with 100% oxygen and throughout data were employed. Pulmonary liquid resistance was taken as each liquid breathing trial. Fig. 6 illustrates group means plus 800 cm of H20/l/sec and pulmonary compliance as 50 ml/cm standard error for arterial pH, Pb , and PCO2 during liquid H20. ventilation. As shown the group mean Po2 remains in a safe Dynamic compliance was simulated with a variable liquid physiological range throughout liquid ventilation (comparable head, open reservoir (inverted bottle with air vent). A change to spontaneous air breathing). The increase in mean Pco2 and in liquid volume in the bottle results in a change in hydro- decrease in pH with time indicate a deficiency in CO2 removal. static pressure or the compliance C, determined by the varia- Klystra [11 and Modell [41 have shown that liquid ventilation tion of AV over AP. Pulmonary resistance was simulated generally results in progressive hypercarbia and acidosis. We with an orifice at the entrance to the bottle. found that demand controlled ventilation results in a more Although an orifice does not exhibit entirely linear char- gradual PCO2 increase and pH decrease with time. acteristics, linearization techniques could be employed to Recovery results from liquid breathing trials are presented in correlate simulation results to design model over the flow Table II. Pulmonary mechanics measurements were taken range of interest (1-5 1 /min). before and after liquid breathing trials. The purpose of the Simulated experimental results and analytical design model measurement was to determine the impact of liquid ventilation results are presented in Fig. 5. The experimental results were on pulmonary mechanics. determined by force oscillating the simulated load througli the It was found that recovery time of pulmonary mechanics, breathing frequency range (.15 to 1.0 Hz). A function gen- from liquid breathing with FC-80 was similar to recovery time erator was used as the liquid breathing system input signal from lung lavage with saline [81, [9] about four to ten days. (open loop) while volume and 'frequency were recorded. A total of 12 experiments were conducted all of which reDesign model results were obtained from (9) which were also sulted in recovery of the test animal from liquid to conscious based upon estimated lung parameters. Although results indi- air breathing. The degree and rate of recovery varied 'n each cate adequate system response up to 0.8 Hz it is necessary to animal and experiment. Animals 2, 3, 4, and 6 died during the ascertain actual liquid filled lung characteristics. The capa- 3 month experimentation or recovery period, however, post bility to determine liquid filled lung characteristics is now mortems revealed that only the deaths of animals 2 and 6 available with the developed liquid breathing system and are could be directly attributed to liquid ventilation. In both of these cases, large amounts of foam were found in the under study. trachea and large bronchi. The lungs were hyperinflated and In Vivo Testing hemorrhagic in appearance. The gross pathological diagnosis Six mongrel dogs were prepared as previously described for of death was mechanical lung damage caused by hyperinflation. 12 liquid breathing system trials. The mean period of time of the system for all trials (exclud- Discussion This study presents a method and a system for implementing ing dog 1, trial 1, and dog 2, trial 1) was 47 minutes ±6.7 SE; with a maximum of 80 minutes, and a minimum of 19 liquid breathing experiments in which the test animal is in
416
IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, SEPTEMBER 1975 Nl
60
x
40 l MEAN kf SE
I
IL
20 1
0L
0
o
20
20c
f a
60
40
f_I_r
100
x--
0
0
a.
ILj
730
Iq
40
60
20
40
60
jx%K
7.40
x-
20
7.20 710
0
TIME ON LBS (MINUTES)
Fig. 6. RECOVERY
OF
TABLE II PULMONARY MECHANICS AFTER LIQUID BREATHING SYSTEM (LBS) EXPERIMENTS PERCENT Post-Liquid
Resistance
Compliance Tidal Volune
240
OF
Breathing
CON1ROL (Days)
2
3
4
_
6
7
8
9
10
11
13
14
19
21
180
180
140
148
198
120
115
80 I
25
27
40
36
17
16
12
44
21
164
152
126
140
8
35
23
13
27
40
68
62
68
84
92
18
12
12
201
14
13
44
64
74
76
82
74
88
78
104
100
12
2
/
100
108
94
110
96 10
2
106
94
90
104
3
116I
20
8
120
100I
2b
7
6
6
5
8
7
4
7
20
Minute Volume
175 125
90 100
95
97
102
90
112
140
90
150
96
110
108
3
30
25
8
6
3
30
Respiratory Rate
320
270
160
120
130
120
116
105
130
100
104
96
110
110I
70
30
41
17
45
10
23
12o 300
14
6
10
5
24
6
cotitrol of liquid delivery. The basis for demand controlled observed in animals 2 and 6 was a result of underestimated assistaince is that the test animal can sense progressive hyper- lung parameters on liquid. Due to the dynamic viscosity of carbia anid acidosis and accordingly adjust his breathing liquid, abnormal alveolar pressure causes hyperinflation of the lungs and mechanical damage. frequency and tidal volume for CO2 clearance. Results- from in vivo anumal studies indicate that demand With the existence of the present system it is now possible controlled assistance does provide stable control of blood gas to ascertain exact values of dynamic resistance and comtensions in the normal range. Although there is still an evident pliance of the liquid filled lung and incorporate them into the trend toward hypercarbia and acidosis, the rate of change- is system control technique. Determination of these values can nmore gradual than previously found [1], 14]. Recovery be accomplished by a method similar to that by Mead [10] results, of pulmonary mechanics measurements were very en- and Von Neergaard [11] except the breathing medium will c6uraging in that recovery time, to control levels (compliance, be liquid and delivery will be from the liquid breathing system. 38.4 ml/ct H20 ± 4.1 SE and resistance, 14.5 cm H20/l/sec) from liquid breathing trials with flurocarbon liquid was as fast ACKNOWLEDGMENT as recovery from lung lavage experiments with saline. Since the liquid breathing system was designed with estiThe authors gratefully acknowledge Dr. S. Dubin for his mated dynamic lung parameters, certain shortcomings were valuable suggestions and surgical abilities. Likewise, it is our anticipated. We believe that the mechanical lung damage as pleasure to thank J. Greiner, F. DuCoin, J. Heacock, and K.
SHAFFER AND MOSKOWITZ: ELECTROMECHANICAL LIQUID BREATHING SYSTEM
Patrick for their technical assistance in performance of
experiments.
REFERENCES [1] J. A. Klystra, C. V. Paganelli, and E. H. Lamphier: "Pulmonary
exchange in dogs ventilated with hyperbarically oxygenated liquid," Journal of Applied Physiology, vol. 21, No. 1, Jan. 1966, pp. 177-184. J. A. Kylstra, "Required properties of a liquid for respiration," Federation Proceedings, vol. 29, September-October 1970, pp. 1724. L. C. Clark and F. Gollan, "Survival of mammals' breathing organic liquid equilibrated with oxygen at atmospheric pressure," Science,vol. 132, June 1966,pp. 1755-1756. J. H. Modell, E. J. Newley, and B. C. Ruiz, "Long term survival of dogs after breathing oxygenated flurocarbon liquid," Federation Proceedings, vol. 29, September-October, 1971. G. D. Moskowitz, S. Dubin, and T. H. Shaffer, "Demand regu-
[6]
[71
gas
[21 [3] [4]
[5]
[8]
[9] [10]
[11]
417
lated control of a liquid breathing system," Journal of the Association for the Advancement of Medical Instrumentation, vol. 5, September-October, 1971. F. H. Raven, Automatic Control Engineering, McGraw-Hill, N.Y., 1961. T. H. Shaffer, "Control analysis and design of demand regulated liquid breathing systems," PH.D. Dissertation, Drexel University, June 1972. S. E. Dubin and R. C. Pfleger, "Pulmonary lavage and respiratory function parameters in beagle dogs," Fission Product Inhalation Program Annual Report, Nov. 1969, pp. 209-213. S. Dubin, A. J. Wilson, and R. C. Pfleger, "Changes in respiratory function of beagle dogs given pulmonary lavage,'" Americali Journal of Veterinary Research, vol. 32, pp. 2033-20,38. J. Mead and J. L. Whittenberger, "Physical properties of human lungs measured during spontaneous -respiration," Jour. Applied Physiology, 5: 779-796,1953. K. Von Neergaard and K. Wirz, "Uber eine Methode zur Messung der Lugenelastizitat am Lebenden Menschen, Insbesonder beim Emphysem," Z. Klin. Med., vol. 105: 35-50, 1927.
IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. BME-22, NO. 5, SEPTEMBER 1975
An Electromechanical Demand Regulated Liquid Breathing System THOMAS H. SHAFFER AND GORDON D. MOSKOWITZ
Abstract-The process of breathing liquid requires consideration of several problems not typical in gas respiration. These problems stem from the physical properties of liquid such as density, viscosity, and low diffusivity of gases which increase respiratory work. The objective of this study was to develop and study the feasibility of a demand controlled system for providing mechanical assistance and ameliorating respiratory work during liquid breathing. The description, design, and analysis of a demand controlled system for liquid ventilation of the lungs is presented. Results from simulated tests and in viw animal studies are also reported. In vio animal studies demonstrate that demand control is an effective method for providing stability in the control of animal blood gases P02 and PCO2 during liquid ventiation.
INTRODUCTION EARLIER studies in pulmonary gas exchange with hyperbaric saline solution indicated that diffusion rather than flow rate was the rate-limiting factor [1], [2]. A typical problem due to diffusion limitations is the removal of carbon dioxide. In the past the latter problem has usually limited liquid ventilation to short periods of time. Clark and Gollan [3] demonstrated that mice and cats could survive in an oxygenated flurocarbon liquid (FC-80)' at ATPS. However, carbon dioxide clearance and respiratory work difficulties were still evident. If liquid ventilation is to have applications for man, it is necessary that long term survival after liquid breathing be demonstrated. Modell [41 found that life could be supported by ventilation with liquids for prolonged periods and demonstrated long term survival after ventilation with liquid. Most previous studies of liquid ventilation indicate difficulty in maintaining physiologically safe blood CO2 levels. These earlier liquid breathing experiments which have demonstrated difficulty with CO2 removal have been conducted under one of two conditions, manual valving of liquid to an intubated animal or total immersion of the animal in the breathing liquid. Studies employing manual valving provide mechanical assistance at a fixed breathing rate which tended to preclude adequate carbon dioxide clearance rates. Total immersion, although providing free selection of breathing rate, offers no mechanical assistance. Manuscript received May 29, 1973; revised July 8, 1974, and February 18, 1975. This work was supported in part by the National Heart and Lung Institute under Public Health Service Research Grant HE13847. T. H. Shaffer was with the Biomedical Engineering and Science Program, Drexel University, Philadelphia, Pa. 19104. He is now with the Department of Physiology, School of Medicine, University of Pennsylvania, Philadelphia, Pa. 19174. G. D. Moskowitz is with the Biomedical Engineering and Science Program, Drexel University, Philadelphia, Pa. 19104. 1 FC-80 product of 3M Company.
The purpose of the study reported here, was to design a system which employs the animal's internal respiratory controller and sensory mechanisms to set respiratory variables such as tidal volume and breathing rate. The system responds to incipient respiratory efforts and provides for anticipated demands. Demand regulation [5] provides optimum mechanical assistance by following the spontaneous breathing effort. This study has been motivated by a potentially wide variety of applications in medical assistance and research for liquid ventilation of the lungs. Potential applications would be in the following areas: 1. lung lavage; 2. lung resuscitation; 3. controlled hypothermia; 4. aquaspace and aerospace. SYSTEM DESCRIPrION AND ANALYSIS The liquid breathing system consists of the following elements (Fig. 1): 1. electrically operated actuator (Fig. 2, H. Beck & Sons, Inc. #14-101); 2. bellows pumps (designed and fabricated by the authors); 3. electronic controller (circuit shown in Fig. 3); 4. oxygen source (24 ft3, compressed gas bottle); 5. solenoid valves; 6. check valves; 7. pressure transducer (statham p23 dB); 8. temperature transducer (thermistor); 9. flow control valves; 10. function generator for controlled respiration. Actuator deflection causes simultaneous displacement of both bellows. In the compression stroke the inspiration pump displaces fresh liquid to the lungs while the expiration pump displaces spent liquid to the regenerator. Solenoid and check valves restrict flow to the desired directions. In the expansion stroke the expiration pump removes liquid from the lungs while the inspiration pump draws liquid from the regenerator. Bypass lines from each pump to the regenerator provide variability of the maximum flow rate to and from the animal. Actuator control (Fig. 3) is obtained through an electronic controller. Incipient intraesophageal pressure changes are sensed with an esophageal balloon catheter and transduced with a low compliance pressure transducer connected to a D.C. amplifier and electronic controller. Deviation of input esophageal pressure from a preset operating point in the system controller provides constant actuator motion. Actuator displacement produces constant liquid flow from the pumps. Pump displacements are proportional to intraesophageal pressure signal changes. A spontaneous decrease of esophageal pressure from the preset point (a decrease of intrapleural pressure) results in flow to the animal and regenerator. An increase in esophageal pressure results in flow from the animal and regenerator. Based on pulmonary function control measurements of the test -animal (airway resistance and lung compliance) a system gain relationship (the ratio of volume output to pressure signal input) is set in the controller.
413
SHAFFER AND MOSKOWITZ: ELECTROMECHANICAL LIQUID BREATHING SYSTEM
Fig. 1.
Z t II-'
SET GAIN
,CA
Fig. 2.
IS - BELLOWS FLOW
C87 TRANSVERSE BELLOWS COMPLIANCE C82- AXIAL BELLOWS COMPUANCE CL-LINE COMPLIANCE R- LINE RESISTANCE RA- ANIMNAL RESISTANCE CA - ANIMAL COMPLIANCE
IA
-
ANIMAL FLOW
i- FLOW TO C81
PT -PRESSURE TRANSDUCER ED -ESOPHAGEAL BALLOON
PB-
BELLOWS PRESSURE
Fig. 3.
Fig. 4.
Besides the demand mode of operation the liquid breati system has a controlled breathing mode of operation. electronically simulating input pressure variations, contro ventilation can be provided for animals with apnea. In mode flow rate, tidal volume, and frequency can be manu set by an operator to meet physiological needs. The animal lung dynamics were based on a first order linear network. (The pulmonary resistance and compha considered constant.) For a step input of intrapleural p sure, Pip, the volume output is:
V(t) = CAPiP [1 - exp (- t/r)]
(1) Bellows Equations: PB= B,C
APB2 = CB2
T, idt
(2)
1
(3)
where T1 and T2 are initial and final time, respectively. (2) Continuity Equation:
(1)
where 7 = RA CA is the lung time constant and RA and CA are pulmonary resistance and compliance, respectively. The system dynamics are described in a voltage-press analog [61 schematic shown in Fig. 4. The component e quations for the linear actuated system are as follows:
ie
IB = i + IA (3)
(4)
Line Equation:
A PL
RL
RLCLS+III
(4) Lung Equation:
(5)
IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, SEPTEMBER 1975
414
TABLE I ANIMAL TRIALS WITH LIQUID BREATHING SYSTEM (LBS) Dog
Sex
1
O~
2
3
4
APA
RAIA
+
'A
21
9 0'
('
T'rial
12
Yes
2
60
Yes
3
80
Yes
24
2
60
3
73
1
40
2
30
25
1
23
15
1
23
2
60
1
20
20
22
Recovery Rcvr
1
(6)
GA S where s is a Laplace operator. Combining (2) through (6) yields the following linear differential equation:
RC dt +IA =IB
Total Time on Lbs. (min.)
1
20
5 6
sWeight (Pounds)
(7)
where R and C are equivalent system compliances and
resistances. Assuming that a step into an amplitude Pip is applied to the maximum effort flow source IB, the flow response of the system is:
(8) IA(t)=IB [1- exp(-t/r1) where r1 = RC, the equivalent time constant of the system, IA (0) = °, IB(O) = 0, and IB(t) = const. Integrating the flow rate yields the volume response of the
system: V1 (t) = IBt + IBr' [exp (- t/lT) - 1] . (9) The difference, 5 V, in animal volume response V(t) and system volume response V(t)1 results in changes of intrapleural pressure. By internally sensing its own physiological requirements (indicated by blood gas levels, for example) the animal is free to regulate his own minute ventilation by adjusting frequency and volume. Note that the liquid breathing system was not designed to simulate air breathing response while breathing liquid but to reduce animal respiratory work for all breathing frequencies. Complete simulation of liquid breathing for air breathing would most definitely result in severe lung damage. Due to the dynamic viscosity of liquid and the development of abnormal alveolar pressure, circulatory interference, or mechanical lung damage would occur for excessive flow rates.
Yes Yes Died 8 days after trial.
Yes Died 24 Hrs. after trial. Died 18 Hrs. after trial. Yes Yes
Died 24 Hrs.
after trial.
METHODS Six mongrel dogs (Table I) were prepared for liquid breathing trials by anesthetization, cannulation, intubation of both an endotracheal tube and esophageal balloon, and hyperventilation with 100% oxygen for 10 minutes. Complete description of animal preparation procedure has been previously reported [7]. Preanesthetic medication consists of fentanyl and droperidol (Innovaret), 1 ml/20 lig. body at 0.5 h prior to trials and anesthesia is accomplished with pentobarbital sodium intravenously. Additional increments of pentobarbital are administered as needed. Before connecting the animal directly to the system, he was filled from a separate liquid reservoir to help remove gas from his lungs and adjust system operating points. Throughout the trials various physiological parameters were closely monitored and recorded, e.g., breathing frequency, blood pressure, esophageal pressure, Po2, PCO2, pH, ECG, and temperature. Blood gas samples were taken at 8 minute intervals and analyzed with Clark and Severinghaus type electrodes. The maximum instantaneous flow output of the system is 5 1/min at 12 BPM, however, none of the animals received the maximum rate. The average ventilation flow rate ranged between 1-3 1/min. Tidal volume is a function of breathing frequency and varied between 100 and 150 ml. Positive instantaneous esophageal pressures were as high as 50 mmHg in one animal but generally averaged around 10 mmHg. RESULTS
Simulated Animal Testing Simulated animal tests were conducted with a dynamic load (estimated) similar to that of a liquid filled lung. The object of the simulated test was to determine the system volume output over the range of physiological breathing
SHAFFER AND MOSKOWITZ: ELECTROMECHANICAL LIQUID BREATHING SYSTEM
415
400
-\g- SIMULATED -
DESIGN MOEL
300
1
C1Z-J
1
-
0
200
o
a I
00
0
.2
4 .6 .8 BREATHING FREQUENCY (HZ)
1.0
Fig. 5.
frequencies. As long as the simulated volume response exceeds minutes. The first trials of dog I and 2 were performed to the design model response the system is considered adequate. determine optimum system operating procedures and thereSince no values of pulmonary resistance and compliance fore are not included in time or blood gas statistics. were available for the liquid (FC-80) filled lung extrapolated Arterial blood samples were taken during spontaneous air values of air breathing and saline breathing (static compliance) breathing, hyperventilation with 100% oxygen and throughout data were employed. Pulmonary liquid resistance was taken as each liquid breathing trial. Fig. 6 illustrates group means plus 800 cm of H20/l/sec and pulmonary compliance as 50 ml/cm standard error for arterial pH, Pb , and PCO2 during liquid H20. ventilation. As shown the group mean Po2 remains in a safe Dynamic compliance was simulated with a variable liquid physiological range throughout liquid ventilation (comparable head, open reservoir (inverted bottle with air vent). A change to spontaneous air breathing). The increase in mean Pco2 and in liquid volume in the bottle results in a change in hydro- decrease in pH with time indicate a deficiency in CO2 removal. static pressure or the compliance C, determined by the varia- Klystra [11 and Modell [41 have shown that liquid ventilation tion of AV over AP. Pulmonary resistance was simulated generally results in progressive hypercarbia and acidosis. We with an orifice at the entrance to the bottle. found that demand controlled ventilation results in a more Although an orifice does not exhibit entirely linear char- gradual PCO2 increase and pH decrease with time. acteristics, linearization techniques could be employed to Recovery results from liquid breathing trials are presented in correlate simulation results to design model over the flow Table II. Pulmonary mechanics measurements were taken range of interest (1-5 1 /min). before and after liquid breathing trials. The purpose of the Simulated experimental results and analytical design model measurement was to determine the impact of liquid ventilation results are presented in Fig. 5. The experimental results were on pulmonary mechanics. determined by force oscillating the simulated load througli the It was found that recovery time of pulmonary mechanics, breathing frequency range (.15 to 1.0 Hz). A function gen- from liquid breathing with FC-80 was similar to recovery time erator was used as the liquid breathing system input signal from lung lavage with saline [81, [9] about four to ten days. (open loop) while volume and 'frequency were recorded. A total of 12 experiments were conducted all of which reDesign model results were obtained from (9) which were also sulted in recovery of the test animal from liquid to conscious based upon estimated lung parameters. Although results indi- air breathing. The degree and rate of recovery varied 'n each cate adequate system response up to 0.8 Hz it is necessary to animal and experiment. Animals 2, 3, 4, and 6 died during the ascertain actual liquid filled lung characteristics. The capa- 3 month experimentation or recovery period, however, post bility to determine liquid filled lung characteristics is now mortems revealed that only the deaths of animals 2 and 6 available with the developed liquid breathing system and are could be directly attributed to liquid ventilation. In both of these cases, large amounts of foam were found in the under study. trachea and large bronchi. The lungs were hyperinflated and In Vivo Testing hemorrhagic in appearance. The gross pathological diagnosis Six mongrel dogs were prepared as previously described for of death was mechanical lung damage caused by hyperinflation. 12 liquid breathing system trials. The mean period of time of the system for all trials (exclud- Discussion This study presents a method and a system for implementing ing dog 1, trial 1, and dog 2, trial 1) was 47 minutes ±6.7 SE; with a maximum of 80 minutes, and a minimum of 19 liquid breathing experiments in which the test animal is in
416
IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, SEPTEMBER 1975 Nl
60
x
40 l MEAN kf SE
I
IL
20 1
0L
0
o
20
20c
f a
60
40
f_I_r
100
x--
0
0
a.
ILj
730
Iq
40
60
20
40
60
jx%K
7.40
x-
20
7.20 710
0
TIME ON LBS (MINUTES)
Fig. 6. RECOVERY
OF
TABLE II PULMONARY MECHANICS AFTER LIQUID BREATHING SYSTEM (LBS) EXPERIMENTS PERCENT Post-Liquid
Resistance
Compliance Tidal Volune
240
OF
Breathing
CON1ROL (Days)
2
3
4
_
6
7
8
9
10
11
13
14
19
21
180
180
140
148
198
120
115
80 I
25
27
40
36
17
16
12
44
21
164
152
126
140
8
35
23
13
27
40
68
62
68
84
92
18
12
12
201
14
13
44
64
74
76
82
74
88
78
104
100
12
2
/
100
108
94
110
96 10
2
106
94
90
104
3
116I
20
8
120
100I
2b
7
6
6
5
8
7
4
7
20
Minute Volume
175 125
90 100
95
97
102
90
112
140
90
150
96
110
108
3
30
25
8
6
3
30
Respiratory Rate
320
270
160
120
130
120
116
105
130
100
104
96
110
110I
70
30
41
17
45
10
23
12o 300
14
6
10
5
24
6
cotitrol of liquid delivery. The basis for demand controlled observed in animals 2 and 6 was a result of underestimated assistaince is that the test animal can sense progressive hyper- lung parameters on liquid. Due to the dynamic viscosity of carbia anid acidosis and accordingly adjust his breathing liquid, abnormal alveolar pressure causes hyperinflation of the lungs and mechanical damage. frequency and tidal volume for CO2 clearance. Results- from in vivo anumal studies indicate that demand With the existence of the present system it is now possible controlled assistance does provide stable control of blood gas to ascertain exact values of dynamic resistance and comtensions in the normal range. Although there is still an evident pliance of the liquid filled lung and incorporate them into the trend toward hypercarbia and acidosis, the rate of change- is system control technique. Determination of these values can nmore gradual than previously found [1], 14]. Recovery be accomplished by a method similar to that by Mead [10] results, of pulmonary mechanics measurements were very en- and Von Neergaard [11] except the breathing medium will c6uraging in that recovery time, to control levels (compliance, be liquid and delivery will be from the liquid breathing system. 38.4 ml/ct H20 ± 4.1 SE and resistance, 14.5 cm H20/l/sec) from liquid breathing trials with flurocarbon liquid was as fast ACKNOWLEDGMENT as recovery from lung lavage experiments with saline. Since the liquid breathing system was designed with estiThe authors gratefully acknowledge Dr. S. Dubin for his mated dynamic lung parameters, certain shortcomings were valuable suggestions and surgical abilities. Likewise, it is our anticipated. We believe that the mechanical lung damage as pleasure to thank J. Greiner, F. DuCoin, J. Heacock, and K.
SHAFFER AND MOSKOWITZ: ELECTROMECHANICAL LIQUID BREATHING SYSTEM
Patrick for their technical assistance in performance of
experiments.
REFERENCES [1] J. A. Klystra, C. V. Paganelli, and E. H. Lamphier: "Pulmonary
exchange in dogs ventilated with hyperbarically oxygenated liquid," Journal of Applied Physiology, vol. 21, No. 1, Jan. 1966, pp. 177-184. J. A. Kylstra, "Required properties of a liquid for respiration," Federation Proceedings, vol. 29, September-October 1970, pp. 1724. L. C. Clark and F. Gollan, "Survival of mammals' breathing organic liquid equilibrated with oxygen at atmospheric pressure," Science,vol. 132, June 1966,pp. 1755-1756. J. H. Modell, E. J. Newley, and B. C. Ruiz, "Long term survival of dogs after breathing oxygenated flurocarbon liquid," Federation Proceedings, vol. 29, September-October, 1971. G. D. Moskowitz, S. Dubin, and T. H. Shaffer, "Demand regu-
[6]
[71
gas
[21 [3] [4]
[5]
[8]
[9] [10]
[11]
417
lated control of a liquid breathing system," Journal of the Association for the Advancement of Medical Instrumentation, vol. 5, September-October, 1971. F. H. Raven, Automatic Control Engineering, McGraw-Hill, N.Y., 1961. T. H. Shaffer, "Control analysis and design of demand regulated liquid breathing systems," PH.D. Dissertation, Drexel University, June 1972. S. E. Dubin and R. C. Pfleger, "Pulmonary lavage and respiratory function parameters in beagle dogs," Fission Product Inhalation Program Annual Report, Nov. 1969, pp. 209-213. S. Dubin, A. J. Wilson, and R. C. Pfleger, "Changes in respiratory function of beagle dogs given pulmonary lavage,'" Americali Journal of Veterinary Research, vol. 32, pp. 2033-20,38. J. Mead and J. L. Whittenberger, "Physical properties of human lungs measured during spontaneous -respiration," Jour. Applied Physiology, 5: 779-796,1953. K. Von Neergaard and K. Wirz, "Uber eine Methode zur Messung der Lugenelastizitat am Lebenden Menschen, Insbesonder beim Emphysem," Z. Klin. Med., vol. 105: 35-50, 1927.
