A dual lumen CO.sub.2 catheter is formed of medical grade dual lumen silicone rubber tubing which has been impregnated with tridodecyl amine. The silicone tubing is gas permeable, and an external reference electrode is used to monitor the potential across the impregnated outer wall of the dual lumen tubing as a function of pH. The analyte to be monitored is selected from the group consisting of CO.sub.2, NH.sub.3, SO.sub.2, NO.sub.2, H.sub.2 S, and HCN, the internal reference and electrolyte solutions contained in the lumens being selected in response to the particular analyte desired to be monitored. In a CO.sub.2 embodiment, the reference lumen is filled with a buffered internal reference solution, and the response lumen is filled with a bicarbonate filling solution. Reference wires formed of Ag/AgCl in respective ones of the lumens forms the internal electrodes.

BACKGROUND OF THE INVENTION 
This invention relates generally to chemical sensors, and more 
particularly, to a catheter-type potentiometric gas sensor suitable for 
continuous in vivo blood gas monitoring. 
The knowledge of gas levels in blood is essential for the accurate 
assessment of the respiratory and acid-base status of a patient. Oxygen 
and carbon dioxide, the respiratory gases, are of fundamental importance 
since their partial pressures, pO.sub.2 and pCO.sub.2, are useful in 
determining cardiopulmonary homeostasis, or the ability of the 
cardiopulmonary system to maintain a delicate balance between the body's 
respiratory CO.sub.2 production and its O.sub.2 consumption. Additionally, 
the partial pressure of carbon dioxide in blood and the dissolved 
bicarbonate anion are the two principle factors which govern the pH of the 
blood. The measurement of two of these three parameters, such as pH and 
partial pressure of carbon dioxide, will completely characterize the 
acid-base state of the blood. 
Commercially available blood gas analyzer systems typically monitor three 
analytes, including blood pH, the partial pressure of carbon dioxide, and 
the partial pressure of oxygen. Such monitoring provides a clinician with 
a complete determination of the respiratory and acid-base status of a 
patient. 
In conventional discrete blood gas analyzers, the partial carbon dioxide 
pressure level in blood is most commonly obtained by drawing a discrete 
arterial blood sample which is then analyzed by a conventional 
Severinghaus carbon dioxide sensor housed in an automated blood gas 
analyzer. At least a two to three minute time lag between the drawing of a 
sample and the recording of the value of the partial pressure of carbon 
dioxide, is inherent in the known discrete sampling arrangement. However, 
in many diagnostic situations, such as emergency, surgical, and critical 
care patients, blood gas levels can change abruptly, illustratively within 
minutes, thereby indicating eminent respiratory or metabolic failure. In 
such situations, a clinician working with a discrete blood gas analyzer is 
required to make frequent blood gas measurements so as to diagnose 
accurately the patient's rapidly changing condition. Thus, a major 
disadvantage of the known method is the high expense associated with the 
maintenance of a sufficient number of blood gas analyzers, and the 
providing of a staff of trained personnel to perform the frequent blood 
gas measurements. More importantly, improper collection and/or handling of 
a blood sample prior to analysis can produce error in the discrete blood 
gas determination. It is evident from the foregoing that the time lag 
which is inherent in the discrete sampling system limits the speed of 
diagnosis, and consequently the implementation of corrective treatment. 
Accordingly, intensive research has been conducted over the past decade, 
devoted to the search for methods which allow for the continuous 
monitoring of arterial blood gas tensions. 
A variety of technologies has been applied to the development of continuous 
blood sensors for the partial pressure of carbon dioxide. The monitoring 
of partial pressures of carbon dioxide can involve invasive and 
noninvasive approaches. Conventional gas sensors have been used in 
noninvasive extracorporeal loops and transcutaneous arrangements. Invasive 
gas sensing probes have been based on mass spectrometry, conductivity, gas 
chromatography, potentiometry, and fiber optic systems. Although each of 
these developments has some merit, none has achieved wide clinical 
application. 
One prior art system is an extracorporeal loop device which continuously 
draws arterial blood which is then circulated through appropriate tubing 
to external sensing devices; the blood then being returned to a vein 
elsewhere in the body of the patient. Typically, the external sensing 
devices incorporate sensors for oxygen, carbon dioxide, and pH 
measurements. It is a problem with all extracorporeal systems that the 
patient must be heparinized to prevent blood clotting in the loop or on 
the sensors, and to prevent loose blood clots from causing vascular 
occlusion or strokes. Heparinization increases the risk of post-operative 
complications in surgical patients since necessary clotting mechanisms are 
inhibited. Additional drawbacks of extracorporeal devices include losses 
of carbon dioxide in the loop tubing, the need for elaborate temperature 
control, and the increased risk of infection. Consequently, extracorporeal 
loops are rarely used for blood gas measurements. 
Non-invasive transcutaneous blood gas sensors were developed in the early 
1970's In this known system, heated conventional gas sensors are placed 
directly on the skin of the patient such that the partial pressure of 
gases diffusing to the skin surface can be measured. This diffusion of 
gases from the subcutaneous arteries to the skin surface is dependent upon 
skin thickness, blood flow, tissue concentration and/or expiration, and 
arterial gas concentration. In infants, the factors balance, such that the 
transcutaneous blood gas values approximate the arterial values. There is, 
however, a wide variation in adult skin thickness which leads to incorrect 
predictions of arterial blood gas levels from transcutaneous measurements. 
Inaccurate predictions of arterial blood gas levels will also occur in 
patients with reduced blood flow from injury or illness. Thus, 
transcutaneous blood gas monitoring is clinically accepted for infants, 
but rarely is applied to adults. 
A known invasive system involves in vivo mass spectrometric blood gas 
analysis. This system was first proposed in 1966 and has since been 
pursued by a number of researchers. The basic configuration requires the 
use of a mass spectrometer, an in vivo sampling probe, and associated 
connecting tubing. A significant limitation of this technique is the 
complexity and cost of mass spectrometer instrumentation relative to that 
of electrochemical sensors. The major drawback to mass spectrometric probe 
measurements is the non-equilibrium flow dependent, or diffusion limited 
nature of the gas sampling probe. The mass spectrometer maintains the 
sampling probe at a negative pressure that withdraws gases from the 
surrounding blood stream. Under low blood flow conditions, the extraction 
of gases by the probe may deplete the adjacent blood of analyte causing 
erroneously low results. Moreover, blood clotting and/or protein build-up 
on the probe can impede the diffusion of gases into the probe, thereby 
yielding false blood gas values. Thus, economic factors and 
non-equilibrium sampling effects account for the limited biomedical 
application of mass spectrometric blood gas analysis. 
Similarly, flow dependent conductivity-based carbon dioxide catheters have 
been described for in vivo applications. These catheters, however, are 
non-selective and suffer the same non-equilibrium disadvantages as the 
mass spectrometric probes. Gas chromatography has also been applied to 
continuous blood gas monitoring. Blood gases diffuse into the body of an 
indwelling catheter probe where they approach their equilibrium 
concentration with a bolus of carrier gas (He) contained therewithin. At 
fixed intervals, illustratively between three and four minutes, the bolus 
is flushed through the gas analyzer unit for separation and measurement. 
Thus, the gas chromatographic probes do not sample continuously, but 
rather an automated analysis is performed every four minutes. The 
disadvantages of this approach include the delicate nature of the probe 
and its associated high failure rate. More importantly, a commercial 
implementation of this approach will yield serious inaccuracies, on the 
order of between 10 and 20 per cent, in blood gas determinations. 
In recent years, fiber optic carbon dioxide sensors have been developed. 
This approach offers several advantages including true equilibrium 
measurements of partial carbon dioxide, lack of electrical connections to 
the patient, and ease of miniaturization of fiber optic devices. Despite 
the promise of fiber optic designs, difficulties with sensor drift appear 
to have limited its utility in vivo applications in measuring the partial 
pressure of carbon dioxide. The drift has been attributed to changes in 
optical path length caused by deformation of the sensor tip while it is in 
the blood vessel. 
The development of intravascular electro-based probes for the measurement 
of partial pressure of carbon dioxide has centered primarily on the 
miniaturization of the known Severinghaus sensor. FIG. 1 is a comparative 
schematic representation of Severinghaus-type ammonia (NH.sub.3) and 
carbon dioxide (CO.sub.2) gas sensors. As shown in the drawing, sodium 
sensor 10 and carbon dioxide sensor 11 are each combination electrodes in 
that they monitor pH levels also. Each of the sensors is provided with a 
respective one of pH electrodes 12 and 13. Electrode 12 is filled with 
NH.sub.3 and electrode 13 is filled with carbon dioxide. Such electrodes 
are arranged in an internal electrolyte and are separated therefrom by 
respective glass membranes 14 and 15. A gas permeable membrane separates 
the internal electrolyte from the substance being monitored, 
illustratively blood (not shown). 
Since Severinghaus sensors are potentiometric devices, measurements are 
made essentially under zero-current conditions so that no analyte carbon 
dioxide is consumed by the measurement process. Thus, true equilibrium 
measurements are made and problems associated with mass transfer of carbon 
dioxide, such as blood flow variations and carbon dioxide diffusion 
limitations, should not cause errors in the partial pressure values of 
carbon dioxide which are determined. In addition, all Severinghaus-type 
carbon dioxide devices are unaffected by anaesthetic gases. 
The design of the system of FIG. 1 employs miniaturized glass pH electrodes 
as internal sensing elements in catheter size devices. However, such 
arrangements are hampered by the fragility, noise, and cost associated 
with the electrodes. The prior art has thrust at the problem associated 
with miniature glass electrodes, by providing sensors which detect the 
partial pressure of carbon dioxide based on quinhydrone and antimony pH 
sensitive electrode systems. These sensors, however, suffer from the 
disadvantages of oxygen sensitivity, instability, and large size. Thus, 
they do not appear to be suitable for in vivo testing. 
FIG. 2 is a schematic representation of a prior art combination sensor 20 
which monitors pH and the partial pressure of carbon dioxide. This known 
probe is comprised of a palladium oxide pH electrode 21, a Ag/AgCl 
reference electrode 22 and a bicarbonate electrolyte 23 housed behind a 
gas-permeable membrane 64 which contains a mobile hydrogen ion carrier 
which makes the membrane permeable to hydrogen ions. The partial pressure 
of carbon dioxide is measured by monitoring the voltage between the two 
internal electrodes, while the sample pH is measured between the Pd/PdO 
electrode 21 and an external reference electrode (not shown). Sensor 20 
measures 0.9 mm in outside diameter. It is a disadvantage of this type of 
sensor that they are poorly flexible, exhibiting a decrease in sensitivity 
when the sensor is bent. This type of sensor also drifts unacceptably 
unless recalibrated every 1.5 hours. The drift is attributed to the PdO 
electrode's sensitivity to redox species. In addition, the sensor is also 
very sensitive to temperature changes, as might occur during fever or 
hypothermia. As a result of this, this combination sensor has seen only 
limited use. 
FIG. 3 is a schematic representation of a catheter sensor 30 which measures 
the partial pressure of carbon dioxide and is based on a tubular polymeric 
membrane internal pH electrode. In this known arrangement, a pH sensitive 
membrane 31 is situated safely within the wall of the internal tubing, 
rather than at the vulnerable tip of the sensor. This protects the sensing 
regions from damage during catheter placement or removal. In addition, 
this geometry allows for sensor size reduction without a corresponding 
decrease in the pH sensitive membrane area and a concomitant increase in 
electrode resistance. Also noteworthy is the heightened rate of 
flexibility afforded by the polymer based internal pH electrode. Finally, 
since the internal electrode is based on a hydrogen ion permselective 
polymer membrane, the sensor is virtually insensitive to sample redox 
species. 
Although this design is promising for continuous in vivo monitoring of the 
partial pressure of carbon dioxide, there are difficulties in its 
fabrication which will limit its application. 
It is, therefore, an object of this invention to provide a simple and 
economical sensor which provides determination of the respiratory and 
acid-base status of a patient. 
It is another object of this invention to provide an implantable sensor for 
continuous monitoring of blood carbon dioxide partial pressures. 
It is also an object of this invention to provide a single catheter implant 
which can monitor pH and the partial pressure of carbon dioxide 
simultaneously. 
It is additionally an object of this invention to provide a blood gas 
analyzer system which does not require trained personnel to be operated. 
It is a further object of this invention to provide a blood gas analysis 
system which reduces the possibility of error resulting from faulty 
collection and/or handling of a blood sample. 
It is still another object of this invention to provide a blood gas 
analysis sensor which eliminates lag time between samplings. 
It is a yet further object of the invention to provide a blood gas 
monitoring arrangement which eliminates the need to heparinize a patient. 
it is also a further object of this invention to provide a blood gas sensor 
which is not plagued by drift. 
It is yet another object of this invention to provide a blood gas sensor 
which is not fragile and does not have a high failure rate. 
SUMMARY OF THE INVENTION 
The foregoing and other objects are achieved by this invention which 
provides an arrangement for in vivo monitoring of the concentration of 
blood gases. The arrangement is constructed to have an elongated body 
which is itself formed of a gas-permeable polymeric material impregnated 
with a pH sensitive material. Such impregnation is effected over a 
predetermined portion of the elongated body, thereby forming a pH 
responsive membrane. In accordance with the invention, the elongated body 
is provided with at least one a lumen therein for accommodating a 
predetermined solution. A sensor electrode is arranged in said lumen of 
the elongated body for communicating electrically with the predetermined 
solution for detecting an electric potential. 
The inventive sensor can be operated in combination with a reference 
electrode external to the elongated body. This external electrode 
facilitates monitoring of the potential across the impregnated outer wall 
of the elongated body, which is formed of a silicone rubber, the potential 
being a function of the pH of the sample being monitored. In a specific 
illustrative embodiment of the invention, the outer wall of the elongated 
body is impregnated with tridodecyl amine. 
The particular solution installed in the lumen is selected in response to a 
predetermined analyte being monitored, which may illustratively be from 
the group consisting of CO.sub.2, NH.sub.3, SO.sub.2, NO.sub.2, H.sub.2 S, 
and HCN. 
In a preferred embodiment of the invention, the sensor electrode is in the 
form of a wire of Ag/AgCl. The sensor electrode is in electrical 
communication with the solution in the lumen. 
In accordance with a highly advantageous embodiment of the invention, the 
elongated body is provided with a further lumen therein, and there is 
additionally provided a further sensor electrode arranged in the further 
lumen for communicating electrically with a second predetermined solution. 
The second predetermined solution is a bicarbonate solution, which 
comprises 0.025 M NaHCO.sub.3 and 0.1M NaCl. 
In the dual lumen embodiment, the elongated body means is formed of a 
gas-permeable polymeric material impregnated with a pH sensitive material 
over a predetermined portion thereof for forming a pH responsive membrane. 
The elongated body has first and second lumens therein for accommodating 
respective ones of first and second predetermined solutions. There is 
additionally provided the use of first and second sensor electrode means 
of Ag/AgCl for communicating electrically with respective ones of said 
predetermined solutions for detecting at least one electric potential.

DETAILED DESCRIPTION 
FIG. 4 is a schematic representation of a dual lumen catheter sensor 40 
constructed in accordance with the principles of the invention. As shown 
in the figure, sensor 40 is constructed using dual lumen tubing 41 having 
a lumen 42 which is filled with a bicarbonate solution, and a lumen 43 
which is filled with a buffered internal reference solution. In one 
specific embodiment, tubing 41 comprises an 11 cm section of Dow Corning 
Silastic Rx50CT medical grade dual lumen tubing, having 1.14 mm outside 
diameter and 0.15 mm wall thickness. In accordance with the invention, the 
medical grade tubing is immersed approximately 2 cm in a hydrogen ion 
carrier impregnating solution, such as tridodecyl amine (TDDA), for two 
minutes or until swelling of the tubing is visible. The impregnating 
solution may be, in certain embodiments, 0.5 g TDDA in 2.5 ml xylene. 
The swollen tubing section is allowed to dry in air, illustratively for an 
overnight period, after which any solution remaining in the tubing lumens 
is expelled with compressed air. A short section of approximately between 
2 to 5 mm of the impregnated end of the tubing is plugged in both lumens 
with silicone rubber adhesive which is allowed to cure for at least one 
hour. Lumen 42 is filled with bicarbonate filling solution, illustratively 
of 0.025M NaHCO.sub.3 and 0.1M NaCl. The buffered internal reference 
solution in lumen 43 is, in one embodiment, comprised 0.80M phosphate 
buffer (pH 7.0) containing 0.10M NaCl. Preferably, the compositions of the 
buffered reference solution and the bicarbonate fill solution are 
optimized to match the osmolarity of blood in order to improve the 
long-term stability of the sensor. 
Each of lumens 42 and 43 contains an associated one of Ag/AgCl reference 
wires 45. In a one embodiment, the Ag/AgCl leads are soldered to a 
transistor socket 45. The area between the top of the silicone rubber 
tubing and the transistor socket is encapsulated in silicone rubber 
adhesive (not shown) to prevent the Ag/Agcl leads from shorting. 
Measurements are then made by connecting the sensor's transistor socket 
leads to a pH meter and monitoring the potential of the electrochemical 
cells. 
The configuration of sensor 40 improves upon the miniaturized Severinghaus 
sensors described hereinabove by forming an internal pH responsive 
membrane in the wall of the dual lumen carbon dioxide permeable rubber 
tube. This design eliminates the need for a separate internal pH 
responsive electrode which greatly simplifies the construction of the 
sensor. Moreover, the elimination of the separate internal pH responsive 
electrode permits reductions in sensor dimensions. 
The sensor design of FIG. 4 may be readily adapted to detect hydrogen ions 
and CO.sub.2 gas simultaneously. This is effected merely by contacting the 
sample with an appropriate external reference electrode, whereby the 
potential across the outer, TDDA-impregnated wall of the dual lumen can be 
monitored as a function of sample pH. Thus, the potential between the two 
inner reference Ag/AgCl wires would track the partial pressure of carbon 
dioxide in the sample while the potential between the Ag/AgCl wire of the 
reference lumen and the external reference electrode would track the 
sample pH. Such an arrangement would facilitate continuous, real-time 
monitoring of two of the three blood gas analytes (blood pCO.sub.2 and 
pH). 
In alternative embodiments, the wall between the lumens in the dual-lumen 
tubing, after being impregnated with the hydrogenion carrier TDDA can 
serve as the pH selective membrane. In such embodiments, the potential 
across this inner tubing wall can be correlated to the pH of the 
bicarbonate filling solution as a function of the partial pressure of 
carbon dioxide in the sample. This embodiment provides reduced fabrication 
complexity and improved potential for miniaturization of the sensor 
dimensions. More significantly, such elimination of the inner pH 
responsive tubing may permit complete automation of fabrication of 
duallumen catheter-based sensors. 
FIG. 5 is a graphical representation of a carbon dioxide calibration curve. 
As shown in this figure, the logarithm of the carbon dioxide concentration 
is plotted against the steady state potential at the output of the sensor. 
The plotted results were obtained from in vitro studies wherein standard 
additions of sodium bicarbonate were made to an acidic sample buffer 
solution at 37.degree. C. The results recorded are the equilibrium 
potential of the dual lumen sensor shown schematically in FIG. 4. 
Reference to FIG. 5 shows that the dual lumen sensor exhibits 
nearNernstian sensitivity (92%) from 5 to 130 torr pCO.sub.2 
(1.6.times.10.sup.-4 to 4.0.times.10.sup.-3 M CO.sub.2). This range of 
Nernstian response covers the entire clinical pCO.sub.2 range which is 
approximately 15-84 torr. 
Although the invention has been described in terms of specific embodiments 
and applications, persons skilled in the art can, in light of this 
teaching, generate additional embodiments without exceeding the scope or 
departing from the spirit of the claimed invention. Accordingly, it is to 
be understood that the drawing and descriptions in this disclosure are 
proffered to facilitate comprehension of the invention and should not be 
construed to limit the scope thereof.