DETAILED DESCRIPTIONS
Fig. 1 shows a system 10 that includes a heart 12 and an Implantable Medical Device (IMD)14 implantable in a chest of a patient. IMD14 includes a pulse generator for generating a current and a voltage detection circuit for detecting a potential difference (i.e., a voltage). IMD14 may be a cardiac pacemaker, a cardiac defibrillator, a combination thereof, or any other type of implantable medical device capable of delivering current stimuli to thoracic tissue and measuring voltages. The IMD14 includes a so-called "can" electrode located on the outer surface of the device 14. In the present illustrative system 10, 3 additional electrodes 16, 18, 20 are shown that may be attached to leads (not shown) connected to the IMD 14. The electrodes 16, 18, 20 are electrically connected to the IMD14 by conductors that penetrate the leads, thereby facilitating the injection of current from the pulse generator between any two of the four electrodes 16, 18, 20 or can electrodes. Similarly, the voltage detection circuit may detect a voltage between any two of the four electrodes 16, 18, 20 or the can electrode. An impedance calculation module within IMD14 may then calculate the impedance by taking the ratio of the measured voltage to the injected current according to ohm's law. Since four electrodes are employed in such impedance measurements, the electrode configuration shown in system 10 may be referred to as a quadrupole configuration. In the tetrapolar configuration shown in fig. 1, electrodes 16 and 18 are positioned over the left ventricle of heart 12, while electrode 20 is positioned in the right atrium of heart 12. Based on their location on the left ventricular lead (not shown in fig. 1), the electrodes 16 may be referred to as proximal left ventricular electrodes, while the electrodes 18 are referred to as distal left ventricular electrodes.
Since the human body includes a certain number of thoracic organs, tissues and fluids, the measurement of thoracic impedance includes contributions from each of the organs, tissues and fluids. For example, the electrical resistivity of the heart muscle, lung, pectoral muscle, thoracic fat, liver, kidney, spleen, stomach, skeletal muscle, bone, cartilage, blood, and other tissues and fluids all contribute to the measurement of thoracic impedance. Thus, changes in the measured thoracic impedance may be caused by changes in the resistivity of these and other organs or tissues.
When measuring impedance to detect or assess a pathological condition (e.g., pulmonary edema or myocardial ischemia), it is desirable to measure impedance using an electrode configuration that is highly sensitive to the region of interest (e.g., the lung or myocardium, respectively). Since resistivity changes in the region of interest have a correspondingly greater effect on the measured impedance, a high sensitivity to the region of interest enables sensitive detection of resistivity changes in the organ or tissue of interest. Thus, a change in impedance can indicate the presence of a pathological condition. Moreover, it is also desirable that such configurations have relatively low sensitivity to organs, tissues and fluids outside the region of interest, such that resistivity changes in these organs, tissues and fluids will have correspondingly less impact on the measured impedance.
Referring again to fig. 1, the system shows two bioelectric lead fields 22, 24. The first bioelectric lead field 22 originates from the can electrode and terminates at the distal left ventricular electrode 18. A second bioelectric lead field 24 (shown in phantom in fig. 1) originates from the right atrial electrode 20 and terminates at the proximal left ventricular electrode 16. As shown in fig. 1, the first bioelectric lead field 22 is closest near the distal left ventricular electrode 18 and near the can electrode, represented by the more closely spaced bioelectric field lines in these regions. Similarly, the second bioelectric lead field 24 is closest proximate the proximal left ventricular electrode 16 and the right atrial electrode 20.
The first and second bioelectric lead fields 22, 24 may be referred to as converging bioelectric lead fields because they originate at different locations relatively far apart from each other (in this case, the left thoracic region and the right atrium, respectively), but terminate in close proximity to each other (in this case, near the left ventricle). Thus, the converging bioelectric lead field lines merge at a high density only in the region of interest. Sensitivity is concentrated in the region of interest-in this case the lateral left ventricle and its adjacent left lung region. The high density of converging bioelectric field lines near the impedance measurement region of interest and the lack of a converging merged lead field in other regions facilitates highly sensitive impedance measurements specifically directed to resistivity changes in the region of interest, as will be described more fully below. This may be in contrast to impedance measurement configurations such as conventional two-and three-pole electrode configurations, as well as conventional four-pole electrode configurations, where the respective bioelectric lead fields are densely concentrated in two regions, resulting in impedance sensitivity not only to regions of interest, but also to regions not of interest (corresponding to other densely concentrated field regions).
The first bioelectric lead field 22 (which may be associated with, for example, a stimulation electrode) may be defined as the field of current density vectors produced by a unit current injection between the can electrode and the distal left ventricular electrode 18. The second bioelectric lead field 24 (which may be associated with, for example, a voltage measurement electrode) may similarly be defined as a field of current density vectors formed per unit of current injection between the right atrial electrode 20 and the proximal left ventricular electrode 16. For a review of bioelectric lead field theory, see j.malmivuo & r.plonsey, biomagnetic: principles and Applications of Bioelectric Fields (Bioelectromagnetism: Principles and Applications of Bioelectric Fields) (1995) (chapters 11.6.2 and 25.2.1 from http:// butler. cc. tut. fi/. malmivuo/bem/bamboo) start at pages 202 and 405, respectively.
The illustrative electrode configuration shown in fig. 1 may be used as an impedance measurement configuration, for example, the pulse generator injects current between the can electrode and the distal left ventricular electrode 18, the voltage detection circuit measures the resulting voltage between the right atrial electrode 20 and the proximal left ventricular electrode 16, and the impedance calculation module calculates the impedance by taking the ratio of the measured voltage to the injected current. Since the convergent bioelectric lead field system as shown in fig. 1 focuses sensitivity mainly near the region of interest, false alarms triggered by measured impedance changes due to resistivity changes in regions of non-interest can be avoided.
Figure 2 shows a convergent bioelectric field electrode configuration that is highly sensitive to changes in lung resistivity and can be used for pulmonary edema detection and assessment. Pulmonary edema can be detected and assessed by making impedance measurements and labeling the impedance changes. Such impedance changes may be predictive of increased fluid in the lungs (i.e., pulmonary edema or the occurrence thereof) and may be treated in time by early detection. Referring to fig. 2, a human body 30 is shown with IMD14 implanted in a left thoracic region of the human body 30. Leads 32, 34 are connected to ports of IMD14 and extend therefrom. Each lead 32, 34 carries one or more conductors therein that electrically connect IMD14 to electrodes 20, 16, 18 implanted within heart 12 of human body 30.
In the illustrative configuration, lead 32 is a right atrial lead and, near the distal end, carries a right atrial electrode 20 located in the right atrium 36 of heart 12. Right atrial lead 32 extends from a port in IMD14, is introduced into the venous system, down the Superior Vena Cava (SVC)38, and into the right atrium 36. Right atrial electrode 20 may be a ring electrode or a tip electrode, or may be located elsewhere along lead 32, such as in SVC38 or anywhere in the innominate vein. Lead 34 is a left ventricular lead that includes a proximal left ventricular electrode 16 and a distal left ventricular electrode 18, both positioned epicardially over a left ventricle 40 of heart 12. Left ventricular lead 34 extends from a port on IMD14, is introduced into the venous system, down SVC18, into right atrium 36, into the coronary sinus, and then into the coronary veins running epicardially over left ventricle 40. The left ventricular electrodes 16, 18 may be ring or tip electrodes, or may be located elsewhere along the lead 34. Although the left ventricular lead 34 is shown in fig. 2 as a bipolar lead, the lead 34 may optionally include additional electrodes, and the lead 34 may also follow a different path through the heart 12. For example, the left ventricular lead 34 may include 3, 4, 5, or more electrodes, and they may be arranged co-linearly. Similarly, right atrial lead 32 may include additional electrodes and may follow a different path through heart 12 than shown in FIG. 2. In one embodiment, the left ventricular electrodes 16, 18 are spaced approximately 15mm apart and the can electrode is spaced approximately 10cm from the right atrial electrode 20. In other embodiments, the distance between the can electrode and the right atrial electrode 20 may be approximately 8-15 cm. Similarly, in other embodiments, the spacing between the left ventricular electrodes is about 25mm, and may be in the range of about 5-30 mm.
IMD14 includes a can electrode located on an external surface of device 14. IMD14 may alternatively or additionally include a button electrode or a header electrode, or other type of electrode. In the tetrapolar configuration shown in fig. 2, a pulse generator internal to IMD14 may inject an electrical current between a can electrode positioned on an external surface of IMD14 and distal left ventricular electrode 18. The injected current flows through at least a portion of the left lung 44 and causes a potential difference (voltage) that is measurable by the right atrial electrode 20 and the proximal left ventricular electrode 16. An impedance calculation module within IMD14 may then calculate the impedance by calculating a ratio of the measured voltage to the injected current. The impedance may be used to assess pulmonary edema in the body 30, as will be described in more detail below. Arrows 42, 46 indicate the general orientation of the two bioelectric lead fields in this configuration. Lead field 42 corresponds to current injection and lead field 46 corresponds to voltage measurement. As shown in fig. 2, the lead fields 42, 46 converge near the left ventricular wall and left lung, but are substantially separated from each other.
Referring now to FIG. 3, another convergent bioelectric field electrode configuration that is highly sensitive to changes in lung resistivity is shown. A left ventricular lead 50 is connected to a port of IMD14 and includes three electrodes 16, 18, 52. The left ventricular lead 50 is a tripolar lead and follows a similar routing path as the left ventricular lead 34 shown in fig. 2. In addition to left ventricular proximal and distal electrodes 16, 18, both located epicardially above the wall of left ventricle 40, lead 50 also includes an electrode 52 located in the brachiocephalic vein (i.e., innominate vein). While a tripolar left ventricular lead 50 is employed in the configuration shown in fig. 3, other configurations may also be used. For example, the brachiocepahlic electrode 52 may be included in place of a right atrial lead, a right ventricular lead, or in place of a dedicated lead terminating in or near the brachiocepahlic vein. The brachiocephalic electrode 52 may alternatively be located anywhere near the brachiocephalic vein. In operation, a pulse generator within the IMD14 may inject an electrical current between the can electrode located on an outer surface of the IMD14 and the distal left ventricular electrode 18. The injected current flows through at least a portion of the left lung 44 and induces a voltage that is measurable by the brachiocepahlic electrode 52 and the proximal left ventricular electrode 16. An impedance calculation module within the IMD14 may then calculate an impedance that may be used to assess pulmonary edema by calculating a ratio of the measured voltage to the injected current. Arrows 42, 54 indicate the general orientation of the two bioelectric lead fields in this configuration. Lead field 42 corresponds to current injection, while lead field 54 corresponds to voltage measurement. The lead fields 42, 54 converge near the anatomical region of interest, but are substantially separated from each other. In one embodiment, the left ventricular electrodes 16, 18 are spaced about 15mm apart and the distance between the can electrode and the brachiocepahlic electrode 52 is about 4 cm. In other embodiments, the distance between the can electrode and the brachiocepahlic electrode 52 is approximately 5-7 cm. Similarly, the spacing between left ventricular electrodes is about 25mm in other embodiments, and may be in the range of about 5-30 mm.
Additional electrodes, such as one or more additional left ventricular electrodes or right atrial electrodes, may be used for current injection and voltage measurement. For example, a brachiocepahlic electrode 52 may be added to the right atrial lead 32 as shown in fig. 2, so that the voltage between the right atrial electrode 20 and the proximal left ventricular electrode 16, and the voltage between the brachiocepahlic electrode 52 and the proximal left ventricular electrode 16 may be continuously measured during the current injection, and the respective impedances calculated as described above. An impedance calculation module within IMD14 may then calculate a weighted average impedance measurement, e.g., by averaging the impedance measurements using an appropriate weighting factor. The measurement may provide a more global assessment of pulmonary edema.
FIG. 4 shows another illustrative convergent bioelectric field electrode configuration that is highly sensitive to changes in lung resistivity, suitable for impedance measurements to assess pulmonary edema. Left ventricular lead 34 extends from a port of IMD14 and follows a routing path as described above with reference to fig. 2, with proximal and distal left ventricular electrodes 16, 18 each located epicardially over left ventricle 40. A defibrillator lead 80 is connected to a port of the IMD14 and contains a defibrillator coil 82 located in the superior vena cava 38 of the person 30. The defibrillation coil 82 is optionally located in the brachiocephalic vein, subclavian vein, or another suitable nearby location. For simplicity, the defibrillation lead 80 is shown in fig. 4 terminating at a defibrillation coil electrode 82. In practice, the defibrillation lead 80 may extend into the right atrium and right ventricle and may include a right ventricular defibrillation coil (not shown). Right atrial or right ventricular tip or ring electrodes, or any combination thereof, may also be included.
In operation, a pulse generator within the IMD14 may inject an electrical current between the can electrode located on an outer surface of the IMD14 and the distal left ventricular electrode 18. The injected current flows through at least a portion of the left lung 44 and induces a voltage that is measurable by the defibrillation coil 82 and the proximal left ventricular electrode 16. An impedance calculation module within the IMD14 then calculates an impedance by taking the ratio of the measured voltage to the injected current, which can be used to assess pulmonary edema. Arrows 42, 84 indicate the general orientation of the two bioelectric lead fields in this configuration. Lead field 42 corresponds to current injection, while lead field 84 corresponds to voltage measurement. Fig. 4 shows a convergent configuration of lead fields 42, 84, converging near the anatomical region of interest, but substantially separated from each other. In one embodiment, the left ventricular electrodes 16, 18 are spaced approximately 15mm apart and the can electrode and the defibrillation coil 82 are spaced approximately 12cm apart. In other embodiments, the can electrode and the defibrillation coil 82 may be spaced approximately 9-20cm apart. Similarly, the spacing between left ventricular electrodes may be about 25mm in other embodiments, and in the range of about 5-30 mm.
In the configuration shown in fig. 4, IMD14 may be a cardiac defibrillator and concurrently perform pacing functions. The leads 34, 80 may contain additional electrodes. It is also possible to combine the measurements obtained from the arrangements shown in fig. 2-4. Such impedance measurements may be averaged with appropriate weighting coefficients to obtain a more reliable or global impedance value that may be used to more globally assess pulmonary edema. Moreover, the roles of the current injection electrode and the voltage measurement electrode can be interchanged, and equivalent results can be achieved. This can be deduced on the basis of reciprocal theory known to the person skilled in the art.
Fig. 5 shows a table 100 of computer simulation results obtained using computer simulation techniques. A three-dimensional computer model obtained using Magnetic Resonance Imaging (Magnetic Resonance Imaging) divides the human thorax into a number of volume units, each corresponding to human tissue. This model was used to model the pulmonary impedance in the normal (baseline) case and in the severe pulmonary edema case. Each small tissue volume unit is assigned an appropriate resistivity according to published tables (e.g., 150 ohm-cm for blood, 1400 ohm-cm for normal lung, 400 ohm-cm for muscle, etc.). The electrodes were then placed at different locations in the model and current was injected. The computer then uses the electric field equations to calculate the voltage potential at each volume unit. The impedance can be calculated using these results by dividing the measured potential by the injected current. When employing the convergent bioelectric field configuration shown in fig. 2-4, computer simulations used a bipolar left ventricular lead and injected current between the distal left ventricular coronary vein electrode and the can electrode of an IMD implanted in the left pectoral region. The voltage between the proximal left ventricular coronary vein electrode and a fourth electrode (located at different locations within or near the heart to simulate various configurations) was then measured. Each row 102 in table 100 represents a different convergent bioelectric field electrode configuration, corresponding to the configurations shown in FIGS. 2-4.
The first row 102a shows data for a configuration in which the fourth electrode is positioned within the brachiocephalic vein and includes a 15mm spacing between the left ventricular electrodes (i.e., 15mm spacing between proximal and distal left ventricular coronary vein electrodes). The second row 102b represents a similar configuration, but with a 25mm spacing between left ventricular electrodes. Rows 102a and 102b represent simulation data for a configuration of the type described above with reference to fig. 3. The third and fourth rows 102c, 102d represent data for a configuration with a fourth electrode in the superior vena cava to simulate the configuration shown in fig. 4. The data in row 102c corresponds to a higher fourth electrode coil implant location within the SVC, while the data in row 102d corresponds to a lower fourth electrode coil implant location within the SVC, i.e., a change that may occur in medical practice to obtain sufficient defibrillation thresholds. The fifth and sixth rows 102e, 102f represent data for a configuration with a fourth electrode in the right atrium to simulate the configuration shown in fig. 2. Row 102e simulates a 15mm left ventricular electrode spacing arrangement, while row 102f simulates a 25mm spacing arrangement. These measurements are taken at end-diastole.
The present model simulates pulmonary edema by gradually decreasing the resistivity of the model volume units corresponding to lung tissue, for example, from 1400 ohm-cm (healthy) to 350 ohm-cm (severe edema). As shown in Table 100, for a given configuration, the impedance measurements for an edematous patient are less than the corresponding measurements for a normal person. For example, in the brachiocephalic, 15mm configuration (row 102a), the impedance measurement at end diastole for a normal person is 34.62 ohms, which is 34.69% different than the 22.61 ohms measurement for an edematous patient. Similarly, the calculated impedance change was 40.67% for the brachiocephalic, 25mm configuration (row 102b), 40.67% and 45.08% for the SVC configuration (rows 102c, 102d, respectively), and 34.44% and 51.60% for the right atrial configuration (rows 102e, 102f, respectively).
Fig. 6 is a bar graph 120 highlighting the beneficial effects of the present invention by comparing the simulated percent change in impedance of the 4 tripolar non-convergent configurations (labeled RV, RV coil, RA, LVCV) disclosed in U.S. patent application 10/303,305 (filed 2005-25/11, applicants, Andres belalazar and Robert Patterson (inventor of the present invention) with the simulated percent change in impedance of the quadrupolar convergent bioelectric field electrode configuration described herein. The vertical axis of the histogram 120 indicates the percent change in measured impedance between its healthy and edematous states for a given patient. The rightmost 6 columns of histogram 120 correspond to the convergent bioelectric field electrode configuration described above with reference to table 100 of figure 5. As shown in fig. 6, each convergent bioelectric field electrode configuration has a significantly increased sensitivity to lung changes compared to the non-convergent configuration. For example, the simulated three-pole configuration measured impedance changes of approximately 9% (RV), 11% (RV coil), 12% (RA), and 24% (LVCV), while the convergent bioelectric field electrode configuration measured impedance changes of 34.69% (brachiocephalic ring, 25mm), 40.67% (brachiocephalic ring, 15mm), 45.08% (SVC high, 25mm), 48.50% (SVC low, 25mm), 34.44% (RA ring, 15mm), and 51.60% (RA ring, 25 mm). Thus, impedance measurements made using these convergent bioelectric field electrode configurations may provide increased sensitivity to changes in lung resistivity, which may allow for better assessment of pulmonary edema.
Fig. 7-9 illustrate convergent bioelectric field electrode configurations that are highly sensitive to myocardial resistivity changes in the left ventricle and may be used to detect and assess myocardial ischemia. Myocardial ischemia or an acute ischemic event (heart disease) can be detected and evaluated by measuring the myocardial impedance and recording the impedance changes. Such impedance changes may be indicative of, for example, an ischemic event that may occur when plaque within a coronary vessel ruptures and causes a thrombotic occlusion. It is well known that myocardial resistivity increases during ischemic events; for example, the myocardial resistivity may double in an acute ischemic event. See Yorocuuhtli Salazazar et al, "Transmural and non-Transmural In Situ Impedance Spectroscopy comparisons of healthy, Ischemic and Healed Myocardium (Transmural Versus Nontranspiral In Situ Electrical Impedance Spectrometry for health, Ischemic, and healthy myocarpium)", 51IEEE Transactions on biomedical engineering 1421, Aug.2004. Timely detection of such impedance changes may allow for thrombolytic or angioplasty intervention, thereby preventing or minimizing permanent myocardial damage from ischemic events and preventing subsequent ventricular fibrillation. Early detection may even save lives. Patients treated with the present method represent a cost reduction for the healthcare system compared to those who sustain more severe cardiac damage. Moreover, impedance measurements using configurations such as those described in FIGS. 7-9 can be used to monitor myocardial tissue following an ischemic event to monitor the healing of the tissue and whether scarring has occurred. This is because the resistivity of scar tissue is about half that of healthy myocardial tissue. See Salazar et al, supra. Thus, abnormally low impedance measurements taken after an ischemic event indicate scarring of the tissue.
Referring now to fig. 7, a human body 30 is shown with an IMD14 implanted in a left pectoral region of the human body 30. IMD14 may be a cardiac pacemaker, a cardiac defibrillator, a combination thereof, or any other type of implantable medical device capable of delivering electrical current stimulation to thoracic tissue and measuring voltage. Leads 32, 200, 34 are connected to ports of IMD14 and extend therefrom. Leads 32, 200, 34 each house one or more conductors therein that electrically connect IMD14 to electrodes 20, 16, 18, 202 implanted within heart 12.
In this illustrative configuration, lead 32 is a right atrial lead with a right atrial electrode 20 positioned in the right atrium 36 of heart 12. Lead 200 is a defibrillator lead and carries a defibrillator coil 202 in the right ventricle 204 of heart 12 near the distal end. The defibrillation lead 200 extends from a port of the IMD14, is introduced into the venous system, down the SVC38, into the right atrium 36, and into the right ventricle 204. Right atrial electrode 20 may be a ring electrode or a tip electrode. Alternatively, right atrial electrode 20 may be located on another lead, such as a right ventricular lead or a left ventricular lead. Lead 34 is a left ventricular lead that includes a proximal left ventricular electrode 16 and a distal left ventricular electrode 18, both positioned epicardially over a left ventricle 40 of heart 12. Left ventricular lead 34 extends from a port in IMD14 and is introduced into the venous system, down SVC18, into right atrium 36, into the coronary sinus, and then into the coronary veins running epicardially over left ventricle 40. The left ventricular electrodes 16, 18 may be ring or tip electrodes, or may be located elsewhere along the lead 34. Although the left ventricular lead 34 is shown as a bipolar lead in fig. 7, the lead 34 may alternatively carry additional or fewer electrodes, and the lead 34 may follow different paths through the heart 12. For example, the left ventricular lead 34 may include 3, 4, 5, or more electrodes, which may be arranged co-linearly. In one embodiment, the left ventricular electrodes 16, 18 are spaced apart by about 10-15mm, or about 1/6 the distance from the apex to the fundus. These dimensions are suitable for hearts with apical-fundus lengths of about 8 cm. Other electrode spacings may be used to adjust the ratio to maintain 1/6 for different sized hearts to facilitate the convergence of the proportional lead field through the left ventricle. The defibrillator 200 may also include additional or fewer electrodes and may follow different paths through the heart 12 as shown in fig. 7.
In the configuration shown in fig. 7, a pulse generator internal to IMD14 may inject an electrical current between defibrillation coil electrode 202 positioned in the right ventricle 204 and the distal left ventricular electrode 18. The injected current passes through at least a portion of the myocardium and induces a potential difference (voltage) that can be measured by the right atrial electrode 20 and the proximal left ventricular electrode 16. An impedance calculation module within IMD14 may then calculate the impedance by calculating the ratio of the measured potential difference to the injected current. The impedance may be used to assess myocardial ischemia within the human body 30, as will be described in more detail below. Arrows 206, 208 represent the general orientation of the two bioelectric lead fields in this configuration. Lead field 206 corresponds to current injection and lead field 208 corresponds to voltage measurement and converges near the anatomical region of interest (left ventricular wall myocardium), yet is substantially separated from each other. In one embodiment, the left ventricular electrodes 16, 18 are spaced approximately 15mm apart and the distance between the right atrial electrode 20 and the defibrillation coil 202 is approximately 10 cm. In other embodiments, the distance between the right atrial electrode 20 and the defibrillation coil 202 is approximately 4-12 cm. Similarly, the left ventricular electrode internal spacing in other embodiments is about 25mm, and varies within a range of about 5-30 mm.
Fig. 8 shows a configuration that includes a defibrillation lead 220 attached to a port of IMD14 and having a proximal defibrillation coil 82 located within the SVC38 and a distal defibrillation coil 202 located within the right ventricle 204. A pulse generator within the IMD14 may inject an electrical current between the distal defibrillation coil electrode 202 positioned in the right ventricle 204 and the distal left ventricular electrode 18. The injected current passes through at least a portion of the myocardium and induces a potential difference (voltage) that can be measured by the proximal defibrillation coil 82 and the proximal left ventricular electrode 16 located within the SVC 38. An impedance calculation module within IMD14 may then calculate the impedance by calculating the ratio of the measured potential difference to the injected current. Arrows 206, 222 represent the general orientation of the two bioelectric lead fields in this configuration. Lead field 206 corresponds to current injection and lead field 222 corresponds to voltage measurement. The lead fields 206, 222 converge near the anatomical region of interest, yet are substantially separated from each other. In one embodiment, the left ventricular electrodes 16, 18 are spaced apart by approximately 15mm, and the distance between the proximal defibrillation coil 82 and the distal defibrillation coil 202 is approximately 10 cm. In other embodiments, the distance between the proximal defibrillation coil 82 and the distal defibrillation coil 202 is approximately 4-12 cm. Similarly, the left ventricular electrode internal spacing in other embodiments is about 25mm, and varies within a range of about 5-30 mm.
The configuration shown in fig. 9 is similar to that shown in fig. 8, but the left ventricular coronary vein electrodes 16, 18 are spaced apart more in fig. 9. In one embodiment, the proximal left ventricular electrode 16 and the distal left ventricular electrode 18 on the left ventricular lead 34 are spaced 25mm apart. Said ratio corresponds to approximately 1/3 in the case of a heart with a fundus-apex length of 8 cm. Other embodiments maintain the same ratio with different heart sizes to achieve similar proportional convergence of the lead field on the left ventricular wall. In other embodiments, the left ventricular electrodes 16, 18 may be spaced apart by 5mm, 8mm, 10mm, 15mm, 20mm, 30mm, or any other suitable distance, such as any distance between 5-30 mm. A pulse generator within the IMD14 may inject an electrical current between the distal defibrillation coil electrode 202 positioned in the right ventricle 204 and the distal left ventricular electrode 18. The injected current passes through at least a portion of the myocardium and induces a potential difference (voltage) that can be measured by the proximal defibrillation coil 82 and the proximal left ventricular electrode 16 located within the SVC 38. An impedance calculation module within IMD14 may then calculate the impedance by calculating the ratio of the measured potential difference to the injected current. Arrows 250, 252 indicate the general orientation of the two bioelectric lead fields in this configuration. Lead field 250 corresponds to current injection and lead field 252 corresponds to voltage measurement. The lead fields 250, 252 converge near the anatomical region of interest, yet are substantially separated from each other, thereby enabling sensitive and specific pathology detection within the region of interest and providing improved clarity of resistivity changes outside the region of interest.
Like the configurations described with reference to fig. 7-9, the convergent quadrupole configurations described above with reference to fig. 2-4 also have advantages for enabling impedance measurements useful for assessing ischemia. The configurations shown in fig. 2-4 are highly sensitive to myocardial resistivity changes in the left ventricle and can therefore be used to detect and assess myocardial ischemia. An impedance calculation module within IMD14 may calculate impedance by calculating a ratio of the measured potential difference to the injected current, and the impedance measurements may be used to detect and assess myocardial ischemia. Furthermore, impedance measurements using the configurations described in fig. 2-4 can be used to monitor myocardial tissue following an ischemic event to monitor whether the tissue heals well and whether scarring occurs.
FIG. 10 is a bar graph 300 highlighting the benefits of the present invention by comparing the simulated left ventricular wall impedance sensitivity percentage for a non-convergent dipolar impedance configuration (bar 302, labeled dipolar LV coil) with a quadrupolar convergent bioelectric field electrode configuration (bar 304) as described above in FIG. 2 and FIGS. 7-9. Bar 304a corresponds to the configuration described in fig. 2, bar 304b corresponds to the configuration described in fig. 7, bar 304c corresponds to the configuration described in fig. 8, and bar 304d corresponds to the configuration described in fig. 9. The vertical axis of the histogram 300 represents the left ventricular wall impedance sensitivity. As shown in fig. 10, each convergent bioelectric field electrode configuration (bar 304) exhibited a significant increase in left ventricular wall impedance sensitivity when compared to the non-convergent configuration (bar 302). For example, the analog dipole configuration (bar 302) has a sensitivity of about 6% for the left ventricular wall myocardium, while the convergent bioelectric field electrode configuration has a sensitivity of about 34% (bar 304a), 35% (bar 304b), 36% (bar 304c), and-11% (bar 304 d). Negative sensitivity indicates that the measured impedance will drop as the left ventricular wall resistivity increases, and vice versa. Thus, impedance measurements made using these convergent bioelectric field electrode configurations are sensitive to changes in left ventricular wall resistivity, allowing for better assessment of ischemia.
Fig. 11 is a block diagram circuit representative of the implantable device 14 of fig. 1-4 and 7-9. Fig. 11 will be described below with reference to the configuration of fig. 2, and then those skilled in the art will understand that the described portions can also be applied to other configurations already described and illustrated in this specification. The apparatus 14 comprises: circuitry for measuring impedance and assessing pulmonary edema and myocardial ischemia, and communications circuitry for interacting with external devices. The impedance measurement circuit 302 includes a current generator 304 that injects a current between any two electrodes; for example, the current generator 304 may inject current between a can electrode located on an outer surface of the IMD14 and the distal left ventricular coronary vein electrode 18 (fig. 2). A switch 306 that directs the current to the appropriate port 308 in a manner known in the art. Current flows from the appropriate port along the wire to the appropriate electrode. The injection current may be an Alternating Current (AC) to prevent undesirable polarization and electrolytic degradation effects at the electrodes, and should be at a flow rate, frequency and duration that will not cause cardiac stimulation. In practice, the frequency of the AC current is about 50-100 kilohertz. Examples of useful current waveforms are sine waves and biphasic pulses (symmetrical or otherwise).
The injection of current between the can electrode and the electrode 18 (see fig. 2) establishes an electric field in the patient. Thus, a voltage potential appears between the right atrial electrode 20 and the proximal left ventricular coronary artery electrode 16. A voltage amplifier 310 is then used to measure the voltage between electrode 20 and electrode 16 in the conductor and on the conductor through port 308 and switch 306. The voltage amplifier 310 may be a signal conditioning unit that measures voltage and may optionally include a demodulator. Alternatively, the roles of the proximal and distal left ventricular coronary vein electrodes 16, 18 may be reversed with appropriate adjustment of the wiring.
The control module 312 receives or stores numerical information of the injected current and the resulting measured voltage. An analog/digital (a/D) converter (not shown) located inside or outside of the control module 312 may be used to interpret this information. A processing unit (not shown), such as a microprocessor, microcontroller, or digital signal processor within the control module 312, may then use the current and voltage information to calculate the impedance by dividing the voltage by the current. As body interstitial fluid increases, tissue impedance decreases. Thus, the impedance ratio can be used to assess pulmonary edema and determine the extent of pulmonary edema in a patient. The algorithm describing the edema value determination will be described later. Conversely, myocardial tissue impedance may increase during an ischemic event. Thus, the impedance values are also used to assess myocardial ischemia, and an ongoing ischemic event may be detected for the patient. Algorithms describing ischemia detection will be discussed later.
The control module 312, as is conventional, may additionally include Read Only Memory (ROM), Random Access Memory (RAM), flash memory, EEPROM memory, etc., which store instructions that are executed by the processing unit as well as digital/analog (D/a) converters, timers, counters, filters, switches, etc. (not shown). Impedance measurements, edema values, and ischemia values may also be stored in memory. These control module components may be integrated into a single device, such as an Application Specific Integrated Circuit (ASIC), or alternatively may be discrete devices. A suitable bus (not shown) enables communication between the various components within the control module 312.
Information from the sensor module 314 may be used to adjust the relationship between the measured impedance and the degree of edema or ischemia. Posture sensor 316 can provide patient orientation information to control block 312, incorporating posture compensation into the assessment of edema or ischemia. Since the tissues and excess fluid in the chest and lungs change with position due to gravity, the measured impedance also changes with the different positions the patient is in. For example, when the patient lies on the right side, the fluid and tissue in the left lung 44 may be gravitated toward the mediastinum near the left ventricular coronary vein electrodes 16, 18, resulting in a lower measured impedance. Thus, based on postural awareness information, the relationship between the impedance measurement and the degree of edema or ischemia may be adjusted to compensate. Similarly, the relationship of the left lying patient may be adjusted in reverse. Various types of position sensors may be used, such as mercury switches, DC accelerometers, or other piezoelectric devices.
An activity sensor 318, which is typically used to assist the pacing application, also provides information to the control module 312. By using these compensation schemes, misjudgment of edema and ischemia caused by postural fluid movement in the patient's body can be avoided. The sensors 316, 318 can each be selectively excluded from the implantable device 14.
Telemetry module 320 may communicate wirelessly using Radio Frequency (RF) transmissions through antenna 322 with a similarly wirelessly equipped external monitoring unit 324. The monitoring unit 324 may be a computer (custom programmer, desktop computer, portable computer, handheld computer, etc.), a telemedicine local station, a wearable device (such as a watch, mobile phone, portable transponder, or any other suitable device), and may be used to program the implantable device 14 or retrieve information (such as impedance measurements, edema values, or ischemia values) from the IMD 14. The communication link may be used to alert a physician or healthcare provider to an acute ischemic event or to detect pulmonary edema, for example, to promptly initiate a medical intervention. Alternatively, monitoring unit 324 may connect to direct 911 using a telephone and summon an emergency team, and may summon a similar rescue via network contact (like the internet), or notify the patient in voice or text form to seek medical attention. In this manner, the patient may be continuously monitored for various pathological tests for 1 day, 24 hours, 1 week, 7 days, and the physician or caregiver is quickly alerted in the event of a pathological test.
The sensing/pacing/defibrillation circuit 330 includes a pacing circuit 332, a defibrillation circuit 334, and a sense amplifier 336 for sensing and/or stimulating (pacing) cardiac events and controlling the heart rhythm. A typical impedance calculation module is not explicitly shown in fig. 11, but may include a plurality of modules, or portions thereof, of fig. 11. Battery 340 provides electrical power to various circuits and IMD14 modules (connections not shown in fig. 11 for simplicity). Alternatively, the impedance measurement circuit 302 can measure the resulting voltage using the normal cardiac stimulation pulse from the sensing/pacing/defibrillation circuit 330 instead of the current injection from the pulse generator 304. The impedance is then calculated. Still alternatively, instead of the impedance measurement circuit 302, the sensing/pacing/defibrillation circuit 330 may be used to implement the current injection and voltage measurement functions necessary to determine impedance.
As previously mentioned, impedance measurements may be affected by resistivity contributions from different organs and/or human tissue. For a given electrode measurement configuration, some organs or human tissue contribute more significantly to the overall impedance measurement, while others contribute less significantly. It is desirable that the target organ/tissue contribute more to the impedance measurement and all other organs/tissues contribute the least to the measurement. For example, when measuring impedance to assess pulmonary edema, it is desirable to see high sensitivity to the lungs, and minimal sensitivity to all other thoracic organs/tissues. Similarly, when measuring impedance to assess myocardial ischemia, it is desirable to see high sensitivity to the myocardium, and minimal sensitivity to all other thoracic organs/tissues.
FIG. 12 is a table 400 showing impedance sensitivity coefficient simulation results compiled using models of the various electrode configurations described above. Each configuration is represented in table 400 by columns 401, 402. Column 401a represents the configuration described in U.S. patent application Ser. No. 10/303,305 (filed 2002, 11/25, and filed by Andres Belalcazar and RobertPatterson (inventor of the present invention) and Rebecca Shult), column 401b represents a well-known configuration, and columns 402a-402b represent the convergent bioelectric lead field electrode configuration described herein. Columns 402a and 402b correspond to the configuration described above with reference to FIG. 3; column 402c corresponds to the configuration described above with reference to FIG. 4; columns 402d and 402e correspond to the configuration described above with reference to fig. 2. For each configuration 401, 402, table 400 shows the quantitative contribution data of various thoracic organs and tissues to the total impedance measurement for that configuration. The higher the coefficient value, the more significant its contribution to the total impedance. For example, when monitoring the lungs, it is desirable that the lung coefficient be as high as possible, while all other coefficients are relatively low. This ensures that the measured impedance changes are lung-specific and that the lung impedance results vary widely as desired. Similarly, for monitoring the left ventricular wall, a high left ventricular wall coefficient is desired, while other coefficients are low.
Table 400 shows that with the convergent bioelectric field electrode configuration (columns 402a-402e), the sensitivity to fat (row 403) and muscle (row 404) near IMD14 is reduced by nearly 10 times (e.g., from 0.2199 and 0.1888 in the tripolar configuration (columns 401a, 401b) to less than 0.022 in the convergent quadrupolar configuration (columns 402a-402 e)). Thus, the convergent bioelectric field electrode configuration is more resistant to postural or other changes in the shoulder and pectoral muscle regions. As expected, the loss of sensitivity results in higher sensitivity to other regions, in particular, to the left ventricular wall and blood (row 406), and to the left lung (row 405). Thus, these configurations are less sensitive to impedance changes caused by arm or shoulder movements (which expand and contract the pectoral muscles in the vicinity of IMD 14), or by pectoral muscle edema, and more sensitive to impedance changes caused by the left lung and left ventricular wall, thus allowing better detection and assessment of pulmonary edema and ischemia.
Table 400 also shows that intra-electrode spacing (15mm and 25mm) of the left ventricular electrodes 16, 18 is an important factor. More distally spaced electrodes provide deeper penetration of the sensing region, which is desirable in the case of lung monitoring and undesirable in the case of left ventricular wall monitoring. For example, the left lung (row 405) sensitivity coefficient for the brachiocephalic configuration with a 25mm left ventricular electrode internal spacing (column 402b) is 0.3196, which is more sensitive to left lung resistivity changes than the coefficient of 0.2548 for a 15mm spacing (column 402 a). Similarly, the right atrium 25mm configuration (column 402e) has a left lung (row 405) sensitivity coefficient 0.4175 compared to a coefficient 0.2493 for 15mm spacing (column 402 d). However, the brachiocephalic 15mm configuration (column 402a) has a left ventricular sidewall (row 406) sensitivity coefficient 0.2514 of-0.0806 compared to a 25mm spacing (column 402b), and the right atrial 15mm configuration (column 402d) has a coefficient 0.3394 compared to a 25mm spacing (column 402e) of-0.1950, both demonstrating better sensitivity to left ventricular wall resistivity. As previously mentioned, the dimensions of 15mm and 25mm are suitable for a typical heart with an apical-to-fundus distance of 8 cm. The intra-electrode spacing can be scaled based on other sizes of the heart to ensure continuous convergent penetration of the lead field through the tissue of interest, thereby enabling assessment and detection of anatomically different sizes.
The table 400 was calculated using the chest calculation module described above to simulate the deployed lead field. The contribution of each different volume unit consisting of an organ or tissue is calculated by the volume integral of the Schmitt-Geselowitz equation as shown in the following equation 1:
(1)Z＝∫ρJLE·JLIdν
in equation 1, Z is an impedance (ohm) obtained by the measuring device by calculating a ratio of the measurement voltage of the acceptance electrode to the injection current of the output electrode. ρ is the resistivity of the tissue at each site (ohm-cm); JLE is the lead field vector (1/cm ^2) of the voltage measuring electrode pair; j. the design is a squareLIA wire field vector (1/cm ^2) for current injection electrode pair; d ν is the volume integral derivative (cm ^ 3). Equation 1 shows that integrating the scalar product of the lead field weighted by the tissue resistivity of each region of each organ volume element yields the scalar product corresponding to the groupThe ohmic value of the contribution to the total measured impedance. The total impedance displayed by the system is divided by the contribution of the organ in ohms to yield the sensitivity coefficients listed in table 400. For further details of the meaning of terms within the above equation, see j&Plonsey, biomagnetic: the principle and application of Bioelectric Fields (Bioelectromagnetism: Principles and Applications of Bioelectric Fields) (1995).
The flow chart of fig. 13 is an example of how an algorithm may be applied to the control block 312 of fig. 11 for edema assessment. The process implemented by the control module processing unit executing the instructions begins at step 500 with the wait time being executed by an edema timer (e.g., located within the processing unit of the control module 312). The procedure may be from about 2 hours to 3 days, as determined by the physician. The values may be programmed to connect to telemetry module 320 (fig. 11) via radio frequency.
After the wait time has elapsed, the control module 312 waits for the next ventricular event in step 510. The ventricular event may be determined by pacing timing control information resident in control module 312 or obtained from a sense amplifier 336 in sense/pace/defibrillation circuit 330 (fig. 11). The occurrence of a ventricular event indicates that the heart is beginning to contract and prompts a wait time of approximately 150ms in step 520 to allow the heart to contract sufficiently. Next, in step 530, the impedance is sampled 10 times at a rate of 25Hz and the impedance samples are stored in the memory buffer Ce (for heart/edema). This facilitates sampling of the impedance before and after the impedance waveform (i.e., the waveform defined by the impedance measurements as the waveform constituent points) to obtain and determine the end-systolic value of the peak, referred to herein as Zes _ s (for Z end-systole for edema detection). Zes _ s is set equal to the maximum of 10 impedance samples in buffer Ce and buffer Re is stored and stored (for breath/edema), and buffer Ce is emptied in step 540.
In step 550, the counter determines whether 48 Zes _ s's have been stored. If not, step 510 and 540 are repeated. In this way, step 510-540 is repeated 47 times, so that the buffer Re is full of 48 end-systolic impedance measurements Zes _ s, sufficient to cover up to about 3 respiratory cycles. Next, in step 560, the pulmonary edema value is set to the median of the 3 smallest impedance values in the buffer Re (i.e., the value corresponding to the end of expiration). The pulmonary edema value may be stored in memory, buffer Re cleared, appropriate timers and counters zeroed (570), and the process ends. The process may then begin again at step 500, waiting until the next edema sampling time. For stability and repeatability of the measurements, measurements need to be taken at the same time within the cardiac and respiratory cycles; for example, measurements are taken at end systole and end expiration (as described above in connection with the flowchart of fig. 13), or at end diastole and end expiration, or at any other point in the cardiac cycle.
The edema value may be compared to a stored edema threshold (which may be programmed on a telemetry link), a warning flag may be set when the edema value exceeds the threshold, or an alarm may be triggered. The stored edema or impedance values may then be transmitted by telemetry module 320 to monitoring station 324 (fig. 11) when monitoring station 324 interrogates device 14. The physician can then analyze the data to determine the trend of the edema values.
The flow chart of FIG. 14 is an example of how an algorithm may be applied to the control block 312 of FIG. 11 to perform ischemia assessment. The process implemented by the control module processing unit executing the instructions begins at step 600 with the execution of a wait time by an ischemia timer (e.g., located within the processing unit of the control module 312). The procedure may be from about 15 seconds to 5 minutes, as determined by the physician. In an embodiment, the wait time may be about 1 minute. The values may be programmed to connect to telemetry module 320 (fig. 11) via radio frequency.
Referring again to FIG. 14, after the wait time has elapsed, the control module 312 waits 610 for the next ventricular event. The ventricular event may be determined by pacing timing control information resident in control module 312 or obtained from a sense amplifier 336 in sense/pace/defibrillation circuit 330 (fig. 11). The occurrence of a ventricular event indicates that the heart is beginning to contract and prompts a wait time of approximately 150ms in step 520 to allow the heart to contract sufficiently. Next, in step 630, the impedance is sampled 10 times at a rate of 25Hz and the impedance samples are stored in a memory buffer Ci (for heart/ischemia). This facilitates sampling of the impedance before and after the impedance waveform (i.e., the waveform defined by the impedance measurements as the constituent points of the waveform) to obtain and determine the end-systolic value of the peak, referred to herein as Zes _ i (for Z end-systole for ischemia detection). Zes _ i is set equal to the maximum of 10 impedance samples in buffer Ci and stored in memory buffer Ri (for breath/ischemia), and buffer Ci is emptied in step 640.
In step 650, a counter determines whether 48 Zes _ i are stored. If not, steps 610-640 are repeated. In this way, the steps 610-640 are repeated 47 times, so that 48 end-systolic impedance measurements Zes _ i are full in the buffer Ri memory, sufficient to cover up to about 3 respiration cycles. Next, in step 660, the lung ischemia value is set to the middle of the 3 minimum impedance values in the buffer Re (i.e., the value corresponding to the end of expiration). The lung ischemia value may be stored in memory, buffer Ri cleared, appropriate timers and counters reset 670, and the process ends. The process may then begin again at step 600, waiting until the next ischemia sampling time. For stability and repeatability of the measurements, measurements need to be taken at the same time within the cardiac and respiratory cycles; for example, measurements are taken at end systole and end expiration (as described above in connection with the flowchart of fig. 13), or at end diastole and end expiration, or at any other point in the cardiac cycle.
The ischemia value may be compared to a stored ischemia threshold (which may be programmed on the telemetry link), an alarm flag may be set when the ischemia value exceeds the threshold, or an alarm may be triggered. The stored ischemia value or impedance value may then be transmitted by telemetry module 320 to monitoring station 324 (fig. 11) when monitoring station 324 interrogates device 14. The physician can then analyze the data to determine the trend of the ischemia values.
In an embodiment, edema and ischemia may be assessed at approximately the same heart rate and posture. One way that this can be done is by starting to perform the steps of the flow chart shown in fig. 13 or fig. 14 only when the heart rate and posture are within a preset programmed range (i.e., only when the heart rate and posture are about the same as when the last edema or ischemia assessment was completed). Alternatively, edema and/or ischemia assessments may be stored in memory along with corresponding heart rate and posture information according to the steps of fig. 13 or 14. The edema and/or ischemia measurements may then be categorized into bins according to heart rate range or posture category and communicated or delivered to the physician corresponding to his/her selected category.
There are many alternative forms of algorithms. For example, a different number of samples may be used, the sampling rate may be varied, the mean value may be substituted for the median value, different latencies may be selected, and different comparison schemes may be performed. For example, a single register may be used in place of the buffer and performed in parallel with the sampling, with the newly sampled measurement compared to the ongoing maximum/minimum values stored in the register. As an alternative to waiting time, the patient may start a series of measurements with a magnetic card. The impedance value and sensor information may be used to assign an edema value or an ischemia value that is different from the measured impedance value. Of course, the time stamp may also be stored with the edema or ischemia values and other relevant information (e.g., posture information, heart rate, or activity level). The telemedicine local station may also initiate the measurement and then transmit the result to the medical center. The impedance measurements may be time-multiplexed with the normal cardiac stimulation events. Alternatively, normal cardiac stimulation pulses may be used in place of dedicated impedance measurement current injections.
Flow chart fig. 15 is a flow chart of how an algorithm may be applied in the control block 312 of fig. 11 to make impedance measurements at different frequencies to discern effects on different pathologies, such as edema and ischemia, which may have confounding effects on single frequency measurements. It is known that the myocardium in ischemic conditions exhibits a significantly different impedance response to different frequencies, such as between 1 khz and 500 khz, while lung tissue and edema fluid do not exhibit much difference to frequencies below 1 mhz. See Salazar et al, supra. The degree of myocardial ischemia has little effect on impedance measurements taken at 500 khz, whereas at 1 khz the degree of ischemia has a significant effect on the impedance Id. Thus, multiple frequency measurements can be made and the results used to classify changes in impedance, for example, due to pulmonary edema and myocardial ischemia.
Referring again to FIG. 15, the process by the control module processing unit executing the instructions begins at step 700 with a first impedance measurement being taken at an injection current frequency of 1 kilohertz and a second impedance measurement being taken at an injection current frequency of 500 kilohertz. The injection current may be a sine wave of 1 kilohertz and 500 kilohertz, respectively, or may be any waveform that causes a response similar in frequency content to the sine wave excitation, for example, a rectangular exponential decay pulse in single phase or biphasic form. Next, step 710 calculates Δ _ Z _1k from the absolute value of the difference between the first impedance measurement value and the first reference value, and calculates Δ _ Z _500k from the absolute value of the difference between the second impedance measurement value and the second reference value. The reference value may be programmed by, for example, a physician at the time of implantation of the device 14, or updated later by the physician at a follow-up visit. Alternatively, the reference value may be updated by the control unit 312 (fig. 11). For example, the control unit 312 may cause an update of the reference value after performing the calculation, such as after determining a running average of the impedance measurements or some other adaptive technique.
If Δ _ Z _1k is less than the first threshold (Thresh _1k) and at the same time Δ _ Z _500k is less than the second threshold (Thresh _500k) (720), then there is no emergency and the process may resume after an appropriate wait time. Otherwise, if at least one of the delta values exceeds its corresponding threshold, edema or ischemia may be present. The process continues from step 730 to determine if a pathological condition has occurred. If Δ _ Z _1k is greater than Thresh _1k and Δ _ Z _500k is greater than Thresh _500k, then the pulmonary edema flag or alarm is set at step 740 and the process ends. However, if the condition of step 730 fails, then only one of the two Δ values has a significant change. As previously mentioned, since high frequency measurements are not sensitive to ischemia, if only low frequency measurements change significantly, the change may only be caused by an ischemic event. Step 750 performs: if Δ _ Z _1k is greater than Thresh _1k and Δ _ Z _500k is less than Thresh _500k, then step 760 activates the acute ischemic event flag or alarm and the process ends. If, on the other hand, only the high frequency measurements change (e.g., a dip in impedance at 500 kHz), the change may be due to an edema event (since ischemia has little effect on the impedance measurements at 500 kHz), and an explanation that the concurrent low frequency measurements have not changed significantly since a period of time may be a concurrent ischemic event. In this case, ischemia causes an increase in impedance at low frequencies, such as 1 khz, while edema causes a decrease. If both events occur simultaneously, the results may cancel out and no significant change is observed during low frequency monitoring. The determining step occurs at step 770: if Δ _ Z _1k is less than Thresh _1k and Δ _ Z _500k is greater than Thresh _500k, then step 780 activates the acute ischemic event flag or alarm and the process ends. The flowchart of fig. 15 describes a process for detecting the occurrence of an edema or ischemia event. Edema clearance and regression of ischemic conditions, or a combination thereof, may be detected in a similar process according to similar principles. Thus, the present invention is able to monitor, detect, distinguish and classify edema and myocardial ischemia.
Many different algorithms may be used. For example, impedance measurements may be made more than twice for different frequencies, and alternative frequencies that may be used are 50Hz, 500Hz, 5KHz, 10KHz, 50KHz, 100KHz, 400KHz, 600KHz, 1MHz, etc. Rather than injecting currents at two different frequencies, a mixed signal current with two frequencies can be injected and an auxiliary filtering circuit within the voltage measurement circuit of device 14 can be used to separate the frequency components. The impedance values obtained by the algorithms described by the flowcharts shown in fig. 13-14 may be used. The telemetry module 320 (FIG. 11) transmits a flag or alarm to a monitoring station 324 (FIG. 11) or to a physician or caregiver. The algorithm may be performed at various suitable time intervals, such as every 15 seconds, every 30 seconds, every minute, every 2 minutes, every 5 minutes, every 7 minutes, every 10 minutes, every hour, and so forth. The measurements and calculations may be stored in memory or downloaded to a monitoring station for reference by a physician.
Some embodiments of the invention have been described. Nevertheless, it will be understood that various modifications may be made without departing from the spirit and scope of the invention. For example, using the principles described above for detecting ischemia, left ventricular wall hypertrophy (wall thickening) can also be monitored, which commonly occurs with hypertension (high blood pressure) or aortic valve stenosis. Hypertrophy may be an indication of cardiac overload (an undesirable condition). The difference in time of occurrence can distinguish ischemic time from hypertrophic time. Ischemic events occur rather rapidly, while hypertrophy typically occurs/resolves over several days. Ischemia can also be ruled out using the multi-frequency techniques described above. The polarity of the current injection and the voltage measurement may be interchanged. Moreover, the roles of the current injection electrode and the voltage measurement electrode may be interchanged.
As is well known to those skilled in the art, impedance is a complex quantity defined in the frequency domain, consisting of magnitude and phase angle. The impedance magnitude is the ratio of the magnitude of the measured voltage to the magnitude of the injected current. The impedance phase angle refers to the degree of phase shift between the injected current and the measured voltage. One way to measure the impedance phase shift is to calculate the amount of time between the peak of the injected current signal and the peak of the measured voltage signal. Any of the configurations shown in fig. 2-4 or 7-9 may measure impedance magnitude and/or impedance phase angle for pathology assessment and detection. Accordingly, other embodiments are within the scope of the following claims.