Left Ventricle

Figure 43: The myofibrils of each cardiac muscle cell are branched and contain a single nucleus. The branches interlock with those of adjacent fibers by adherens junctions which act to prevent scission of the cardiomycytes during the high-shear, forceful contractions of the heart. The muscle is richly supplied with mitochondria which are largely confined to the spaces between the fibrils. The fibrils are covered with a membrane, the Sarcolemma, which is frequently invaginated to form the Transverse tubules. These invaginations of the plasma membrane or sarcolemma, are called transverse tubules and they reach deep into the myofibrils and  bring the action potential deep into the fibers. Specialized intercellular junctions, the Intercalated discs, facilitate rapid transmission of the electrical signals which initiate myocyte contraction. The myofibrils are formed by myosin and actin fibers aligned in a distinct pattern which is visible under light microscopy as the A-, H- and I- bands.

      Yajima stain was used to prepare the Control (Figure 44), FGP  and FIG cardiac tissue for light microscopy. The FGP cardiac muscle showed increased interstitial space, probably indicative of interstitial edema. In many areas the sarcolemma appeared to be separated from the cytoplasm of the myocyte and, occasionally, appeared to have disintegrated into debris in the interstitial spaces (Figure 45). The myofilaments appeared maximally relaxed with widened I-bands . The mitochondria were grossly swollen and contained numerous amorphous matrix densities. The sarcolemma was fragmented beneath an intact basement membrane and there was increased space between the capillary endothelium/basement membrane and intact areas of the sarcolemma of the cardiomyocytes. The cell nuclei  were unremarkable.

Figure 44: Control-1, Left  Ventricle, Yajima, 100x. Control cardiac muscle demonstrated crisp, well defined membranes and the normal density and pattern of myofibril structure. Capillary endothelium appeared intact and the capillary basement membrane was well anchored to adjacent myocytes and appeared intact.

Figure 45: FGP-1 Left Ventricle, Yajima, 100x. In the FGP animals the myocardium exhibited increased interstitial space (IIS) as well as the presence of debris in the IIS which appeared to be disrupted sarcolemma (yellow arrows). The capillary basement membrane was often observed to be separated from the sarcolemma of the adjacent myocytes and endothelial cell nuclei were sometimes observed devoid of plasma membranes or cytoplasm (red arrow).The occasional naked myocyte nucleus could also be observed (green arrow).

The same changes were also present in the FIG group with the added presence of a “ragged” or rough appearance of the myofibrils where they were silhouetted against interstitial space (Figure 46). There also appeared to be holes or spaces, possibly as a result of edema, in the fabric of the myofibrils that were not present in the myocardium of either the control, or the FGP animals.

Most surprising was the general absence of contraction band necrosis in the FIG group, possibly as a consequence of the protective effect of reasonably prompt post-cardiac arrest refrigeration. No microscopic evidence of fracturing, either gross or microscopic, was noted in the myocardium of either the FGP, or  the FIG groups.

Figure 46: FIG-2 Left Ventricle, Yajima, 100x. Separation and fragmentation of the sarcolemma were observed in the FIG myocardium to a greater extent than that seen the in myocardium from the FGP animals (yellow arrow). Additionally, the fibers of myofibrils had a more ragged appearance and consistently displayed open spaces in the bands which were not seen in the myocardium of either the Control or the FGP animals (red arrows).

 Figure 47: The myofibrils of both the FGP and FIG animals appeared maximally relaxed with a marked increase in the thickness of the I-band. Intact red blood cells (RBCs) were observed in the FIG animals and represent incomplete blood washout (red cell trapping) despite perfusion with large volumes of washout, cryoprotectant and fixative solution (~8-10 L) over a time course of ~140 minutes of perfusion.

Cerebral Cortex

 Figure 48: The cerebral cortex consists of six distinct layers, beginning with the first layer, the Molecular Layer (Stratum zonale), which consists of finely branched medullated and non-medullated nerve fibers. The molecular layer is largely devoid of neuronal cell bodies. Those neuronal cell bodies which are present are the cells of Cajal which possess irregular cell bodies and typically have four or five  dendrite that terminate within the molecular layer and a long nerve fiber process, or neuraxon, which runs parallel to the surface of the cortical convolutions.

 The second layer of the cortex consists of a layer of small Pyramidal cells with the apices of the pyramids being directed towards the surface of the cortex. The apex of the small Pyramidal cells terminates in a dendron, which reaches into the molecular layer, giving off several collateral horizontal branches. The final branches in the molecular layer take a direction parallel to the surface. Smaller dendrites arise from the lateral and basal surfaces of these cells, but do not extend far from the body of the cell. The neuronal axon (neuraxon) always arises from the base of the small Pyramidal cells and passes towards the central white matter, thus forming one of the nerve-fibers of the white matter. In its path, the neuraxon gives off a number of collaterals at right angles, which are distributed to the adjacent grey matter.

The third cortical layer consists of Pyramidal neurons which are characterized by the presence of cells of the same type as those of the preceding layer, but of a larger size. The nerve-fiber process becomes a medullated  fiber of the white matter.

 The fourth layer is comprised of  Polymorphous neurons which  are irregular in outline and give off several dendrites which branch into the surrounding grey matter. The neuraxons of the Polymorphous neurons give off a number of collaterals, and then become a nerve-fiber of the central white matter. Scattered through these three layers are the cells of Golgi, whose neuraxon divides immediately with the divisions terminating in the immediate vicinity of the Polymorphous neuron cell-bodies. Some cells are also found in which the neuraxon, instead of extending into the white matter of the brain, passes towards the surface of the cortex; these are called cells of Martinotti.

 The fifth cortical layer contains the largest pyramidal neurons which send outputs to the brain stem and spinal cord and comprise the the pyramidal tract. Layer 5 is particularly well-developed in the motor cortex.

 Layer 6 consists of pyramidal neurons and neurons with spindle-shaped cell bodies. Most cortical outputs leading to the thalamus originate in layer 6, whereas most outputs to other subcortical nuclei originate in layer 5.

The cortical blood supply is via the pia mater which overlies the cerebral hemispheres.

Bodian stain was used to prepare the control, FGP, and FIG brain tissue  samples for light microscopy. Three striking changes were apparent in FGP cerebral cortex histology: 1) marked  dehydration of both cells and cell nuclei, 2) the presence of  tears or cuts at intervals of 10 to 30 microns throughout the tissue on a variable basis (some areas were spared while others were heavily lesioned), and 3) the increased presence (over Control) of irregular, empty spaces in the neuropil as well as the occasional presence of large peri-capillary spaces (Figures 54,56, and 57). These  changes were fairly uniform throughout both the molecular layer and the second layer of the cerebral cortex. Changes in the white matter paralleled those in the cortex, with the notable exception that dehydration appeared to be more pronounced (Figure 55).

Other than the  above changes, both gray and white matter histology appeared remarkably intact, and only careful inspection could distinguish it  from control (Figures 52, 58, 59 and 60). The  neuropil appeared normal (aside from the aforementioned holes and tears) and many long  axons and collaterals could be observed traversing the field. Cell membranes appeared crisp, and apart from appearing dehydrated, neuronal architecture  appeared comparable to control. Similarly, staining was comparable to that observed in Control  cerebral  cortex. Cell-to-cell connections appeared largely intact.

The histological appearance of FIG brain differed from that  of FGP animals in that ischemic changes such as the presence of pyknotic and fractured nuclei were much in evidence and cavities and tears in the neuropil appeared somewhat more frequently. The white matter of the FIG animals presented a macerated appearance, in addition to exhibiting the rips or tears observed in the white matter of the FGP brains (Figure 61).

Both FGP and FIG brains  presented occasional  evidence of microscopic fractures.

Figure 49: Control-1, 1st (molecular) cell layer, cerebral cortex, Bodian, 40x. Cells of Cajal (N) and a dense weave of axons (A) are visible. The tissue is perforated by numerous capillaries (C) and a  small venue containing many red blood cells (RBCs).

Figure 50: Control-1, 2nd cell layer, cerebral cortex, Bodian, 40x, showing a pyramidal neuron (N, lower left) multiple capillaries (C) and the interwoven connections of dendrites that comprise the neuropil.

Figure 51: Control-1, white matter, cerebral cortex, Bodian, 40x. Myelinated axons (MA) appear both in cross section (yellow arrows) and laterally (green arrows). Unmyelinated axons are present inside the black circle. Glial cell nuclei (GN) are scattered throughout the tissue.

 Figure 52: FGP-1, Cerebral Cortex, 1st cell layer, Bodian, 40x. Two large capillaries (LC) are present, one with a red blood cell present (right). Neurons (N, cells of Cajal) are present in normal density and the neuropil appears intact. This section appears indistinguishable from that of the Control animal.

Figure 53: FGP-1, Cerebral Cortex, 2nd cell layer, Bodian, 40x. This area of FGP cerebral cortex shows injury typical of that seen in both FGP and FIG animals. There are a number of large tears in the neuropil (red arrows) approximately 10 to 30 microns across. A pyramidal neuron is present in the lower left of the micrograph and it appears somewhat dehydrated. There are a number of naked glial cell nuclei (yellow arrows), as well some nuclei with what appears to adherent cytoplasm visible at the margins of the tears in the neuropil.  

Figure 54: FGP-1, Cerebral Cortex, 2nd cell layer, Bodian, 40x. In this area of the 2nd layer of the cerebral cortex the neuropil presents a somewhat “moth eaten” appearance, with numerous tears and vacuoles in evidence (red arrows). One large tear appears to be a pericapillary ice hole (yellow arrow).

Figure 55: FGP-3, Cerebral Cortex, white matter, Bodian, 100x. There are numerous open spaces in the white matter that appear to be ice holes (red arrows). The density of the tissue appears markedly increased over that of the Control white matter, possibly as a result of glycerol-induced dehydration. This apparent dehydration is also evident in the increased density of the axoplasm seen in the myelinated axons (green arrows).

Figure 56: FIG-3, Cerebral Cortex, 1st cell layer, Bodian, 40x. Extraordinarily normal appearing Molecular layer of the FIG cerebral cortex. The neuropil appears intact with the exception of what appear to be scattered tears or ice holes (red arrows).

Figure 57: FIG-2, Cerebral Cortex, 1st cell layer, Bodian, 40x. Large tears are evident (red arrows) and naked glial cell nuclei and fragmented cytoplasm are apparent (nn). Several intact capillaries are in evidence (C) as well as what appears to be two capillaries that have been separated from the neuropil and appear largely surrounded by open (pericapillary) space (green arrows). A mass of debris appears to occupy some of the luminal space of what appears to have been a capillary (Cd).

Figure 58: FIG-2, Cerebral Cortex, 2nd cell layer, Bodian, 40x. Remarkably intact neuropil with several capillaries, including several capillaries sectioned oblique to the plane of the tissue (OC). A neuron (N) with what appears to be a crisp plasma membrane is present at the upper right of the micrograph.

Figure 59: FIG-2, Cerebral Cortex, 2nd cell layer, Bodian, 40x.Normal appearing cerebral cortex in an FIG animal. There are multiple intact neurons with normal appearing dendrites (D) and axons (A). An intact large capillary (LC) is present and appears free of red cells.

Figure 60: FIG-2, Cerebral Cortex, 2nd cell layer, neuropil, Bodian, 100x. Normal appearing layer 2 of the cerebral cortex with intact neurons (N), axons (A), and neuropil. A capillary (C)with intact endothelial cells and an endothelial cell nucleus (EN) is also visible (left, center).

Figure 61: FIG-2, Cerebral Cortex, white matter, Bodian, 40x. Severely injured white matter typical of that seen in FIG animals. The tissue presents a macerated appearance (black circles) with numerous rips and tears, possibly as a result of ice formation (red arrows). The capillaries (C) are separated from the tissue parenchyma (yellow arrow) and what appears to be a naked endothelial cell nuclei projected into the intraluminal space of one capillary (green arrow).

END OF PART 3

Histology was evaluated in two animals each from the FIG and FIGP groups, and in one control animal.  Only brain histology was evaluated in the straight-frozen control animal.

Liver

      The histological appearance of the liver in all three groups  of animals was  one  of  profound injury.  Even in the FGP group, the cellular integrity of the liver appeared grossly disrupted.  In liver tissue prepared using Yajima stain, the sinusoids and spaces of Disse were filled with flocculent debris, and it was often  difficult or impossible to  discern cell membranes (Figures 30-32). The collagenous  supporting structures of the bile canaliculi were in evidence and the nuclei of the hepatocytes appeared to have survived with few alterations evident at the light level, although frequent pyknotic nuclei were noted in the FIGP group (Figures 31 & 34).  Indeed, the nuclei often appeared to be floating in a sea of amorphous material (Figure 34).  Not surprisingly, the density of  staining of the cytoplasmic material was noticeably reduced over that  of  the fixative-perfused control. Few intact capillaries were noted.

FGP  liver  tissue prepared with PAS stain  exhibited  a  similar degree  of disruption (Figure 32).  However, quite remarkably, the borders of  the hepatocytes  were defined by a clear margin between  glycogen granule containing cytoplasm  and non-glycogen containing membrane or other material (membrane debris?) which failed to stain with Yajima due to gross physical disruption, or altered tissue chemistry (Figure 35).

Figure 28: Control-1 Liver, Yajima, 100x. Liver sections from the Control animal demonstrated normal morphology as can be seen in the image above.

Figure 29: Control-1 Liver, PAS, 100x. Liver sections were prepared with both Yajima and PAS stain in order to allow visualization of structures that neither stain discloses alone; in this case, most importantly, the presence or glycogen granules in the hepatocytes of the Control animal. Note the presence of normal intralobular architecture with crisp cell membranes in evidence, normal appearing sinusoid spaces, and residual sinusoidal red blood cells (RBCs) not washed out during fixative perfusion.

Figure 30: FGP-1 Liver, Yajima, 100x. The livers of FGP animals demonstrated extensive histological disruption. The sinusoids were all but obliterated and appeared filled with debris (ds) and the cytoplasm was extensively vacuolated (v). 

Figure 31: FIG-2 Liver, Yajima, 100x. As was the case with the FGP animals, the sinusoids were barely discernable and appeared filled with cellular debris (cd). In addition to extensive cytoplasmic (cv) and nuclear vacuolization (nv), pyknotic nuclei (pn) were also present. Cell membranes were difficult to discern and in many areas, frank cell lysis appears to have occurred with flocculent cellular debris (cd) appearing to fill the sinusoids.

Figure 32: FGP-1, Liver, PAS, 100x. The intensely red-stained granules present in the cytoplasm of the hepatocytes are glycogen deposits selectively stained by PAS. There is extensive cytoplasmic (cv) and nuclear vacuolization (nv) and the sinusoids appear filled with flocculent cellular debris (d). Indeed, it is only possible to discern the outlines of the original individual hepatocytes from the pattern of the intracellular glycogen granules disclosed by the PAS stain.

Figure 33: FIG-2, Liver, Yajima, 100x, well preserved area. While the bulk of the hepatic parenchyma exhibited the severe injury seen in Figures 30-32, there were frequently observed islands of comparatively well preserved tissue visible in both the FGP and FIGP sections suggesting that freezing injury is occurring non-homogenously.

Figure 34: FIG-2, Liver, PAS, 100x, necrotic area. There were patchy areas of frank necrosis visible in the livers of the FIGP animals that were not present in the livers of the FGP animals. This area, adjacent to  a central vein, shows extensive cell lysis with heavy vacuolization of the cytoplasm (v) and many pyknotic nuclei (pn) in evidence.

Figure 35: FIG-2, Liver, PAS, 100x. Note the presence of a few scattered glycogen granules (GG). Interestingly, in this comparatively well preserved area of FIGP liver it is possible to see some remaining deposits of glycogen that were not consumed during the long post-arrest ischemic interval. The absence of pyknotic nuclei and the relative absence of large intracellular vacuoles is also remarkable.

 Kidney

Figure 36: The functional unit of the kidney is the nephron, consisting of the glomerulus and the uriniferous tubule ( the renal corpuscle: a).The capillary tuft of the nephron, the glomerulus, is enclosed within a double cell layered structure; Bowman’s capsule. Bowman’s capsule and the capillary tuft it encloses comprise the glomerulus. Bowman’s capsule and the glomerular capillary tuft constitute the renal (or Malpighian) corpuscle (b).

PAS  stain  was used to prepare the control, FGP and  FIGP  renal tissue for light microscopy.  The histological appearance of FGP renal tissue was surprisingly good (Figures 329, 40 & 41).  The glomeruli and tubules  appeared grossly intact  and stain uptake was normal.  However,  a  number  of alterations  from  the appearance of the control were  apparent.  The capillary tuft of the glomeruli appeared swollen and the normal  space between the capillary tuft and Bowman’s capsule was absent.  There was also marked interstitial edema, and marked cellular edema as evidenced by the obliteration of the tubule lumen by cellular edema.

By contrast, the renal cortex of the FIGP animals, when  compared to  either  the control or the FGP group, showed a  profound  loss  of detail, absent intercellular space, and altered staining (Figures 40 & 42). The  tissue appeared frankly necrotic, with numerous pyknotic nuclei and  numerous large  vacuoles  which peppered the cells.  One  striking  difference between FGP and FIGP renal cortex was that the capillaries, which were largely  obliterated in the FGP animals, were consistently  spared  in the FIGP animals. Indeed, the only extracellular space in evidence in this  preparation was the narrowed lumen of the  capillaries,  grossly reduced in size apparently as a consequence of cellular edema.

Both ischemic and non-ischemic sections showed occasional evidence of  fracturing, with fractures crossing and severing tubule cells  and glomeruli (Figure 41).

Figure 37: Control, Renal Cortex, PAS 40x. Three glumeruli are present (G) adjacent to crisp, well defined proximal (P) and distal (D) convoluted tubules. The intertubular capillaries (C) show normal diameter with lumens free of red cells or debris. There is normal capsular space between Bowman’s capsule (BC, yellow arrows) and the glomerular capillary tuft and vascular pole (VP) are also normal in appearance.

Figure 38: Control-1, PAS 40x. Collecting ducts (CD), distal tubules (D) and a glomerulus  (G) are present in 6this micrograph of renal apical column. At right, a glomerulus is present with normal Bowman’s space (BS) and the macula densa (MD) in evidence.

Figure 39: FGP-1, Renal Cortex, PAS, 40x. The intertubular space (ITS) is great expanded and the tubule cells are heavily vacoulated (V) and lack definition. The intratubular space (IS) is no longer in evidence and the architecture of the glomerluar capillary tuft (GT) is radically altered and there is an absence of the normal architecture of Bowman’s space (yellow arrows). The intertubular capillaries appear to have been reduced to debris (D) visible in the intertubular spaces.

 Figure 40: FIG-2 Renal Cortex, PAS, 40x. There is massive cellular edema present with almost complete obliteration of Bowman’s space. The tubule (T) lumens are no longer visible and the tubule cells are extensively vacoulated with many pyknotic nuceli in evidence. Individual tubular cell membranes are impossible to resolve. The afferent glomerular arteriole (AA) appears intact (red arrow).

Figure 41: FIG-2 Renal Cortex, fracture present (arrows), PAS, 40x. Two renal tubules, possibly a proximal and distal convoluted tubule (T) are dissected by a fracture as is the macula densa (MD) of the glomerulus (G). Remarkably, there is still a small amount of intertubular space present in this micrograph.

Figure 42: FIG-2 Renal Cortex, PAS 40x. vacuolization (black arrows) and extensive vacuolization (blue arrows) accompanied by necrotic changes, such as the frequent presence of pyknotic nuclei (red arrows).

END OF PART 2

Experimental Studies to Determine the Effectiveness of LAPC under Laboratory Conditions

Experimental Studies to Determine the Effectiveness of LAPC under Laboratory Conditions

 [This section is an edited version of an article authored by Steven B. Harris, Michael G. Darwin, Sandra R. Russell, Joan M. O’Farrell, Mike Fletcher and Brian Wowk entitled, Rapid (0.5°C/min) minimally invasive induction of hypothermia using ~4ºC perfluorochemical lung lavage in dogs, which first appeared in Resuscitation, 2001. 50: p. 189-204.)]

 1. Introduction

The potential utility of profound and ultra-profound hypothermia (0-5^oC ) to arrest deleterious neurological changes has long been understood in both biology and medicine.[168],[169],[170],[171],[172] In 1959 Benson, et al., reported good outcome using profound hypothermia (10-22^oC ) as a treatment following cardiac arrest. [173] This work was followed up by a number of clinicians [168],[174],[175],[176] who also reported favourable results. However, due to coagulopathy, arrhythmias, and the increased incidence of pneumonia and sepsis associated with such deep and prolonged cooling, post arrest hypothermia failed to gain acceptance and was abandoned. It was not until the work of Safar, et al., [160],[53],[82] that the utility of mild therapeutic hypothermia (MTH) (DT = −2 to −3°C) as an active treatment for the post-resuscitation syndrome was rigorously demonstrated, and subsequently validated by others. [177],[178],[55],[179-181]  As noted in Section 1, while CPB offers the most rapid core cooling possible, it is logistically unsupportable as currently practiced. Additionally, CPB carries the added risks of anticoagulation, further activation of the immune-inflammatory cascade, RBC aggregation, and the danger of gas embolism, as chilled, nitrogen-saturated blood is rapidly re-warmed as it perfuses warm tissues.[182]

As was also previously noted, less invasive modalities with the potential for in-field application, such as surface cooling and lavage of body viscuses with a balanced salt solution, are only effective in achieving cooling rates in the range of 0.10–0.15°C/min. The seemingly straightforward  experimental technique of ‘tidal liquid ventilation’ (TLV) with chilled, oxygenated PFC uses the ~20 m² surface area of the lungs for heat exchange, but thus far has been no more effective in inducing hypothermia than surface cooling with ice bags or chilled water blankets.[155] After preliminary experiments demonstrated the technical adequacy of LAPC at achieving heat exchange in range of 0.25 to 0.35°C/min [183] a comprehensive study was undertaken by 21^st Century Medicine Inc., (21CM) and Critical Care Research, Inc., (CCR), beginning in 1999, to define and validate this cooling modality in a canine model. The goals of this research were to, a) demonstrate the fundamental safety and efficacy of the technique, b) determine the optimum cycle and volume of liquid and gas fractional tidal liquid ventilations (FTLVs), and c ) attempt to determine safe airway pressures and define liquid and gas ventilation strategies that minimized or eliminated baro- and volutrauma.

This technique, developed at 21CM/CCR, was initially called ‘mixed-mode liquid ventilation cooling’ and was later renamed ‘gas-liquid ventilation’ (GLV). However, neither of these names adequately describes the technique, and this author (Darwin) has chosen to the use the term liquid assisted pulmonary cooling (LAPC) instead. In previous studies where large fractional tidal liquid volumes and shorter cycles of FTLV were used, the performance of LAPC deteriorated towards that seen when TLV-cooling (or warming) was used. In practice however, certain significant differences remained and understanding these differences proved essential to optimization of the technique.

In LAPC the critical elements of gas ventilation are retained allowing for flexibility in selecting ventilation parameters independently for heat and gas-exchange, allowing for liquid-mediated heat-exchange to be easily undertaken using existing ventilation systems[1]. The combination of gas and liquid FTLVs may also play a role in the surprisingly good thermal efficiency of LAPC as compared with TLV.

The following study explored the performance of LAPC using a prototype automated FTLV device, and discussed the basic mechanics and intrinsic limitations of heat-exchange using FTLV.

2. Materials and methods

These experiments were approved by 21st Century Medicine, Inc.’s Institutional Animal Care and Use Committee and were in compliance with the Animal Welfare Act and the National Research Council’s Guide for the Care and Use of Laboratory Animals. Fifteen mongrel dogs weighing 13.8–25.7 kg were used (Table 1). Dogs were pre-medicated with I.M. acepromazine (1.0 mg/kg) and atropine (0.02 mg/kg) prior to induction of general anesthesia using sodium pentobarbital (30 mg/kg I.V., with maintenance dosing). Anesthetized dogs were intubated with a reinforced 10.0 mm I.D. (Willy Rusch AG, Kernen, Germany) endotracheal tube (E.T.), and ventilated on room air using a Bennett MA1 or Siemens Servo 900 C ventilator. Ventilator parameters, unless otherwise noted, were 12 gas-breaths/min, gas tidal-volume of 15 ml/kg, I:E ratio of 1:3, and a maximal positive inspiratory pressure (PIP) limit of 26 cm H₂O (2.5 kPa). Gas pressures were measured at the E.T. adapter. Gas minute-volume (V_g) was adjusted to maintain PaCO₂ between 35 and 40 torr. Animals were maintained at ~37.5°C prior to LAPC, using a temperature-controlled water blanket. Rectal and bilateral tympanic temperatures (Ttym) were monitored continuously using a type-T thermocouple system (Cole-Parmer, Vernon Hills, IL) with a response time constant (to) of 5 s.

Combination pressure, blood sampling, and temperature-probe catheters were constructed from rigid polyethylene pressure-monitoring catheters, threaded centrally with 0.05 in. O.D. Teflon™-sheathed type-T thermocouples (to=0.3 s, Physitemp Instruments, Clifton, NJ). In order to reduce the risk of catheter-associated clot formation, I.V. sodium heparin was given to adjust activated clotting times to 300–500 s, prior to central line placement. Femoral vessels were isolated surgically, and arterial and venous catheters placed and advanced to a level above the renal vessels, as confirmed by X-ray. During surgery, bupivacaine (0.5%) was infiltrated into wounds to mitigate post-operative pain.

In one dog (Trial I-2), a femorally-placed pulmonary artery thermodilution catheter replaced the venous combination catheter. Blood and ventilator pressures were acquired through a Hewlett Packard 78532-B monitor/transducer system.

Immediately prior to LAPC, dogs were assessed for adequacy of general anesthesia, and then given pancuronium bromide (2 mg) to inhibit shivering and spontaneous breathing. FIO₂ was increased to 100% and external temperature control discontinued. To serve as a cannula for both delivery and removal of PFC liquid, a 19-Fr. flat-wire reinforced Bio-Medicus® venous catheter (Medtronic, Eden Prairie, MN) was introduced through the suction port of the E.T. adapter, and advanced ~45 cm to approximately the level of the carina (Figures 2-1 and 2-2) as confirmed by X-ray. This cannula was connected to the LAPC apparatus described below. LAPC was performed using the PFC liquid ‘FC-75’ (3M Corporation, St. Paul, MN), a perfluorinated butyl-tetrahydrofuran isomer mixture.[184]

Figure 2-1 (right): PFC delivery and withdrawal catheter threaded through the endotracheal tube with the tip positioned at the level of the carina.

A two reservoir circuit (Figure 2-3) was used to deliver and remove PFC from the lungs (FTLV) via the cannula, in cycle periods of 37 s (Trial I) or 16 s (Trial II). During timed PFC infusions (t_in=20 s for Trial I, or 10 s for Trial II), PFC was pumped through the cannula by a continuously-operating Travenol CPB roller-pump (Sarns, Ann Arbor, MI). A bypass loop, open during suction, allowed the roller-pump to divert (recirculate) PFC flow back into the storage reservoir whenever flow was not directed by line clamps V1–V3 into the animal. PFC was pumped continuously through an in-line 0.2 μ ‘pre-bypass’ filter (Pall PP 3802, Pall, East Hills, NY), a primary heat-exchanger (Torpedo-T, Sarns, Ann Arbor, MI), and a combination silicone membrane oxygenator/heat-exchanger (SciMed II-SM35, SciMed Life Systems, Minneapolis, MN). The oxygenator was supplied with 5–6.5 l/min O₂ (maximal device design rate), and the reservoir PFC was allowed to circulate and equilibrate with heat-exchangers and O₂, before LAPC was initiated. The circuit tubing was constructed of S-50 HL TYGON® 3/8 and 1/2 in. I.D. class VI tubing (Norton/Performance Plastics, Akron, OH) with the exception of a length of silicone tubing (Masterflex® 96410-73, Barrant Co., Barrington, IL) used in the roller-pump head in order to allow flexibility at low temperatures. PFC suction was driven by a vacuum pump (model 107CAB18B, Thomas Compressors, Sheboygan, WI), and suction reservoir negative pressure was limited to −35 torr by a vacuum relief valve. Figure 2-2: Detail of LAPC PFC introduction and removal catheter and ET Tube and gas ventilator configuration. A Biomedicus CPB venous return catheter was threaded through the suction port of a standard 16 mm respiratory ET tube swivel connector.

Figure 2-3: The LAPC system. The LAPC system was connected to a catheter inserted into the suction port of the E.T. adapter. PFC flows were directed by manual or mechanical clamps at V1–3. During the suction phase, FC from the lungs was removed into a sealed ‘suction’ reservoir, for later addition to the primary circuit (via adjustment of V4 and V5), while ‘infusion ready’ PFC was re-circulated through a bypass loop. Negative pressure was limited by a vacuum relief valve (VrV). Photo (right) A) Suction Reservoir, B) Storage Reservoir, C) Solenoid Valves (V1-V3), D) SciMed Oxygenator & Heat Exchanger, E)Sarns Roller Pump, F) PFC Suction/Delivery Catheter, G) Pump Controller, H) Heat Exchanger Return Line with weighted water diffuser (yellow), I) Thermocouple Probes.

Figure 2-4: Gas ventilator and respiratory monitoring equipment used in the LAPC experiments; a) Novametrix CO₂SMO respiratory function monitor and capnograph, b) Siemens Servo 900 C ventilator, c) Korr Medical, Inc., automated device used to perform rapid-cycle LAPC, d) LAPC apparatus.

  Concurrent FC-75 FTLV and gas ventilation (LAPC) was performed for 18 min in Trials I and II (n=12). This time was chosen, on the basis of preliminary work (data not shown), to achieve rapid systemic-cooling of greater than 5°C. For Trials I and II, the PFC recirculation rate within the LAPC device (=PFC infusion rate, ˙V _inf), was set at 50 ml/kg per min rate, V_FTLV) was set at 50 ml/kg per min.

Immediately after a timed FTLV, PFC was removed as rapidly as possible. Infusion of PFC for the next FTLV began immediately after suction was discontinued. In LAPC experiments, PFC was chilled to ~4°C prior to FTLV (Table 1), whereas in normothermic (control) dogs, isothermic PFC was delivered to the dog within ~2°C of tympanic temperature (T_tym). The PFC inflow and outflow temperature was measured continuously by a thermocouple inserted into the PFC path at the base of the delivery/removal cannula. Temperature data was collected throughout LAPC, and for 22 min after LAPC was completed. Arterial blood gas (ABG) samples were taken from the femoral arterial line before the start of LAPC, and every 2 min during LAPC. Following the post-LAPC equilibration period, monitoring devices were removed and incisions closed.

Table 1:

Table 2:

2.3. Trial I (manually-controlled LAPC)

Trial I was designed to investigate the variability in the response of individual animals to LAPC and to investigate the physiological effects of the LAPC technique with and without cooling (i.e., ~4ºC PFC vs. isothermic PFC FTLV). Either isothermic (near-body temperature) or ~4ºC PFC FTLV was administered using a manually-controlled system (V1–V5 in Fig. 1 represent CPB tubing-occluders in this Trial). One FTLV (period t_c ~37 s) was composed of a timed FTLV (t_inf = 20 s), followed by PFC suction (t_s ~17 s). Suction was stopped when PFC liquid return became sparse, or gas pressure in the ventilator circuit fell below −5 cm H2O (−0.5 kPa). Five dogs received ~4ºC FTLV (Trial I-1–5), while two controls received the same protocol using isothermic FTLV (Trial I-6 and 7).

 2.4. Trial II (machine-controlled LAPC)

Trial II assessed the utility of using an automated device (custom manufactured by Korr Medical, Inc., Salt Lake City, UT) to perform rapid-cycle LAPC. Computer-controlled solenoid clamp-valve occlusion of circuit lines at V1–V3 allowed smaller FTLV volumes (V_FTLV) and smaller t_c. While t_inf was decreased to 10 s in Trial II, V_FTLV remained constant, and the effective PFC FTLV rate (V_FTLV) remained in the range of V_FTLVfor Trial I. Table 1 gives relevant trial parameters. In Trial II, suction of PFC from the lungs began immediately after infusion, and was automatically stopped whenever a ventilator circuit pressure of −5 cm H₂O was reached (t_s ~6 s, giving t_c ~16 s). Three dogs received ~4ºC PFC (Trial II-1–3), while two controls (Trial II-4 and 5) received isothermic PFC.

 2.5. Animals A, B and C

 Selected data from three dogs in an earlier method development series was used. These dogs had been prepared as above, then manually given 1, 15 and 21 FTLVs, respectively with ~4ºC PFC, at much slower rates than in Trials I and II (Table 1). Data from these animals allowed independent measurements of FTLV volume heat-contents and temperatures, and thus the heat capacities and heat transfer efficiencies, by a more thorough thermal accounting method (Table 2, Appendix A).

2.6. Data collection and correction, statistical methods, graphical display and presentation

 Temperature and pressure data were collected using a PCI E series data acquisition board and LabVIEW™ software (National Instruments, Austin, TX). Graphical analysis and display of temperature data, and curve fitting, was done using the software package Origin™ (Microcal Software, Northhampton, MA). Statistical comparison of Trial group values was done using GraphPad Prism (GraphPad Software, San Diego, CA). Group means are reported ± standard

Figure 2-5 (above): Body temperature changes observed during LAPC (Method of Trial I). In this illustrative experiment from Trial I (I-4), FTLVs of ~4ºC FC-75 were infused (~20 s) and removed (~17 s) from the lungs. LAPC was performed for 18 min (hatched bar), then stopped to allow thermal equilibration (22 min). Arterial temperature (˜Tart), central venous temperature (˜T_ven), tympanic temperature (˜T_tym), and rectal temperature (˜T_rec) are shown. Inset: Enlarged view of temperature changes recorded during the first two cycles of PFC infusion (gray bar) and removal (yellow bar).

 deviation (SD) except as otherwise noted. For each animal, the T_tym from whichever probe cooled most rapidly, was used (right probe in 12/15 dogs). In order to facilitate comparison of cooling rates between sites in the same animal, temperatures at all probe sites were corrected to the baseline aortic temperature (T_art), as measured immediately prior to the start of LAPC. For ease of description, LAPC-cooling is presented in terms of thermal-deficit (‘cold’) moving from the lungs into successive body compartments. A compartmental analysis of thermal transfer in this model, and a glossary of notation and equations used, is given in Appendix A.

 3. Results

 LAPC allowed FTLV of dogs during concurrent gas ventilation. Suction from the submerged catheter tip at the carina allowed collection of PFC even during forced gas inspiration. It was discovered that a long suction catheter was necessary to insure that adequate suction pressure could be used to withdraw PFC throughout the liquid removal sequence, without prolonged exposure of the gas filled portion of the airways to the negative pressure of the suction system/reservoir. Additional protection of the airways against excessive negative pressure during the relatively brief time after liquid no longer filled the suction line was provided by incorporating a negative pressure relief valve on the suction reservoir. Suction in this manner was efficient, although FTLV volume measurements showed that the lungs retained ~12 ml/kg PFC (approximately FRC) between FTLVs.

The PFC pump circulation/infusion rate (˙V _inf), measured volumetrically preceding and following LAPC, was stable to within 1% over the duration of LAPC, and was not significantly different between trials (P = 0.28). The V_FTLV, calculated as tinf  _inf/t_c, was 30.7±2.3 ml/kg per min (Trial I) and 36.4 ± 3.2 ml/kg per min (Trial II). The ˙V_FTLV was significantly (P = 0.023) larger in Trial II because machine-controlled suction made more efficient use of available non-infusion time, resulting in faster net PFC removal.

Figure 2-6: Thermal equilibration after LAPC. Mean Tart and T_ven values (Fig. 2) are shown for Trial I, dogs 1–5. To highlight equilibration changes, T_tart curve nadirs (n=5) were superimposed before calculation of means, and T_ven data (n = 4) for each dog was adjusted with its corresponding Tart curve. Incompatible T_ven data from a pulmonary artery thermodilution catheter in I-2 has been omitted. Inset: The sigmoidal mean (N = 5) T_art recovery during the first ~12 s after final LAPC. FTLV is approximated by linear fitting.

Figure 2-7: Body temperature changes during manual and mechanical LAPC (Trial I vs. Trial II). The relative rates of core body cooling in dogs undergoing 18 min (hatched region) of manual (Trial I, solid squares) or machine-driven (Trial II, open circles) ~4ºC LAPC, were assessed by comparing changes in group mean T_tym. Symbols represent the mean and SEM (n = 5 for manual, and 3 for machine groups).

 3.1. Thermal results of LAPC

3.1.1. Cooling time delay

 Figure 2-5 illustrates LAPC cooling in a representative dog (I-4) from Trial I. The T_art began to decrease 3–6 s after the start of each PFC FTLV. Since this delay included circulation delay from lungs to aorta, the transfer of thermal-deficit from newly-introduced PFC to pulmonary blood was very rapid. The venous temperature (T_ven) began to decrease 10.4 ± 6.9 s after T_art decline, representing the minimum systemic circulation time. Though exhibiting delay, damping, and broadening behavior (presumably due to peripheral heat-exchange and varying systemic blood-return path lengths), T_ven transients from FTLVs mirrored T_art transients. T_tym temperatures, presumably reflecting brain and viscera temperatures, were non-oscillatory. The T_tym did not begin to decrease until ~24 s after the start of LAPC. This decrease occurred in three phases: an initial phase lasting ~100 s, an exponential phase lasting for ~900 s, and a final linearly-decreasing phase lasting until the end of LAPC. Core cooling as measured by T_tym continued for about 120 s after the end of LAPC (Figure 2-5), then exhibited a marked rebound effect [185] with exponential dampening (t ~20 min, Figures 2-5–2-7). These phases of cooling and equilibration were consistent with a five-compartment thermal model, in which the three compartments representing animal tissues corresponded roughly with (1) the blood and vasculature; (2) the classical thermal core; and (3) the classical thermal periphery (Figure 2-8). Modelling equations and estimation of compartment sizes are given in the Appendix A.

3.1.2. Cooling rate

 Crude cooling rates were determined numerically from appropriate T vs. t graph segments. The mean cooling rate from LAPC initiation, or DT_tym/Dt, reached a maximum value in Trial I at −0.49±0.09°C/min (t=6.6 min). The differential cooling rate d (DT_tym)/ dt = dT_tym/dt reached a maximum (max) value of -0.59 ± 13°C/min at t ~100 s, near the end of the initial heat exchange development region. (This value is comparable to analytic d (DT_tym)/dt  (max) from (Eq. (1)) = DT_k/t_o  = −0.63°C/min). Corresponding cooling rates in Trial II were DT_tym / Dt (max) = −0.33 ± 0.02°C/min (at t = 7.3 min) and dT_tym / d_t (max) = −0.37 ± 0.06°C/min (at t = 100 s).

3.1.3. Mean cooling power

 The mean heat removal rate (cooling-power) P over the entire duration of LAPC, for each animal, was estimated from DT_e according to P = m C_m DT_e/t (total).

Here t (total) is the entire LAPC application time = ~1080 s. (Note: for this calculation, the more accurate Trial I mean C_m is used for all Trial II animals.) The mean cooling power of Trial I was 336 ± 60 watts, while that of Trial II (using the Trial I value of C_m) was 207 ± 49 watts (P = 0.02). Variation in animal size was the major source of intra-group variability.

Figure 2-8: Heat transfer among body compartments during LAPC. Heat transfer during LAPC in the dog may be modeled using 5 thermal compartments. Heat transfer between compartments (which is by blood circulation, except as noted) is shown in the box diagram as double-headed arrows. The pair of arrows connecting Compartments 2 and 3 represents the different processes of lung equilibration with (1) pulmonary artery flow; and (2) with the complete blood volume and selected viscera.

 3.2. Gas exchange

 ABG measurements demonstrated that infusion of ~4ºC PFC stabilized PaO₂ and PaCO₂ during LAPC. In contrast, LAPC using isothermic PFC failed to maintain baseline PaO₂ or PaCO₂ levels (Figure 2-9). In Trial II-4, hypercarbia during the first 13 min of isothermic LAPC was abolished by increasing the tidal volume from 15 to 25 ml/kg (final V_g = 375 ml/kg per min). In Trial II-5, V_g was pre-set to 375 ml/kg per min in an attempt to avoid hypercarbia, and no significant ABG changes were observed.

Figure 2-9: LAPC does not maintain normocarbia at isothermic temperatures without alteration of the gas ventilation parameters. Animals in Trials I and II underwent LAPC using either ~4ºC (˜) or isothermic (r£ ™¯) FC-75. Both arterial PaO₂ (Panel A) and PaCO₂ (Panel B) levels were affected by FTLV temperature. ~4ºC FTLV data from Trials I and II were very similar in magnitude, and therefore, have been combined (n=8). Isothermic LAPC is shown as four separate experiments (Trial I-6 and 7, and Trial II-4 and 5). Gas tidal volume was increased from 15 to 25 ml/kg in Trial II-4 at t=13 min, and at t=0 in Trial II-5, normalizing PaO₂ and PaCO₂ in both animals. Declining PaO₂ in Trial I-7 was due to inadvertent failure to pre-oxygenate PFC.

Figure 2-10: Effect of LAPC on VCO₂ and EtCO₂ during 20 minutes of ~4ºC FTLV. LAPC allows superior CO₂ removal due to gas ventilation because ^·V _PFC = no more than 30 mL/kg/min of liquid, allowing gas ventilation of at least 200 mL/kg/min, resulting in a maximum ^·VCO₂ removal rate of  >8 mL/kg/min or a minimum of 400% of basal metabolic rate.

3.3. Clinical observations and gross pathology

 With the exception of one dog, animals subjected to LAPC displayed mild tachypnea and increased expiratory sounds, but otherwise exhibited unremarkable recovery from anesthesia, including the ability to walk and drink. The exception was an eosinophilic animal (Trial II-1) which had normal oxygenation during LAPC, but developed severe hypoxemia shortly after LAPC.

Chest X-ray pre- and post-procedure showed no (comparatively) remarkable features. This dog was sacrificed at 9 h.

Necropsy revealed a mass of D. immitis (heart- worm) embolized into the pulmonary arterial circulation, possibly as a result of local chilling of the parasite mass due to LAPC (this animal had been heartworm seronegative). Necropsies performed on nine remaining Trial I and II animals sacrificed 24 h post-procedure revealed diffuse spongy, resilient hyperinflated non-collapsible lungs (HNCL) seen in animals exposed to a high-vapor pressure PFC at high PIP pressures.[186]  HNCL was most prominent in the anterior, least dependent areas of the lung lobes. This trapped intra-alveolar PFC was thought to be the cause of broncho-constriction and wheezing found in post-LAPC animals. There was also evidence gross, dependent-lung damage evidenced by pulmonary edema with consolidation in both isothermic and ~4^oC PFC-FTLV animals.

Other organ systems in this series were grossly normal. Two animals in Trial II (II-3 and II-4) were not sacrificed, and were held for long term evaluation. They were neurologically normal at 1 year post-LAPC.

3.4 Impact on Hemodynamics

 FTLV with both ~4ºC and isothermic PFC resulted in an almost immediate modest decrease central venous pressure (CVP) which persisted for the duration of the FTLV and recovered to pre-FTLV values at the conclusion of each FTLV cycle (Figure 2-11 – 2-13).

Figure 2-11: Impact of LAPC on HR.

In animals subjected to FTLV with ~4ºC PFC there was an immediate, transient reduction in heart rate (HR) and mean arterial blood pressure (MAP) and a corresponding increase CVP during the FTLV cycle. In animals undergoing ~4ºC FTLV these effects could be attributed to the acute, cyclical chilling of the coronary blood supply during each FTLV with chilled PFC. The temperature of the blood entering the coronary os was 10^o to 15^oC colder than systemic blood (as measured by pulmonary artery catheter), and this would be expected to have an immediate depressive effect on myocardial contractility due to the transient hypothermia the myocardium would experience as a consequence of perfusion with chilled blood. In fact, consonant with this interpretation, HR decreased steadily during ~4ºC FTLV, recovering progressively less after each FTLV cycle, as systemic hypothermia was induced (Figure 21).

Figure 2-12: Cyclical variation in CVP is response to FTLVs with isothermal PFC.

Figure 2-13: Effect of FTLV with ~4ºC PFC on MAP and CVP over the course of 6 minutes of LAPC.  MAP is transiently markedly depressed and CVP is concurrently increased in response to loading with ~4ºC PFC; this effect is reversed when the PFC load is suctioned from the lungs; although there is increasing depression of MAP in response to the induction of systemic hypothermia. Cardiac output (CO) was not measured in these studies and mathematical analysis of the MAP waveforms generated during isothermic FTLV were not done. Thus, the precise extent to which FTLV with isothermic PFC (i.e., without the thermal-metabolic effects of chilled blood on the myocardium as occurs in LAPC) impairs cardiac output or coronary perfusion pressure is unknown.

Interestingly, the cyclical increase and recovery of the CVP remained constant during both ~4ºC and isothermic FTLV. Initially, the effect of FTLV on CVP was thought to be due to compression of the thoracic vena cavae by the relatively dense PFC load in the lungs. It was hypothesized that this effect might be more pronounced in the dog due to the V-shape of the canine thorax with the cavae resting at the bottom of the thoracic ‘trough’ in a dependent position under the lungs. To test this hypothesis, isothermic FTLV was carried out with a dog in the prone position. Pronation had no effect on the transient, cyclical depression of HR and increase in the CVP associated with FTLV. It thus seems possible that loading of the lungs with dense PFC liquid results in increased pressure on the thoracic vasculature, in particular on the thoracic venous vasculature, in much the same way gas PEEP reduces thoracic venous capacitance and raises CVP; reducing right ventricular preload and right ventricular output (cardiac output). These effects would seem the most likely explanation for the reduction in MAP observed during maximal PFC loading during both ~4ºC and isothermic FTLV.

CO was not measured during these LAPC studies nor was mathematical analysis of the aortic pressure waveforms undertaken to determine with precision the degree to which FTLV depressed MAP. Crude analysis of MAP during isothermic FTLV suggests that cyclical PFC loading is responsible for ~15-20% reduction in MAP over baseline.  Mean CVP is increased ~35% over basal levels during FTLV. To what extent CO will be impacted as a result of FTLV with PFC during CPR will have to be determined experimentally (see discussion in 4.5.3. Overcoming Increased Intrathoracic Pressure and Preserving CO, below).

4. Discussion

 4.1. Apparent effect of temperature on gas exchange

 Isothermic LAPC in this model was surprisingly poor at removing CO₂, considering that the CO₂ carrying capacity in FC-75 decreases by only ~23% from 0 to 40°C (extrapolated from [184]). A useful observation was that pO₂ values decreased even in isothermic animals, indicating an influence on total ventilation and possibly also reflecting decreased CO as evidenced by the (average) decline in MAP and increase in systemic vascular resistance during LAPC.

Capnographic analysis of LAPC in Trials I and II (data not shown) indicated that isothermic LAPC had a much larger negative effect on pressure-limited total gas ventilation V_g, as compared to ~4ºC LAPC using the same technique and the same gas ventilator settings. Since LAPC at a ˙V_FTLVr of 30–36 ml/kg per min relies on gas ventilation V_g for ~50% of total alveolar ventilation, a differential loss of pressure-limited V_g with temperature appeared to be the basis of CO₂ retention in isothermic LAPC. The mechanism of the implied differential change in lung compliance may be related to the depressive effects of FTLV on CO, and presumably, on perfusion. Thus, gas ventilation adjustments similar to those in Trial II-4 and 5 may be required if LAPC is used as a re-warming technique, and it may additionally be necessary to abolish gas PEEP, or even apply continuous negative airway pressure [187],[188] to counteract the PEEP-like effects of PCF loading and to generally improve CO and coronary perfusion pressure (CPP) during CPR.[189]

4.2. Thermal transfer efficiency and kinetics

 The optimal LAPC cooling (or warming) protocol remains unknown. However, the finding that the thermal equilibration of non-dead space PFC and local pulmonary blood flow proceeds very rapidly (t_o <12 s) suggests that FTLV infusion times need to be no longer than this time scale. When PFC lung dwell times exceed this duration, the FTLV load is in place longer than is required to transfer the most labile part of its thermal potential to the pulmonary blood and parenchyma.

Figure 2-14: Relative insensitivity of PFC dwell time to heat exchange was demonstrated by progressively shortening the duration of FTLVs and increasing their frequency. Even at the maximum achievable rate of FTLV at ˙V_FTLV  rates of 30 ml/kg per min and ˙V_FTLV  rates of 50 ml/kg per min no deterioration in the efficiency of heat exchange was observed. This suggests that thermal equilibration at the level of the alveolus is practically instantaneous and that the shortest FTLV dwell time should be used for optimum heat exchange.

Since FTLV ventilation rates (˙V_FTLVr) in the present study are already at least a third of the maximal rates possible in TLV, it seems probable that PFC infusion rates and pressures, rather than heat transfer rates from PFC to lung, will be the fundamentally limiting factor to power transfer in LAPC. These observations suggest that, as least to ˙V_{FTLV r} rates of 30 ml/kg per min and ˙V_FTLV  rates of 50 ml/kg per min, the total cooling power (cooling rate) in LAPC will be greatest if no PFC dwell time is allowed, and all available time during the FTLV cycle is used to either introduce, or remove, PFC.

4.3. Question of diffusion dead space in LAPC

 Mammalian lungs depend on simple gas diffusion for CO₂ transport through the acinar airways during normal tidal ventilation. An intractable problem in experimental TLV has been that simple diffusion is not sufficient to similarly move CO₂ through liquid PFC at physiologic CO₂ partial pressure gradients. This limitation appears in TLV as a ‘CO₂ diffusion dead space’ which effectively lowers alveolar ventilation. In part due to such extra physiologic dead space, TLV of adult humans has been estimated to require liquid minute-volumes near 70 ml/kg per min.[190] This value is at the upper bound of realistically attainable liquid flow rates [191],[192] and leaves little leeway for treating hypercarbia, hypermetabolic states, or lung disease. Such difficulties are not a theoretical limitation in LAPC, however, since LAPC does not require high liquid flow rates for ventilation. In the most rapid-cooling LAPC protocol used in this trial, ˙V_FTLV was 31 ml/kg per min—a low baseline value which permitted the addition of 10 times this minute-volume of gas ventilation (see Fig. 4, Trial II-4, 5). Moreover, since normal gas minute volumes were required to maintain normocapnia in Trial I, there is as yet no evidence for any CO₂ diffusion limitations caused by intrapulmonary PFC in LAPC. Possible reasons for this are discussed below.

Thermal-diffusion limits in TLV have not been studied per se, but their presence is suggested by the results of Shaffer and co-workers.[155] In their cat TLV model using a ˙V _FTLV of 75 ml/kg per min, a decrease in PFC inspiration temperature from 20 to 10°C (increasing the thermal gradient by a factor of 1.6) increased the cooling rate from −0.13 to −0.15°C/min. This small rate change represented a significant loss of efficiency. By contrast, in the present LAPC study using PFC at 4°C, there was no evidence of a thermal-diffusion limit at rates up to 4 FTLVs/min. Notably, in Trial I, where 100% of the V_FTLVr, and 40% of the ˙V_FTLVr of the cat TLV model was used, cooling rates for LAPC were more than three times those reported for cats subjected to TLV at 4.5 liquid breaths/min at 10°C.[155]

The possible quantitative presence of a thermal diffusion limit for LAPC at 4 FTLVs/min may be evaluated using a modified version of the concept of gas-exchange dead space (V_D). The respiratory system of a dog undergoing LAPC heat-exchange may be considered, by analogy with gas exchange dead space (V_D), to also contain a ‘thermal exchange dead space’ (V_Dtherm). Each thermal FTLV volume ˙V_FTLV of PFC then also contains a V_Dtherm , which by definition does not participate in heat-exchange. Thus, cycle thermal transfer efficiency E_f may be expressed as (V_FTLVr − V_Dtherm)  / ˙V_FTLV, and any measured value of mean E_f may be expressed as an equivalent mean V_Dtherm = ˙V_FTLVr  (1− E_f). For Trial I (E_f = 0.6, Appendix A for calculation), mean V_Dtherm was then seen to be 7.5±1.6 ml/kg, and in Trial II (using Eq. (6) E_f  Value = 0.40), V_Dtherm was 5.3±0.8 ml/kg (P = 0.072). The absence of an increase in V_Dtherm in Trial II vs. Trial I indicated that the size of V_Dtherm in these LAPC protocols was non-dynamic at time-scales of one FTLV cycle, providing evidence against the presence of a ‘thermal diffusion dead space’ (analogous to a CO₂ diffusion dead space) at these FTLV rates. In absolute terms, it may be useful to compare calculated V_Dtherm in the LAPC dog model to the expected physiologic gas-exchange dead space, VDCA, which in healthy animals is close to the dog anatomic VD = ~6.5 ml/kg.[193]

In thermal diffusion, as in gas diffusion, diffusion physiologic dead space would be expected to significantly add to anatomical dead space. However, the sum of mechanical-VD in the LAPC circuit (~1.5 ml/kg) plus the anatomic VD for dogs is found to be more than the calculated V_Dtherm in either trial in this study, leaving little room for a large heat-diffusion contribution to V_Dtherm. For these reasons it is suggested that the loss of cooling power observed in Trial II was not due to heat diffusion limitations, but instead due to a loss of efficiency effect similar to that seen with low tidal volumes in ordinary gas ventilation. In these terms, low FTLV volumes in LAPC result in an increase in ‘thermal dead space ventilation’ at the expense of PFC flow involved in active heat-exchange, resulting in a larger ‘wasted’ FTLV V_Dtherm / ˙V_FTLV. VD in heat transfer (V_Dtherm) that is analogous to VD in gas-transfer; in as much as all dead space is ‘diffusion dead space’ at long-enough time-scales. However, some of the mechanisms for diffusion modification of V_Dtherm are unique. By contrast with gas molecules, heat diffuses rapidly through device tubing into the PFC in the LAPC circuit dead space, and heat also diffuses directly through the tracheal wall into the anatomic-VD. Thus, heat diffusion from dead space liquid at sufficiently slow FTLV rates might be expected to have a pronounced effect on E_f in LAPC, due to slow heat-diffusion reduction in V_Dtherm.

Some evidence for such a process was found, though at FTLV dwell times too long to be of interest for rapid cooling. At the relatively small tc of Trials I and II, the calculated V_Dtherm was found to be ~VDCA; but in animal B, with a much longer t_c of 7 min, the V_Dtherm was only 2.6 ml/kg. The limit of this process was reached in animal A, in which the V_Dtherm of a single retained ‘breath’ of highly-oxygenated PFC fell to nearly zero after 10 min. Disappearance of V_Dtherm by thermal equilibration, estimated from individual cycle E_f variations in animals B and C, was estimated to occur with a half-time of ~5 min (data not shown).

This process was slow enough to be neglected when the duration of FTLV intervals (t_c) was less than several minutes.

Thus, at the FTLV rates of Trials I and II, a full-sized V_Dtherm of ~6 ml/kg appeared, and accounted for significant loss of cooling power at low ˙V_FTLV (e.g. Trial II where V_FTLV was only 8.8 ml/kg). The characteristic size of V_Dtherm at all but the slowest FTLV rates (~1 FTLV per 5 min) implies that the only thermally-efficient solution for performing LAPC at faster rates is maintenance of [˙V_FTLVr  / VDCA] or [˙V_FTLVr / V_Dtherm] ratios >3, in order to avoid excessive ‘wasted’

V_Dtherm  ventilation. This requires a ˙V_FTLV of ~18 ml/kg in dogs. In humans, where the anatomic-VD is <3 ml/kg, less than half the value for dogs, both the V_Dtherm, and therefore, most-efficient ˙V_FTLV values, might also be expected to be correspondingly less. In any case, it is clear that rapid-cooling LAPC techniques cannot wait for the relatively slow thermal equilibration of PFC within the anatomical VD, since equilibration in the remaining non-VD parts of the lung is so rapid (i.e. less than Trial II t_c of 16 s).

4.3.2. Possible synergy of combined gas and liquid ventilation in assisting mass (CO₂) and heat transfer

 The absence of expected heat-diffusion and gas-diffusion limitations in LAPC suggests that some assistive process for both gas and heat transfer through PFC in the peripheral lung may occur in LAPC. The authors’ fluoroscopic observations (made with the non-brominated and relatively radiolucent FC-75) have been that each gas breath in PLV produces a flash of fine bubbles which spread uniformly throughout the lung. As compared to the more familiar behavior of water, the low surface tension of PFCs (15 dyne-cm for FC-75, about 1/5th that of water) lowers the energy barrier to producing small bubbles in forced gas/liquid flows. Such

bubbles moving within small airways may induce eddies and turbulence in laminar liquid flows at small scales, contributing significantly to heat and mass (CO2) transport though PFC liquid by means other than diffusion.

It is hypothesized that the lack of bubble-induced turbulence in TLV may account for the large diffusion-dead-space for heat and CO₂ which seems to be present in TLV at even low breathing rates – an effect which is apparently absent in both PLV and LAPC.

4.4. Potential development of clinical LAPC

 Rapid cooling of the CNS has now become a primary goal in the clinical management of the post resuscitation syndrome [49] and potentially in the acute management of spinal cord injury.[194],[195],[196] Based on their work, Safar, et al., have noted that clinical implementation of mild resuscitative hypothermia, which was highly effective in the dog model of SCA, will depend upon the development of truly rapid MTH.[197]  A recent editorial in Stroke [83] commented on the striking ability of the combination of  (33-35^oC ) [198] and pharmacological pre-treatment to ameliorate ischemic brain damage in the rat middle cerebral artery occlusion model, and  then addressed similar concerns: “A problem for use of this technique for acute stroke therapy is that the time required to induce hypothermia in patients is likely to be considerably longer than for rats. […]. To substantially increase the rate of hypothermia induction in humans, it will almost certainly be necessary to use some sort of invasive procedure, such as a heat-exchanger, to cool the circulation.”

The technique of LAPC may eliminate the need for such invasive measures. For example, in the cited trial [199], rats were cooled from 37 to 33°C (−4°C) over 40 min, using surface cooling with ice packs. By contrast, the present study demonstrates cooling of the canine body core and brain by −4°C in less than 10 min.

4.5. Challenges Ahead

 4.5.1. PFC Selection

Development of clinical LAPC awaits identification of suitable PFCs for various LAPC applications. For example, the pharmaceutical PFC perfluorooctylbromide (Perflubron™ Alliance Pharmaceuticals) would presumably not be suitable for rapid LAPC cooling due to its freezing point of +6°C, but might be useful for slower cooling or for LAPHE facilitated re-warming. Some industrial PFCs have pour-points low enough to make them potentially useful as LAPC rapid-cooling media; however, most of these agents also have unacceptably high vapor pressures at 37°C or are not chemically defined in terms of chain length or even precise chemical composition (see Section 3: Perfluorchemicals). PFCs with such high vapor pressures exacerbate barotrauma by causing HNCL.

High vapor pressure PFCs may also increase the danger of long lasting lung collapse as a result of PFC, secondary to pneumothoraces, entering the pleural space and vaporizing (‘perfluorothorax’). FC-75, (formerly FX-80) was the first PFC used in liquid ventilation [10], but its relatively high vapor pressure (31.5 torr) makes it an undesirable LAPC agent. Assuming that a PFC with the desired biophysical properties is identified and produced to medical standards, LAPC should be easily scalable to adult humans. For example, the viscosity of FC-75 is similar to water [11], and under standard suction a 19 Fr. An adult pulmonary toilet catheter will remove FC-75 at ~2 l/min. As in the system described, a LAPC system may interface with a conventional gas ventilator system via a simple liquid-carrying catheter which extends through the endotracheal tube adapter suction port.

 4.5.2. Coronary Perfusion during LAPC

The effect of LAPC on coronary perfusion in the setting of ROSC following cardiac arrest due to myocardial infarction (MI) or in the presence of coronary artery disease (CAD) is unknown but is a possible cause of concern. One possible adverse effect is the potential compromise of the coronary circulation in CAD due to perfusion of the heart with profoundly chilled blood (i.e., blood temperature  @10°C below systemic temperature) leaving the pulmonary circulation and entering the coronary os. One of the authors (Darwin) has observed the onset of severe, acute angina in two hemodialysis patients with stable angina who were inadvertently dialyzed using very cold dialysate (Q_B of 250 ml/min at a blood temp of ~10° to 15°C) which resolved only when the dialysate temperature was increased to 30°C or above.

It is well established that exercise in cold environments, with associated inhalation of cold air, can trigger angina and lead to cardiac arrest in patients with coronary artery disease.[200],[201]  There has been considerable debate as to whether the cause of angina in this setting is due to cold air inhalation per se, or to the effects of a cold environment. It has been suggested that exposure to a cold environment is the primary factor in inducing cold weather angina, presumably by increasing peripheral vascular resistance resulting in an increase in cardiac workload at any given level of exercise [202],[203] in the same way that the cold pressor test produces acutely increased afterload and thus increased left ventricular wall stress.[204],[205],[206],[207] Hattenhauer and Neill [208] studied 33 male patients with coronary artery disease,11 of whom gave a positive history of angina when exposed to cold winter air. Seventeen of these patients were subjected to inhalation of cold air at -20°C for 4 min, and this resulted in angina at rest in four of these patients.

Cold air inhalation also produced angina in four of the seven patients who were paced at a heart rate which was well below the threshold for angina at room temperature. Cold air inhalation did not significantly increase myocardial oxygen consumption, or alter coronary blood flow as determined by the xenon clearance method. In a separate arm of the Hattenhauer and Neill study, cold air inhalation for 90 s in 18 patients produced no detectable-constriction of coronary arteries visualized angiographically. The investigators concluded from these findings that cold air inhalation induced angina could not be explained by an increase in cardiac work and myocardial oxygen consumption.

This study also demonstrated that there was no evidence of large or intermediate coronary artery or coronary arteriole constriction in response to the inhalation of very cold air. As a consequence, the authors proposed that cold air inhalation induced angina might be the result of constriction of minute coronary collaterals, or to other vessels compromising blood flow to potentially ischemic regions of the diseased myocardium. The work of Lassvik and Aveskog [209], confirmed the effects of cold air inhalation in inducing angina and failed to demonstrate any decrease in workload at either the onset of angina or at maximal workload during inhalation of moderately cold air (-10°C) in a room at 20°C. Dodds, et al., [210] evaluated 12 male patients with stable angina inhaling cold air (-8.8°C) during exercise to investigate if the vasoconstrictor peptides endothelin-l (ET-1) and angiotensin-II (AT-lI) played a material role in the etiology of this phenomenon. They concluded that neither ET-1 nor ATII had any significant role in the pathophysiology of cold air inhalation induced angina.

In contrast to Hattenhauer and Neill, and Lassvik and Aveskog, these investigators documented decreased myocardial oxygen consumption during peak exercise in cold air inhalation and concluded that the cause of this was a centrally operating mechanism such as a reduction in coronary flow. These investigators noted that the response to cold air inhalation was biphasic, and posit that the initial response; earlier onset of angina during exercise while breathing cold air, was due to activation of cold receptors in the upper airways stimulating a systemic increase in peripheral vascular resistance; and thus the observed accompanying rise in blood pressure and increased cardiac workload (i.e., the cold pressor test response). The secondary response; reduction in total exercise time, which occurred at a significantly lower rate-pressure product compared with the same patients breathing ambient temperature air, was thought to be due to peripheral reflex responses. Thus, these investigators hypothesize that the early onset of angina during exercise is due to sympathetic stimulation from inhaled cold air, but as exercise continues, central mechanisms play an increasing role in the pathophysiology of cold air inhalation induced angina.

The implications of cold air induced angina for LAPC in the setting of coronary artery disease are troubling and certainly bear careful investigation in animal models of compromised myocardial circulation. Patients in all of these studies developed significant ST depression (³ 1 mm) concurrent with the onset of cold air inhalation induced angina, suggesting clinically significant myocardial ischemia. What is unclear is whether the concomitant profound reduction in myocardial temperature seen in LAPC (both transient and long-term) will be protective against any perturbation in myocardial perfusion induced as a result of the sympathetic or central effects of cold FTLV.

Such a protective effect is suggested by the well established finding that ‘cold’ (~34°C) hemodialysis (HD) is protective against both intra-dialytic hypotension and angina [211], [212],[213],[214], as a result of sympathetic stimulation from the return of ‘cool’ blood to the systemic circulation. [215],[216],[217] In addition to its protective effect on hemodynamics during aggressive ultrafiltration, cold HD also attenuates the hypoxemia leucopoenia, [218] and the production of Complement 5a induced by blood exposure to the dialyzer membrane.[218],[219] It should also be noted that catecholamine administration (mimicking profound sympathetic stimulation) in the form of epinephrine is still a mainstay treatment of ventricular fibrillation and aystole in SCA.[59, 220],[221]

Finally, it is essential to point out that mild and even moderate induced hypothermia not only do not interfere with defibrillation from cardiac arrest, but actually greatly facilitate conversion of VF to perfusing NSR.[222],[223],[224],[225],[226]

 4.5.3. Potential for Regional (Myocardial) Overcooling

Under the low flow conditions of CPR, the possibility exists that due to thermal compartmentalization myocardial temperature could be reduced to below the fibrillation threshold or to adversely affect contractility. In the laboratory and the clinical setting moderate (28-32 ^oC ) and deep (10-22 ^oC) hypothermia are known to induce both benign and malignant cardiac arrhythmias; [227],[228] and  ventricular fibrillation is the most common cause of death in accidental hypothermia. [229],[230] In addition to the arrythymogenic effect of deep hypothermia (J waves, prolonged PR, QRS, QT intervals, and atrial arrhythmias) there is evidence that defibrillation becomes increasingly problematic as myocardial temperature decreases below 30^oC.[231],[232],[233]

The temperature of blood leaving the pulmonary circulation under normal flow conditions (during spontaneous circulation) within 5 minutes of the start of LAPC can reach temperatures of 17^-20^oC (see Figure 2-14, FC Suction Reservoir Temperature; this temperature closely approximates the PA blood temperature during the FTLV cycle). This is benign under high flow conditions because heat is being rapidly and continuously transferred between body compartments (Figures 2-5 – 2-7). However, under low flow conditions, and especially in the presence of pharmacologically induced severe peripheral vasoconstriction, rapid, selective core over-cooling could occur.

Disproportionate cooling of the body core happens under normal conditions in humans given large volumes of intravenous fluid chilled to 4^oC, and this is the basis for cold IV fluid induced hypothermia for cardiac arrest.[234],[235],[236],[237] This surprisingly durable core cooling, well beyond that predicted on the basis of calculations for the whole body, occurs because the thermal distribution volume in humans given rapid  cold IV infusions turns out to be much lower than total body volume. The result is that chilled IV fluids are ~3 times more effective in inducing hypothermia than suggested by all-compartment equilibrium calculations.[234]  Several explanations have been offered for this apparent thermodynamic inconsistency; most plausibly that peripheral vasoconstriction and inherently slower kinetics of heat exchange between central and peripheral compartments act to keep core temperature below the all-compartment equilibrium for at least 60 min after the conclusion of the cold IV infusion [238],[239] which is long enough for external cooling to begin contributing to cooling and for endovascular cooling to be initiated under ideal conditions.

These findings provide additional reason for vigilance in avoiding excessive myocardial or core cooling during CPR and suggest that a surrogate for myocardial temperature be sought and that the use of warmer PFC FTLVs be explored; trading off rapidity of temperature systemic reduction against the danger of over-cooling the heart.

 4.5.4. Overcoming Increased Intrathoracic Pressure and Preserving CO

Following the development of active compression decompression CPR (ACD-CPR) by Cohen, et al., in 1992 [240] the critical importance of maintaining negative intrathoracic pressure during the decompression phase of the CPR duty cycle has become increasingly understood.[241],[242]  There is a rapidly growing body of both animal and clinical CPR research documenting improved survival and decreased neurological morbidity when the intrathoracic pressure is kept negative during the  decompression (release of chest compression) phase of CPR by the use of inspiratory impedance threshold devices and active ACD-CPR.[188],[243],[244],[241] Similarly, there is accumulating evidence that the increased intrathoracic pressure that results from excessive positive pressure ventilation (PPV) during CPR dramatically reduces CO and causes increased morbidity and mortality.[245],[246],[247],[248],[249],[250]

In 2004, Yannopoulos, et al., reported the development of a device which allows for the continuous application of negative intrathoracic pressure (Figure 2-15) by applying controlled suction to the airway.[187]  This device, called the intrathoracic pressure regulator (ITPR) allows PPV to be delivered as needed during ACD-CPR, while maintaining negative intrathoracic pressure at all times when PPV is not being administered. The device effectively transforms the thoracic cavity into a low negative pressure (vacuum) chamber; increasing venous return from the body and consequently increasing preload and cardiac output. The ITPR also markedly increases CPP while at the same time decreasing intracranial pressure (ICP). In CPR ICP is typically elevated from the basal value of 12-16 mm Hg to 22-30 mm Hg (as the result of pressure transmission by blood in non-valved veins and by transmission of intrathoracic pressure via the cerebrospinal fluid) further compromising already inadequate cerebral perfusion.[251] Reduction of ICP during CPR has been shown to improve both survival and neurological outcome in an animal model of CPR.[252]

Figure 2-15: Prototype ITPR (Advanced Circulatory Systems, Inc.) in position in a typical bag-vale resuscitator – ET tube set-up.

 The ITPR has been shown to dramatically improve gas exchange, hemodynamics, blood flow, vital organ perfusion, and short-term survival rates during VF cardiac arrest in a porcine model of SCA and CPR [187] The ITPR is able to not only overcome the high intrathoracic pressures. associated with CPR (45 to 55 mmHg or ~61 to 75 cmH₂0 [253]) but to both create and sustain negative intrathoracic pressure (determined indirectly by measuring the ET tube pressure) continuously during prolonged periods of ITPR-CPR, even in the presence of induced hypovolemia.[189]   In hemorrhaged (hypovolemic) pigs, Yannopoulos, et al., were able to sustain CPP at >15 mm Hg (the accepted threshold for successful defibrillation in human SCA) and the isovolemic VF animals in the study maintained CPP at >25 mm Hg throughout the full 15 minutes of ITPR-CPR. In both groups, ETCO₂ was consistently maintained >25 mm Hg, and the 1-hour survival was 100%, as contrasted with 10% in control animals receiving AHA standard CPR (P = 0.0001).

By comparison, after 3 minutes of conventional CPR the control animals had a mean coronary perfusion pressure of <15 mm Hg and all had developed pseudo-respiratory alkalosis indicative of the V/Q mismatch of standard CPR.[254]  Blood gases in VF animals were strikingly preserved during ITPR-CPR; paO₂, which was 96±2 mm Hg at baseline, was   214±12.37 mm Hg after 10 min and 198±6.75 mm Hg after 15 min of ITPR-CPR. These findings would seem to suggest that ITPR-CPR may be reducing or eliminating the pulmonary edema that accompanies CPR and the high intrathoracic (and thus pulmonary arterial and venous pressures) generated during CPR.  ITPR is similarly effective at improving both hemodynamics and survival in a swine model of severe hypovolemic hypotension.[255]

The use of an ITPR during LAPC offers the prospect of reducing or abolishing the undesirable hemodynamic effects of FTLV with PFC. These effects, while not clinically significant in the healthy animal with spontaneous circulation undergoing LAPC (and the ability to dynamically respond to alterations in preload, CVP and SVR), may be unacceptable in the setting of cardiac arrest and CPR.

While it is clear that FTLV with PFC reduces MAP and elevates the mean CVP,[2] the extent to which these effects will reduce CO during CPR is unknown and will necessarily require further experimentation in animal models of SCA and CPR that closely approximates those experienced in humans under clinical conditions. The mechanics of CO and perfusion in CPR are radically different than those that pertain under conditions of spontaneous circulation.[256],[257],[258] In the healthy beating heart, modest increases in CVP (~5-15 mm Hg) result in increased cardiac output via the Frank-Starling mechanism (Starling’s Law of the Heart). Increased CVP increases left ventricular diastolic end pressure (LVEDP) increasing ventricular volume and stretching the ventricular myofibrils. Myofibrillary stretch results in sarcomere extension thereby increasing the affinity of troponin C for calcium, causing a greater number of cross-bridges to form within the myofibrils; this increases the contractile force of the heart increasing CO. This biochemical mechanism is not operational during cardiac arrest and CPR. Indeed, as outlined in the discussion below, little of the mechanics of perfusion under normal physiological conditions pertain in the setting of CPR.

4.5.5. Consideration of the Mechanics of Blood Flow in CPR and Implications for LAPC

During CPR in humans ventricular volume is not a primary determinant of CO, at least not in an appreciable fraction of patients undergoing CPR.[258] Despite the fact that it has been forty-eight years since the invention of closed chest CPR – the mechanics of blood flow during CPR in humans have not been definitively established – and a great deal of controversy surrounds the subject. Broadly, three mechanisms of antegrade blood flow have been proposed in (human) CPR:

• The Direct Cardiac Compression (Cardiac Pump) Theory posits that mechanical compression of the ventricles between the sternum and vertebral column creates a pressure gradient between the ventricle and the aorta (or pulmonary artery in the case of the right ventricle); as a consequence the mitral and tricuspid valves are closed, and blood is moved antegrade out of the ventricle. The ventricle then refills during the decompression phase of CPR and the process is repeated with each compression cycle.[259]  The cardiac pump theory was most definitively challenged with the publication of case data documenting the ineffectiveness of CPR in patients with flail chest. This could be reversed only be restoring elastic recoil to the chest by binding it; indicating that it was the generation of a net negative intrathoracic pressure upon recoil of the chest during the decompression phase of CPR that was essential for blood flow.[260]
• The Thoracic Pump Theory asserts that chest compression increases intrathoracic pressure forcing blood to flow from the thoracic to the extrathoracic circulation[3]. Retrograde flow from the right heart to the systemic veins is prevented by the venous valves; with the heart serving only as a passive conduit, and having no function as a pump.[261], [262] In the thoracic pump paradigm blood circulates because increased intrathoracic pressure is transmitted more or less equally to all of the intrathoracic vascular structures [256] and the atrioventricular valves remain open during the compression phase of CPR.[263]  However, these intravascular pressures are not equally transmitted to the extrathoracic arterial and venous beds, thus creating an extrathoracic arterial-venous pressure gradient resulting in antegrade blood flow during the period of high thoracic pressure. During the decompression phase of CPR, blood flows into the lungs as a result of the extrathoracic venous-to-intrapulmonary pressure gradient.  Angiographic [256] and echocardiographic studies in dogs documented static ventricular volumes during CPR [262], and case reports of human patients with cardiac tamponade recovering successfully after closed chest CPR both supported the thoracic pump mechanism as the explanation for antegrade flow in CPR.[264]
• In the Lung Pump Theory, as in the thoracic pump theory, the increased intrathoracic pressure during the compression phase of CPR is similarly transmitted equally to all intrathoracic vascular structures and there is insignificant regurgitation of blood from the pulmonary artery into the right ventricle and vena cavae until the pulmonary valve closes. [265] Thus, blood under pressure within the pulmonary vasculature will flow out of the lungs via the left side of the heart.[263],[266] During the decompression phase of the cycle the intrathoracic pressure drops below that present in the extrathoracic vasculature and blood flows into the thorax via the vena cavae and aorta. The pulmonary vascular volume is replenished by blood flowing from the right side of the heart through the open tricuspid and pulmonary valves. [266], [267] Retrograde aortic flow is prevented by competent closure of the aortic valve.[268],[265]

Transthoracic echocardiographic studies reported in the early 1980s during CPR in humans supported the thoracic pump theory.[268],[265] However,more recent studies employing transesophageal echocardiography (TEE) have concluded that because the mitral valve was observed to close during the compression phase, and open during the decompression phase, and the left and right ventricular volumes decreased during the compression phase, the  mechanism of antegrade flow during CPR was consistent with the direct cardiac compression theory.[269],[263]

Perhaps the most likely explanation of these seemingly incompatible findings is that all three mechanisms are responsible for antegrade flow, but not in every patient. In other words, the mechanics of forward flow may differ from patient to patient.  In fact, in the 1981 paper in which they proposed the thoracic pump theory of antegrade blood flow in CPR, Weisfeldt and Chandra stated, “It is not essential in the human to think about these mechanisms in an exclusive fashion. Direct cardiac compression is useful when possible; when it is not potent enough to maintain cerebral perfusion, manipulation of intrathoracic pressures would likely have a favorable additive effect on carotid blood flow.”[261]

In support of this view is the study by Ma, et al., which evaluated 17 patients undergoing CPR using TEE to measure both pulmonary and trans-mitral flow.[270]  Five of the 17 patients demonstrated closure of the mitral valve during the compression phase of CPR, with associated mitral regurgitation and forward aortic flow occurring which is consistent with the cardiac pump theory.

In the remaining 12 patients, the mitral valve remained open during both the compression and decompression phases of CPR with maximal antegrade mitral flow occurring during the compression phase of CPR. Eight of these 12 patients demonstrated antegrade mitral flow during the compression phase that was accompanied by antegrade pulmonary vein flow; consistent with the classic ‘thoracic pump’ mechanism. The last 4 patients in this study evidenced retrograde pulmonary vein flow concurrent with antegrade mitral flow during the compression phase; which is most consistent with lung pump theory as the primary cause of antegrade flow, at least in these patients.

5.0 Review of the Literature on LAPC

From the foregoing, it should be easy to understand the difficulty of establishing by experiment, let alone extrapolating on any theoretical basis, what the hemodynamic impact of FTLV or LAPC will be under clinical conditions in humans undergoing CPR. Similarly, the effect of LAPC on regional myocardial blood flow, myocardial irritability, coronary electrophysiology, and susceptibility to defibrillation will require further laboratory and, likely, experimental clinical investigation as well. Since publication of the 21CM/CCR LAPC research in 2001 there have been 4 subsequent studies to evaluate various aspects of the utility and safety of LAPC for application in CPR, or as a therapy in MI.

The first of these studies was published by Hong, et al., in 2002 and announced, “Our study is the first to demonstrate an induction of hypothermia by adopting PLV (no need of an extracorporeal circuit) and 0°C PFC.” [271] failing to cite the previously published work of either Darwin, et al., or Harris, et al. [272] These investigators attempted to validate the utility of LAPC using FTLV to achieve rapid reduction in core temperature and also studied the physiological impact of the procedure. The cooling rate achieved appears to have been ~0.11°C/min and this slow rate was almost certainly an artifact of the small volume of PFC used for each intrapulmonary exchange (~20 ml) and the small number of exchange (10 – 12 exchanges over 38 minutes). In fact, the cooling rate was faster in the animals in this study that were cooled externally by repeated application of ice water slush (0.17°C/min).

The most interesting result of this investigation were that MAP, mean PA pressure, HR, and CO, CBC and lactate of the LAPC treated animals were not significantly different than in the surface cooled animals. One variable that was altered in the LAPC group was an increase in pulmonary vascular resistance which appeared immediately after liquid loading and did not begin to return to baseline until halfway through the LAPC period. It is also remarkable that 1 animal in both the LAPC and the surface cooled groups (n = 7 for both groups) developed pulmonary hypertension and lethal pulmonary edema. The authors speculate that the observed pulmonary hypertension could have been a result of pulmonary venous constriction and/or increased pulmonary microvascular blood sludging due to the profound local cooling of lung vasculature resulting from instillation of 0°C PFC. They also raise the possibility that the observed elevation in PAP may have resulted from the mechanical effects of the PFC load; but offer no explanation for this.

Otherwise, hemodynamics were unaffected by LAPC. The paCO₂ became progressively elevated during the hypothermic interval in the LAPC group, perhaps due to inadequate PEEP or inappropriate parameters of mechanical ventilation in the face of evaporating PFC (which was not replenished during the course of the experiments).

The second study to be published was by Ko, et al., in 2002.[273] The purpose of this study was to evaluate the effect of PLV on pulmonary blood flow under low blood flow conditions designed to simulate those encountered during CPR in order to further validate the use of LAPC as an adjunct to cardiopulmonary cerebral resuscitation. Isolated perfused rat lungs were subjected to 20 min of PLV with room temperature Perflubron™ and segmental (i.e. pre-capillary, capillary, and post-capillary) hemodynamics were studied at a perfusate flow rate of 6 ml/min (~5% normal cardiac output [274]). Lungs received either gas ventilation or 5 or 10 ml/kg PLV. Segmental pressures and vascular resistances were determined, as was transcapillary fluid flux. PLV at both the 5 and 10 ml/kg Perflubron™ dose produced no detectable changes in pulmonary blood flow or in transcapillary fluid flux and the investigators concluded that, “These data support further investigation of this technique as an adjunct to cardiopulmonary resuscitation.”

Figure 2-16: Dramatic reduction in myocardial infarction size in rabbits rapidly cooled to ~34 by the use of TLV LAPC: infarct volume was 4.0 ± 5% in the LAPC groups vs. 37.7 ± 1.3% in the normothermic gas ventilated group (p = 0.001). Redrawn from Tissier, et al.[275]

Unfortunately, this study did not explore the effect of instilling cold PFC into the lungs, nor did it employ FTLV. Because it was an ex vivo study it was not possible to determine the effects of PFC loading, cold or isothermic, on the sympathetic response, coronary blood flows or other physiological parameters of concern in LAPC.

In 2007 a study by Tissier, et al., [275] used a rabbit model of evolving MI to evaluate the efficacy of TLV LAPC administered during the ischemic interval in achieving rapid core cooling to reduce infarct size and provide myocardial protection [276],[277],[278] even after substantial delay  [279] and when induced during reperfusion. [280],[281] It is well established that intra- and post-ischemic myocardial hypothermia dramatically reduces infarct size in animal models of MI and this observation has recently been extended to humans in the clinic.[282]

These investigators note that in the US the time to coronary revascularization following infarction is 120 min in 41.5% of patients admitted during off-hours and that 27.7% of patients presenting during normal business hours still averaged delays to revascularization in excess of 120 min.[283]  Rapid induction of hypothermia in these patients could be effective in providing substantial myocardial salvage during the delay between presentation and revascularization. This study demonstrated a remarkable reduction in infarct size in the LAPC treated group; 4.0 ± 5% vs. 37.7 ± 1.3% in the normothermic, gas ventilated animals (p = 0.001). The cooling rate achieved with TLV was 1.32°C/min; 6.6°C in 5 minutes. The effectiveness of TLV LAPC is not consistent with results obtained in larger animals in this author’s experience. It may be that the relatively short distances between the large and small airways and the very small diameter of even the largest airways in the rabbit as compared with the dog or human, may have improved the efficiency of heat exchange.

The most recent study by Staffey, et al., [284] is more on-point and provides considerable reassurance that cold PFC loading of the lungs during CPR and subsequent resuscitation is not only not deleterious, but in fact is markedly beneficial. In this study swine were subjected to 11 min of VF and treated either with a static intrapulmonary infusion of PFC chilled to -12 C to ~75% of total lung volume (40 ml/kg), static infusion of isothermic PFC (33^oC) under the same conditions, TLV with -15^oC PFC at 6-cycles per minute during 10.5 minutes of the arrest interval, and a control group that was arrested with no intervention during the ischemic interval. The animal were then given AHA standard CPR with defibrillation and advanced cardiac life support (ACLS). The specific objectives of this study were to determine what the effects LAPC; tidal and static, administered during the period of cardiac arrest would be on hemodynamics, CPP, gas exchange, defibrillation, and hemodynamic stability during a 1 hour period of evaluation post ROSC. The endpoint was ROSC for 1 hour without ionotropic support.

The global objective of the study was to determine if LAPC could be used to rapidly induce cardiopulmonary hypothermia during the arrest period as opposed to cerebral or systemic hypothermia. Work by Rhee, et al., and Boddicker et al., also using swine models, had previously demonstrated that systemic hypothermia induced before resuscitation from cardiac arrest resulted in more rapid and consistent defibrillation from VF as well as earlier and more stable ROSC. Since the purpose of this study was to cool only the thoracic viscera to determine the effect of local cardiopulmonary hypothermia on resuscitation; ‘targeted cardiopulmonary intra-arrest moderate hypothermia (28-32^oC),‘ LAPC was not continued beyond the period of cardiac arrest.

Both static and TLV LAPC succeeded in reducing the pulmonary artery temperature to the desired temperature of 34^oC by 6 and 10 min post arrest, respectively. Eighty two percent of the animals in both LAPC groups were resuscitated successfully and survived the 1 hour evaluation period without pharmacological support, as contrasted with only 27% of the controls animals (p = 0.03).  Interestingly, 73% of the isothermic  TLV swine achieved and maintained ROSC as opposed to 27% of the swine in the control group (p = 0.09), suggesting that the presence of PFC in the lungs during either the arrest or resuscitation interval may improve the odds of successful defibrillation.

Figure 2-17: Results of the study by Staffey, et al., to evaluate the effect of cold and isothermic PFC infusion and TLV on resuscitability of swine after 11 min of electrically induced VF cardiac arrest. PFC loading, and in particular cold PFC loading or TLV improved 1 hour stable hemodynamic survival by 73% versus 27% in controls. Infusion of isothermic PFC to near vital capacity also showed a trend towards increased survival. Redrawn from Staffey, et al.,[284].

Chamberlain et al., have recently proposed that dilation of the right ventricle (RV) that occurs after the first ~5 min of cardiac arrest results in compression of the left ventricle (LV) within the confines of the pericardium. The effect of this would be to reduce myocyte stretch and reducing the contractile strength of the left ventricle in response to defibrillation (see discussion of the Frank-Starling curve above). They posit that an arrested heart that has a reduced contractile state, and in which left ventricular volume is reduced by the right ventricle over-distended with venous blood will not be able to initiate effective contraction even if coordinated electrical activity is restored. Distension of the RV occurs post-arrest due to centrally mediated sympathetic contraction of the vasculature, as well as movement of blood into the venous circulation as arterial and venous pressure equilibrate; this is a commonplace finding at both necropsy and autopsy in normovolemic subjects.

Unloading the RV prior to defibrillation results in an immediate improvement in successful ROSC independent of any known metabolic that might accrue from ~10 to ~15 sec of coronary perfusion that might also result. Chamberlain, et al., believes that it is the decompression of the left ventricle and restoration of a more physiologic morphology that facilitates or even enables defibrillation under these conditions. Instillation of a large volume of dense PFC may antagonize distension of the RV and prevent LV volume loss and compression of the LV to below the threshold required to established effective contractile activity in response to defibrillation. PFC to vital capacity (VC) should effectively preclude this post-arrest pooling of blood in the thoracic caval and pulmonary vessels and may act to preserve left ventricular morphology during prolonged cardiac arrest.

Especially encouraging findings from the Staffey, et al., study were the absence of any noticeable adverse hemodynamic impact from either isothermic PFC loading or from cold PFC loading or cold TLV. MAP, CVP, CPP, pH and blood gases were not statistically different between the four groups of animals in the study.

These four studies of LAPC aimed at answering questions bearing on the feasibility of LAPC in MI, SCA and CPR are very encouraging. They indicate that LAPC is being seriously considered as a potential therapeutic application in humans. That this research is being undertaken in independent academic and medical research both in the US and elsewhere is especially heartening and would seem to indicate that the enormous therapeutic potential of LAPC-induced hypothermia has been successfully communicated.

6. Conclusions

LAPC is capable of inducing hypothermia in a fraction of the time that it takes to prepare a patient for cooling via CPB. In addition, automated LAPC need not have the spatial and technical restrictions of the hospital setting. Although relatively simple methods of continuous arterio-venous shunt heat-exchange that do not require a blood pump or carry the most of the risks attendant to CPB have been described which might be potentially applicable in the field [285], these techniques also have the drawback of requiring skilled personnel for cannulation of a major artery and vein. Intracaval heat exchange catheters can reduce core temperature,  but only at rates of ~1.46 ± 0.42°C/h [286]; far too slowly to achieve the maximum benefit from post-reperfusion hypothermia

Figure 2-18: Large negative excursions in airway pressure occurred during suctioning of PFC at the end of most FTLVs when this operation was carried out manually (A). This occurred because when the PFC suction catheter was no longer filled with PFC, evacuation of gas from the airways occurred very rapidly owing to the much lower viscosity of gas compared to PFC. It was impossible for the operator to anticipate when the last of the bulk liquid would be removed, or to react rapidly enough when this occurred. Computerized sensing and control eliminated this potential source of baro-injury (B).

By contrast, LAPC may be a candidate for a much wider range of emergency field-uses in civilian and military settings since the primary technical skill required to initiate LAPC in the field is endotracheal intubation; a skill possessed by paramedical personnel throughout Canada, the U.S. and Europe. LAPC has also potential as a very rapid treatment for heatstroke and malignant hyperthermia. While  not the subject of this report, LAPHE clearly also holds promise for core rewarming in severe hypothermia, although an absolute maximal PFC temperature of 42°C would in theory limit the re-warming rate to about one-third of that possible in cooling.

Computer control of both gas and FTLVs is effective at eliminating excessive positive and negative airway pressures during LAPC (Figure 2-16) and computer control can easily be extended to encompass cooling rate, depth, and duration and can also be extended to control gas ventilation, as necessary.

Whether used inside or outside hospital, successfully implemented LAPC might more generally serve as a neuroprotective bridge [287] in order to gain time for more technically sophisticated supportive or definitive treatment (e.g. neurovascular thrombolysis or interventional thrombectomy, emergency CPB, spinal cord decompression or definitive management of hemorrhagic shock in trauma or surgery).

Although some modalities of liquid ventilation have been clinically evaluated [288], the safety parameters of rapid and cold liquid delivery to the lungs remain to be determined. As noted in this study, LAPC can cause pulmonary injury. The mechanism of such damage suggest by the location of lesions indicates that both barotrauma (dependent lung) and volu-trauma (nondependent lung) are the primary, if not the sole factors. It has been observed that LAPC causes little permanent lung injury in long term survival animals. Similar pathology seen in lungs exposed to either isothermic or ~4ºC LAPC in the present study (data not shown) suggest that thermal/chilling-injury per se is not the major insult. Although more subtle biochemical and immune problems secondary to hypothermia itself are suggested by reports from some longer duration studies of MTH (pneumonia and sepsis), it is not clear that the short duration of treatment necessary to achieve the benefit of post-resuscitative MTH will pose such problems. It is hypothesized that the pulmonary injury observed in LAPC may be reduced with better control of LAPC pressure and volume limits, and by use of PFC liquids having more physiologically suitable properties.

A significant, unresolved concern is the potential negative impact of LAPC on myocardial perfusion and the danger of overcooling the heart during the low flow conditions of CPR with possible adverse effects on achieving ROSC and maintaining a stable rhythm following defibrillation. As already noted, these questions can only be resolved with further study.

Acknowledgements

 The authors thank Saul Kent and William Faloon for support, and Casey Brechtel for helpful discussions. Several of the authors have applied for LAPC device patents. This trial was funded by a grant from the Life Extension Foundation (Hollywood, FL).

Appendix A

 (The abbreviations contained in this appendix are also reproduced in the table of abbreviations and acronyms at the beginning of this document.)

 A.1. Abbreviations and notation

 In the text, volumes (V) are given in ml/kg, and flows (dV/dt =V’= ˙V ) in ml/kg per min. Since all V and ˙V are expressed in mass-specific (per kg animal) terms, derived quantities DQ and Cm are automatically mass-specific. C_m and C_vf are given in calories / (g or ml) per °K for easy comparison with water.

PFC = perfluorochemical; hydrogen-free organic molecule in which most of the peripheral atoms are fluorine.

TLV = tidal liquid ventilation is a modality in which liquid completely fills the lungs and ventilator.

PLV = partial liquid ventilation is a modality in which all gas exchange is via gas ventilation, with ~1/2 FRC of PFC liquid present in the lungs to recruit dependent lung in ALI or ARDS.

LAPC = liquid assisted pulmonary cooling is a heat-exchange modality in which ventilation occurs via both gas and liquid ventilation proceeding concurrently.

T_tym = tympanic temperature

T_art = arterial temperature

T_ven = central venous temperature

T_rec = rectal temperature

DT_enet DT_tymp = resulting from LAPC, after equilibration at t=40 min

T_FTLV  = FTLV cycle infusion time

t_s = FTLV cycle suction time

t_c = FTLV cycle period (=tinftc +ts)

V_FTLV = single-cycle PFC FTLV infusion volume=tinf V_ inf

 V_S = single-cycle PFC FTLV suction-return volume

V_D  = ventilatory dead space (any type)

V_DCA = expected gas ventilation VD=sum of circuit (mechanical) VD plus anatomic VD

V_DTherm = thermal or heat-exchange VD (ml/kg, in reference to liquid PFC infusion)

˙V _inf = PFC infusion rate (set to _50 ml/kg per min in Trials I and II)

˙V_FTLV = effective PFC FTLV rate=LAPC liquid FTLV minute-ventilation (ml/kg per min) = V_FTLV/ t_c

˙V_g = gas minute-ventilation (ml/kg per min) m animal mass Ch heat capacity

 C_T = total heat capacity of the animal (= mC_m)

m = mean mass-specific heat capacity of the animal ( = DQ_T / DT_e)

C_vf = volume-specific heat capacity of FC-75 (mean of 0 and 25°C values) = 0.45 cal/ml per °K

DQT = total heat removed during LAPC (kJ/kg animal) = S DQ_c

DQc = heat removed during one FTLV cycle?????

E_f = mean cycle heat transfer efficiency=mean of [DQ_c / (theoretic DQ_c (max)] for all cycles in a single experiment

n = number of FTLV cycles in LAPC experiment

S = sum entire quantity following, for all cycles i = 1 through n

 T_inf   = PFC infusion temperature

T_S = PFC suction removal temperature (time-averaged PFC suction flow temperature)

T_SM= PFC mixed suction return-volume temperature (temperature of mixed VS)

 A.2. Thermal kinetics

 During LAPC cooling and equilibration, the blood and tympanic temperature changes in the animals were modelled by a simple five compartment model (Figure 2-8). During the initial ~100 s of LAPC (value used as empiric time mark), full development of heat-exchange behavior is established between the lungs, blood volume, and the thermal core of the animal, as suggested by the characteristic half-times for equilibration of these systems (see below).

Modelling of cooling during LAPC:

After the initial ~100 s of cooling, the data for tympanic DT(t) = DT_tym during LAPC in Trial I and II were modelled by a single time-constant exponential decline.

Mean DTtym data for each trial from times t=100 to 1080 s were fit using (Eq. (1)).

DT (t) = T₁₀₀ + DT_k [1− exp (−t / t_o) ],

DT(t), total T_tym change from baseline T_tym at start of LAPC; t, =time in seconds after empiric time mark, 100 s after start of LAPC; T₁₀₀, observed DT at empiric time mark, 100 s after start of LAPC; DT_k, observed temperature-interval constant, specific to each LAPC method; to, observed natural-base time-constant, in sec (t_o = halftime / ln 2).

Best-fit values for Trial I data were: T₁₀₀ = −0.52 ± 0.02°C; DT_k = −11.2 ± 0.02°C; and to = 1064 ~3 s. Trial II values were T₁₀₀ = −0.24 ± 0.02°C; DT_k = −8.14 ± 0.02°C; and t_o = 1107 ± 5 s. The relatively long time-constant associated with this thermal phase, which was similar in the two trials, presumably reflects the long time-constant (~700 s, see below) associated with heat transfer from the thermal core of the animal to the thermal periphery; thus the exponential phase represents full development of heat exchange between the LAPC cooling device and the entire animal. The final linear segments of cooling occurring after this phase, measured at −0.29°C/min (Trial I) and −0.21°C/min(Trial II), represent the final relatively simple state which exists after heat exchange equilibrium between cooling device and animal has been fully established.

Cooling in blood and tympanic sites during LAPC, and thermal evolution in these sites during equilibration phase after LAPC was discontinued, was in accordance with a five-compartment thermal model (Figure 2-8). In this model, the tissues of the animal are divided into three thermal compartments, corresponding loosely with the vascular system, the thermal core, and the thermal periphery.

Modelling of equilibration after LAPC:

Perfusion-driven convection is the major heat transfer mechanism in very rapid systemic cooling processes. This fact allowed T_art and T_ven changes to be used to quantify some features of heat transfer between body thermal compartments during the equilibration period after LAPC. The mean T_art curve in Trial I increased nearly linearly (R² = 0.9976) for 12 s after the end of LAPC, rising at a rate of 7.9°C/min. After this initial 12 s, T_art departed from linearity (Figure 2-6 inset), and was modelled by the sum of three exponential terms with respective time constants (t₀) of 12 ± 0.4, 102 ± 2 and 701 ± 8 s. These t₀ times differed to a large enough extent that their respective influences could be considered to be controlling over discrete time periods of about twice their value. Thus, the four equilibration phases seen after the end of LAPC lasted approximately 12, 24, 200, and 1400 s (23 min), respectively and represented 34, 14, 25, and 27% of the 5.1°C rise in T_art during equilibration after LAPC.

These data may be interpreted as follows: during each phase of the equilibration process, one or more thermal compartments in the animal equilibrated with the next-most closely-connected compartment (Figure 2-8). Afterwards, the newly captured compartment(s), as part of a larger unit bound together by blood mediated convection, equilibrated with the next-most closely connected compartment, and so on. The 12 s linear first equilibration phase (Figure 2-7, inset) most likely represents development of heat transfer from the lungs to local pulmonary blood flow. This phase was not associated with blood recirculation since it was seen as a rise in T_art but not T_ven. The second equilibration phase (duration ~24 s) was characterized by an increase in dT_ven / dt to the value of d_Tart  /dt, indicating that the lungs, blood-volume, and certain other well-perfused viscera, such as the kidneys, were now evolving into a single thermal system. Since the observed to for this phase was 12 s, less than the animal’s mean circulation time (= cardiac output / blood volume ~30 s), this process appeared to be driven by blood circulation via the most rapid paths (e.g. renal circulation). Such short paths for circulatory heat transfer were evident in the relatively small lag times (10.4 ± 6.9 s) noted between T_art and T_ven changes in these animals.

During the first two equilibration processes, the pulmonary circulation added thermal potential to the blood-volume more rapidly than it could be removed by the systemic circulation. By the end of the second equilibration phase, however, lung-to-blood heat transfer no longer dominated, and the gap between T_art and T_ven was set by the magnitude of heat transfer from the circulating blood volume to the tissues that comprise the ‘thermal core’ of the animals. In this third equilibration phase (duration ~200 s), the viscera and blood-volume, as a unit, equilibrated with the remainder of the ‘well-perfused’ tissues of the body (thermal core, comprising about 70% of the animal’s heat capacity). Heat capacities for thermal compartments are calculated below. The t_o for this process is seen most directly in the ~2 min. delay between maximal DT_ven and maximal DT_tym (Figure 2-7).

Finally, heat transfer within well-perfused tissues fell to a new minimum, and the T_art to T_ven gap decreased to a value set by the fourth equilibration phase (duration ~23 min) during which the well-perfused tissues equilibrated, as a unit, with a succession of the more poorly-perfused compartments, e.g. gut contents, fat, and other tissues comprising the thermal ‘periphery’ [12]. These processes could be consolidated into a single exponential term. Due to the long time-scale, heat transfer during phase four was probably partly conductive. Estimates of basal metabolism in the anesthetized, non-shivering dog (@90 J/kg per minute) indicate that as much as 0.6°C of warming per 20 min in this model may be due to metabolism.

A.3. Thermal accounting

 Heat transfer efficiency:

Although machine-LAPC allowed 2.3 times the FTLV frequency of the manual method, and resulted in a larger ˙V_FTLV by a factor of 1.2, the cooling magnitudes and rates for machine-LAPC significantly (P < 0.001) fell short of those obtained with manual-LAPC (Fig. 3). The strategy of increasing FTLV frequency (1/t_c) and decreasing ˙V_FTLV, in order to arrive at approximately the same FTLV rate (˙V_FTLV), therefore, significantly decreased the fraction of thermal potential which was transferred from each FTLV (= heat transfer efficiency, E_f). Unexpectedly, when the E_f for each of the 8 LAPC cooled animals of Trials I and II (Table 2) was calculated using (Eq. (2)), the value did not differ (P = 0.46) between trials; nor did total heat removed per kg (DQT), as calculated using (Eq. (3)), differ (P=0.14) between Trials. Both of these quantitative methods were therefore inaccurate for some dogs. Calculation of whole-animal mass-specific heat capacities C_m (= DQT / DT_e) suggested that the Trial II values of ~DQT and E_f principally were inaccurate, since the mean C_m value of 0.70 ± 0.1 cal/g per °K for Trial I was consistent with the C_m reported in the literature for mice and humans [289], whereas C_m values calculated for Trial II using (Eq. (2)) were unrealistic, being greater than the C_m of water.

E_f  (method 1) = 1/nS (T_S−T_inf) / (T_ven−T_inf),     (2)

DQ_T (method 1) = ˙V_FTLV C_vf ST_S−T_inf.                     (3)

To independently check the accuracy of Trial I values, we computed DQT and C_m for three dogs from an earlier study (Tables 1 and 2: dogs A, B, and C) that had been given ~4ºC PFC with FTLV times sufficiently long to allow the volumes and mixed-temperatures of suction-return liquid to be measured for each FTLV.

This allowed computation of DQ_T and E_f by a more detailed method (Method 2, Eqs. (4) and (5)), which used the extra thermal data (not available for Trials I and II) to more directly estimate FTLV heat transfer.

DQT (method 2)

= C_vf S V_FTLV (T_ven−T_inf) − V_S (T_ven−T_SM),                              (4)

When this was done, the mean Cm for dogs A, B, and C was found to be 0.68 ± 0.06 cal/g per °K, consistent with the Cm in Trial I (P=0.80). Method 1 (Eqs. (2) and (3)) required the assumptions that PFC suction volume equalled infusion volume, that suction flows remained constant, and that thermal hysteresis was negligible. These assumptions apparently held true at the larger ˙V_FTLV and t_c values of Trial I, but not for the smaller values of Trial II.

With this information, a new E_f for Trial II was estimated using method 3 (Eq. (6)), which employed an estimate for the heat required for the observed DT_e, vs. the total PFC thermal-deficit theoretically available.

This estimate required a presumed value of C_m. However, if the mean Trial II Cm was assumed to be the same as that of Trial I, then the true Trial II Ef could be calculated (Eq. (6)) to be 0.40 ± 0.06.

This value agreed with the rough estimation that since Trial II achieved only 73% of the DTe of Trial I, despite using 1.19 times more total PFC (Table 1), the

E_f in Trial II was expected to be about 73%/1.19 = 61% that of Trial I.

Thermal compartment size:

Heat removal for each cycle (DQ_c) in Trial I was calculated from the individual terms of Eq. (3), and individual-cycle mean cooling power calculated as DQ_c/t_c. The latter parameter was useful since thermal compartments in the dog are relatively isolated at short time scales (Figure 2-8), and thus the ratio of cooling-power to cooling rate (Eq. (7)) at a given probe site was expected to give the heat capacity (C_h) of the system of thermal compartments that were in equilibrium with each other, and with the site, at the time of the measurement:

C_h (Comp N), total heat capacity of thermal compartments N = 2 + 3, or N = 2 + 3 + 4.

Both t_c and dT/dt values were chosen at a time t, of interest when compartment system N has not yet equilibrated with slower half-time compartment(s). Thus, in

Trial I, near the end of FTLV cycle c1 (t = 30 s), the FTLV thermal-deficit had equilibrated within thermal Compartments 2 + 3 (PFC/viscera/blood-volume), but had not yet significantly reached Compartments 4 or 5.

If the cooling rate of T_ven at t = 30 s (−1.9 ± 0.8°C/min) was then taken as the cooling rate of the system of PFC/viscera/blood-volume, the C_h for this compartment system could be estimated from (Eq. (7)) as 20 ± 9% of C_T, the total heat capacity of the animal (C_T = m C_m). Subtracting the C_h contribution of lung PFC (=m ˙V_FTLV C_vf) allowed estimation of the remaining tissue C_h for Compartment 3 as ~19 ± 9% of C_T.

Similarly, the Ch of Compartments 2, 3, and 4 together, was estimated at t = ~140 s (cycle 4) as 71 ± 17% of C_T, corresponding to the classical whole-body ‘thermal core.’ The Compartment 5 C_h was then calculated to be the remainder (100−71%) = 29 ± 17% of C_T, corresponding to the classical ‘thermal periphery.’

Table of Contents

Introduction………………………………………………………………………………………………………. 6

Table of Abbreviations, Symbols and Acronyms………………………………………….. 9

Section I: Introducing Liquid Assisted Pulmonary Cooling………………………… 14

Liquid Assisted Pulmonary Cooling in Cardiopulmonary Cerebral Resuscitation, Introduction

………………………………………………………………………………………………………………….. 15

Hypothermia as an Active Therapeutic Agent

………………………………………………………………………………………………………………….. 17

The Benefits and Limits of ‘Delayed’ MTH: Real World Experience

………………………………………………………………………………………………………………….. 26

The Problem of Heat Exchange

………………………………………………………………………………………………………………….. 29

The Pathophysiology and Biophysical Limitations of External Cooling

………………………………………………………………………………………………………………….. 33

Consideration of Invasive Core Cooling Methods

………………………………………………………………………………………………………………….. 35

Exsanguinating Trauma Resulting Cardiac Arrest

………………………………………………………………………………………………………………….. 37

The Lungs as Heat Exchangers

………………………………………………………………………………………………………………….. 38

A Brief Précis of the History and Development of Liquid assisted Pulmonary Cooling (LAPC) for Induction of Hypothermia during CPR

………………………………………………………………………………………………………………….. 41

What LAPC Can Potentially Deliver

………………………………………………………………………………………………………………….. 46

Section 2: Experimental Studies to Determine the Effectiveness of LAPC under Laboratory Conditions ……………………………………………………………………………………………………… 51

1. Introduction

………………………………………………………………………………………………………………….. 52

2. Materials and methods

………………………………………………………………………………………………………………….. 53

2.3. Trial I (manually-controlled LAPC)

………………………………………………………………………………………………………………….. 59

2.4. Trial II (machine-controlled LAPC)

………………………………………………………………………………………………………………….. 59

2.5. Animals A, B and C

………………………………………………………………………………………………………………….. 59

2.6. Data collection and correction, statistical methods, graphical display and presentation

………………………………………………………………………………………………………………….. 59

3. Results

………………………………………………………………………………………………………………….. 60

3.1. Thermal results of LAPC

………………………………………………………………………………………………………………….. 62

3.1.1. Cooling time delay

………………………………………………………………………………………………………………….. 63

3.1.2. Cooling rate

………………………………………………………………………………………………………………….. 63

3.1.3. Mean cooling power

………………………………………………………………………………………………………………….. 63

3.2. Gas exchange

………………………………………………………………………………………………………………….. 64

3.3. Clinical observations and gross pathology

………………………………………………………………………………………………………………….. 66

3.4 Impact on Hemodynamics

………………………………………………………………………………………………………………….. 67

4. Discussion

………………………………………………………………………………………………………………….. 69

4.1. Apparent effect of temperature on gas exchange

………………………………………………………………………………………………………………….. 69

4.2. Thermal transfer efficiency and kinetics

………………………………………………………………………………………………………………….. 70

4.3. Question of diffusion dead space in LAPC

………………………………………………………………………………………………………………….. 71

4.3.2. Possible synergy of combined gas and liquid ventilation in assisting mass (CO₂) and heat transfer

………………………………………………………………………………………………………………….. 73

4.4. Potential development of clinical LAPC

………………………………………………………………………………………………………………….. 73

4.5. Challenges Ahead

………………………………………………………………………………………………………………….. 74

Acknowledgements…………………………………………………………………………………… 81

Appendix A……………………………………………………………………………………………….. 81

Section 3: Perflurochemicals………………………………………………………………………… 96

The Perflurochemicals

………………………………………………………………………………………………………………….. 97

Physical Chemistry and Synthesis

………………………………………………………………………………………………………………….. 97

Physical Properties

………………………………………………………………………………………………………………….. 98

Commercially Available PFCs

………………………………………………………………………………………………………………….. 99

Toxicology

………………………………………………………………………………………………………………… 102

Environmental Impact and Future Availability

………………………………………………………………………………………………………………… 104

Section 4: History of Liquid Assisted Ventilation and Implications for LAPC

………………………………………………………………………………………………………………………109

History of liquid Ventilation

………………………………………………………………………………………………………………… 110

Partial Liquid Ventilation (PLV)

………………………………………………………………………………………………………………… 114

Unanticipated Effect of Lung Protective Ventilation Strategies

………………………………………………………………………………………………………………… 119

Defective Translational Research Models

………………………………………………………………………………………………………………… 119

Failure to establish a Dose-Response Curve

………………………………………………………………………………………………………………… 120

Gas Trapping and Selection of Appropriate PFCs

………………………………………………………………………………………………………………… 120

The PFC-Air Interface and Shear Effects in Small Airways

………………………………………………………………………………………………………………… 121

PLV and the Law of Laplace

………………………………………………………………………………………………………………… 122

The Best as the Enemy of the Good

………………………………………………………………………………………………………………… 127

Possible Implications for SCA and LAPC

………………………………………………………………………………………………………………… 127

__________________________________________________________

This work is dedicated to David W. Crippen, MD, FCCM, who, literally, made it possible. The computers this work was written on, the myriad books and journals consulted, and the liaison with many of the professionals required for such work were all directly facilitated by Dave Crippen. But, these are the least important elements he contributed. By far the most critical ingredients were and are his faith and friendship. For these he has my unending thanks.

___________________________________________________________

 
Intellectual Property Rights: The intellectual property rights to the technology of liquid assisted pulmonary cooling (heat exchange) (LAPC or LAPHE) using combined gas and fractional tidal liquid ventilation are controlled by Critical Care Research, Inc., under US Patent 6,694,977, “Mixed-mode liquid ventilation gas and heat exchange.” Inquiries regarding access to this technology should be directed to Critical Care Research, Inc., 10743 Civic Center Drive, Rancho Cucamonga, CA 91730-3806. Telephone: (909)987-3883.

Liquid Assisted Pulmonary Cooling in Cardiopulmonary Cerebral Resuscitation

By Michael G. Darwin

Introduction

Liquid assisted pulmonary cooling (LAPC), or liquid assisted pulmonary heat exchange (LAPHE) will not likely be applied in the West any time in the foreseeable future for many reasons; not a few of them having to do with regulation and consumer perception of what constitutes acceptable risk versus benefit.

LAPC relies completely upon the unique properties of the perflurochemicals (PFCs). PFCs are other-worldly molecules; neither soluble in water or lipids, twice as dense as water, but much less viscous; and chemically so inert as to make gold seem a highly reactive metal by comparison. It is impossible to work with the PFCs suitable for introduction into the lungs of mammals without immediately appreciating how truly amazing they are. PFC chemistry exists in a world of its own outside of conventional chemistry, and in fact, conventional chemists refer to PFC chemists as ‘unnatural’ chemists; with the all the double entendre that moniker implies. PFCs do not occur in nature, and if you want to find a marker for intelligent, industrial life elsewhere in the universe, one approach might be to look for the spectra of PFCs in the atmosphere of the candidate planet; they occur only as a result of deliberate, intelligent industrial activity. Thus, they are what I call, ‘the thinking species’ molecule.’

There is an ancient Chinese curse: “May you live in interesting times and come to the attention of important people.” The medical-molecular equivalent is: “May you be synthesized in interesting times and come to the attention of important medical (and environmental) bureaucrats.” This perfectly describes the fate of the medically useful liquid PFCs. They are dramatic molecules and they are at once both therapeutically paradigm changing and physiologically outrageous. There is a strong emotional reaction – one part wonder and one part shock – at seeing a patient’s lungs filled up with liquid, any liquid. Bulk liquid does not belong in mammalian airways, and in every clinical instance where it occurs, or is introduced into the lungs – in pulmonary edema, drowning and inhalation of organic liquids – it is an unmitigated disaster. The ‘breathable’ PFCs thus attracted a lot of attention.

So, PFCs had this strike against them, and they came about as putative therapeutic agents in ‘interesting’ times in the most Chinese sense of the word. The explosion of technological innovation that has occurred since the Industrial Revolution has brought not only powerful advantages, but also often damaging and not infrequently frightening downsides. We live in an age when people, at least in the West, have become acutely sensitized to the dark side of technology and increasingly unwilling to accept risk unless it is precisely quantified, small, and will almost certainly not affect them.

The PFCs present some unfortunate challenges in this regard. First, while they are among the most chemically inert substances known to man, they are biologically active in fentogram quantities. Most of the medically useful PFCs abolish white blood cell chemotaxis in picogram to fentogram concentrations.  While their very inertness and lack of chemical reactivity make them indispensable as a liquid ventilating medium, these same properties mean that, should they escape the lungs due to pneumothorax (and enter the mediastinum or other closed body viscuses or capsules) they will likely remain there for the rest of the patient’s life. This may seem ample reason to eschew these molecules as therapeutic agents. However, I believe these facts must be taken in the context of how we use and abuse similar (and in some cases nearly identical) molecules both in medicine and in our daily lives.

Introducing liquid into patients’ lungs will never be a trivial matter, or something undertaken lightly. Such therapies and their enabling molecules will not be broadly used (or abused) in medicine. Indeed, their use is necessarily confined to the emergency medical and critical care setting in intubated and mechanically ventilated patients. Foreseeable off-label medical use of neat PFCs is virtually nil. And yet, these molecules will ‘seem unlikely to be approved for medical use in the West in foreseeable future.

All of the existing molecules are either unsuitable for use in the induction of hypothermia (such as Perflubron™ which freezes at +6 deg C) or are chemically heterogeneous, as I explain in this manuscript. It would thus be necessary to synthesize and purify new classes of molecules and then vet them individually for safety and efficacy. The cost of doing this is more than any Western drug company could bear when weighed against even the most optimistic projections of financial return on the investment. There is very little place for a drug that a patient will use once during a lifetime and that, if it works, will leave him more than able to sue all involved in its application to him should he develop some adverse effects, even if these occur only after decades of additional, healthy and productive life which would not have been possible otherwise.

I hasten to add that I do not think the PFCs will be free of long term side effects. Most of the chronic survival dogs in our studies that underwent PFC ventilation and/or cooling have since grown old and died. We saw nothing unusual; they died of heart failure, cancer and the usual things old dogs die of, and seemingly at the usual rates, and at the usual ages. But, our sample size was ‘microscopic;’ and I think it likely that large populations of people treated with PFCs may have a statistically significant increase in conditions related to the immunomodulating and immunosuppressive effects of these molecules; possibly more infections, more serious infections, and more neoplasms. My answer to that is, ‘so what?’ We accept adverse effects in oncology, chronic hemodialysis, chronic circulatory support (left ventricular assist devices) and solid organ transplantation that are, by comparison, the stuff of nightmares. We know and accept that children who are given radiation and chemotherapy have a good chance of surviving many childhood cancers, but very often at the expense of lowered IQs and a lifetime complicated by greatly increased risk of infection and of cancers unrelated to their original diagnosis. The liquid ventilating PFCs are benign compared to OTC aspirin which claims ~50,000 lives each year; mostly in treating nothing more life threatening than the pain of osteoarthritis, headaches and joint and muscle pain. And yet, the PFCs will not be approved for liquid ventilation or heat exchange applications in the foreseeable future; not here in the US, and not anywhere else in the West.

In the Russian Federation these very same molecules are approved and are being used parenterally in the form of any oxygen carrying ‘blood substitute’ called Perftoran.[1] It isn’t the best such oxygen therapeutic possible (Alliance’s Oxygent™ was far superior[2]), but it works and it is an acceptable risk; at least Russians physicians and regulators think so (and for what it is worth, I agree). The Chinese, vastly more Westernized than the Russians, have still to approve the first PFC-based blood substitute (or liquid ventilating media, for that matter). Nevertheless, it will be in one of these places, or someplace like them, not in the US, Australia or Europe, where LAPC/LAPHE is first clinically applied. So, the information present here will necessarily be of mostly theoretical interest. Hopefully it will find distribution in parts of the world where it will be of some practical value as well.

 Mike Darwin,

22 October, 2008

TABLE OF ABBREVIATIONS, SYMBOLS & ACRONYMS

Section 1:

Introducing Liquid Assisted Pulmonary Cooling

Liquid Assisted Pulmonary Cooling in Cardiopulmonary Cerebral Resuscitation

By Michael G. Darwin

Introduction

Each year in the United States there are ~450,00 deaths from myocardial infarction (MI) [1] (with 310,000of these deaths occurring before the patient reaches the hospital) as a result of a non-perfusing arrhythmia, principally ventricular fibrillation.[2] This mode of sudden cardiac arrest[3] (SCA) is also responsible for the majority of the 190,000 in-hospital deaths from MI, which typically occur within the first 24 hours following admission.[3]   Especially tragic is that 50% of these deaths occur in persons ~60 years of age or less.[4]  An estimated additional 20,000 incidents of SCA occur as a result of asphyxiation, drowning, electrocution, and genetic or developmental predisposition to lethal arrhythmias (Wolf-Parkinson’s White Syndrome, congenital thickening of the interventricular septum, and idiopathic arrhythmic disease) and other non-atherosclerosis causes. This latter category of SCA typically occurs in individuals whose mean age is less than 35.[5],[6]

At this time the principal treatments for SCA consist of initiation of manual, ‘bystander’ cardiopulmonary resuscitation, so-called Basic Cardiac Life Support (BCLS or BLS) followed by ‘definitive’ treatment of the arrhythmia beginning with defibrillation and the application of Advanced Cardiac Life Support (ACLS or ALS).[7]

 Figure 1-1 (right): Mortality from sudden cardiac arrest (SCA)in 2004  as a result of myocardial infarction compared to death from other ‘high profile’ causes of mortality in the US.

ACLS consists of the application of an algorithm of manual CPR, electrical defibrillation and pharmacologic therapy aimed at restoring a perfusing cardiac rhythm and adequate blood pressure and cardiac output to sustain life until definitive treatment of the underlying cause of the cardiac arrest can be achieved (e.g., coronary revascularization, implantation of an automatic defibrillator, or life-long anti-arrhythmic therapy).

 Figure 1-2 (right): Probability of survival as a function of time following cardiac arrest.[8]

As is shown in Figure 1-6 below, the time to survival without neurological deficit following cardiac arrest in the absence of BCLS declines rapidly following a sigmoid curve with survival without neurological deficit being ~80-90% following 1 minute of arrest time, and less than 10% following 9 minutes of arrest.[8] Put another way, 50% of patients will experience significant morbidity or death following 4 minutes of circulatory arrest (Figure 1-2).

What is not shown in this graph is that the effect of immediate bystander CPR on survival is negligible in most studies [9],[10] with the primary benefit being observed in patients who’s time from the initiation of BCLS to successful cardiac resuscitation was greater than 8 minutes.[11] There is evidence in the literature that morbidity is improved with prompt by-stander CPR [12] providing that EMS response is also rapid, although this remains controversial.[11],[13]  A corollary of this is that the overall survival rate following SCA, with or without serious neurological morbidity, ranges between 1% (New York City, NY) [14] to 17% (Seattle, WA).[15] The mean survival (defined as survival to discharge from the hospital) in the United States as a whole is generally agreed to be at best 15%  [16] with ~70% of these patients experiencing lasting neurological morbidity (ranging from ‘mild’ cognitive impairment to total incapacitation in the Persistent Vegetative State (PVS).[17],[18],[19]

The primary cause of non-survival in patients experiencing SCA is failed cardiac or cerebral resuscitation. Arguably, it is failed cerebral resuscitation, since most underlying causes of refractory cardiac arrest could be treated by ‘bridging’ supportive technologies such as emergency femoral-femoral cardiopulmonary bypass (CPB) until myocardial revascularization and hemodynamic stabilization were achieved.[20] When emergency CPB is applied to patients who are candidates for good neurological outcome, the survival rate is increased.[21],[22],[23],[24] However, these technologies are not typically used on patients who are unsuccessfully resuscitated (restoration of adequate cardiac rhythm and perfusion) because of the justified perception that irreversible brain damage would have occurred during the prolonged period of cardiac arrest or CPR/ACLS.[21]  Similarly, it is for this reason that most attempts to achieve cardiopulmonary resuscitation in hospitalized patients who are not hypothermic or intoxicated with sedative drug are terminated after 15 minutes.[25],[26]

Within medicine it is widely understood that ‘CPR doesn’t really work’ and that if the return of spontaneous circulation (ROSC) is not achieved within ~ 5 minutes of cardiac arrest, the chances for survival are slim, and the chances for survival absent neurological impairment are slimmer still.[8] The principal reasons that conventional CPR is not effective are that it fails to supply an adequate amount of flow at an adequate pressure. Cardiac output (CO) is typically ~1/3^rd of the at-rest requirement (~1.5 versus ~4.5 liters per minute), and mean arterial pressure (MAP) is typically 25 mm Hg to 45 mm Hg; well short of the 60 mmHg required to sustain cerebral viability.[27],[28]

The condition of the typical sudden cardiac arrest (SCA) patient and the circumstances under which he experiences cardiac arrest are far from the ideal of a patient who is a candidate for emergency cardiopulmonary bypass (CPB) in hospital. The typical SCA patient is middle aged or elderly, often suffering from one or more co-morbidities (diabetes, obesity, COPD, hypertension), and if subjected to prolonged CPR will invariably have impaired gas exchange due accumulation of fluid in both the parenchyma and the air-spaces of the lungs (pulmonary edema with alveolar flooding). This occurs because closed chest CPR quickly causes pulmonary edema.[29],[30]  As previously noted, even when the SCA patient is a ‘good’ candidate for salvage; someone who is relatively young and free of co-morbidities, CPR will likely prove futile due to cerebral ischemia-reperfusion injury and the post-resuscitation syndrome.

Over the past 25 years a vast number of therapeutic interventions have shown great promise in animal models of regional and global cerebral ischemia in the laboratory.[31],[32],[33],[34] In the last 6 years alone, over 1000 experimental papers and over 400 clinical articles on pharmacological neuroprotection have been published.[35],[36] However, with one exception, none of these interventions has been successfully applied clinically despite many attempts. [37],[38],[39],[40],[41],[42],[43],[44] The sole exception to this frustrating debacle has been the introduction of mild therapeutic hypothermia (MTH) as the standard of care for a select (and very small) minority of SCA patients.[45],[46],[47],[48],[49],[50],[51]

Hypothermia as an Active Therapeutic Agent

Since the demonstration by Safar, et al., of the neuro-salvaging effects of mild systemic hypothermia after prolonged cardiac arrest in dogs [52],[53] there has been an explosion of translational research which has lead to a transformation in our understanding and application of mild hypothermia.[54], [55] Once seen solely as a protective tool which conferred benefit by reducing metabolism, it has become clear that mild hypothermia (33°C–35°C) [56] has therapeutic effects which appear to be primarily anti-inflammatory and anti-apoptotic in nature, and which operate independently of hypothermia’s effect on metabolic rate.[57],[58] Table 1-1 reviews some of the known pro-inflammatory factors inhibited or moderated by mild therapeutic hypothermia (MTH) and documents the supporting literature.

 Table 1-1: Inhibition of Injury Cascades by Mild Therapeutic Hypothermia (MTH)

Reproduced with modifications from Zhao, H., Steinberg, GK, Sapolsky, RM., General versus specific actions of mild-moderate hypothermia in attenuating cerebral ischemic damage. J Cerebr Blood Flow Metab, 2007. 27: p. 1879-1894.

Table 1-1:  Intraischemic hypothermia delays or attenuates both ATP depletion (Ibayashi et al, 2000; Sutton et al,1991; Welsh et al, 1990) and anoxic depolarization (Bart et al, 1998; Nakashima and Todd, 1996; Takeda, et al, 2003), it also blocks glutamate release (Busto et al, 1989b; Patel et al, 1994; Winfree et al, 1996), suppresses inflammation (Kawai et al, 2000; Wang et al, 2002), maintains the integrity of the BBB (Dietrich et al, 1990; Huang et al, 1999; Kawanishi, 2003), reduces free radical production (Maier et al, 2002), inhibits protein kinase C translocation (Cardell et al, 1991; Shimohata et al, 2007a, b; Tohyama et al, 1998), inhibits matrix metalloproteinase expression (Hamann et al, 2004), and blocks both necrosis and apoptosis. Intraischemic hypothermia also preserves the base-excision repair pathway, which repairs oxidative damage (Luo et al, 2007). In addition to those cascades directly associated with neuronal injury, hypothermia further blocks astrocyte activity and inhibits white matter injury (Colbourne et al, 1997; Dempsey et al, 1987; Kimura et al, 2002). Similarly, postischemic hypothermia blocks free radical generation (Horiguchi et al, 2003), attenuates inflammation (Horstmann et al, 2003; Ohta et al, 2007), prevents BBB permeability (Preston and Webster, 2004), and suppresses caspase activities (Van Hemelrijck et al, 2005). Indeed, a browse through the literature gives an overwhelming impression that hypothermia seems to block every damaging event associated with necrosis or apoptosis. One reason for this impression of pan-inhibition may lie in the causality of ischemic damage.  For example, is the inflammatory response the cause of tissue damage or is it induced by brain injury? If it is the latter, then since hypothermia prevents tissue damage, it certainly also prevents the inflammatory response.

No doubt, the commonly cited ‘obstacles’ of lack of institutional protocols, lack of physician education about the benefits and guideline changes, as well as the inevitable inertia that accompanies any paradigm shift in treatment are playing a significant role in the failure of MTH to become the practiced standard of care for the post resuscitation syndrome.[64],[60] However, what is not being said, or considered, is that while MTH as currently practiced represents a large relative improvement in outcome, the benefits are still modest in absolute terms. Only a miniscule subgroup of SCA patients currently can benefit from MTH; and even in its best clinical implementation MTH still fails to rescue ~60% of that sub-group of SCA patients to whom it is applied.[65],[66],[67],[68],[69] This is in stark contrast to what can be achieved with MTH in ameliorating post-ischemic encephalopathy in the laboratory, where post-resuscitation MTH consistently provides rescue with stunning efficacy.[70],[71]

Figure 1-3: The impact of a delay of 10 min in inducing MHT is a dog model of cardiac arrest followed by 3 min of systemic ischemia, 7 minutes of mechanical CPR and 50 minutes of advanced life support. Hypothermia to 34^oC was induced beginning at 10 min post arrest in the early hypothermia group˜ and at 20 min post arrest in the delayed hypothermia group ¢.  In the early hypotherrmia group group, 5 of 7 surviving dogs were functionally normal (OPC 1 or 2), 1 had OPC 3, and 1 had OPC 4 (coma) at 96 hours of recovery. Histologically, 4 of 8 dogs in this group were normal (HDS 0), 1 had HDS 16, 1 had 22, and 1 had 98. The only surviving dog in the DH group was functionally normal at 96 hours (OPC 1, NDS 0) with an HDS of score of32 (mild injury) Due to early mortality only two other dogs in the delayed hypothermia group were evaluated histologically and their  HDS scores 38 and 45, respectively. Dogs in this study were scored by ‘overall performance categories’ (OPC; 1=normal, 2=moderate disability, 3=severe disability but conscious, 4=coma, and 5=death) Neurological function and  neurological deficit scores (NDS; 0% to 10%=normal, 100%=brain death). [72],[73] Histological damage scores were obtained by neuropathological examination of 19 distcrete brain regions for severity and extent of ischemic neuronal changes, infarcts, and edema.  A total brain histological damage score (HDS)  >40 represented moderate damage, and HDS >100 represented severe damag.[74] Redrawn from Nozari, A., et al., Critical time window for intra-arrest cooling with cold saline flush in a dog model of cardiopulmonary resuscitation. Circulation, 2006. 113(23): p. 2690-6.

The primary obstacle to realizing this bonanza in translation research has been the practical impossibility of achieving systemic cooling within the narrow therapeutic window demonstrated in animal models of SCA and resuscitation.[75],[67],[65],[66],[67],[68],[69] If the clinical outcome of MTH was even half that achievable in the laboratory, widespread application would likely have been rapid and uniform; there is rarely resistance to the ‘miraculous’ if it is simple, easy to understand, biophysically well characterized and highly cost-effective. MTH applied immediately post ROSC would be all of these things.[4]

Figure 1-4: The results of the Nozari, et al., [71] study on the effects of delayed MTH are presented graphically at right with the addition of historical controls from the literature treated similarly, but with no hypothermia (no survivors). This graphic illustrates the potency of truly rapid post arrest hypothermia in ncreasing survival.

The data in Figures 1-3 and 1-4 exemplify what is possible when MHT is induced within its optimum therapeutic window of 0-15 min post ROSC versus a delay of even 10 minutes. In this study by Nozari, et al., of PeterSafar’s group, [71] VF was electrically induced in 17 dogs all of whom were subjected to a period of 3 minutes of no flow beginning when the MAP dropped below 30 mm Hg, followed by 7 minutes of mechanical CPR and 50 minutes of advanced life support during which time VF was maintained and mechanical CPR was continued. Nine animals were treated with rapid (early) induction of MTH to 34^oC starting at 10  min post arrest (EH group) (concurrent with the start of ALS to simulate the time course of arrival of EMS paramedics) using a combination of cold IV saline and veno-venous heat exchange. Induction of hypothermia was not

Figure 1-5: Results of a study of 48 cardiac arrest patients treated with MTH via endovascular cooling. A strong correlation was found between rapidity of cooling and both neurological outcome and serum neuron specific enolase levels. Left: Time course of MTH among patients with good and poor neurological outcome. The curves indicate the course of mean body core temperature during MTH among patients with good (¢) and those with poor (p) outcome as well as in the entire (▬) patient group. Right: Correlation between time to coldest temperature (minutes) and the maximum NSE values (μg/L). Normal serum NSE is 9.6±0.7 μg/L. Redrawn from Wolff B, et al., Early achievement of mild therapeutic hypothermia and the neurologic outcome after cardiac arrest, Int J Cardiol (2008).

begun until 20 min post arrest in the delayed hypothermia group (DH group) which consisted of 8 dogs. Target core temperature was achieved at 6.0±2.7 minutes after the initiation of cooling (3.5 minutes after the start of veno-venous cooling) in both groups. The delay from arrest to reaching ~34^oC  was 16.6 min in the EH group and 25.4 minutes in the DH group.

After 60 minutes of VF, ROSC was achieved with cardiopulmonary bypass for 4 hours, and intensive care was given for 96 hours. In the early hypothermia group, 7 of 9 dogs survived to 96 hours, 5 with good neurological outcome. By contrast, in the delayed hypotherrmia group 7 of 8 dogs died of multiple organ failure within 37 hours (P=0.012);  3 animals in secondary VF that was resistant to CPR with antiarrhythmic treatment and repeated defibrillations. Only one dog in the EH group died, and that animal succumbed to single organ failure; pulmonary edema with hemoptysis. This study extends the previous work by this group documnting an optimum therapeutic window for MHT (in dogs) of ~10-15 min.[76],[76],[77] The therapeutic window of MTH after cardiac arrest has been demonstrated to be similarly short in other species.[78],[79],[80],[81]

The dramatic efficacy of MTH in the laboratory made quick converts of the pioneering researcher-clinicians who forged ahead with the application of MTH to SCA in the clinic precisely because it was dramatic; indeed it was as close to the miraculous as interventions in medicine come. The real barrier to translating that ‘miracle’ to everyday practice has been the seemingly intractable problem of achieving cooling over the same time course that has proven so effective in the research setting.

The problem is that the optimum therapeutic window for the treatment of cerebral ischemia-reperfusion injury appears to be in the range of 0 to 15 minutes post ROSC. One of the first follow-up studies on MTH carried out by Safar, et al., demonstrated that in a standardized model of cardiac arrest in dogs a delay in the application MTH of as little as15 min after ROSC abolished most of the benefit.[77],[82] While the work of Bernard, et al., [46] and that of the Hypothermia after Cardiac Arrest Study Group [48] demonstrated that delays in cooling of up to 2-3 hours post ROSC in humans still have sufficient clinical utility to justify the routine application of MTH in a selected group of SCA patients, this benefit is marginal when contrasted to that achievable in the laboratory when MTH is rapidly induced during the first 15 minutes after ROSC.[77] Thus, the optimum clinical benefit of MTH in ischemic and very likely traumatic, CNS injury requires the ability to achieve very rapid core cooling.[83]

Figure 1-6: Survival after cardiac arrest declines rapidly as a function of time to ROSC, exhibiting the sigmoidal curve shown at left, above (¢>¢), with essentially all patients failing to survive with normal mentation after arrest intervals of @ 10 min. Application of MTH (¢) within the window of @ 15 min offers the promise of squaring the survival curve in SCA of @ 10 min duration yielding a survival rate of ~65-70% with little or no neurological deficit. Application of deep (10-22^oC) or profound hypothermia (5-9^oC) (¢) may allow survival after intervals of as long as 1-2 hours of CPR. Graphic by M.G. Darwin

In the clinical arena the time to reach the target core temperature under ‘good’ circumstances is in the range of 3-4 hrs; not 10 to 30 min, as is the case in the laboratory. Even with such long delays in cooling the adverse effect of delay is still present. Wolf, et al., recently published a study of 49 out of hospital cardiac arrest patients who were treated with MTH (32.0-34.0°C; with a target temperature of 33.0°C) of 24 h duration using endovascular cooling.[69]  The study endpoints were neurological outcome on discharge from hospital and serum neuron specific enolase (NSE) levels (a sensitive and specific marker of neuroinjury) at 24 h intervals to 3 days (Figure 1-5).

Figure 1-7: The graph above shows the hypothesized relative effect on survival of effectively administered CPR started at 5 min post-arrest followed by defibrillation at 6 min and ACLS at 8 min post arrest (~30% survival). The light blue shaded area of this graph shows the expected improvement in survival if MTH is induced at the start of ACLS (8 min post-arrest) and target temperature is reached by 15 min post ROSC. The dark blue shaded area shows the potential of Emergency Preservation resuscitation (EPR) using moderate (10-22^oC ) or profound (5-9^oC) hypothermia to not only square the curve of survival with CPR, but to facilitate survival in patients who would otherwise not benefit from either BCLS or ACLS (i.e., refractory to defibrillation, hypovolemic, etc.). Graphic by M.G. Darwin

 As is the case in laboratory studies of MTH in cardiac arrest in dogs, Wolff, et al., found that neurological outcomes were binary, with no patients who survived experiencing moderate degrees of disability; patients either recovered well with no or /mild neurological impairment, or experienced severe disability (n = 1) coma or PVS (n = 6). Twenty-eight patients were discharged with a good outcome and a strong correlation was found between good outcome and the time interval from the start of cooling to the lowest temperature (p =.035) and a less robust correlation with the time to reach target temperature (p=.071). Similarly,

NSE levels were found to correlate well with the time required to reach the lowest temperature achieved in each patient (Figure 1-5).

Even with delays in the start of cooling that averaged 2.5 hrs; and a mean time to reach target core temperature of 6.8 hrs, additional injury accruing from slowness in cooling was still clinically and biologically apparent. Despite the homogeneity of the patients, their arrest times, and their course of treatment, ~60% of the patients in this study either did not survive, or were comatose or PVS.

The Benefits and Limits of ‘Delayed’ MTH: Real World Experience

To understand the benefits and limits of MTH when it is aggressively and competently implemented with currently available technology, it would be hard to find a better example than that of Wake County, SC. Wake County, is located in the northeast central region of North Carolina and is part of the Research Triangle metropolitan area, which consists of Raleigh, Durham, Chapel Hill, and surrounding urban and suburban areas. The area serviced by the Wake County Emergency Medical System (Wake EMS) has a population of 832,970 (as of 2007). The Wake County EMS operates 35 ambulances from 23 locations with 825 ALS personnel; ambulances are staffed with two paramedics 95% of the time and there is always one paramedic responding.[84] In 2002 the Wake EMS answered more than 50,000 medical requests for service. Based on the latest national data Wake County ranks third in the US for recovery from “survivable” cardiac arrests (primarily ventricular fibrillation-ventricular tachycardia). Nationally, the average survival rate is 17% for patients presenting with these arrhythmias.

 Figure 1-8: Improvement in overall survival of patients in cardiac arrest in response to the phased introduction of CPR per AHA 2005 Guidelines, use of an impedance threshold device (ITD) and in-field induction of MTH.[85]

 Beginning in January 2004 Wake EMS initiated a study to evaluate the efficacy of CPR as they then practiced it, and to evaluate the effectiveness of the impending change in the American Heart association (AHA) guidelines for CPR, the introduction of an impedance threshold valve (the ResQPod™) and Wake EMS’ planned implementation of the ILCOR Guidelines for post-arrest MTH.[85],[86] From January 2004 until April of 2005, Wake EMS personnel employed the then extant AHA guidelines, which mandated an emphasis on intubation and a 15:2 compression-to-ventilation ratio; with interruption of chest compressions for ventilation. This period constituted the baseline of the study, and data were collected per protocol; not gathered retrospectively.

During the baseline period survival to discharge from hospital was 2.4% for all patients given CPR and 12.1% for patients with ventricular fibrillation-ventricular tachycardia (VF-VT) arrhythmias. In April 2005 Wake EMS implemented continuous cardiac compression CPR with a 30:2 compression-to-ventilation ratio with emphasis on no, or very minimal, interruption of chest compressions.  After 12 months, the overall survival rate had risen to 4%; and had more than doubled to 21.8% for patients who presented in VF-VT.

In April 2006, Wake EMS added the use of an impedance threshold device (ITD) to improve cerebral and coronary perfusion during CPR. Introduction of the ITD resulted in an increase in overall survival to 4.5% and an increase in the survival of patients with VF-VT to 28.5%.

  Figure 1-9: Dramatic improvement in neurologically intact survival as a result of the phased introduction of CPR per AHA 2005 Guidelines, use of an impedance threshold device (ITD) and in-field induction of MTH.[85]

 The final phase of the investigative protocol began in October of 2006 when Wake EMS added in-field induction of MTH to the two previous interventions. MTH was induced using a combination of external cooling employing ammonium nitrate-water eutectic ‘instant cold packs’ applied to the axilla and groin, and  cold IV saline (1-2^oC, 30mL/kg to a maximum of 2 liters) given rapidly via two large bore catheters and/or intraosseous infusion. Criteria for induction of hypothermia were that the patient have ROSC and show no return of consciousness (Glasgow Coma Score (GCS) <8). Induction of hypothermia was initiated two to three minutes after ROSC. There was heavy emphasis on avoiding over-ventilation and on attempting to maintain end-tidal CO₂ (EtCO₂) at a minimum of 40 mm Hg.  Patients undergoing MHT were sedated with etomidate, paralyzed with vercuronium and given a titrated dopamine drip to maintain mean arterial pressure (MAP) between 90-100 mm Hg. The mean time to target temperature (34^oC) in this study was extraordinarily short: 68 minutes (95% CI 47 to 88); compared to the 2-3 hours typically required to induce MTH.

With the combination of continuous compression CPR, use of the ITD and prompt application of MTH, survival rates for the 12 months from October of 2006 to October 2007 had increase to 6.7% overall and 37.4% for patients with VF-VT. The odds of overall survival increased three-fold (95% CI 1.7 to 5.0) and the odds of survival for patients in VF-VT increased 4.3-fold (95% CI 2.2 to 8.6) from the beginning of the study (Figures 1-8 and 1-10).The probability of a good neurological outcome increased from 20% at baseline to 80% at the conclusion of the study (Figure 1-9). In a multivariate analysis, the odds ratios for survival for each phase of implementation were as follows:

Figure 1-10 (right): Multivariate odds for all factors in outcome evaluated during the Wake EMS study. MTH was by far the most powerful intervention. As in most previous studies of survival factors associated with CPR age and residence (home versus extended care or assisted living facility) had only modest impact on survival.[85]

• New CPR protocol: 2.13 (95% CI 1.12 to 4.04)
• Addition of impedance threshold device: 2.33 (95% CI 1.09 to 5.00)
• Addition of early hypothermia: 3.99 (95% CI 2.19 to 7.27)
• Patients who received bystander CPR 1.79-fold (95% CI 1.18 to 2.72) more likely to survive.

Figure 1-11 (right): The Engel-15 portable, (compressor-type) refrigerator/freezer has a 14 L capacity, weighs 11.5 kg and can maintain 12-13 liters of saline at 1-2^oC at ambient temperatures as high as 40^oC . It retails for ~$380 US. [Photo courtesy of Engel, Ltd., Australia]

Interestingly, all three elements of the Wake EMS protocol were implemented at a cost of less than $200 per patient. Due to budget constraints, Wake EMS chose simple, inexpensive commercial products for refrigeration of IV fluid and implementation of external cooling, as opposed to more costly products developed specifically for medical application, such as the EMCOOLS surface cooling system (Emergency Medical Cooling Systems, AG, Austria). Saline was kept at the requisite temperature of 1-2^oC with a compact, 12V operated, consumer travel refrigerator/freezer (Figure: 1-11, Engel-15: http://www.i-m-d.com/) and surface cooling was with generic ammonium nitrate cooling packs.

The Wake County EMS program is extraordinary in every way. It represents the best application of the best available technology by arguably some of the best medical and paramedical personnel in the US. The mean time to target temperature of 68 minutes is unprecedented in any clinical study employing MTH. Of the 359 patients who participated in the study (all comers) after MTH was in place; 25 survived. In the subgroup of 93 patients who presented with VF-VT; 34 survived, with 78% or 27 patients being discharged with a good neurological outcome. Put another way 92% of patients who presented under the most favourable circumstances (VF-VT), treated with the best currently available interventions, at the fastest rate of cooling so far reported, failed to survive or did so with profound neurological debility.

The primary difference between the survivors and the profoundly disabled or dead was the development of the post-resuscitation syndrome and the primary reason for this complication was not comorbidity, or delay in paramedical assistance, but rather delay in the rapidity of cooling which, if achieved within the first 15 min post ROSC, would have offered the prospect of neurologically intact survival in the range of 70-80% in patients presenting with VF-VT, and 30-40% in all comers.

These interventions, remarkable achievements that they are, do not escape from the harsh reality that the 400% increase in survival from cardiac arrest in Wake County, when expressed in absolute terms, means that the number of lives saved increased from ~5 to 25 – out of 395 SCA patients; a huge relative gain, but a comparatively small increase in the absolute number and percentage of lives saved, and minds salvaged. The true life saving potential of MTH remains elusive by virtue of its exceedingly small therapeutic window.

 The Problem of Heat Exchange

Because of this minute therapeutic window, there is a pressing need to achieve rapid and durable core cooling of patients during CPR by simple, easily accessible means.  External cooling is only effective at reducing core temperatures by 0.15 to 0.25ºC/min in the average patient undergoing CPR (Figure 1-12) and this is achieved only by complete immersion of patients in a stirred ice water bath.

The Efficacy of External Cooling in Four Cryopatients[5]

 Figure 1-12:  Comparison of the cooling rates of four cryopatients. Immediately following pronouncement of medico-legal death patients were given closed chest mechanical cardiopulmonary support and placed in a stirred ice water bath for induction of hypothermia. Epinephrine was administered as per ACLS guidelines; thus peripheral vasoconstriction would be expected to be comparable to that seen in the typical SCA patient undergoing cardiac resuscitation.

The most effective external cooling achieved by a commercial system using direct, whole body surface cooling employing circulation of ice water (ThermoSuit,™ Life Recovery Systems, Kinnelon, NJ)  is probably the work of Janata, et al., using human human-sized swine.[87]

They were able to achieve core cooling at a rate of 0.3^oC/min; however it is important to note that the animals in this study were not in cardiac arrest while undergoing CPR in the presence of profoundly peripherally vasoconstricting agents, such as epinephrine or vasopressin; as would usually be the case during ACLS in humans [59] and which is known to further slow surface cooling.[88]

 Figure 1-13: The Life Recovery Systems ThermoSuit™ employs direct ice water contact with the patient’s skin to achieve the maximum possible rate of cooling by external means. The system consists of an inflatable insulating and water containment patient enclosure inside of which the patient rests on a mat of Dacron bonded polyester ‘wool’ which acts to diffuse and film water pumped over the dorsal surface of the patient’s body. Water at 2-4^oC is thin-filmed over the ventral surface of the body by a thin, transparent blanket with many hundreds of small perforations through which water under pressure pours out and over the patient. Cold water is recirculated over crushed or cubed ice in an insulated reservoir containing a disposable liner and pumps. Cooling is computer controlled via a thermistor which can be placed in any desired anatomical location. All patient contact items are single-use and disposable (including, as previously mentioned, the pumps) http://www.life-recovery.com/.

The obvious problems with this system are its bulk (Figure 1-13), likely high cost, lack of ease in field deployment (again related to its bulk and weight) and the intrinsic physiological problems associated with the induction of hypothermia via external cooling. As extensively discussed in Section Two, the mammalian body consists of multiple thermal compartments transiently ‘isolated’ from each other by differences in blood flow and heat conductivity.[89] Broadly, these compartments can be classified as strongly and weakly circulated (perfused); corresponding to the body core and periphery. The core tissues receive ~63% of the resting cardiac output (CO) but constitute only ~19% of the total body mass. By contrast, the peripheral tissues receive ~37% of the basal CO and constitute ~81% of the body’s mass (Figure 1-14).

 Figure 1-14: The parenchymatous organs that comprise the visceral core of the body receive an aggregate of ~63% of resting the cardiac output while comprising only ~19% of the body mass. By contrast, the peripheral tissue mass which accounts for ~81% of body mass receive only ~19% of the cardiac output. External cooling profoundly chills peripheral tissues before significantly reducing core temperature. Values for organ and tissue masses were obtained from: IAEA. Compilation of anatomical, physiological and metabolic characteristics for a Reference Asian Man. Volumes 1 and 2. Report IAEA-TECDOC-1005, (Vienna, Austria: International Atomic Energy Agency) (1998), Boecker, BB. References values for Basic Human Anatomical and physiological characteristics for use in radiation protection. Radiation Protection Dosimetry.105(1–4): 571–574;2003, de la Grandmaison GL, Clairand I, Durigon M. Organ weight in 684 adult autopsies: new tables for a Caucasoid population. Forensic Sci Int. 2001 Jun 15;119(2):149-54 and Heymsfield SB, Gallagher D, Mayer L, Beetsch J, Pietrobelli A. Scaling of human body composition to stature: new insights into body mass index. Am J Clin Nutr 2007;86:82–91.Values for organ and tissue blood flows were obtained from: Williams, LR, Leggett, RW. Reference values for resting blood flow to organs of man. Clin Physiol Meas. 10:187-212;1989. Graphic by M.G. Darwin

 The Pathophysiology and Biophysical Limitations of External Cooling

The objective of MTH is to provide protection against ischemia-reperfusion injury to the brain, heart, kidneys and liver; the visceral organs that constitute the strongly circulated core of the body. The peripheral tissues (skin, skeletal muscle, connective tissues and bone) are at once much more resistant to ischemia and less well perfused.  External cooling rapidly chills the ischemia-resistant peripheral tissues cooling them profoundly, while failing to provide protection to the vulnerable parenchymatous organs in the body core. This is not only undesirable in terms of its inefficiency; it also poses a number of hazards and risks.[90] Hypothermia is therapeutic in ischemia-reperfusion because it down-regulates the immune-inflammatory response; a response that is vital for host defense, wound healing and hemostasis. Hypothermia, like any major medical intervention that perturbs fundamental physiological processes, carries with it serious risks, as well as benefits. In both animals and humans, hypothermia is markedly immunosuppressive [91],[92] and interferes with the both the biochemistry of the clotting cascade and the production of platelets and clotting proteins.[93],[94]

In humans perioperative minimal hypothermia (MinH) (36^oC) increases the rate of wound infections [95] and prolongs hospitalization. [96] These effects occur in part due to the regional thermoregulatory vasoconstriction MinH induces; which in turn leads to reduced oxygen delivery to injured  tissues, [97] inhibition of oxidative killing by neutrophils, [98] and reduced collagen deposition.[96]  MinH induces significant suppression of mitogenic responses to Concanavalin A (con A), phytohemagglutinin (PHA), and pokeweed mitogen (PWM) and these changes are known to persist for at least 48 h. The mitogens PHA and ConA activate T cells, whereas PWM stimulates both T and B cells, thus indicating that the suppressive effects of MinH involve a variety of lymphocyte subpopulations. Hypothermia of as little as 1^oC significantly inhibits production of  interlukins (IL-1б, IL-2, IL-6) and TNFα in post-surgical patients, [99] and this suppression of cytokine production persists for least 24 hours after even a brief post-operative hypothermic interval.[96] The inhibition of pro-inflammatory cytokine production by IL-1б and TNFα induce tissue factor which is critical to angiogenesis, collagen elaboration and fibroblast activation; all essential processes in wound repair and hemostasis.[100],[101],[102] Significantly, many of these of adverse effects of post-operative MinH can be prevented by maintaining normothermia in the perioperative period.[96]

In the Hypothermia After Cardiac Arrest Study Group, patients treated with MTH experienced twice the incidence of sepsis. This finding is consistent with other studies where MinH and MTH were found to double the rate of post operative wound infection.[96]  A recent meta-analysis of MTH for traumatic brain injury (TBI) found that the incidence of pneumonia was also doubled for patients undergoing MTH.[103]

External cooling causes greater perturbation in hemodynamics than does central cooling, [104] resulting in increased systemic vascular resistance and decreased cardiac index; a phenomenon observed in all of the patients in the Bernard, et al., study that employed MTH for post-arrest cerebral resuscitation.[46]  This is particularly undesirable in the setting of MI, CHF and cardiogenic shock. For these reasons, as is the case with any potent therapy, careful attention must be paid to the dose-response curve, and overshoot or excessive regional cooling must be minimized or avoided.[105]

In patients cooled with ice packs and non-feedback controlled cooling blankets, there have been persistent and extensive problems with overshoot. In one study, 63% of patients overshot to <32^oC, 28% to <31^oC, and 13% were inadvertently cooled <30^oC.[105] By contrast, patients in the European Resuscitation Council Hypothermia After Cardiac Arrest Registry who were cooled endovascularly had less overshoot (mean lowest temperature 32.9^oC, IQR; 32.6^oC to 33^oC) when compared to patients cooled by other methods (mean lowest temperature 32.4^oC, IQR: 31^oC to 32.9^oC).[106]

In part, the problem of overshoot in external cooling can be corrected by servo control of patient cooling. However, the fundamental problems of external cooling remain. Excessive cooling of poorly perfused peripheral tissues will inevitably result in ‘after-drop’ of core temperature as thermal equilibration occurs. This process is idiosyncratic and inherently difficult to model or predict. In large measure the speed and character of thermal equilibration between peripheral and central tissues will depend upon highly variable factors such as body morphology and composition, body surface area, cardiac output, regional blood flow distribution, and the administration of vasoactive medication such as ionotropes; with their attendant peripheral vasoconstriction.

When the tissues of non-hibernating (or unprepared hibernating mammals) are cooled to £20ºC a wide range of deleterious changes occur. The saturated fats which comprise cell membrane lipids undergo phase change, resulting in red and white blood cell rigidity; with accompanying inability to deform and pass through capillaries. Red cell aggregation also occurs and this, in association with reduced flow as a result of vasoconstriction, results in blood sludging and failure of the microcirculation.[107],[108]  Profound hypothermia, either local or systemic, results in hemoconcentration as a consequence of translocation of vascular water and electrolyte to the interstitial space.[109],[104],[110],[111],[112] This hemoconcentration further exacerbates regional ischemia in deeply chilled tissues.

Independent of injury from the freezing of water, moderate, profound or ultraprofound hypothermia is known to cause cellular damage which is referred to as ‘chilling injury.’ Chilling injury appears to be a multifactorial process in which alterations of membrane structure (reorganization of lamellar lipid sheets with lateral phase separation between regions of gel phase and regions of liquid crystal phase result in loss of membrane integrity), [113],[114] failure of ion pumping (with consequent disruption of cellular ionic homeostasis), [115],[116] depolymerisation of some elements of the cytoskeleton, [117],[118] generation of free radicals, [119],[120] and metabolic disruption due to selective inactivation of critical enzymes [121] all appear to play a role.

Cooling of tissues to 5ºC for as little as 1 hour has been shown to cause microvascular endothelial damage similar to that observed in ischemia-reperfusion injury; loss of endothelial cell tight junctions, infiltration of capillary and venule walls with leukocytes, and frank extravasation of red cells from injured vessels.[122] A possible reason for the similarity in the histological appearance of chilling injury with ischemia-reperfusion injury may be due to the fact that both types of injury appear to be caused, at least in part, by reactive oxygen species and by disruption of the cytoskeleton.

The molecular changes induced by moderate, profound and ultraprofound hypothermia may also directly compromise endothelial cell integrity. For example, chilling of several types of epithelial cells has been shown to result in disassembly (depolymerisation) of the intracellular microtubules resulting in compromise of the polarized membrane expression and function of some transport proteins in these cells.[123] These functions are slow to return to normal (@20 hr) and are associated with prolonged dysfunction of allografts that have undergone cold preservation storage.[124],[125]

Deep cooling of the peripheral tissues may also result in immunosuppresion in chilled limbs and skin, and possibly impaired hematopoiesis due to localized moderate hypothermia in bone marrow in the cranium, sternum, vertebrae, and to a lesser extent, in the pelvis. [In children, the long bones are the principal repository of hematocytoblasts, and this marrow would also be disproportionately chilled during external cooling.] In contrast to the anti-inflammatory effect of MTH, systemic hypothermia to £28ºC, either accidental or induced has been shown to increase levels of pro-inflammatory cytokines.

Core cooling is thus the gold standard for the induction of systemic hypothermia (mild, moderate, profound or ultraprofound) and external (peripheral) cooling should be used only where there is no other alternative for achieving truly rapid cooling, or maintaining it for the required 24-48 hours following induction.

 Consideration of Invasive Core Cooling Methods

Figure 1-15 (right): Typical (idealized) cooling and re-warming curve achievable with maximum extracorporeal (cardiopulmonary bypass) cooling.

Extracorporeal cooling via cardiopulmonary bypass (CPB) [126] or high flow veno-venous heat exchange [127],[128]  allows for cooling rates of 0.8ºC/min to 1.0ºC/min (Figure 1-15) However, it cannot be applied rapidly enough given existing logistic and regulatory constraints. CPB requires complex hardware and highly skilled personnel who must maintain their clinical reflexes by practicing perfusion on a regular (preferably daily or at least weekly) basis. Even in centers of excellence, with a highly skilled, rapid-response CPB team at the ready (including a well-practiced surgeon or cardiologist), the soonest CPB can be initiated after cardiac arrest is typically 15-20 minutes.[129],[130]

The use of emergency CPB applied in this time frame is largely confined to patients undergoing cardiac catheterization and/or revascularization (i.e., angioplasty or stent placement) who arrest in the cardiac catheterization lab. When such patients experience cardiac arrest that is refractory to treatment with drugs and defibrillation – they are usually otherwise healthy – they have normally functioning lungs, normal fluid balance and fluid distribution (are neither dehydrated nor edematous from fluid overload), and have failure of a only a single organ –  the heart. Even under such ‘ideal’ conditions, CPR is often inadequate to maintain brain viability during the brief interval between cardiac arrest and the start of CPB.

While veno-venous extracorporeal cooling is less technically demanding, it still requires skill-intensive vascular access under field conditions, and reported rates of cooling are modest; ~0.012ºC/min [131],[87]. Both of these techniques may require anticoagulation and certainly require coagulation monitoring; which again is a barrier to in-field application.

Administration of large volumes (40mL/kg) of chilled intravenous fluid has also been used to reduce core temperature in the field in patients undergoing CPR. However this method is extremely slow (~0.058ºC/min) [132],[133],[134], and is sharply constrained by the maximum volume of fluid that can be administered.

 Figure 1-16 (left): Cryopatient undergoing ECMO and blood washout in the home. Even under ideal circumstances cardiopulmonary bypass takes in excess of 30 minutes to initiate.

 The effectiveness of intravascular fluid administration in achieving durable core cooling is also a function of body composition. Obesity, which has become epidemic in the developed world, is now rapidly becoming a global problem with 1.1 billion adults worldwide classified as overweight,and 312 million of them as obese by the World Health Organization (WHO). [135] Vascular volume does not increase proportionally to increase in body weight; in fact, the vascular volume to body weight ratio falls toward an asymptotic value of approximately 45 ml/kg in the obese human.[136] Thus, the use of cold intravenous infusions will necessarily be even less effective in the obese than might be expected since the maximum volume of fluid that may be safely given does not increase linearly with body weight. Fat is also an extraordinarily good insulator, and obesity often unfavorably alters body surface to volume area. Both of these factors combine to dramatically reduce the speed and overall effectiveness of external cooling.

The limits of efficiency of core cooling that can be achieved by irrigating the peritoneal or pleural spaces with cold liquid are in the range of 0.05 to 0.3ºC/min [137],[138] and both of these techniques are invasive, require considerable technical skill, and carry with them risks of serious iatrogenic consequences.

As noted earlier, the ideal way to address these needs would be extracorporeal membrane oxygenation (ECMO) and cooling. However, the practicality of rapidly bringing this demanding technology to bear in the EMS setting is negligible. Surgery (or time-consuming percutaneous vascular access; at or beyond the optimum therapeutic window of |15 min post ROSC) is required to access the circulation and this cannot be accomplished under field conditions. Experience with emergent initiation of ECMO in patients presenting for experimental cryopreservation (Figure 1-16) are probably representative of the time required to achieve extracorporeal support under field conditions for patients with SCA. Such patients typically experience a delay of 60 to 80 minutes (and not infrequently longer) after the initiation of closed chest CPR before the application of CPB, even under optimum circumstances.[139],[140],[141]

Thus, what is needed for patients who have suffered prolonged (³5 min) cardiac arrest, or in whom ROSC cannot be effectively established, is the ability to rapidly induce hypothermia via core cooling simply and noninvasively.  In addition to moderating the injury cascade initiated by ischemia-reperfusion, mild (33-35^oC) or moderate (28-32^oC) hypothermia can indirectly act to improve gas exchange and hemodynamics by bringing the patient’s cerebral and systemic metabolic demands closer to those that can be delivered by CPR.

Exsanguinating Trauma Resulting in Cardiac Arrest

Closely related in pathophysiology to prolonged normothermic ischemia secondary to SCD is cardiac arrest secondary to exsanguinating trauma. It is estimated that ~20,000 US civilians a year die as result of hemorrhage from abdominal and thoracic injuries [36], or from poly-trauma. In developing nations this problem is even more severe as a disproportionate amount of trauma occurs in rural settings remote from tertiary care facilities; and with no helicopter or other airlift infrastructure available to shorten this interval.

Similarly, approximately 20% of the battlefield casualties who fail to reach tertiary care facilities die from intractable hemorrhage on the battlefield or during transport.[142],[143],[144],[145],[146] The US military, under the auspices of DARPA, has been funding a multimillion project to achieve ~30 minutes of battlefield ‘suspended animation’ using chilled, drug containing crystalloid solutions, to solve this major cause of war-related mortality [37] and preliminary studies have been positive.[147],[128],[147]

Cardiac arrest secondary to exsanguinating trauma offers a unique opportunity for intervention in the pathophysiological cascade of cardiac arrest. Because blood loss and deterioration of the patient to the agonal state occur over a time course of minutes, to an hour or longer, it is possible to begin induction of hypothermia before cardiac arrest occurs; potentially allowing for the opportunity to prevent many of the foreseeable irreversible pathological events which occur during ischemia as a result of cardiac arrest.

Beyond inhibition or moderation of the injury cascade attendant to ischemia-reperfusion injury, deep (15-27^oC) profound (5-14^oC) or ultra-profound (5-0^oC) hypothermia offers the prospect of acting  as a bridge to definitive therapy in cases of uncontrolled hemorrhage in both the military and civilian settings, as well as in cases of cardiac arrest that are refractory to prompt defibrillation and that require prolonged CPR until CPB can be initiated and coronary revascularization or application of long-term circulatory support (i.e., left ventricular assist device (LVAD) or total artificial heart (TAH)) can be implemented. This was the vision of one of the fathers of CPR and the discoverer of MTH, Dr. Peter Safar. Safer envisioned a time in the near tomorrow when patients who were not salvageable in the field would be placed into what he termed a ‘metabolic lock-box’ and preserved until definitive therapy could be applied in hospital. His formal name for this paradigm was ‘emergency preservation and resuscitation’ (EPR). [148]

As is the case with truly effective implementation of MTH in cardiac arrest or neurotrauma the barrier to EPR is safe, rapid and easily implemented core cooling. One possible solution to this problem is to use the lungs as a heat exchanger.

The Lungs as Heat Exchangers

Possessing the surface area of a tennis court, and being obligated, as they are, to accept and thin film all of the cardiac output over much of that surface area, the lungs commend themselves as extraordinarily suited to serve as heat exchangers, as well as mass exchangers. The principal obstacle to utilizing the lungs for this purpose is the extreme poverty of gas as a heat exchange medium. While the specific heat of room air (gas) is 1.4 J g⁻¹ K⁻¹ c_p at 23ºC [149] compared to 4.18 J g⁻¹ K⁻¹ c_p for water (liquid) at 25ºC [150], this is not the rate-limiting factor for heat transfer using air or oxygen (0.918 J g⁻¹ K⁻¹ c_p at 25ºC). While water possesses an extraordinarily high specific heat, this is atypical and is exceeded only by liquid ammonia (4.70 J g⁻¹ K⁻¹ c_p at 25ºC) and hydrogen (gas) (14.30 J g⁻¹ K⁻¹ c_p at 25ºC). The limitation of gases in facilitating heat exchange is their low density compared to liquids. Dry air has a density of ρ_SATP = 1.168 kg/m³ compared to 997.05 kg m^-3 (25.0°C) for liquid water. This is roughly a 1,000 to 1 difference in density and is the primary reason why air, oxygen and gases in general transfer heat at rates too low to be useful for rapidly reducing body temperature.[151],[152]

Yet another constraint on the use of gases for heat exchange in the lungs of most vertebrates is their tidal, bi-directional scheme of ventilation which results in admixture of inspired and expired gases and sharply constrains the maximum flow rate of gas per unit of time (minute volume) ‘through’ (i.e., in and out of) the lungs. By contrast, the lungs of birds employ a unidirectional, flow-through approach to ventilation (circulatory lungs) which is more efficient at gas exchange and allows for higher peak minute volumes.[153]

The final constraint on using gases is the necessity of keeping the temperature of the gas (or other heat exchange medium) at or above 0ºC in order to avoid freezing damage to the tissues or provoking broncho-constriction which interferes with both gas and heat exchange.[154]   This sharply limits the ΔT, which will also decay steadily as the temperature of the patient decreases towards 0ºC.

The only solution to the heat exchange limitation imposed by gases is to replace the tidal gaseous breath, or some fraction of it, with an appropriate liquid. This is a formidable challenge because the respiratory systems of air breathing animals have evolved complex and delicate systems to facilitate gas exchange; and these systems are incompatible with the presence of bulk liquid in the airways.

Figure 1-17: Anatomy of the acinus. Anatomically, alveoli are not independent structures but rather have common walls and are organized around and attached to the terminal bronchiole. This unit of structural organization is the acinus. The acinus is reinforced with both comparatively rigid (collagen) and elastic (elastin) fibers and the arrangement of the alveoli in a honeycomb-like manner within the acinar unit provides added structural stability and give the acinus emergent properties similar to those of foam. The structural features of the acinus, acting in concert with surfactant (which lowers alveolar surface tension providing additional architectural stability), is termed interdependence. Interdependent structural stabilization of the alveoli maintains alveolar volume and surface area fairly constant during ventilation under normal conditions. In lung injury, even the collapse of a single alveolus in an acinus causes shear stress – not only in the walls of the collapsed alveolus, but also in the walls of the adjacent alveoli (see Figure 4-11). The shear stress that develops under such conditions may exceed 140 cmH₂0, and the stabilizing interdependence of the acinus is lost, with the alveoli increasingly behaving as independent units that change volume dramatically with each tidal cycle.

The smallest gas exchange units of the mammalian respiratory system, the alveoli, are a mere ~125 µ in diameter and, absent molecular reinforcement in the form of surfactant, are inherently unstable and collapse. Surfactant has evolved to function as a semi-aqueous thin film lining the alveolus and other small acinar structures (such as the terminal bronchioles and the alveolar ducts) (Figure 1-17) and it is dissolved, damaged, or inactivated when exposed to water, and to other molecules typically found in body fluids such as proteins and lipid particles. Surfactant has both hydrophilic and lipophylic components which impose the seemingly impossible constraint that any liquid introduced into the lungs must be both hydro- and lipophobic, as well as virtually completely chemically inert.

 A Brief Précis of the History and Development of Liquid assisted Pulmonary Cooling (LAPC) for Induction of Hypothermia during CPR

So far, only one class of compounds, the perfluorocarbons (PFCs), meets these requirements. With the demonstration in 1966 by Clark and Gollin that mice could survive extended periods of ‘liquid breathing’ with the PFC FX-80 (FC-75; perflurotetrahydrofuran) the feasibility of using liquids as gas exchange media began to be explored. Despite the (in retrospect) painfully obvious implication that PFCs could be used in the lungs for heat exchange as well as gas exchange, the idea did not occur to the author until sometime between the winter of 1993 and the spring of 1994. This happened while rereading Guyton’s Textbook of Medical Physiology during the course of several transcontinental flights across the United States (US). The author had been aware of Leland Clark’s ‘fluorocarbon breathing mice’ since shortly after these experiments were carried out in 1966; they were the subject of extensive media coverage, including dramatic footage shown on science-oriented television programs of the time (such CBS’ ‘The 21^st Century’). Yet, it was not until re-reading Guyton’s chapter on respiratory physiology that the idea of using the lungs as a liquid-based heat exchanger became clear. Having spent literally 22-years obsessed with the problem of how to rapidly, non-invasively, and relatively simply, induce hypothermia in humans under field conditions; in hindsight it seems utterly obvious to use tidal liquid ventilation (TLV) employing a refrigerated PFC as a way of achieving rapid, homogenous, systemic cooling.

Figure 1-18: Mice can briefly tolerate breathing the PFC FC-77 but soon succumb to exhaustion from the work of breathing and hypothermia from the ambient temperature liquid.

An investigation of the literature disclosed the feasibility of TLV based on the work of Shaffer and his colleagues. [155]  Unfortunately, acquiring a PFC suitable for liquid assisted ventilation was very problematic. The only commercially available PFCs at that time were the 3M Fluorinet™ fluids and Rimar-101 (the latter manufactured by the Miteni Corporation in Milan, Italy). Neither company distributed their products through the conventional chemical houses[6]. The reason for this was that both 3M and Miteni specifically embargoed the sale of their PFCs for any research application involving biological systems – and in particular any application that involved liquid assisted ventilation (LAV) or O₂ carrying blood substitutes (oxygen therapeutics).

By January of 1995, the author and Respiratory Therapist Michael Fletcher had constructed the first TLV ventilator for research into inducing MTH during CPR, with the idea very much in mind that this technology had tremendous potential for broad medical application (post resuscitation syndrome, stroke, subarachnoid hemorrhage (SAH) and possibly traumatic brain and ischemic (and traumatic) spinal cord injury (TBI and SCI).

Figure 1-19: The first TLV apparatus for cardiopulmonary cerebral resuscitation-related research (left) used blood as the gas exchange medium. The device used a bubble oxygenator and heat exchanger and relied on gravity to provide ‘exhalation’ drainage from the lungs. The inspiratory roller pump and the timing and pressure controller functioned well, but blood proved to be too injurious to the lungs t serve as a gas exchange medium. These experiments made the need for an expiratory pump obvious. Photo: circa February, 1995.

Because of the unavailability of PFCs, and because of the prior liquid breathing research by Johannes Klystra  [156],[157],[158] which demonstrated that mammals could survive after breathing a balanced salt solution under hyperbaric conditions (where gas solubility in water is dramatically improved), we decided to use bovine blood as the gas exchange medium in a non-survival canine model of TLV. Blood, unlike saline, has the proven capacity to deliver and exchange an adequate amount of O₂ and CO₂under normobaric conditions. A review the literature (then and now) indicated that this had not been attempted, or apparently even proposed. These two initial experiments using bovine blood in an acute canine model of TLV demonstrated that blood was not suitable for use, and that TLV was a demanding technique that required an exhalation pump in addition to an inspiration pump. The prototype device taught us a great deal, and in fact, the control unit from that first TLV ventilator was used until 2000 to conduct much of the early 21^st Century Medicine, Inc., and Critical Care Research, Inc., work on liquid assisted pulmonary cooling (LAPC).

 Figure 1-16: Prototype tidal liquid ventilator using the ‘sweep flow’ (continuous liquid delivery/removal to the carina) technique circa 1995, and the active compression-decompression, high impulse Thumper™ mechanical CPR device (Figure 1-17) used to facilitate tidal movement of PFC between large and small airways (below). [Photo (left) courtesy of Charles Platt, 1995.]

By March of 1995 several PFCs had become available in sufficient quantity for use in a canine model of TLV. A continuous flow approach was being used to deliver chilled PFC to the carina, with tidal movement of liquid in and out of the small airways being achieved by active compression and decompression of the chest wall using the suction cup of an Ambu CardioPump™ on a specially constructed, pneumatically driven Thumper™ CPR device. This technique was christened ‘sweep-flow liquid ventilation.’ [159]

Figure 1-17: Custom fabricated Michigan Instruments, Inc., Thumper™ active compression decompression cardiopulmonary resuscitator

Dynamically sensing and controlling the amount of PFC in the animals lungs proved problematic with this system, and despite a very high predicted minute alveolar ventilation rate (with 80 to 100 liquid ‘breaths’ per minute from chest compression-decompressions), and minimal anatomical dead-space (the tip of the liquid delivery catheter was at the level of the carina), hypercarbia developed after the first 10 to 15 minutes of ventilation.

While survival of animals following TLV using PFCs was 100%, with a low incidence of adverse effects (providing TLV was discontinued prior to lethal levels of hypoxemia), the logistics of applying TLV to a field-setting under emergency conditions was deemed insurmountable. This was particularly true in the mid-1990s due to limitations on both portable computing power and the lack of adequately miniaturized and lightweight sensing and servo-control technology.

During the fall of 1996, the author learned of the development and initial clinical application of a new technique of liquid assisted ventilation (LAV) called partial liquid ventilation (PLV) which was being used to treat respiratory distress syndrome in (primarily) premature infants. This work was rapidly extended to include adults, and in the Spring of 1996 the author attended a three-day conference and workshop: the ‘Basics of Liquid Ventilation Management of Severe ARDS, ECMO, Liquid Breathing and PCIRV Conference’ in Ann Arbor, MI.[7]  In contrast to TLV, PLV was effected by simply adding PFC to the patient’s respiratory system until it was filled to functional residual capacity (FRC, ~30 ml/kg) while continuing conventional mechanical ventilation with gas; PLV was the essence of simplicity PFC was added to the endotracheal tube with a syringe until the desired volume was given or a meniscus of PFC appeared at the desired level in the ET tube. What is more, as the author had already observed while rounding at the University of Michigan Medical Center’s ICU, due to their unique physical properties, PFCs had the ability to rapidly reverse the hypoxemia and hypercarbia that accompany respiratory distress and its attendant pulmonary edema. Thus, PLV offered the added promise of restoring adequate gas exchange in fulminating pulmonary edema, as well as serving as a safe and effective heat exchange medium.

Early in 1997, due to the technical and logistic difficulties imposed by TLV, the author began to investigate the use of multiple lavages of the lungs in dogs using PFC chilled to ~4ºC. In this technique the animals’ lungs were filled with ~4 ºC PFC to vital capacity (VC, 60-70 ml/kg) following which the PFC was suctioned out and this process was repeated with additional loads of ~4ºC PFC until the desired core temperature was reached.[159]  While technically straightforward and very simple in application, this approach required large amounts of chilled PFC and had the obvious potential to interfere with gas exchange under conditions of hyper-metabolism, oxygen debt, or hypercarbia. Additionally, the time required for loading and suctioning out 60-70 ml/kg of PFC did not appear to be consonant with maximizing heat exchange, given that much of the PFC load would necessarily remain in the large and intermediate sized airways, where heat exchange is very slow compared to in the alveoli.

At this point it became apparent that a practical, field-applicable method of inducing hypothermia by PFC lung lavage would involve a hybrid approach, wherein both liquid and gas tidal ventilation would have to proceed simultaneously, or on a rapidly alternating basis. This work was largely set aside at that time (1997) due to a research commitment to develop a multi-modal method of dealing with ischemia-reperfusion injury attendant to prolonged cardiac arrest. This model involved the use of multiple drugs, as well as the rapid (~15 minutes) induction of mild post-insult hypothermia using the Safar, et al., [160] model canine model of cardiac arrest and resuscitation, which employed CPB for the induction of hypothermia.

 Figure 1-18 (above): The alveolar capillary unit is the fundamental repeating structural unit of the lung. Above is a schematic rendering of an alveolus as it would appear magnified ~3,000 times. The polyhedral alveolar air spaces (AAS) are separated from each other by the alveolar septa which are ~15µ thick. The Type I epithelial cells overlay the capillaries (C) and constitute the air side of the airway-blood barrier (ABB). The capillaries are the yellow spaces with the red blood cells (RBCs) inside them. The Type II epithelial cells are heavily vesiculated and secrete surfactant which maintains the structural integrity of the alveoli and prevents them from collapsing during exhalation. The total thickness of the ABB is only ~0.5µ which means that heat exchange between blood and PFC at the alveolar wall is extremely rapid – the rate limiting factor on heat exchange is, in fact, the rate at which PFC can be moved into and out of the alveolus.

With the success of this post-resuscitation hypothermia and multi-drug research cerebral resuscitation model in 1998, it became important to demonstrate that very rapid, in-field induction of hypothermia was technically and logistically possible. A decision was made to return to the LAPC research in order to validate the feasibility of this approach for immediate, simple, post-arrest induction of MTH (~34 ºC).The first results of the work were presented in1999 at the 28th Educational and Scientific Symposium of the Society of Critical care Medicine in San Francisco, CA. [161]  At the time, it was anticipated that completion of the research would take approximately 6 months.

The team that began this work consisted of the author, Steven B. Harris, M.D., Sandra Russell, BS, Michael Fletcher, R.T., Joan O’Farrell and Brian Wowk, Ph.D.  This project proved more formidable than initially anticipated and ultimately consumed over 2 years of full-time effort in order to develop the core elements of a workable technique, and achieve a basic understanding of the physiology and pathophysiology of fractional tidal liquid assisted pulmonary cooling (LAPC); the technique that was ultimately developed. This team worked together seamlessly to establish the effective volumes of PFC to be loaded and unloaded (i.e., fractional liquid tidal volumes), optimum dwell times for PFC, tolerable peak and mean airway pressures for ventilation during liquid cycling, workable gas ventilation strategies, and the basic mechanics of both heat and gas exchange using this technique.

 What LAPC Can Potentially Deliver

Properly implemented, LAPC would offer the following advantages:

• Ease of application, requiring far less highly skilled personnel than are needed for CPB,
• Technically less demanding, requiring fewer total personnel than CPB,
• Effective at achieving a rate of heat exchange in the brain comparable to or better than that achievable with CPB,
• Effective at achieving good gas exchange even with patients with severe lung disease/injury (pulmonary edema, acute lung injury (ALI) and acute respiratory distress syndromes (ARDS),
• Relatively inexpensive to use,
• Improves oxygenation and ventilation,
• Reduces interfacial surface tensions (i.e., between liquid and gases and liquids, gases and alveolar membranes),
• Reduces ventilation pressures,
• Recruits alveoli,
• Redistributes pulmonary blood flow and mitigates V/Q mismatch,
• Mitigates barotraumas and volutrauma,
• Lavages airway and alveolar debris,
• Reduces inflammation and lung injury.
• Allow CPR  to serve as a viable bridge to CPB and eventual definitive therapy to restore spontaneous circulation or supplement or replace it as necessary (i.e., LVAD or TAH)

The physiological considerations underlying LAPC are these:

1)    All of the blood that flows out from the heart to the body flows through the lungs first, where it is oxygenated and carbon dioxide (CO₂) is removed. In order to facilitate gas exchange, the architecture of the lungs has evolved to provide an enormous surface area combined with an ultra-microscopic barrier between blood and air. The alveoli, the gas exchange units of the lungs, are polyhedral compartments ~250 µ in diameter which branch out from the terminal bronchioles (which are about twice the diameter of an alveolus; 250µ) (Figure 1-18, above). Each adult human lung contains ~300 x 10⁶ alveoli with an aggregate surface area of 140 m² [162] The alveolar epithelium is comprised of two types of cells; Type I epithelial cells that constitute the alveolar side of the air-blood barrier (ABB) and Type II epithelial cells which are intercalated in the epithelial lining and secrete surfactant. The Type I epithelial cells are so thinned out where they cover the capillary endothelial cells (which are even thinner) that the composite thickness of both cell layers and the basal lamina is ~0.5µ. These unique anatomical features of the lungs, combined with the physiological fact that they process all of the cardiac output, makes them ideal as a heat exchanger.

2) As previously noted, air, oxygen and other gases make extremely poor heat exchange media since they are roughly a thousand times less dense than water, they will remove heat at only roughly one thousandth the rate.

3) Mammals can survive briefly while spontaneously breathing perfluorochemicals (PFCs) and the usual cause of death when small mammals, such as mice, are left immersed in PFCs is hypothermia due to loss of heat from the large surface area of the lungs to the room temperature PFC in which they are immersed.[163] The severely injured lungs of human patients with pulmonary edema and/or acute lung injury (ALI) or acute respiratory distress syndrome (ARDS) not only tolerate extended (days to weeks) partial filling with PFCs of the appropriate vapor pressure, density, and viscosity, their ability to carry out gas exchange is improved, and there is evidence that long-term damage from ARDS may be reduced.[164],[165],[166],[167]

This technique of non-invasive in-field induction of hypothermia using the lungs is called liquid assisted pulmonary heat exchange (LAPHE), or liquid assisted pulmonary cooling (LAPC) when used for inducing hypothermia.

FOOTNOTES