Lung surfactant (LS) is a unique mixture of lipids and proteins that lines the alveoli and lowers the surface tension in the lungs, thereby insuring a negligible work of breathing. The surface tension control imposed by LS is compromised during Acute Respiratory Distress Syndrome (ARDS), which afflicts 150,000 with a 40% mortality rate in the US. In a related syndrome, the absence of LS due to prematurity leads to neonatal Respiratory Distress Syndrome (NRDS), which is routinely treated with animal-derived replacement LS. However, replacement LS loses its ability to reduce surface tension and is said to be “inactivated” when used to treat ARDS. A common factor among ARDS patients is elevated levels of serum proteins in the alveolar fluid. In vitro, LS mixed with serum proteins shows an ARDS-like depression of LS activity; serum inactivation is one reason why treatment of ARDS with replacement LS is unsuccessful. Several hydrophilic polymers have recently been shown to enhance the ability of replacement LS to resist serum inactivation both in vitro and in vivo. The use of polymer additives to enhance LS performance might offer a scalable and cost effective treatment for ARDS. This work aims to provide the physical and chemical principles explaining both serum inactivation and inactivation reversal by hydrophilic polymers. This will help medical professionals understand the origins of ARDS and enhance efforts to identify biocompatible polymers to formulate replacement LS resistant to inactivation. From the materials side, this work will contribute to a fundamental understanding of competitive adsorption at liquid-vapor and solid-liquid interfaces. We have developed an in vitro model that approximates LS behavior in the alveoli during respiration using the compression/expansion cycles of a Langmuir trough, with albumin as the prototypical surface active serum protein. Utilizing fluorescence and confocal microscopy, we have simultaneously visualized the competitive adsorption of LS and albumin to the air-water interface. Our results show healthy LS is inactivated by the addition of albumin; images of the interface show that an albumin film adsorbed to the interface creates a barrier to LS adsorption. This competitive adsorption of albumin to the alveolar air-liquid interface can be modeled as an energy barrier to LS adsorption, which can be quantitatively analyzed using a variation of the classical Smolukowski description of colloidal stability.

LS adsorption can be enhanced by the addition of polyethylene glycol (PEG) via a depletion attraction between the LS aggregates and the interface. The depletion attraction effectively pushes LS aggregates toward the interface due to increased polymer entropy induced by the elimination of the “excluded volumes” of the LS aggregates and the interface. Addition of ~1% wt. 10 kDa PEG is sufficient to reverse inactivation and restore LS to the interface. LS adsorption increases exponentially with PEG concentration as predicted by the Asakura and Oosawa model of depletion attraction. The PEG response is molecular weight dependent, with PEG 6-35 kDa yielding optimal inactivation reversal while the depletion attraction lacks sufficient range for smaller polymers. At the concentrations necessary for reversing inactivation, PEG greater than 35 kDa reaches overlap concentration resulting in decreasing performance and a breakdown of the simple Asakura and Oosawa model of depletion attraction. Fluorescence images detail the transition from an albumin covered interface to a LS covered interface during successful inhibition reversal. After LS breaks through the albumin film, the interface shows a coexistence of extended albumin and LS domains and finally the albumin domains forced from the interface at sufficiently high surface pressure.

Freeze-fracture transmission electron microscopy images show that PEG and albumin do not adsorb to the surfactant aggregates, nor do these macromolecules penetrate the interior water compartments of the surfactant aggregates. This results in an osmotic pressure difference that dehydrates the bilayer aggregates, causing a decrease in the bilayer spacing as shown by small angle x-ray scattering and an increase in the ordering of the bilayers as shown by freeze-fracture electron microscopy. Small angle x-ray diffraction shows that the relationship between the bilayer spacing and the imposed osmotic pressure for replacement LS is a screened electrostatic interaction with a Debye length consistent with the ionic strength of the solution.

Utilizing in situ grazing incidence x-ray diffraction (GIXD) and x-ray reflectivity (XR), we have examined the surface ordering of clinical replacement LS at the air-liquid interface and the effect of albumin and PEG. XR measurements confirm that albumin imposes a steric barrier to LS adsorption, inhibiting the LS characteristic GIXD peaks. In the LS free system, scattering experiments show no evidence of PEG surface ordering while LS on a PEG subphase shows only a subtle lateral condensation of the LS. However, the addition of PEG to albumin inhibited LS restores the LS characteristic XR and GIXD peaks and progressive cycling shows the LS replacing the albumin on the interface. These scattering results are consistent with fluorescence images of the interface which show a coexistence of ~1000μm albumin and LS regions until all albumin regions are eventually expelled from the interface.

In addition to the PEG generated osmotic pressure, the bilayer perturbing properties of chitosan, a biocompatible cationic polysaccharide derived from deacetylated chitin can also enhance LS adsorption to the air-liquid interface. Recent results demonstrate that chitosan reverses albumin induced LS inhibition at significantly lower concentrations (0.01 mg/mL) than PEG (10 mg/mL) suggesting another promising therapy for ARDS. LS is ~70 wt % dipalmitoylphosphatidylcholine (DPPC) and freeze fracture transmission electron microscopy images of DPPC vesicles treated with chitosan show distinct morphological changes. Untreated DPPC vesicles are ~50 nm in diameter and uniformly distributed through the solution while a low concentration of chitosan (0.005 mg/mL) causes aggregation of vesicles. At higher chitosan concentrations (0.5 mg/mL), images show numerous ~50 nm vesicles entrapped inside a larger vesicle, similar to “vesosomes” used for drug delivery. These morphological changes induced by chitosan suggest that it enhances LS adsorption by facilitating the bilayer to monolayer transition at the interface. These results demonstrate the utility of applying classic materials science principles to competitive adsorption problems of biomedical interest. More broadly, this materials science approach will benefit a broad class of technological problems involving competitive adsorption to interfaces.