Polymer lung surfactants

Abstract

To provide a method of treating a pulmonary disorder of a mammal caused by deficiency of functional lung surfactant.SOLUTION: The present invention provides a method comprising administering to a mammal a therapeutically effective amount of a polymer lung surfactant composition, wherein the composition comprises a synthetic biocompatible amphiphilic homopolymer or copolymer, and monomers of the homopolymer or the copolymer are selected from the group consisting of: ethylene glycol (EG), ethylene oxide (EO), vinyl alcohol (VA), alkyl oxazoline (AO), D,L-lactic acid or D,L-lactide (LA), glycolic acid or glycolide (GA), ε-caprolactone (CL), styrene (PS), alkyl methacrylate (AMA), and alkyl acrylate (AA).SELECTED DRAWING: None

Description

Technical Field

[0001] Cross - Reference to Related Applications This U.S. patent application is related to and claims the benefit of priority of U.S. Provisional Patent Application No. 62 / 374,325, filed on August 12, 2016, and the entire content of the U.S. Provisional Patent Application is incorporated herein by reference. Statement Regarding Governmental Budgets

[0002] The present invention was made with government support under Grant No. CBET - 1264336 awarded by the National Science Foundation. The government has certain rights in the invention.

[0003] Broadly speaking, the present disclosure relates to polymeric lung surfactants, and more particularly to polymeric lung surfactant materials that meet all the performance requirements as surfactants and are easier to handle than current respiratory distress syndrome therapeutics.

Background Art

[0004] This section introduces aspects that may help to facilitate a better understanding of the present disclosure. Therefore, these explanations should be read from this perspective and should not be understood as an admission as to what the prior art is.

[0005] Newborns born before the completion of a 40 - week gestation period are considered "preterm infants" (if born before 37 weeks of gestation) or "premature infants" (if born before 34 weeks). One of the major health risks associated with being preterm / premature is that lung development may not be sufficient. This is a cause of high neonatal mortality. Newborns born before 37 weeks of gestation are born without alveolar structures and have a low production of lung surfactant. Therefore, preterm and premature infants struggle to breathe and, without appropriate treatment, will die within a few days. This respiratory insufficiency is called respiratory distress syndrome (RDS). It was formerly also known by the incorrect name of hyaline membrane disease because the cause of this disease was previously misunderstood to be a virus.

[0006] When RDS was misnamed hyaline membrane disease, it was the leading cause of death among infants in the United States at that time, with a higher mortality rate than pneumonia and influenza. However, today, thanks to three established treatment methods by skilled physicians, the mortality rate due to RDS has significantly decreased. The three treatments are carried out step by step, and if the previously implemented treatment is effective, the next treatment is not carried out. The first treatment for RDS is a preventive treatment in which the mother is administered steroids 24 hours before delivery to increase the production of the infant's own lung surfactant. Clinical data on steroid treatment with betamethasone has shown that the incidence of RDS has effectively decreased from 25.8% to 9%. The second treatment, surfactant replacement therapy (SRT), includes intratracheal injection of lung surfactant extracted from animals into the lungs of infants immediately after birth. The development of effective SRT has been the main driving force in reducing RDS-related mortality, and due to its high effectiveness, it is also included in the World Health Organization's list of essential medicines. The third treatment method is mechanical ventilation, which applies nasal continuous / non-continuous positive airway pressure therapy to infants to increase the oxygen concentration in the lungs. Treatment by mechanical ventilation is the oldest treatment method for treating RDS. In its first clinical trial, it was shown that the mortality rate decreased from 80% to 20%. However, oxygen toxicity and mechanical damage to the lungs still remain harmful effects. Among the three treatment methods, SRT is the most reliable treatment method that directly addresses the root cause of RDS, and no harmful effects have been reported so far. The improvement in the treatment of RDS is considered to be due to the progress of SRT technology.

Summary of the Invention

Problems to be Solved by the Invention

[0007] Although successful in reducing the RDS-related mortality rate among premature and low-birth-weight infants in the country, nevertheless, due to the high treatment cost and complex treatment procedures, especially in developing countries on a global scale, RDS still remains one of the main causes of neonatal death. Very Although effective, the cost of SRT is extremely high, and in some countries, the cost of SRT treatment drugs alone exceeds the GNP per capita. The economic imbalance affecting the use of SRT is evident from Figure 1, which shows that SRT is not fully utilized in the countries of Central Asia and Africa. It is possible to develop a lower-cost RDS treatment drug that does not require skilled doctors or advanced neonatal intensive care units (NICUs) and can be used for treatment with simpler procedures, thus solving the above problems and reducing the main cause of neonatal death worldwide. Note that even in the United States, in some rural areas, there is a shortage of medical facilities such as skilled doctors and necessary NICU facilities, so preterm and premature infants are at risk of RDS-related death. It should be noted that in areas where SRT is not practical, treatment mainly depends on mechanical ventilation. Therefore, there is still an unmet need for better SRT technology. Although effective, the cost of SRT is extremely high, and in some countries, the cost of SRT treatment drugs alone exceeds the GNP per capita. The economic imbalance affecting the use of SRT is evident from Figure 1, which shows that SRT is not fully utilized in the countries of Central Asia and Africa. It is possible to develop a lower-cost RDS treatment drug that does not require skilled doctors or advanced neonatal intensive care units (NICUs) and can be used for treatment with simpler procedures, thus solving the above problems and reducing the main cause of neonatal death worldwide. Note that even in the United States, in some rural areas, there is a shortage of medical facilities such as skilled doctors and necessary NICU facilities, so preterm and premature infants are at risk of RDS-related death. It should be noted that in areas where SRT is not practical, treatment mainly depends on mechanical ventilation. Therefore, there is still an unmet need for better SRT technology. Although effective, the cost of SRT is extremely high, and in some countries, the cost of SRT treatment drugs alone exceeds the GNP per capita. The economic imbalance affecting the use of SRT is evident from Figure 1, which shows that SRT is not fully utilized in the countries of Central Asia and Africa. It is possible to develop a lower-cost RDS treatment drug that does not require skilled doctors or advanced neonatal intensive care units (NICUs) and can be used for treatment with simpler procedures, thus solving the above problems and reducing the main cause of neonatal death worldwide. Note that even in the United States, in some rural areas, there is a shortage of medical facilities such as skilled doctors and necessary NICU facilities, so preterm and premature infants are at risk of RDS-related death. It should be noted that in areas where SRT is not practical, treatment mainly depends on mechanical ventilation. Therefore, there is still an unmet need for better SRT technology. Although effective, the cost of SRT is extremely high, and in some countries, the cost of SRT treatment drugs alone exceeds the GNP per capita. The economic imbalance affecting the use of SRT is evident from Figure 1, which shows that SRT is not fully utilized in the countries of Central Asia and Africa. It is possible to develop a lower-cost RDS treatment drug that does not require skilled doctors or advanced neonatal intensive care units (NICUs) and can be used for treatment with simpler procedures, thus solving the above problems and reducing the main cause of neonatal death worldwide. Note that even in the United States, in some rural areas, there is a shortage of medical facilities such as skilled doctors and necessary NICU facilities, so preterm and premature infants are at risk of RDS-related death. It should be noted that in areas where SRT is not practical, treatment mainly depends on mechanical ventilation. Therefore, there is still an unmet need for better SRT technology. Although effective, the cost of SRT is extremely high, and in some countries, the cost of SRT treatment drugs alone exceeds the GNP per capita. The economic imbalance affecting the use of SRT is evident from Figure 1, which shows that SRT is not fully utilized in the countries of Central Asia and Africa. It is possible to develop a lower-cost RDS treatment drug that does not require skilled doctors or advanced neonatal intensive care units (NICUs) and can be used for treatment with simpler procedures, thus solving the above problems and reducing the main cause of neonatal death worldwide. Note that even in the United States, in some rural areas, there is a shortage of medical facilities such as skilled doctors and necessary NICU facilities, so preterm and premature infants are at risk of RDS-related death. It should be noted that in areas where SRT is not practical, treatment mainly depends on mechanical ventilation. Therefore, there is still an unmet need for better SRT technology. Although effective, the cost of SRT is extremely high, and in some countries, the cost of SRT treatment drugs alone exceeds the GNP per capita. The economic imbalance affecting the use of SRT is evident from Figure 1, which shows that SRT is not fully utilized in the countries of Central Asia and Africa. It is possible to develop a lower-cost RDS treatment drug that does not require skilled doctors or advanced neonatal intensive care units (NICUs) and can be used for treatment with simpler procedures, thus solving the above problems and reducing the main cause of neonatal death worldwide. Note that even in the United States, in some rural areas, there is a shortage of medical facilities such as skilled doctors and necessary NICU facilities, so preterm and premature infants are at risk of RDS-related death. It should be noted that in areas where SRT is not practical, treatment mainly depends on mechanical ventilation. Therefore, there is still an unmet need for better SRT technology. Although effective, the cost of SRT is extremely high, and in some countries, the cost of SRT treatment drugs alone exceeds the GNP per capita. The economic imbalance affecting the use of SRT is evident from Figure 1, which shows that SRT is not fully utilized in the countries of Central Asia and Africa. It is possible to develop a lower-cost RDS treatment drug that does not require skilled doctors or advanced neonatal intensive care units (NICUs) and can be used for treatment with simpler procedures, thus solving the above problems and reducing the main cause of neonatal death worldwide. Note that even in the United States, in some rural areas, there is a shortage of medical facilities such as skilled doctors and necessary NICU facilities, so preterm and premature infants are at risk of RDS-related death. It should be noted that in areas where SRT is not practical, treatment mainly depends on mechanical ventilation. Therefore, there is still an unmet need for better SRT technology. Although effective, the cost of SRT is extremely high, and in some countries, the cost of SRT treatment drugs alone exceeds the GNP per capita. The economic imbalance affecting the use of SRT is evident from Figure 1, which shows that SRT is not fully utilized in the countries of Central Asia and Africa. It is possible to develop a lower-cost RDS treatment drug that does not require skilled doctors or advanced neonatal intensive care units (NICUs) and can be used for treatment with simpler procedures, thus solving the above problems and reducing the main cause of neonatal death worldwide. Note that even in the United States, in some rural areas, there is a shortage of medical facilities such as skilled doctors and necessary NICU facilities, so preterm and premature infants are at risk of RDS-related death. It should be noted that in areas where SRT is not practical, treatment mainly depends on mechanical ventilation. Therefore, there is still an unmet need for better SRT technology. Although effective, the cost of SRT is extremely high, and in some countries, the cost of SRT treatment drugs alone exceeds the GNP per capita. The economic imbalance affecting the use of SRT is evident from Figure 1, which shows that SRT is not fully utilized in the countries of Central Asia and Africa. It is possible to develop a lower-cost RDS treatment drug that does not require skilled doctors or advanced neonatal intensive care units (NICUs) and can be used for treatment with simpler procedures, thus solving the above problems and reducing the main cause of neonatal death worldwide. Note that even in the United States, in some rural areas, there is a shortage of medical facilities such as skilled doctors and necessary NICU facilities, so preterm and premature infants are at risk of RDS-related death. It should be noted that in areas where SRT is not practical, treatment mainly depends on mechanical ventilation. Therefore, there is still an unmet need for better SRT technology. Although effective, the cost of SRT is extremely high, and in some countries, the cost of SRT treatment drugs alone exceeds the GNP per capita. The economic imbalance affecting the use of SRT is evident from Figure 1, which shows that SRT is not fully utilized in the countries of Central Asia and Africa. It is possible to develop a lower-cost RDS treatment drug that does not require skilled doctors or advanced neonatal intensive care units (NICUs) and can be used for treatment with simpler procedures, thus solving the above problems and reducing the main cause of neonatal death worldwide. Note that even in the United States, in some rural areas, there is a shortage of medical facilities such as skilled doctors and necessary NICU facilities, so preterm and premature infants are at risk of RDS-related death. It should be noted that in areas where SRT is not practical, treatment mainly depends on mechanical ventilation. Therefore, there is still an unmet need for better SRT technology. Although effective, the cost of SRT is extremely high, and in some countries, the cost of SRT treatment drugs alone exceeds the GNP per capita. The economic imbalance affecting the use of SRT is evident from Figure 1, which shows that SRT is not fully utilized in the countries of Central Asia and Africa. It is possible to develop a lower-cost RDS treatment drug that does not require skilled doctors or advanced neonatal intensive care units (NICUs) and can be used for treatment with simpler procedures, thus solving the above problems and reducing the main cause of neonatal death worldwide. Note that even in the United States, in some rural areas, there is a shortage of medical facilities such as skilled doctors and necessary NICU facilities, so preterm and premature infants are at risk of RDS-related death. It should be noted that in areas where SRT is not practical, treatment mainly depends on mechanical ventilation. Therefore, there is still an unmet need for better SRT technology.

Brief Description of Drawings

[0008] (Figure 1) Countries where surfactant replacement therapy (SRT) is currently being implemented are shown in green. Countries where patients can access at least one commercially available pulmonary surfactant treatment drug are defined as "SRT implemented". The data was collected using Medtrack on March 24, 2015. Countries where surfactant replacement therapy (SRT) is currently being implemented are shown in green. Countries where patients can access at least one commercially available pulmonary surfactant treatment drug are defined as "SRT implemented". The data was collected using Medtrack on March 24, 2015. Countries where surfactant replacement therapy (SRT) is currently being implemented are shown in green. Countries where patients can access at least one commercially available pulmonary surfactant treatment drug are defined as "SRT implemented". The data was collected using Medtrack on March 24, 2015. Countries where surfactant replacement therapy (SRT) is currently being implemented are shown in green. Countries where patients can access at least one commercially available pulmonary surfactant treatment drug are defined as "SRT implemented". The data was collected using Medtrack on March 24, 2015.

[0009] (Figure 2) The Young-Laplace pressure of spherical alveoli is calculated as a function of surface tension and radius. The Young-Laplace pressure of spherical alveoli is calculated as a function of surface tension and radius.

[0010] (Figure 3) A commercially available pulmonary surfactant obtained by repeating the compression-expansion cycle. Surface tension - relative area isotherm of Survanta (Abb Vie). The subphase was a solution containing 150 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3 (pH 7.0 - 7.4, 25 °C) was used. The monolayer was compressed / expanded at a rate of 50 mm / min , and the compression - expansion cycle was set to 7.18 minutes per cycle. The data shown in the figure are the last 10 compression - expansion cycles out of a total of 50 consecutive cycles performed after spreading 4 mg of Survanta .

[0011] (Figure 4) Chemical structure of the PEG - based block copolymer tested for use as a lung surfactant . (Table 1) Table 1. Molecular characteristics of the investigated polymeric lung surfactant candidate materials 1 Lactide: glycolide = molar ratio 70:55 2 Lactide: glycolide: caprolactone = molar ratio 30 :29:41 JPEG2025102844000001.jpg45163

[0012] (Figure 5) Constant - pressure compression surface tension - area isotherms at 25 °C of monolayers of PLGA - PEG and PLGACL - PEG spread on the surface of Milli - Q pure water (resistivity 18 MΩ·cm) with (a) chloroform or (b) water. The surface tension during compression at a rate of 3 mm / min was measured. The hydrodynamic diameters of the PLGA - PEG and PLGACL - PEG micelles were 144.1 and 53.0 nm, respectively (measured by DLS).

[0013] ​(Figure 6) (a) Isotherms of the surface tension - area at constant compression of the PtBMA-PEG monolayer developed with chloroform or water at 25 °C. (b) Isotherms of the surface tension - area at constant compression of the PtBMA-PEG monolayer developed with water at three temperatures of 10 °C, 25 °C, and 40 °C. In all measurements, Milli-Q pure water (resistivity 18 MΩ·cm) was used as the subphase, and the compression rate of the monolayer was 3 mm / min. The hydrodynamic diameter of the PtBMA-PEG micelles was 26.3 ± 3.9 nm (measured by DLS).

[0014] (Figure 7) Surface tension - isotherms at 25 °C of the PS-PEG monolayer developed with chloroform or water on Milli-Q pure water (resistivity 18 MΩ·cm) under continuous compression at a rate of 3 mm / min. The numbers in parentheses represent the number-average molecular weight (g / mol) of each block.

[0015] (Figure 8) TEM images of (a) PS(1560)-PEG(5000), (b) PS(2993)-PEG(5000), (c) PS(5610)-PEG(5000), and (d) PS(13832)-PEG(5000) micelles formed in bulk aqueous solution. The dried micelle samples were negatively stained with uranyl acetate. (Table 2) Table 2. Diameters of PS-PEG micelles determined by TEM (Figure 8) or DLS JPEG2025102844000002.jpg32163

[0016] (Figure 9) Isotherms of the surface pressure - area at constant compression at 25 °C of the monolayers developed with (a) chloroform or (b) water for four types of PS-PEG materials. Milli-Q pure water (resistivity 18 MΩ·cm) was used as the subphase. The compression rate of the monolayer was 3 mm / min.

[0017] (Figure 10) (a) Longitudinal relaxation decay curve of PEG protons at 25 °C. The curve shown by the solid line fits a single exponential decay function (G(t) = exp(-t / T1)). (b) Transverse relaxation decay curve of PEG protons at 25 °C. In the spectra of PS(5610)-PEG(5000) micelles and PS(13832)-PEG(5000) micelles, two PEG peaks (a sharp peak at approximately 3.61 ppm and a broad peak at approximately 3.56 ppm) were observed. The decay curves of these peaks were separately fitted to a single exponential decay function. The open symbols represent the broad PEG peak, and the filled symbols represent the sharp PEG peak. In the spectra of PEG(5000) and PLGA(2385)-PEG(5000), one PEG peak was observed. The decay curve of PEG(5000) was fitted to a single exponential function, and the decay curve of PLGA(2385)-PEG(5000) was fitted to a double exponential function (G(t) = a·exp(-t / T 21 )+(1 - a)·exp(-t / T 22 )) (Table 3) Table 3. Best-fit T1 and T2 values obtained in Figure 10. †Percentage of PEG segments contributing to the sharp and broad PEG peaks out of the total number of available PEG segments in the system, estimated based on the pyridine internal standard . ‡Coefficient a of the first term of the double exponential decay function . For comparison, the predicted T2 value of the PEG(5000) melt at 100 °C is also shown . JPEG2025102844000003.jpg54163

[0018] (Figure 11) Surface tension - area isotherm of the monolayer of water-expanded PS(5610) -PEG(5000) in repeated compression - expansion cycles. The subphase used was a solution containing 150 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3, with a pH of 7.0 - 7.4, at 25 °C. The compression / expansion rate of the monolayer was 50 mm / min It took 7.18 minutes to perform one compression-expansion cycle. In this measurement, a total of 50 consecutive compression-expansion cycles were used.

[0019] (Figure 12) Surface tension-area isotherms of (a) Survanta (4 mg) and (b) PS(5610)-PEG(5000) micelles (10 mg) spread with and without addition of BSA (30 mg) during repeated compression-expansion cycles. The data shown are for the last 10 of a total of 50 consecutive cycles performed after spreading 4 mg of Survanta or PS-PEG micelles. A separate 30 mg of BSA was spread on the monolayer of Survanta (4 mg) or PS-PEG micelles (10 mg). The subphase used was a solution containing 150 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3 (pH 7.0 - 7.4, 25 °C). The monolayer was compressed / expanded at a rate of 50 mm / min, with each compression-expansion cycle taking 7.18 minutes. added and without addition, and (b) PS(5610)-PEG(5000) micelles (10 mg) spread with and without addition of BSA (30 mg) in water. added and without addition. The data shown are for the last 10 of a total of 50 consecutive cycles performed after spreading 4 mg of Survanta or PS-PEG micelles. A separate 30 mg of BSA was spread on the monolayer of Survanta (4 mg) or PS-PEG micelles (10 mg). The subphase used was a solution containing 150 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3 (pH 7.0 - 7.4, 25 °C). The monolayer was compressed / expanded at a rate of 50 mm / min, with each compression-expansion cycle taking 7.18 minutes. -PEG micelles. The data shown are for the last 10 of a total of 50 consecutive cycles performed after spreading 4 mg of Survanta or PS-PEG micelles. A separate 30 mg of BSA was spread on the monolayer of Survanta (4 mg) or PS-PEG micelles (10 mg). The subphase used was a solution containing 150 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3 (pH 7.0 - 7.4, 25 °C). The monolayer was compressed / expanded at a rate of 50 mm / min, with each compression-expansion cycle taking 7.18 minutes. of the 50 consecutive cycles. A separate 30 mg of BSA was spread on the monolayer of Survanta (4 mg) or PS-PEG micelles (10 mg). The subphase used was a solution containing 150 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3 (pH 7.0 - 7.4, 25 °C). The monolayer was compressed / expanded at a rate of 50 mm / min, with each compression-expansion cycle taking 7.18 minutes. 0 mg) was spread separately with BSA (30 mg) on the monolayer. The subphase used was a solution containing 150 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3 (pH 7.0 - 7.4, 25 °C). The monolayer was compressed / expanded at a rate of 50 mm / min, with each compression-expansion cycle taking 7.18 minutes. on the monolayer of Survanta (4 mg) or PS-PEG micelles (10 mg). The subphase used was a solution containing 150 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3 (pH 7.0 - 7.4, 25 °C). The monolayer was compressed / expanded at a rate of 50 mm / min, with each compression-expansion cycle taking 7.18 minutes. 50 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3 solution (pH 7.0 - 7.4, 25 °C) was used. The monolayer was compressed / expanded at a rate of 50 mm / min, with each compression-expansion cycle taking 7.18 minutes. / expanded at a rate of 50 mm / min, with each compression-expansion cycle taking 7.18 minutes.

[0020] (Figure 13) Rate of surface tension reduction of PS(5610)-PEG(5000) micelles and Survanta measured 100 seconds after spreading each compound. The subphase used was a subphase solution containing 150 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3 (pH 7.0 - 7.4, 25 °C). -PEG(5000) micelles and Survanta. The subphase used was a subphase solution containing 150 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3 (pH 7.0 - 7.4, 25 °C). The subphase used was a subphase solution containing 150 mM NaCl, 2 mM CaCl2, and 0.2 mM NaHCO3 (pH 7.0 - 7.4, 25 °C). containing subphase solution (pH 7.0 - 7.4, 25 °C) was used.

[0021] (Figure 14) Body weight of mice measured as a function of time after intratracheal injection of PS(4418)-PEG(5000) micelle polymer surfactant at different doses. Each dose group -PEG(5000) micelle polymer surfactant at different doses. Each dose group It consists of one mouse.

[0022] (Figure 15) H&E stained tissue section of mouse organs taken at the second week after intratracheal injection of 1.6 mg of PS(4418)-PEG(5000) micelle polymer pulmonary surfactant (1600μm × 1200μm).

[0023] (Figure 16) Ex vivo PV lung mechanics of fetal lungs of rabbits on the 27th day of pregnancy after injection of PS(4418)-PEG(5000) pulmonary surfactant at various doses. Each group consisted of 5 fetal rabbits.

Mode for Carrying Out the Invention

[0024] To facilitate understanding of the principles of the present disclosure, reference is made to the embodiments shown in the drawings and this is described using specific terms. However, it should be understood that this is not intended to limit the scope of the present disclosure.

[0025] In response to unmet needs, we have developed a polymeric pulmonary surfactant as a possible solution to these problems. Our newly developed fully synthetic polymer-based pulmonary surfactant can be manufactured at a much lower cost and enables the use of an aerosol delivery method that is much simpler, non-invasive, and does not require the assistance of a physician.

[0026] Surface Tension and Pulmonary Surfactant

[0027] To develop new SRT therapies, it is necessary to understand how lung function is impaired when the lungs are underdeveloped. The lungs of premature and immature infants differ from those of fully developed healthy infants in that they do not have a fully developed alveolar structure and produce little or no pulmonary surfactant. Although an underdeveloped alveolar structure has an adverse effect on the infant's breathing, this is not the main cause of RDS. In fact, even when the gestation period is complete, mammals such as mice are born with an incompletely developed alveolar structure and do not develop RDS. 1 This suggests that the main reason for RDS is not a structural defect but a deficiency of a functional compound, namely pulmonary surfactant. Pulmonary surfactant is a surfactant composed of 80% lipid and 20% protein that plays an important role in the biophysical function of the lungs by reducing the surface tension of the alveolar lining fluid (ALF; also called epithelial lining fluid, ELF). 2 When there is a deficiency of pulmonary surfactant in the lungs, the reduction of the surface tension of the ALF is hindered, and this situation causes a large pressure difference between the air sacs (alveoli) of the lungs.

[0028] During the respiratory cycle, the radius (R) of the alveoli changes, and accordingly, the pressure (ΔP) inside the alveoli also changes according to the following Young-Laplace equation.

[0029] JPEG2025102844000004.jpg9163

[0030] Figure 2 shows the calculation results of how ΔP changes as a function of γ and R. When there is no surfactant in the lungs, the surface tension of the ALF is constant regardless of R (about 0.069 N / m at 37°C). Therefore, as shown by the red solid line in Figure 2, when the alveoli contract until they are at their smallest (diameter 0.02 - 0.03 mm), the pressure can become very high (about 9,000 N / m 2)。Under such high pressures, large alveoli continue to grow, while small alveoli constantly become smaller and ultimately the alveolar structure is completely destroyed. If pulmonary surfactant is present, as the radius of the alveoli decreases, the surface tension also continuously decreases (down to less than 0.01 N / m) (see the green dotted line in Figure 2). This mechanism prevents pressure fluctuations in the breathing cycle (and lung destruction due to these pressure fluctuations).

[0031] The decrease in the surface tension of the ALF as the radius of the alveoli decreases is due to the repulsive force between lipid molecules in the pulmonary surfactant system. Since lipids are amphiphilic, they are easily adsorbed at the interface between air and the ALF. When the radius of the alveoli decreases during exhalation, reducing the alveolar surface area, the lipid molecules adsorbed on the surface are compressed laterally. As shown in Figure 3, the pressure generated by compressing the lipids causes the surface tension to decrease. When the surface tension decreases to near zero, the lipids adsorbed on the surface become unstable and begin to desorb into the sub-phase of the ALF. This desorption of lipids causes a fluctuation in the profile of the surface tension near 0 mN / m as shown in Figure 3. to occur in the profile 12 During inhalation, the surface area of the alveoli increases, creating an uncovered surface. Due to the action of surfactant proteins B and C (SP-B and SP-C respectively), the lipids that have desorbed and are in the sub-phase are adsorbed back onto the surface of the ALF. 13 Compression and expansion cycles, the continuous desorption and re-adsorption of lipids is only possible in the presence of SP-B and SP-C proteins, and such continuous desorption and re-adsorption is an important process that should be reproducible with any SRT therapeutic agent. Without the presence of SP-B / SP-C, the desorbed lipids generated during compression do not re-adsorb onto the surface of the ALF, so the initial clinical trials of SRT using lipid-only formulations were mostly unsuccessful. 14,15 。The first synthetic lung surfactant, dipalmitoyl phosphatidylcholine (DPP C) containing 85% by weight, 9% by weight of hexadecanol, and 6% by weight of tyloxapol, Exosur f (GlaxoSmithKline) has been discontinued due to its low effectiveness in the treatment of RDS. The main reason for the insufficient performance was that this formulation contained no protein components at all.

[0032] The current trend in SRT research is the development of synthetic compounds to replace natural surfactant proteins. To date, there is only one lung surfactant product in this category at the clinical evaluation stage, which is Surfaxin (formerly Discovery Labs; now Wi ndtree Therapeutics). This formulation contains a synthetic polypeptide (KL4) designed to mimic the function of natural surfactant proteins. However, Surfaxin has the drawback of low stability against temperature fluctuations compared to lung surfactants extracted from animals (such as Survanta, Curosurf, In 16,17 fasurf, etc.). Once Surfaxin is heated before use (for example, at 44°C for 15 minutes in a dry block heater and then cooled to body temperature, etc.), it cannot be returned to the refrigerated state. In the clinical trials of Surfaxin, it was shown to be slightly improved in reducing morbidity and mortality compared to Survanta (Abb Vie) extracted from cows or Curosurf (Chiesi 18 Pharma) extracted from pigs. 1 9,20 ​​However, in the "Surfaxin in Therapy Against RDS (STAR)" trial, the enrollment rate was low, so it ended early (only half of the expected data was obtained), and the "Safety and Assessment of Effectiveness of Lucinactant versus Exosurf in a Clinical Trial ( SELECT)" trial has been noted to be criticized for biased evaluation. 21,2 2 Therefore, it cannot be said that Surfaxin has better treatment results than RDS therapeutics extracted from animals, nor can the opposite be said. To date, there is no single lung surfactant product that has been shown to have clearly superior results compared to other lung surfactant products. In April 2015, Discovery Labs (renamed Windtree Therapeutics in April 2016) 23 halted the production of Surfaxin to restructure its mainstay around the aerosol formulation Aerosurf. As a result, all currently marketed SRT therapeutics are animal-derived products.

[0033] Cost comparison between different SRT therapeutics is not straightforward. The overall treatment cost, including the number of endotracheal injections, the duration of mechanical ventilation, the required preparations, dosage, etc., depends on several factors. 18 The cost of the lung surfactant drug itself accounts for only about 2 - 3% of the total SRT treatment cost in the United States. Most of the treatment cost is related to the complex procedures associated with the administration of the lung surfactant. 24,25 Therefore, the use of SRT technology on a global scale ​​​​​​​The two main factors that hinder, namely high cost and complex delivery procedures, are interrelated. To solve this problem that has persisted for decades, it is probably necessary to "think outside the box". For this reason, in our laboratory, we have studied the behavior of synthetic polymer surfactants (especially PEG-based amphiphilic block copolymers) at the air / water interface over the past few years. Polymer lung surfactants can be produced at low cost and in large quantities, are easily aerosolized (in both liquid and dry powder forms), and are chemically stable (the polymer preparation has a long shelf life), making them attractive as SRT therapeutic agents. During the course of this study, we developed design criteria for polymer lung surfactants. For a polymer

[0034] lung surfactant to be a successful candidate, it should (1) be biocompatible / biodegradable, (2)

[0035] The chemical structures of the tested candidate polymer lung surfactant materials are shown in Figure 4. Their molecular characteristics are also summarized in Table 1. The first two materials included in the list are widely known biodegradable block copolymers that have obtained FDA approval, namely, poly(lactic-co-glycolic acid-block-ethylene glycol) (PLGA-PEG) and its derivative poly(lactic-co-glycolic acid-co-caprolactone-block-ethylene glycol) (PLGACL-PEG). The other candidate block copolymers under investigation are poly(tert-butyl methacrylate-block-ethylene glycol) (PtBMA-PEG) and poly(styrene-block-ethylene glycol) (PS-PEG). The PtBMA-PEG and PS-PEG materials are biocompatible, especially when prepared as micellar solutions. In the micellar solution, the PtBMA block and the PS block form a hydrophobic core domain, and the PEG chains form a hydrophilic corona layer. PS-PEG micelles have been studied in great detail as potential drug delivery systems. Pharmacological studies, including toxicity, biodistribution, and pharmacokinetic tests, have suggested that PS-PEG micelles are very safe for use in biomedical applications 26~30 . There is very little pharmacological data currently available for PtBMA-PEG micelles. However, considering the similarity between PS-PEG micelles and PtBMA-PEG micelles, PtBMA-PEG micelles are also thought to have high biocompatibility. Data obtained from similar systems, such as poly(methyl methacrylate-co-methacryloxysuccinimide-graft-polyethylene glycol)) ("PMMA-co-PMASI-g-PEG") micelles, also support this finding 31 . JPEG2025102844000005.jpg44163

[0036] Table 1. Molecular characteristics of the investigated polymer lung surfactant candidate materials 1 Lactide:glycolide = molar ratio of 70:55 2 Lactide:glycolide:caprolactone = molar ratio of 30:29:41

[0037] In our study, these block copolymer pulmonary surfactants were prepared in aqueous micellar solutions. A solvent exchange method was used to obtain spherical micellar structures. 32,33 The polymeric micelle preparation was very stable at room temperature and did not require any pretreatment steps before use. This alone is a major advantage compared to conventional lipid-based pulmonary surfactant formulations, which generally require specific storage and pretreatment procedures before use to obtain reproducible results. This advantage in terms of handling alone could already help effectively reduce the overall cost of treatment.

[0038] Criterion 2 for polymeric pulmonary surfactants: Extremely low surface tension (or high surface pressure) at high compression

[0039] First, we conducted our research using FDA-approved biodegradable block copolymers. We focused on PLGA-PEG, a polymer that is generally It forms a film that spreads easily across the air / water interface, called a monolayer. A simulated lung test environment was created in vitro using PLGA-PEG. When the PLGA-PEG polymer spreads beyond the point of total coverage onto the water surface, it forms an insoluble brush-coated film. At this time, the PLGA segments form a slightly glassy insoluble polymer membrane. The PEG segments penetrate into the water subsurface (forming a brush layer). When strongly compressed in the lateral direction, the PLGA-PEG is compressed by the PEG of the PLGA. The combined effect of the cross transition and the repulsive force of the PEG brush reduces the surface tension of water to nearly zero. Become smaller 35,36 Langmuir microscopy of PLGA-PEG monolayers under various compression conditions The morphological and surface mechanical properties are detailed in ref. 34,35 .

[0040] Figure 5(a) shows the surface tension-area isotherms obtained from Langmuir monolayers formed by PLGA-PEG and its non-glassy analog, PLGACL-P EG. Chloroform was used as the spreading solvent for the preparation of the monolayers. At high compression, the monolayer of PLGA-PEG had a low surface tension of about 8 mN / m. The monolayer of PLGACL-PEG had a much higher surface tension, about 40 mN / m, even at maximum compression. This is due to the fact that PLG ACL is a non-glassy polymer. In actual therapeutic applications, chloroform cannot be used as the spreading solvent, and the formulation must be water-based . In aqueous solution, PLGA-PEG and PLGACL-PEG exist in the form of micelles. The surface tension-area profiles of the micelle monolayers of PLGA-PEG and PLGACL-PEG (prepared using water as the spreading solvent) are shown in Figure 5(b). In these water-spreading situations, the surface tension behaviors of the monolayers of PLGA-PEG and PLGACL-PEG were very similar to each other. In both cases, the smallest observed surface tension was about 50 - 60 mN / m at the highest compression level tested . The reason for the significant difference in the surface tension isotherms between the monolayer of PLGA-PEG spread with chloroform and the monolayer of PLGA-PEG spread with water is that in the water-spread monolayer system, the PL GA-PEG polymer remains in micellar form, whereas in the situation of spreading with chloroform, the polymer forms a molecularly extended ("anchor brush") monolayer . . The surface tension-area profiles of the micelle monolayers of PLGA-PEG and PLGACL-PEG (prepared using water as the spreading solvent) are shown in Figure 5(b). In these water-spreading situations, the surface tension behaviors of the monolayers of PLGA-PEG and PLGACL-PEG were very similar to each other. In both cases, the smallest observed surface tension was about 50 - 60 mN / m at the highest compression level tested . In both cases, the smallest observed surface tension was about 50 - 60 mN / m at the highest compression level tested .

[0041] The reason for the significant difference in the surface tension isotherms between the monolayer of PLGA-PEG spread with chloroform and the monolayer of PLGA-PEG spread with water is that in the water-spread monolayer system, the PL GA-PEG polymer remains in micellar form, whereas in the situation of spreading with chloroform, the polymer forms a molecularly extended ("anchor brush") monolayer . This is because in the water-spread monolayer system, the PLGA-PEG polymer remains in micellar form, while in the case of spreading with chloroform, the polymer forms a molecularly extended ("anchor brush") monolayer 。The method of spreading with this water was tested with various PLGA-PEG samples with significantly different total molecular weights (3.5 - 28.6 kg / mol) and block compositions (28.4 - 74.3 wt% PEG). However, for all samples, the surface tension did not become sufficiently small under high compression (typically always > 45 mN / m). PLGA-PEG micelles spread with water always remained as they were after the initial spreading at the air / water interface (unpublished results / manuscript in preparation).

[0042] Other hydrophobic chemical substances were tested to achieve a surface tension lower than that obtained using PLGA-PEG and PLGA-CL-PEG with polymer lung surfactants spread with water. Assuming that the micelles spread with water would fuse to form an anchor brush type monolayer, the resulting monolayer would have a significantly lower surface tension. To verify this hypothesis, PtBMA-PEG was used. PtBMA has a tendency to wet the water surface. 37,38 The wettability (i.e., spreadability) of the polymer can be quantified as follows using the spreading coefficient (S).

[0043] S = γ 空気-水 - γ 空気-ポリマー - γ ポリマー-水

[0044] The more positive the value of S, the stronger the wettability. The already known interfacial tension values, namely, γ = 72 mN / m (25 °C), γ 空気-水 = 30 mN / m, γ 空気-PtBMA = 30 mN / m, γ PtBMA-水Using γ = 18 mN / m, the spreading coefficient of PtBMA is calculated to be S = 24 mN / m > 0. The spreading coefficient of PLGA is about 10 nN / m. These numbers indicate that PtBMA-PEG has a stronger tendency to spread on the water surface than PLGA-PEG. Therefore, the PtBMA-PEG micelle monolayer is more likely to spontaneously transform into a laterally uniform anchor brush-type monolayer.

[0045] The surface tension properties of the PtBMA-PEG monolayer spread with chloroform and the PtBMA-PEG monolayer spread with water were investigated. The results are shown in Fig. 6(a). Interestingly, in the case of PtBMA-PEG, even the monolayer spread with water had an extremely low surface tension (less than 10 mN / m at high compression), similar to the monolayer spread with chloroform. A (pseudo) Plateau transition was observed in both the isotherms for the case of spreading with chloroform and the case of spreading with water, at around a surface tension of about 48 mN / m. This may be due to the PtBMA anchor block. The PtBMA anchor block forms a layer of continuous film. This interpretation is supported by the fact that the surface tension at the Plateau transition is the same as the value estimated from the spreading coefficient of PtBMA, i.e., (γ ポリマー-空気 +γ ポリマー-水 ) = γ 空気-水 - S = 72 - 24 = 48 mN / m. This is clear evidence that when the micelle monolayer is laterally compressed, the core domain of the PtBMA-PEG micelles collapses and fuses into the film. This interpretation is also supported by the observation that the Plateau transition becomes more prominent as the temperature increases (Fig. 6(b)). At high temperatures, the mobility of the PtBMA segments increases, promoting the micelle fusion process.

[0046] For PS-PEG, the water wetting property is opposite to that of PtBMA-PEG. PS has a spreading coefficient of S = -8 mN / m with a negative value 39~40, it is prone to dewetting phenomenon on the water surface. Since PS is prone to dewetting in this way, even when developed with chloroform, it still hinders the formation of a uniform anchor brush monolayer of PS-PEG copolymer in the lateral direction. In this regard, PS-PEG is distinguished from other polymers (PLGA-PEG, PLGACL-PEG, PtBMA-PEG) studied. In the literature, it has been demonstrated that when developed with an organic solvent, generally chloroform, PS-PEG forms surface micelles on the water surface. 41~47 . Also, it has been shown that the PS-PEG polymer developed with chloroform has a low surface tension (less than 10 mN / m) during high compression. Before our investigation, it was unknown whether the surface tension would also be low in the monolayer of PS-PEG micelles developed with water. The answer is shown in Figure 7.

[0047] Due to the dewetting of PS at the air / water interface, monolayers of PS-PEG developed with either chloroform or water contain non-fused micelles. However, the surface micelles formed by spreading a chloroform solution of PS-PEG on the water surface are assumed to have an anisotropic molecular form due to the asymmetry of the air / water interface, while generally an isotropic spherical micelle form is obtained in the bulk aqueous solution (Figure 8(c)). Such a difference in form seems to create a significant difference in the shape of the surface tension-area isotherm (Figure 7). Most importantly, the monolayer of PS-PEG micelles developed with water was found to meet the requirement of having a low surface tension, which may be used in SRT. In this preparation, it was possible to make the surface tension extremely low (about 0 mN / m) during high compression (Figure 7).

[0048] We investigated how varying the size of the PS block affects the morphology and surface tension properties of the block copolymers. Three new PS-PEG samples with different molecular weights of the PS block, namely, PS(1560)-PEG(5000), PS(2993)-PEG(5000), and PS(13832)-PEG(5000), were prepared with a constant PEG molecular weight of 5,000 g / mol. As shown in Figure 8 (and summarized in Table 2), the molecular weight of the PS block significantly affected the size of the micelles.

[0049] JPEG2025102844000006.jpg32163

[0050] Table 2. Diameters of PS-PEG micelles determined by TEM (Figure 8) or DLS

[0051] Four PS-PEG materials (both developed in chloroform and in water) The surface tension-area isotherm was measured for each sample. The data is shown in Figure 9. To determine the surface area per chain, the isotherm data were plotted as logarithmic surface pressure versus logarithmic surface area per chain. Here, the surface pressure (π) is defined as follows:

[0052] π=γ 空気-水 -γ 空気-ポリマー-水 =γ 空気-水 -(γ 空気-ポリマー +γ ポリマー -水 )

[0053] Here, γ 空気-ポリマー-水 is the surface tension at the polymer-coated air / water interface Due to the limited size of the Langmuir trough, it is difficult to measure a large monolayer area. To construct a complete surface pressure-area isotherm for each type of monolayer, the range of monolayer areas is It was necessary to perform the measurements multiple times (2-3 times) at different ranges. The isotherms obtained from different regions of the monolayer expanded in [solvent] were closely and mutually overlapping without separation (Figure 9). On the other hand, as also shown in the figure, the curves obtained from different regions of the monolayer expanded in water were segmented. This suggests that in the method of expanding in water, some of the material (PS-PEG micelles) was taken into the subphase and lost. When the polymer was expanded from a chloroform solution, the loss of material to the subphase was negligible. The monolayer of PS-PEG expanded from a chloroform solution showed a similar isotherm profile with a surface pressure of less than about 10 mN / m regardless of the molecular weight of the PS block (Figure 9(a)). When compressed above the surface pressure level of 10 mN / m, the surface pressure of the monolayer expanded in chloroform increased more rapidly as the molecular weight of PS increased. We consider this observation to be due to the fact that the larger the molecular weight of the PS segment, the larger the core domain of the PS-PEG surface micelle. The mechanical surface properties of PS-PEG surface micelles expanded in chloroform have been studied by other researchers in the past. Unlike the case of expansion in chloroform, the surface pressure of the monolayer of PS-PEG expanded in water did not show a monotonic form with respect to the molecular weight of the PS block. One notable observation was that in the system expanded in water, the surface pressure was maximum when the molecular weight of the PS block was intermediate. The most rapid increase in surface pressure during compression was observed in the monolayer of PS(5610)-PEG(5000) micelles expanded in water (Figure 9(b)). Interestingly, among all the systems tested, the PS(13832)-PEG(500) monolayer expanded in water had the lowest surface pressure at the same compression level. This is thought to be related to the fact that when the molecular weight of the PS block is too small, the hydrophobic interaction between PS segments is weak, and when the molecular weight is too large, the entanglement of PS segments becomes significant, both of which affect the formation and stability of the surface micelle. In the case of expansion in chloroform, the loss of material to the subphase was negligible. The monolayer of PS-PEG expanded from a chloroform solution showed a similar isotherm profile with a surface pressure of less than about 10 mN / m regardless of the molecular weight of the PS block (Figure 9(a)).

[0054] When compressed above the surface pressure level of 10 mN / m, the surface pressure of the monolayer expanded in chloroform increased more rapidly as the molecular weight of PS increased. We consider this observation to be due to the fact that the larger the molecular weight of the PS segment, the larger the core domain of the PS-PEG surface micelle. The mechanical surface properties of PS-PEG surface micelles expanded in chloroform have been studied by other researchers in the past. Unlike the case of expansion in chloroform, the surface pressure of the monolayer of PS-PEG expanded in water did not show a monotonic form with respect to the molecular weight of the PS block. One notable observation was that in the system expanded in water, the surface pressure was maximum when the molecular weight of the PS block was intermediate. The most rapid increase in surface pressure during compression was observed in the monolayer of PS(5610)-PEG(5000) micelles expanded in water (Figure 9(b)). Interestingly, among all the systems tested, the PS(13832)-PEG(500) monolayer expanded in water had the lowest surface pressure at the same compression level. 43,44 This is thought to be related to the fact that when the molecular weight of the PS block is too small, the hydrophobic interaction between PS segments is weak, and when the molecular weight is too large, the entanglement of PS segments becomes significant, both of which affect the formation and stability of the surface micelle. Unlike the case of expansion in chloroform, the surface pressure of the monolayer of PS-PEG expanded in water did not show a monotonic form with respect to the molecular weight of the PS block. One notable observation was that in the system expanded in water, the surface pressure was maximum when the molecular weight of the PS block was intermediate. The most rapid increase in surface pressure during compression was observed in the monolayer of PS(5610)-PEG(5000) micelles expanded in water (Figure 9(b)). Interestingly, among all the systems tested, the PS(13832)-PEG(500) monolayer expanded in water had the lowest surface pressure at the same compression level. This is thought to be related to the fact that when the molecular weight of the PS block is too small, the hydrophobic interaction between PS segments is weak, and when the molecular weight is too large, the entanglement of PS segments becomes significant, both of which affect the formation and stability of the surface micelle. Interestingly, among all the systems tested, the PS(13832)-PEG(500) monolayer expanded in water had the lowest surface pressure at the same compression level. It should be noted that the original text seems to have some parts where the solvent name is missing in . I assumed it to be "chloroform" based on the context in subsequent paragraphs. Also, the last sentence in the original seems incomplete in terms of what is being described about the PS(13832)-PEG(500) system. The translation attempts to make sense of the overall text flow and scientific content as much as possible with the given information.0) The maximum surface pressure was the lowest in the monolayer (Fig. 9(b)). PS(1383) spread with water 2) The maximum surface pressure of -PEG(5000) (10 - 20 mN / m) was comparable to that of PLG A-PEG and PLGACL-PEG spread with water.

[0055] The three types of monolayer systems (PS-PEG, PLGA-PEG, and PLGACL-PEG micelle monolayer) to be studied and spread with water have similar morphologies (i.e., these monolayers are all composed of micelles that cannot merge), but it is interesting to note that only the PS-PEG micelles can generate a high surface pressure (above 40 mN / m) under high compression. These results indicate that the PS-PEG micelles are much more hydrophobic (than the PLGA-PEG or PLGACL-PEG micelles), so they are more strongly fixed at the air / water interface, which makes the PS-PEG micelles more stable and less likely to sink, and thus the mechanical resistance of the PS-PEG micelle monolayer to lateral compression is higher. This explanation seems unlikely if we consider that the corona of the micelles is composed of the same polymer (PEG) regardless of which hydrophobic material (PS, PLGA, or PLGACL) is used to make the block copolymer. That is, if these different types of micelles share the same PEG corona structure, the question is how the PS-PEG micelles become more hydrophobic than the PLGA-PEG / PLGACL-PEG micelles. We believe that such a difference actually occurs because PS is more hydrophobic than PLGA (or PLGACL) (γ PS-水 = about 41 mN / m, γPLGA-water = about 25 mN / m) 34It is considered to be caused by the fact. As a result (unlike PLGA-PEG micelles and PLGACL-PEG micelles that form a brush layer with completely hydrated PEG chains), in order to minimize the unfavorable state in which the PS material is exposed to water molecules, the PEG chains in the PS-PEG micelles exist in a disrupted state, which ultimately makes the PS-PEG micelles more hydrophobic overall than PLGA-PEG micelles or PLGACL-PEG micelles. Therefore, the hydrophobicity of the micelles is an important criterion for designing / selecting materials suitable for use as lung surfactants.

[0056] Further studies were conducted to confirm the disrupted conformation of the PEG chains in the PS-PEG micelles. Specifically, for this purpose, in situ NMR spin relaxation techniques were used to measure the mobility of the micelle PEG brush segments in PS-PEG micelles and PLGA-PEG micelles. The disrupted PEG corona chains of the PS-PEG micelles are considered to have significantly lower mobility than the completely hydrated PEG corona chains of the PLGA-PEG micelles. From the NMR spin relaxation measurement results, it is possible to determine two real-time constants, namely the longitudinal relaxation time (T1) and the transverse relaxation time (T2). T1 is related to the chemical structure of the chain segment ("fast mode"), and T2 is related to the conformation of the chain segment ("slow mode"). 48 It is considered that the PEG T1 values are the same between the micelles of PS-PEG and PLA-PEG, but the T2 values are significantly different. The NMR measurements were carried out on four representative systems: PS(5610)-PEG(5000) micelles, PS(13832)-PEG(5000) micelles, PLGA(2385)-PEG(5000) micelles, and PEG(5000) homopolymer in heavy water. For the PS-PEG micelles, two separate PEG proton peaks were observed (a sharp ("hydrated PEG") peak at about 3.61 ppm and a broad ("disrupted PEG") peak at about 3.56 ppm). These two peaks were analyzed separately for T1 and T2. The results are shown in Figure 10.

[0057] As shown in Fig. 10, since all four types of samples tested showed the same PEG T1 value, the validity of the measurement was confirmed. Conversely, the measured PEG T2 values differed significantly among the samples. To obtain a measure of PEG mobility, the T2 value of a 5 kg / mol PEG homopolymer melt at 100 °C was calculated. Under these conditions, the Rouse time of PEG is 281.52 ps, which corresponds to T2 = 0.3838 s. 49 . Since the hydrated PEG chains have high mobility, it is considered that the T2 value becomes longer than 0.3838 s. By fitting the transverse decay curve to the single exponential function G(t) = exp(−t / T2), the T2 value of the hydrated free PEG(5000) chains was estimated to be 0.6604. On the other hand, since the mobility of the PEG segments varies depending on how close the PEG segments are to the graft surface, the transverse decay curve of the PLGA(2385)-PEG(5000) micelles is represented by the double exponential function G(t) = a·exp(−t / T 21 )+(1 - a)·exp(−t / T 22 ). T 21 corresponds to the PEG segments away from the graft surface, which contributed significantly to the overall signal intensity (a = 0.9071). T 22 corresponds to the PEG segments close to the graft surface. The T value of the PLG 21 A-PEG micelles was higher than the T 21 value of the PEG melt and slightly lower than the T value of the hydrated PEG(5 21 000). This indicates that the PEG corona chains of the PLGA-PEG micelles are actually fully hydrated. In the case of the PS-PEG micelles, two separate PEG peaks were observed in the NMR spectrum. These two PEG peaks were separately fitted to single exponential functions. The T2 value obtained from the decay curve of the sharp PEG peak of the PS-PEG micelles was comparable to the T value obtained for the PLGA-PEG micelles 21 , indicating that , it is suggested that the sharp PEG peak corresponds to the hydrated PEG segments of the PS-PEG micelles. However, the T2 value obtained from the broad PEG peak of the PS-PEG micelles was very small and even smaller than the T2 value obtained from the PEG melt. This clearly indicates that in the PS -PEG micelles (due to the strong hydrophobicity of the PS material), a significant portion of the PEG segments exist in a disrupted state. Also, this result is consistent with past reports on the disrupted PEG brush structure in poly(butadiene-block-ethylene glycol) PB-PEG) micelles (γ = 45.9 mN / m) based on low-temperature TEM and small-angle neutron scattering (SANS) experiments and considerations in self-consistent field (SCF) theory PB - 水 = 45.9 mN / m), as well as 50,51 .

[0058] JPEG2025102844000007.jpg54156

[0059] Table 3. Best-fit T1 and T2 values obtained in Figure 10. †Based on the pyridine internal standard estimated, among the total number of available PEG segments in the system, the proportion of PEG segments contributing to the sharp PEG peak and the broad PEG peak. ‡Coefficient a of the first term of the biexponential decay function . For comparison, the predicted T2 value of PEG(5000) melt at 100 °C is also shown.

[0060] Furthermore, it is very interesting to note that the PS(13832)-PEG(5000) micelles have more hydrated PEG segments than the PS(5610)- PEG(5000) micelles. The hydrated PEG segments and disrupted PE of the PS-PEG micelles The absolute concentration of the G segment was determined using the NMR signal from pyridine added as an internal standard. PS(13832)-PEG(5000) micelles (34.1 ± 1 .6%) were found to have a considerably higher proportion of hydrated PEG segments than PS(5610)-PEG(5000) micelles (11.6 ± 1.6%) (Table 3). These results clearly support the idea that, for some reason not yet clear, PS(13832)- PEG(5000) micelles are less hydrophobic than PS(5610)-PEG(5000) micelles and therefore do not bind as strongly to the air / water interface. Indeed, as shown in Figure 9, in the monolayer of PS(13832)-PEG(5 000) micelles spread on water, only relatively low surface pressures were found even under high compression (about 10 mN / m).

[0061] Overall, our investigations have now led to the identification of a very promising class of candidate materials with surface tension / pressure properties that could potentially be used as lung surfactants, namely PS(5610)-PEG(5000) block copolymers prepared in the form of aqueous micelles. The micellar aqueous solution of PS(5610)-PEG(5000) exhibits excellent colloidal stability over long periods. In PS(5610)-PEG(5000) micelle samples, it was confirmed that the same surface pressure - area profile was reproduced after storage at room temperature for longer than 3 months (data not shown).

[0062] The monolayers of PS-PEG and PtBMA-PEG spread with water are equally promising from the perspective of their surface tension properties. However, pharmacological data for evaluating the in vivo safety of the PtBMA-PEG micelle material are not available at present. Therefore, the PS-PEG micelle system is more suitable for use in SRT. During consecutive compression-expansion cycles, the dynamic hysteresis properties of the monolayer system of PS(5610)-PEG(5000) spread with water were further analyzed. The results are shown in Fig. 10. The monolayer of PS-PEG showed significant surface tension hysteresis in the compression-expansion cycle and a very low minimum surface tension (about 0 mN / m) at maximum compression, similar to that observed with the commercial lung surfactant Survanta (see Fig. 3 for comparison). After the first three compression-expansion cycles, the hysteresis loop of the compression-expansion isotherm showed a closed shape (Fig. 11). This suggests that there was little loss of the polymer material to the subphase during repeated compression-expansion cycles. It is noteworthy that this combination of hysteresis and reversibility properties was easily achieved in the PS-PEG monolayer system without adding proteins. We consider this to be an advantageous aspect of using polymer agents for the control of the air / water interface. The PS-PEG micelles do not appear to desorb from the water surface. Therefore, unlike natural lung surfactants, polymer lung surfactants do not require proteins to function.

[0063] Polymer Lung Surfactant Criterion 3: Protein Resistance

[0064] Respiratory failure caused by inactivated lung surfactant induced by an increase in serum albumin protein (or other surfactants) is called acute respiratory distress syndrome (ARDS). 52,53 Therapeutic agents developed for the treatment of nRDS are not effective for the treatment of ARDS because the injected lung surfactant is inactivated. The protein resistance properties of the PS-PEG micelle lung surfactant were evaluated.

[0065] First, we tested how a commercially available lung surfactant, Survanta (Abb Vie), would react to the addition of the surfactant protein bovine serum albumin (BSA). As shown in Figure 12(a), BSA inactivated Survanta. When BSA was added, Survanta lost its ability to lower the surface tension to near zero, and the minimum surface tension increased to approximately 26 mN / m. However, the activity of PS-PEG micelles to reduce surface tension was hardly affected by the addition of BSA (Figure 12(b)). Lung surfactants of PS-PEG micelles may also be useful for the treatment of ARDS.

[0066] Polymer Lung Surfactant Criterion 4: Rapid spreading on water

[0067] The surface tension-lowering effect of the SRT therapeutic agent must occur within a few minutes after the first administration. Otherwise, the patient will face a life-threatening situation. We tested whether PS-PEG micelles would spread naturally on the water surface in an extremely short time. As shown in Figure 13, we confirmed that the surface tension-lowering effect of PS-PEG micelles reached the steady-state level within approximately 100 seconds after the polymer was spread on the water surface. This is much shorter than the induction time (about 800 seconds) observed with Survanta. This spreading rate does not necessarily have to exactly match the rate at which a therapeutic effect is obtained. Still, the fast spreading rate of PS-PEG micelle lung surf actant is an advantageous feature that enhances the feasibility for application in SRT. Characteristic. Safety of PS-PEG micelles intratracheally injected into adult mice

[0068] To evaluate the safety of PS-PEG lung surfactant intratracheally administered to adult mice a preliminary study was conducted. Three polymer dosage levels of 0.64 mg, 6.5 mg and 64 mg of polymer per kg of mouse body weight were tested (0.6 mg / ml, 6 mg / ml , 20 microliters of a 60 mg / ml PS-PEG micelle solution was injected into each mouse . The polymer used was PS(4418)-PEG(5000), and the overall molecular weight and block composition were selected based on the above-described in vitro results. To confirm that this PS(4 418)-PEG(5000) material meets the four performance criteria of the above-described polymeric pulmonary surfactant, separate experiments were conducted. For the polymeric pulmonary surfactant, the trachea was surgically incised and injected. From the time of injection, mice were monitored for signs of toxicity ( weight loss and behavioral symptoms) for 14 days. The body weight of the mice is shown in Figure 14. The initial weight loss observed in the first 1-2 days is presumably due to the surgical incision. When the mice recovered from the surgery, they showed normal behavior and their body

[0069] weights began to increase steadily. No signs of toxicity were observed in the mice during the 14-day period at all the polymer doses tested. On the 14th day after injection, the mice were sacrificed and the major organs (brain, heart, kidney, liver, lung, spleen) were collected for histological analysis. Blind histological evaluation of H&E-stained organ specimens was performed by a histopathology expert. Representative images of the H&E-stained tissue sections are shown in Figure 15. On the 14th day after injection, the mice were sacrificed and the major organs (brain, heart, kidney, liver, lung, spleen) were collected for histological analysis. Blind histological evaluation of H&E-stained organ specimens was performed by a histopathology expert. Representative images of the H&E-stained tissue sections are shown in Figure 15. The effectiveness of PS-PEG micelles intratracheally injected into premature rabbit lungs

[0070] Pressure (P)-volume (V) measurements using an ex vivo lung model (excised animal lungs) are a common method for evaluating the effectiveness of RDS

[0071] therapeutic agents. Since the reproducibility of lung PV mechanics is high in independent tests, ex vivo PV tests are used for quality control of lung surfactant products extracted from animals . Since the reproducibility of lung PV mechanics is high in independent tests, ex vivo PV tests are used for quality control of lung surfactant products extracted from animals . Since the reproducibility of lung PV mechanics is high in independent tests, ex vivo PV tests are used for quality control of lung surfactant products extracted from animals An FDA-approved method for quality control (QC) / quality assurance (QA). 2 . In this study, As an animal model, fetuses of New Zealand white rabbits on the 27th day of pregnancy, which is equivalent to preterm birth, were used . A commercially available bovine-extracted lung surfactant product, Newfacten (Yuhan Corporation), was also tested as a positive control.

[0072] The same polymer as that used in the in vivo toxicity test in mice, namely PS(4418 )-PEG(5000), was also used in this ex vivo efficacy test on rabbit fetuses. For the lungs of rabbit fetuses, PS-PEG micelle solutions at 0.6 mg / ml, 6 mg / ml, and 60 mg / ml were achieved by injecting 1.5 ml per kg of body weight respectively, and three polymer dosages of (0, )6, 60, 96 mg per kg of body weight of rabbit fetuses were used (the body weight of rabbit fetuses was 20 - 30 g). 96 mg / kg was at a polymer concentration of 64 mg / ml (which is the highest polymer concentration achievable with our current solvent exchange procedure), and was the maximum possible dosage achievable with an optimal liquid installation volume of 1.5 ml per kg of body weight. .

[0073] The PV profiles of the lungs of rabbit fetuses after PS-PEG (or Newfactan) injection are shown in Figure 1 6. At first glance, PS-PEG lung surfactant seems to be less efficient in increasing lung compliance and reducing the work of breathing at the same mass dosage compared to the commercially available formulation Newfacta n. However, these results are promising in demonstrating the dose-dependent effect of PS-PEG on improving lung compliance. Furthermore, ​​​The dose-dependent slope of PS-PEG indicates that PS-PEG is indeed treating RDS at its optimal dose. These results suggest that PS(4418)-PEG(5000) has great potential for use in treating cancer. Therefore, further studies are needed to identify the optimal dose of PS(4418)-PEG(5000) required to achieve a sufficient therapeutic effect.

[0074] It should also be noted that recent studies suggest that viscosity plays an important role in determining the performance of pulmonary surfactants, particularly as it controls fluid distribution during intratracheal instillation. 56,57 Our PS(4418)-PEG(5000) preparation has a bulk shear viscosity of 1.06 cp at 37 °C at a polymer concentration of 64 mg / ml, which is not far from the bulk shear viscosity of pure water, 0.6 cp. It is also noteworthy that the viscosity of Exosurf (Glaxo-Wellcome), a discontinued synthetic pulmonary surfactant product, was 3 cp. 56,58,59 Computational fluid dynamics (CFD) simulations suggest that low viscosity can lead to non-uniform distribution of pulmonary surfactant within the lungs, whereas animal-derived pulmonary surfactant products with high shear viscosities (approximately 30 cp) (e.g., Survanta, Infasurf, Curosurf) generally distribute quite uniformly within the lungs upon instillation. 56 Thus, the low viscosity of the current PS(4418)-PEG(5000) preparation also suggests room for improvement.

[0075] Our data suggest that polymeric pulmonary surfactants have great potential for use in SRT. Since the first animal-derived drugs for the treatment of RDS were developed in the 1980s, little further progress has been made in this field. Aerosol delivery and replacement with synthetic proteins have been the main drivers of research, but efforts have been limited in fruition. 60~63。Testing of fully synthetic polymer materials in SRT indicates a fundamental shift towards researching pulmonary surfactants. Polymer pulmonary surfactants may open the door to new therapeutic options for the treatment of RDS that have not been achievable with conventional lipid-based SRT therapeutics to date. Polymer pulmonary surfactants can be easily aerosolized in liquid or powder form. In the pharmaceutical delivery literature, polymers such as PS, PLGA, and PLGA-PEG have been frequently used as excipients for drug delivery to the lungs. 64~69 。In addition, polymer pulmonary surfactant preparations can also be used for co-delivery with other therapeutic agents that, when delivered alone, carry a risk of causing inactivation of the pulmonary surfactant.

[0076] Perhaps the greatest advantage will be the ability to achieve the reversible and hysteretic surface tension lowering effects observed in PS-PEG and PtBMA-PEG systems using a wider range of chemical options, creating opportunities to further improve this technology and generate new applications.

[0077] For the first time, the concept of using fully synthetic polymer materials as SRT therapeutics has been proposed and its feasibility demonstrated. Polymer pulmonary surfactants may address the limitations in current animal-derived lipid-based RDS therapeutics, namely high manufacturing / treatment costs and complex delivery procedures. Polymer pulmonary surfactants will have a much longer shelf life and will not require complex pretreatment steps before use. Unlike lipid-based pulmonary surfactants, the dynamic surface activity characteristics of polymer pulmonary surfactants are not impaired even in the presence of competing surfactant proteins. In preliminary animal experiments, intratracheally administered polymer pulmonary surfactants were found to be tolerable (excretable from the body) without damaging the major organs of mice, and a dose-dependent destiffening effect on the lung compliance characteristics of fetal rabbits could be produced. Further research is warranted to optimize the preparation for maximum therapeutic effect and to evaluate in detail the short- and long-term toxicity of the materials.

[0078] Experimental procedure

[0079] Synthesis of PLGA-PEG and PLGACL-PEG

[0080] The PLGA-PEG material and the PLGACL-PEG material were synthesized by ring-opening polymerization using a tin catalyst Purified poly(ethylene glycol) monomethyl ether (PEG-OH , M n = 5,000 g / mol, Sigma Aldrich) was used as a macroinitiator, and tin( II) 2-ethylhexanoate (Sigma Aldrich) was used as a catalyst. The polymerization reaction was carried out at 1 30 °C. Before use, the D,L-lactide (Lactel) and glycolide (Sigma Al drich) monomers were recrystallized twice from toluene (Sigma Aldrich) and tetrahydrofuran (Sigma Aldric h). The ε-caprolactone (Sigma Aldrich) monomer was used as received . The synthesized PLGA-PEG and PLGACL-PEG products were precipitated in 2 -propanol (Sigma Aldrich) and dried under vacuum before storage at the use / refrigeration temperature .

[0081] Synthesis of PtBMA-PEG and PS-PEG

[0082] The PtBMA-PEG material and the PS-PEG material were synthesized by reversible addition-fragmentation chain transfer (RA FT) polymerization. 4-Cyano-4-[(dodecylsulfanylthiocarbonyl ) sulfanyl] pentanoic acid (Sigma Aldrich) was used as a RAFT agent. First, by uteric esterification, the RAFT agent was purified poly(ethylene glycol) monometh Thiol ether (PEG-OH, M n = 5,000 g / mol, Sigma Aldrich) was conjugated to it 70 . PEG-OH (1 g, 0.2 mmol), RAFT agent (161.4 m g, 0.4 mmol) and 4-dimethylaminopyridine (Sigma Aldrich, 4.89 mg , 0.04 mmol) were mixed in 10 ml of dichloromethane (Sigma Aldrich) and kept at 0 °C under magnetic stirring. Separately prepared dicyclohexylcarbodiimide (82.5 mg, 0.4 mmol) in dichloromethane (5 ml) was added dropwise to the above mixture, and the reaction was carried out at 0 °C for 5 minutes , then at 20 °C for 3 hours to prepare "PEG-RAFT". The as-synthesized PEG-RAFT product was first filtered through filter paper to remove insoluble urea by-products, and then precipitated twice in hexane for further purification. The RAFT polymerization reaction was carried out at 70 °C by mixing PEG-RAFT, styrene (Sigma-Aldrich) or tert-butyl methacrylate (Sigma- Aldrich) without inhibitor and azobisisobutyronitrile (Sigma-Aldri ch), a free radical initiator, in dioxane (Sigma Aldrich). The obtained PtBMA-PEG and PS-PEG products were precipitated twice in hexane and dried under vacuum . The synthesis procedure was adopted from the following reference Benaglia, M. et al. Universal (Switchable) RAFT Agents. Journal of the American Chemical Society 131, 6914-6915, doi:10.1021 / ja901955n (2009) Analysis of the polymer

[0083] The number average molecular weight (M n ) of the polymer was determined by using a Bruker ARX NMR spectrometer (500 MHz) 1 with 1H NMR spectroscopy. 1 For 1H NMR measurements, the polymer sample was prepared at a polymer concentration of 5 wt% in deuterated chloroform. Also, the polydispersity index (PDI) of the polymer was measured by size exclusion chromatography (SEC) using an Agilent Technologies 12000 series instrument equipped with a refractive index detector and three 5 μm PLgel MIXED-C columns. Tetrahydrofuran was used as the mobile phase (kept at 35 °C and flowed at a rate of 1 ml / min). Calibration was performed using polystyrene standards (Agilent Easi Cal). equipped with a refractive index detector and three 5 μm PLgel MIXED-C columns. The surface tension-area isotherms of Survanta and the polymeric surfactant were measured using a KSV 5000 Langmuir trough (51 cm × 15 cm) with symmetric double barriers. The total surface area of the trough was 780 cm and the volume of the subphase was 1.3 L. For surface tension measurements, a filter paper Wilhemly probe was used. Before each measurement, the trough and barriers were washed three times with ethanol and MilliQ pure water. The water surface was also aspirated to remove surface active contaminants. When the water surface was completely clean, the reference value of the surface tension did not change during the blank compression test. The lung surfactant sample was spread on the water surface using a Hamilton microsyringe, i.e., by forming a microliter-sized droplet at the tip of the syringe needle and bringing it into contact with the water surface. Calibration was performed using polystyrene standards (Agilent Easi Cal).

[0084] Surface tension - area isotherm

[0085] A KSV 5000 Langmuir trough (51 cm × 15 cm) with symmetric double barriers was used to measure the surface tension - area isotherms of Survanta and the polymeric surfactant. The total surface area of the trough was 780 cm and the volume of the subphase was 1.3 L. 2 For surface tension measurements, a filter paper Wilhemly probe was used. Before each measurement, the trough and barriers were washed three times with ethanol and MilliQ pure water. The water surface was also aspirated to remove surface active contaminants. When the water surface was completely clean, the reference value of the surface tension did not change during the blank compression test. The lung surfactant sample was spread on the water surface using a Hamilton microsyringe, i.e., by forming a microliter-sized droplet at the tip of the syringe needle and bringing it into contact with the water surface. Preparation of polymer micelles Preparation of polymer micelles

[0086] Spherical polymer micelles were prepared using the solvent exchange method. First, 200 mg of the polymer was dissolved in 4 ml of acetone (Sigma Aldrich). Next, 36 ml of Milli-Q pure water (resistivity 18 MΩ·cm) was added dropwise to the polymer solution at a rate of 0.05 ml / min using a syringe pump, and the mixture was continuously stirred vigorously for 24 hours. To remove acetone, the solution was transferred to a dialysis bag (Spectra / Por 7, molecular weight cut-off 50 kDa) and dialyzed against 1 L of Milli-Q pure water for 3 hours. The reservoir was replaced with fresh Milli-Q water every hour. The micelle preparation procedure was adopted from the following reference.

[0087] Zhang, L. & Eisenberg, A. Multiple Morphologies and Characteristics of "Crew-Cut" Micelle-like Aggregates of Polystyrene-b-poly(acrylic acid) Diblock Copolymers in Aqueous Solutions. Journal of the American Chemical Society 118, 3168-3181, doi: 10.1021 / ja953709s (1996) Analysis of Polymer Micelles

[0088] The hydrodynamic diameter of the block copolymer micelles was measured at 25 °C by dynamic light scattering (DLS) using a Brookhaven ZetaPALS instrument. The scattering intensity was measured at a scattering angle of 90° using a 659 nm laser. The hydrodynamic diameter was determined from the measured diffusion coefficient using the Stokes-Einstein equation. For DLS measurements, the sample was diluted to ensure single scattering and filtered through a 0.2 μm syringe filter to remove contaminants.

[0089] The polymer micelles were imaged using a transmission electron microscope (TEM). Preparation of the TEM specimen was carried out by placing 20 μl of a polymer micelle solution at 0.01 - 0.05 mg / ml on a carbon-coated copper TEM grid (hydrophobically treated using an O2 plasma cleaner). 10 μl of a 2% uranyl acetate solution was added to the sample solution already on the TEM grid, and the mixture was blotted using filter paper and dried. The sample thus prepared was imaged using a FEI Tecnai 20 TEM instrument at 200 kV. These TEM images were analyzed using Gatan Digital Micrograph software. NMR Spin Relaxation Measurements

[0090] NMR spin relaxation measurements were performed using a Bruker Avance-III-800 spectrometer equipped with a sample temperature control unit. Micelle samples of PLGA-PEG and PS-PEG were prepared using the solvent exchange method (described above) using D2O as the final solvent (instead of H2O). PEG homopolymer samples were prepared by directly dissolving PEG in D2O. For all samples, the polymer concentration was set at 0.5 wt%. An inversion recovery sequence was used for T1 relaxation measurements, and a Carr-Purcell-Meiboom-Gill (CPMG) pulse sequence was used for T2 relaxation measurements. Nonlinear least squares regression was used to fit the data to a first- or second-order exponential decay function. Preliminary Evaluation of the Tolerance of Intratracheally Injected Polymeric Lung Surfactant in Adult Mice

[0091] In this study, 7-week-old female BALB / C mice (purchased from Jackson Laboratory) were used. Before injecting the polymeric lung surfactant, the mice were anesthetized by injecting 80 μl of 0.75 mg / ml Prevail (VetOne) behind the neck. Next, the mice were anesthetized with isoflurane at a flow rate of 2 l / min. Once the mice were completely sedated, a tracheotomy was performed and the polymeric lung surfactant solution was injected. Per kilogram of the mouse's body weight, the polymer Three doses of -0.64 mg, 6.5 mg, and 64 mg were tested. The liquid injection volume was maintained at 20 μl, corresponding to 1.07 ml of liquid per 1 kg of body weight. After polymer injection, the mice were sutured, and GLUture (Abbott Laboratories) was applied to the surgical site. Experienced animal technicians monitored the mice for signs of toxicity (weight loss and behavioral symptoms). After 14 days, the mice were sacrificed, major organs were collected, perfused with PBS, and fixed in 10% formalin. To prepare lung tissue, an additional step of inflating the excised lung with 10% formalin before fixation was added. For microscopic examination, the collected organ tissues were sliced and stained with H&E.

[0092] The pressure-volume mechanics of fetal rabbit lungs after administration of polymer lung surfactant The PV mechanics of fetal rabbit lungs on gestational day 27 of New Zealand white rabbits were tested ex vivo after administration of polymer lung surfactant. The fetuses were obtained by cesarean section. Polymer lung surfactant (or Newfacten) was injected into the fetal rabbit lungs by a single intratracheal instillation of 1.5 ml of liquid per 1 kg of body weight. The doses of polymer lung surfactant tested were 6 mg, 60 mg, and 96 mg of polymer per 1 kg of body weight, and the dose of Newfacten was 60 mg / kg. After instillation of the lung surfactant, a 10-minute wait was allowed before PV analysis.

[0093] Based on the above description, we now have neonatal, acute, or adult respiratory distress syndrome, including those caused by a deficiency of functional lung surfactant in mammals, including humans.

[0094] The PV mechanics of fetal rabbit lungs on gestational day 27 of New Zealand white rabbits were tested ex vivo after administration of polymer lung surfactant. The fetuses were obtained by cesarean section. Polymer lung surfactant (or Newfacten) was injected into the fetal rabbit lungs by a single intratracheal instillation of 1.5 ml of liquid per 1 kg of body weight. The doses of polymer lung surfactant tested were 6 mg, 60 mg, and 96 mg of polymer per 1 kg of body weight, and the dose of Newfacten was 60 mg / kg. After instillation of the lung surfactant, a 10-minute wait was allowed before PV analysis. (or Newfacten) was injected into the fetal rabbit lungs by a single intratracheal instillation of 1.5 ml of liquid per 1 kg of body weight. The doses of polymer lung surfactant tested were 6 mg, 60 mg, and 96 mg of polymer per 1 kg of body weight, and the dose of Newfacten was 60 mg / kg. After instillation of the lung surfactant, a 10-minute wait was allowed before PV analysis. The doses of polymer lung surfactant tested were 6 mg, 60 mg, and 96 mg of polymer per 1 kg of body weight, and the dose of Newfacten was 60 mg / kg. After instillation of the lung surfactant, a 10-minute wait was allowed before PV analysis.

[0095] Based on the above description, we now have neonatal, acute, or adult respiratory distress syndrome, including those caused by a deficiency of functional lung surfactant in mammals, including humans. Neonatal, acute, or adult respiratory distress syndrome, including those caused by a deficiency of functional lung surfactant in mammals, including humans. ​​A method for treating a lung disease can be disclosed. This method involves administering to an animal or human subject a therapeutically effective amount of a polymeric lung surfactant composition. This polymeric lung surfactant composition contains an effective amount of a biocompatible or biodegradable amphiphilic synthetic homopolymer or copolymer, and the monomers of this homopolymer or copolymer are ethylene glycol (EG), ethylene oxide (EO), vinyl alcohol (VA), alkyl oxazoline (AO), D,L-lactic acid or D,L-lactide (LA), glycolic acid or glycolide (GA), ε-caprolactone (CL), styrene (PS), alkyl methacrylate (AMA), alkyl acrylate (AA), and are selected from the group consisting of these.

[0096] Another method for treating lung diseases caused by a deficiency of functional lung surfactant in mammals including humans, including neonatal, acute or adult respiratory distress syndrome, this method involves administering synthetic blocks to an animal or human subject as a single therapeutic agent or in combination with other therapeutic agents.

[0097] Another method for treating lung diseases caused by a deficiency of functional lung surfactant in mammals including humans, including neonatal, acute or adult respiratory distress syndrome, is to administer a synthetic random copolymer to an animal or human subject as a single therapeutic agent or in combination with other therapeutic agents. This includes administering to an animal or human subject.

[0098] Another method for treating lung diseases caused by a deficiency of functional lung surfactant in mammals including humans, including neonatal, acute or adult respiratory distress syndrome, this The method involves administering a synthetic homopolymer, either as a single therapeutic agent or in combination with other therapeutic agents, to an animal or human subject for administration. including administration to an animal or human subject for administration.

[0099] Another method for treating lung diseases, including neonatal, acute, or adult respiratory distress syndrome caused by a deficiency of functional lung surfactant in mammals including humans, the method comprising administering to an animal or human subject a polymeric lung surfactant composition comprising a poly(styrene-block-ethylene glycol) (PS-PEG) block copolymer.

[0100] Another method for treating lung diseases, including neonatal, acute, or adult respiratory distress syndrome caused by a deficiency of functional lung surfactant in mammals including humans, the method comprising administering to an animal or human subject a polymeric lung surfactant composition comprising a poly(tert-butyl methacrylate-block-ethylene glycol) (PtBMA-PEG) block copolymer.

[0101] Another method for treating lung diseases, including neonatal, acute, or adult respiratory distress syndrome caused by a deficiency of functional lung surfactant in mammals including humans, the method comprising administering to an animal or human subject a polymeric lung surfactant composition comprising a poly(D,L-lactic acid-block-ethylene glycol) (PLA-PEG) block copolymer.

[0102] A method for treating lung diseases, including neonatal, acute, or adult respiratory distress syndrome caused by a deficiency of functional lung surfactant in mammals including humans, the method comprising administering a polymeric lung surfactant composition to an animal or human subject for administration by intratracheal instillation in the form of an aqueous solution to the lung of the animal or human subject for administration.

[0103] A method for treating lung diseases including neonatal, acute or adult respiratory distress syndrome caused by a lack of functional lung surfactant in mammals including humans, the method comprising administering a polymeric lung surfactant composition to a subject animal or human in the form of droplets or a dry powder type aerosol to the lungs of the patient while continuously applying airway positive pressure.

[0104] Exemplary polymeric lung surfactant surfactant compositions used in the above treatment method have a preparation comprising about 0.02 to 20% by weight of an amphiphilic block copolymer dispersed in micellar form in physiological saline, and the amphiphilic block copolymer compound has a hydrophilic block (such as PEG) with an average molecular weight in the range of about 50 Da to about 500 kDa and a hydrophobic block (such as PS) with an average molecular weight in the range of about 50 Da to about 500 kDa.

[0105] Note that any of the above methods can be used for the treatment of neonatal, acute or adult respiratory distress syndrome. Furthermore, these methods can also be used for the treatment of bronchopulmonary dysplasia.

[0106] Those skilled in the art will recognize that numerous modifications are possible to the above specific embodiments. These embodiments should not be limited to the specific limitations described. Other embodiments may be possible. Furthermore, several publications related to the present disclosure are listed below and cited herein. The entire contents of these references are incorporated herein by reference. [References] JPEG2025102844000008.jpg205156 JPEG2025102844000009.jpg196163 JPEG2025102844000010.jpg205163 JPEG2025102844000011.jpg215163 JPEG2025102844000012.jpg206163 JPEG2025102844000013.jpg206163 JPEG2025102844000014.jpg206163 JPEG2025102844000015.jpg205156 JPEG2025102844000016.jpg205156 JPEG2025102844000017.jpg63156

Claims

**Claim 1** A method for treating a lung disease of a mammal caused by a deficiency of functional lung surfactant in the mammal, comprising administering to the mammal a therapeutically effective amount of a polymeric lung surfactant composition. **Claim 2** The method according to claim 1, wherein the lung disease is acute respiratory distress syndrome in an infant or an adult caused by a deficiency of functional lung surfactant. **Claim 3** The method according to claim 1, wherein the polymeric lung surfactant composition comprises a synthetic biocompatible amphiphilic homopolymer or copolymer, and the monomer of the homopolymer or copolymer is selected from the group consisting of ethylene glycol (EG), ethylene oxide (EO), vinyl alcohol (VA), alkyl oxazoline (AO), D,L-lactic acid or D,L-lactide (LA), glycolic acid or glycolide (GA), ε-caprolactone (CL), styrene (PS), alkyl methacrylate (AMA), and alkyl acrylate (AA). **Claim 4** The method according to claim 1, wherein the polymeric lung surfactant composition comprises a synthetic block copolymer, random copolymer or homopolymer. **Claim 5** The method according to claim 4, wherein the polymeric lung surfactant composition comprises a poly(styrene-block-ethylene glycol) (PS-PEG) block copolymer. **Claim 6** The method according to claim 4, wherein the polymeric lung surfactant composition comprises a poly(tert-butyl methacrylate-block-ethylene glycol) (PtBMA-PEG) block copolymer. **Claim 7** The method according to claim 4, wherein the polymeric lung surfactant composition comprises a poly(D,L-lactic acid-block-ethylene glycol) (PLA-PEG) block copolymer. **Claim 8** A method for treating a lung disease of a mammal caused by a deficiency of functional lung surfactant in the mammal, comprising administering to the mammal, in the form of an aqueous solution, a therapeutically effective amount of a polymeric lung surfactant composition by intratracheal instillation to the lung of the mammal. **Claim 9** A method for treating a lung disease of a mammal caused by a deficiency of functional lung surfactant in the mammal, comprising ​ ​ A method comprising administering to the mammal a therapeutically effective amount of a polymeric lung surfactant composition in the form of droplets or dry powder type aerosols to the lung of the mammal by continuously applying positive airway pressure.

10. A method for treating a lung disease of a mammal caused by a lack of functional lung surfactant in the mammal, comprising administering to the mammal a therapeutically effective amount of a polymeric lung surfactant composition in combination with other therapeutic agents, wherein the lung disease includes acute respiratory distress syndrome in infants or adults.

11. A polymeric lung surfactant composition comprising a biocompatible or biodegradable amphiphilic synthetic polymer or copolymer, wherein the monomers of the polymer or copolymer are ethylene glycol (EG), ethylene oxide (EO), vinyl alcohol (VA), a lkyl oxazoline (AO), D,L-lactic acid or D,L-lactide (LA), glycol lic acid or glycolide (GA), ε-caprolactone (CL), styrene (PS), a lkyl methacrylate (AMA), alkyl acrylate (AA), selected from the group consisting of A polymeric lung surfactant composition.

12. The polymer according to claim 11, wherein the polymer is a poly(styrene-block-ethylene glycol) (PS-PEG) block copolymer.

13. The polymer according to claim 11, wherein the polymer is a poly(tert-butyl methacrylate-block-ethylene glycol) ( PtBMA-PEG) block copolymer.

14. The polymer according to claim 11, wherein the polymer is a poly(D,L-lactic acid-block-ethylene glycol) (PLA-PEG ) block copolymer.

15. A composition used as a lung surfactant, comprising about 0.02 to 20% by weight of an amphiphilic block copolymer dispersed in micelle form in an aqueous solution or physiological saline.

16. The composition according to claim 15, wherein the amphiphilic polymer comprises a plurality of hydrophilic blocks having an average molecular weight in the range of about 50 Da to 500 kDa and a plurality of hydrophobic blocks having an average molecular weight in the range of about 50 Da to 500 kDa, and.

17. The hydrophilic block is PEG (polyethylene glycol), and the hydrophobic block The composition according to claim 16, wherein K is PS (polystyrene).

18. The composition according to claim 15, wherein the amphiphilic polymer is poly(D,L-lactic acid-block-ethylene glycol) (PLA -PEG) block copolymer.

19. The composition according to claim 15, wherein the amphiphilic polymer is poly(tert-butyl methacrylate-block-ethylene glycol -l) (PtBMA-PEG) block copolymer.

20. The composition according to claim 15, wherein the amphiphilic polymer is a polymer or copolymer, and the monomer of the polymer or copolymer is selected from the group consisting of ethylene glycol (EG), ethylene oxide (EO), vinyl alcohol (VA), alkyloxazoline (AO), D,L-lactic acid or D,L- lactide (LA), glycolic acid or glycolide (GA), ε-caprolactone (C L), styrene (PS), alkyl methacrylate (AMA), alkyl acrylate ( AA). The composition according to claim 15.

21. A polymeric lung surfactant preparation containing about 0.02 to 20% by weight of an amphiphilic block copolymer dispersed in micellar form in an aqueous solution or physiological saline, wherein the amphiphilic block copolymer compound comprises a plurality of hydrophilic blocks (such as PEG) with an average molecular weight in the range of about 50 Da to about 500 kDa, and a plurality of hydrophobic blocks (such as PS) with an average molecular weight in the range of about 50 Da to about 500 kDa. A polymeric lung surfactant preparation.

Images