Lung Biology in Health & Disease Volume 221 Inhalation Aerosols: Physical and Biological Basis for Therapy, Second Edition

Inhalation Aerosols DK604x_C000a.indd 1 09/26/2006 8:04:50 AM LUNG BIOLOGY IN HEALTH AND DISEASE Executive Editor ...

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Inhalation Aerosols

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LUNG BIOLOGY IN HEALTH AND DISEASE

Executive Editor Claude Lenfant Former Director, National Heart, Lung, and Blood Institute National Institutes of Health Bethesda, Maryland

1. Immunologic and Infectious Reactions in the Lung, edited by C. H. Kirkpatrick and H. Y. Reynolds 2. The Biochemical Basis of Pulmonary Function, edited by R. G. Crystal 3. Bioengineering Aspects of the Lung, edited by J. B. West 4. Metabolic Functions of the Lung, edited by Y. S. Bakhle and J. R. Vane 5. Respiratory Defense Mechanisms (in two parts), edited by J. D. Brain, D. F. Proctor, and L. M. Reid 6. Development of the Lung, edited by W. A. Hodson 7. Lung Water and Solute Exchange, edited by N. C. Staub 8. Extrapulmonary Manifestations of Respiratory Disease, edited by E. D. Robin 9. Chronic Obstructive Pulmonary Disease, edited by T. L. Petty 10. Pathogenesis and Therapy of Lung Cancer, edited by C. C. Harris 11. Genetic Determinants of Pulmonary Disease, edited by S. D. Litwin 12. The Lung in the Transition Between Health and Disease, edited by P. T. Macklem and S. Permutt 13. Evolution of Respiratory Processes: A Comparative Approach, edited by S. C. Wood and C. Lenfant 14. Pulmonary Vascular Diseases, edited by K. M. Moser 15. Physiology and Pharmacology of the Airways, edited by J. A. Nadel 16. Diagnostic Techniques in Pulmonary Disease (in two parts), edited by M. A. Sackner 17. Regulation of Breathing (in two parts), edited by T. F. Hornbein 18. Occupational Lung Diseases: Research Approaches and Methods, edited by H. Weill and M. Turner-Warwick 19. Immunopharmacology of the Lung, edited by H. H. Newball 20. Sarcoidosis and Other Granulomatous Diseases of the Lung, edited by B. L. Fanburg

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21. Sleep and Breathing, edited by N. A. Saunders and C. E. Sullivan 22. Pneumocystis carinii Pneumonia: Pathogenesis, Diagnosis, and Treatment, edited by L. S. Young 23. Pulmonary Nuclear Medicine: Techniques in Diagnosis of Lung Disease, edited by H. L. Atkins 24. Acute Respiratory Failure, edited by W. M. Zapol and K. J. Falke 25. Gas Mixing and Distribution in the Lung, edited by L. A. Engel and M. Paiva 26. High-Frequency Ventilation in Intensive Care and During Surgery, edited by G. Carlon and W. S. Howland 27. Pulmonary Development: Transition from Intrauterine to Extrauterine Life, edited by G. H. Nelson 28. Chronic Obstructive Pulmonary Disease: Second Edition, edited by T. L. Petty 29. The Thorax (in two parts), edited by C. Roussos and P. T. Macklem 30. The Pleura in Health and Disease, edited by J. Chrétien, J. Bignon, and A. Hirsch 31. Drug Therapy for Asthma: Research and Clinical Practice, edited by J. W. Jenne and S. Murphy 32. Pulmonary Endothelium in Health and Disease, edited by U. S. Ryan 33. The Airways: Neural Control in Health and Disease, edited by M. A. Kaliner and P. J. Barnes 34. Pathophysiology and Treatment of Inhalation Injuries, edited by J. Loke 35. Respiratory Function of the Upper Airway, edited by O. P. Mathew and G. Sant’Ambrogio 36. Chronic Obstructive Pulmonary Disease: A Behavioral Perspective, edited by A. J. McSweeny and I. Grant 37. Biology of Lung Cancer: Diagnosis and Treatment, edited by S. T. Rosen, J. L. Mulshine, F. Cuttitta, and P. G. Abrams 38. Pulmonary Vascular Physiology and Pathophysiology, edited by E. K. Weir and J. T. Reeves 39. Comparative Pulmonary Physiology: Current Concepts, edited by S. C. Wood 40. Respiratory Physiology: An Analytical Approach, edited by H. K. Chang and M. Paiva 41. Lung Cell Biology, edited by D. Massaro 42. Heart–Lung Interactions in Health and Disease, edited by S. M. Scharf and S. S. Cassidy 43. Clinical Epidemiology of Chronic Obstructive Pulmonary Disease, edited by M. J. Hensley and N. A. Saunders 44. Surgical Pathology of Lung Neoplasms, edited by A. M. Marchevsky 45. The Lung in Rheumatic Diseases, edited by G. W. Cannon and G. A. Zimmerman

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46. Diagnostic Imaging of the Lung, edited by C. E. Putman 47. Models of Lung Disease: Microscopy and Structural Methods, edited by J. Gil 48. Electron Microscopy of the Lung, edited by D. E. Schraufnagel 49. Asthma: Its Pathology and Treatment, edited by M. A. Kaliner, P. J. Barnes, and C. G. A. Persson 50. Acute Respiratory Failure: Second Edition, edited by W. M. Zapol and F. Lemaire 51. Lung Disease in the Tropics, edited by O. P. Sharma 52. Exercise: Pulmonary Physiology and Pathophysiology, edited by B. J. Whipp and K. Wasserman 53. Developmental Neurobiology of Breathing, edited by G. G. Haddad and J. P. Farber 54. Mediators of Pulmonary Inflammation, edited by M. A. Bray and W. H. Anderson 55. The Airway Epithelium, edited by S. G. Farmer and D. Hay 56. Physiological Adaptations in Vertebrates: Respiration, Circulation, and Metabolism, edited by S. C. Wood, R. E. Weber, A. R. Hargens, and R. W. Millard 57. The Bronchial Circulation, edited by J. Butler 58. Lung Cancer Differentiation: Implications for Diagnosis and Treatment, edited by S. D. Bernal and P. J. Hesketh 59. Pulmonary Complications of Systemic Disease, edited by J. F. Murray 60. Lung Vascular Injury: Molecular and Cellular Response, edited by A. Johnson and T. J. Ferro 61. Cytokines of the Lung, edited by J. Kelley 62. The Mast Cell in Health and Disease, edited by M. A. Kaliner and D. D. Metcalfe 63. Pulmonary Disease in the Elderly Patient, edited by D. A. Mahler 64. Cystic Fibrosis, edited by P. B. Davis 65. Signal Transduction in Lung Cells, edited by J. S. Brody, D. M. Center, and V. A. Tkachuk 66. Tuberculosis: A Comprehensive International Approach, edited by L. B. Reichman and E. S. Hershfield 67. Pharmacology of the Respiratory Tract: Experimental and Clinical Research, edited by K. F. Chung and P. J. Barnes 68. Prevention of Respiratory Diseases, edited by A. Hirsch, M. Goldberg, J.-P. Martin, and R. Masse 69. Pneumocystis carinii Pneumonia: Second Edition, edited by P. D. Walzer 70. Fluid and Solute Transport in the Airspaces of the Lungs, edited by R. M. Effros and H. K. Chang 71. Sleep and Breathing: Second Edition, edited by N. A. Saunders and C. E. Sullivan 72. Airway Secretion: Physiological Bases for the Control of Mucous Hypersecretion, edited by T. Takishima and S. Shimura

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73. Sarcoidosis and Other Granulomatous Disorders, edited by D. G. James 74. Epidemiology of Lung Cancer, edited by J. M. Samet 75. Pulmonary Embolism, edited by M. Morpurgo 76. Sports and Exercise Medicine, edited by S. C. Wood and R. C. Roach 77. Endotoxin and the Lungs, edited by K. L. Brigham 78. The Mesothelial Cell and Mesothelioma, edited by M.-C. Jaurand and J. Bignon 79. Regulation of Breathing: Second Edition, edited by J. A. Dempsey and A. I. Pack 80. Pulmonary Fibrosis, edited by S. Hin. Phan and R. S. Thrall 81. Long-Term Oxygen Therapy: Scientific Basis and Clinical Application, edited by W. J. O’Donohue, Jr. 82. Ventral Brainstem Mechanisms and Control of Respiration and Blood Pressure, edited by C. O. Trouth, R. M. Millis, H. F. Kiwull-Schöne, and M. E. Schläfke 83. A History of Breathing Physiology, edited by D. F. Proctor 84. Surfactant Therapy for Lung Disease, edited by B. Robertson and H. W. Taeusch 85. The Thorax: Second Edition, Revised and Expanded (in three parts), edited by C. Roussos 86. Severe Asthma: Pathogenesis and Clinical Management, edited by S. J. Szefler and D. Y. M. Leung 87. Mycobacterium avium–Complex Infection: Progress in Research and Treatment, edited by J. A. Korvick and C. A. Benson 88. Alpha 1–Antitrypsin Deficiency: Biology • Pathogenesis • Clinical Manifestations • Therapy, edited by R. G. Crystal 89. Adhesion Molecules and the Lung, edited by P. A. Ward and J. C. Fantone 90. Respiratory Sensation, edited by L. Adams and A. Guz 91. Pulmonary Rehabilitation, edited by A. P. Fishman 92. Acute Respiratory Failure in Chronic Obstructive Pulmonary Disease, edited by J.-P. Derenne, W. A. Whitelaw, and T. Similowski 93. Environmental Impact on the Airways: From Injury to Repair, edited by J. Chrétien and D. Dusser 94. Inhalation Aerosols: Physical and Biological Basis for Therapy, edited by A. J. Hickey 95. Tissue Oxygen Deprivation: From Molecular to Integrated Function, edited by G. G. Haddad and G. Lister 96. The Genetics of Asthma, edited by S. B. Liggett and D. A. Meyers 97. Inhaled Glucocorticoids in Asthma: Mechanisms and Clinical Actions, edited by R. P. Schleimer, W. W. Busse, and P. M. O’Byrne 98. Nitric Oxide and the Lung, edited by W. M. Zapol and K. D. Bloch

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99. Primary Pulmonary Hypertension, edited by L. J. Rubin and S. Rich 100. Lung Growth and Development, edited by J. A. McDonald 101. Parasitic Lung Diseases, edited by A. A. F. Mahmoud 102. Lung Macrophages and Dendritic Cells in Health and Disease, edited by M. F. Lipscomb and S. W. Russell 103. Pulmonary and Cardiac Imaging, edited by C. Chiles and C. E. Putman 104. Gene Therapy for Diseases of the Lung, edited by K. L. Brigham 105. Oxygen, Gene Expression, and Cellular Function, edited by L. Biadasz Clerch and D. J. Massaro 106. Beta2-Agonists in Asthma Treatment, edited by R. Pauwels and P. M. O’Byrne 107. Inhalation Delivery of Therapeutic Peptides and Proteins, edited by A. L. Adjei and P. K. Gupta 108. Asthma in the Elderly, edited by R. A. Barbee and J. W. Bloom 109. Treatment of the Hospitalized Cystic Fibrosis Patient, edited by D. M. Orenstein and R. C. Stern 110. Asthma and Immunological Diseases in Pregnancy and Early Infancy, edited by M. Schatz, R. S. Zeiger, and H. N. Claman 111. Dyspnea, edited by D. A. Mahler 112. Proinflammatory and Antiinflammatory Peptides, edited by S. I. Said 113. Self-Management of Asthma, edited by H. Kotses and A. Harver 114. Eicosanoids, Aspirin, and Asthma, edited by A. Szczeklik, R. J. Gryglewski, and J. R. Vane 115. Fatal Asthma, edited by A. L. Sheffer 116. Pulmonary Edema, edited by M. A. Matthay and D. H. Ingbar 117. Inflammatory Mechanisms in Asthma, edited by S. T. Holgate and W. W. Busse 118. Physiological Basis of Ventilatory Support, edited by J. J. Marini and A. S. Slutsky 119. Human Immunodeficiency Virus and the Lung, edited by M. J. Rosen and J. M. Beck 120. Five-Lipoxygenase Products in Asthma, edited by J. M. Drazen, S.-E. Dahlén, and T. H. Lee 121. Complexity in Structure and Function of the Lung, edited by M. P. Hlastala and H. T. Robertson 122. Biology of Lung Cancer, edited by M. A. Kane and P. A. Bunn, Jr. 123. Rhinitis: Mechanisms and Management, edited by R. M. Naclerio, S. R. Durham, and N. Mygind 124. Lung Tumors: Fundamental Biology and Clinical Management, edited by C. Brambilla and E. Brambilla 125. Interleukin-5: From Molecule to Drug Target for Asthma, edited by C. J. Sanderson 126. Pediatric Asthma, edited by S. Murphy and H. W. Kelly

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127. Viral Infections of the Respiratory Tract, edited by R. Dolin and P. F. Wright 128. Air Pollutants and the Respiratory Tract, edited by D. L. Swift and W. M. Foster 129. Gastroesophageal Reflux Disease and Airway Disease, edited by M. R. Stein 130. Exercise-Induced Asthma, edited by E. R. McFadden, Jr. 131. LAM and Other Diseases Characterized by Smooth Muscle Proliferation, edited by J. Moss 132. The Lung at Depth, edited by C. E. G. Lundgren and J. N. Miller 133. Regulation of Sleep and Circadian Rhythms, edited by F. W. Turek and P. C. Zee 134. Anticholinergic Agents in the Upper and Lower Airways, edited by S. L. Spector 135. Control of Breathing in Health and Disease, edited by M. D. Altose and Y. Kawakami 136. Immunotherapy in Asthma, edited by J. Bousquet and H. Yssel 137. Chronic Lung Disease in Early Infancy, edited by R. D. Bland and J. J. Coalson 138. Asthma’s Impact on Society: The Social and Economic Burden, edited by K. B. Weiss, A. S. Buist, and S. D. Sullivan 139. New and Exploratory Therapeutic Agents for Asthma, edited by M. Yeadon and Z. Diamant 140. Multimodality Treatment of Lung Cancer, edited by A. T. Skarin 141. Cytokines in Pulmonary Disease: Infection and Inflammation, edited by S. Nelson and T. R. Martin 142. Diagnostic Pulmonary Pathology, edited by P. T. Cagle 143. Particle–Lung Interactions, edited by P. Gehr and J. Heyder 144. Tuberculosis: A Comprehensive International Approach, Second Edition, Revised and Expanded, edited by L. B. Reichman and E. S. Hershfield 145. Combination Therapy for Asthma and Chronic Obstructive Pulmonary Disease, edited by R. J. Martin and M. Kraft 146. Sleep Apnea: Implications in Cardiovascular and Cerebrovascular Disease, edited by T. D. Bradley and J. S. Floras 147. Sleep and Breathing in Children: A Developmental Approach, edited by G. M. Loughlin, J. L. Carroll, and C. L. Marcus 148. Pulmonary and Peripheral Gas Exchange in Health and Disease, edited by J. Roca, R. Rodriguez-Roisen, and P. D. Wagner 149. Lung Surfactants: Basic Science and Clinical Applications, R. H. Notter 150. Nosocomial Pneumonia, edited by W. R. Jarvis 151. Fetal Origins of Cardiovascular and Lung Disease, edited by David J. P. Barker

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152. Long-Term Mechanical Ventilation, edited by N. S. Hill 153. Environmental Asthma, edited by R. K. Bush 154. Asthma and Respiratory Infections, edited by D. P. Skoner 155. Airway Remodeling, edited by P. H. Howarth, J. W. Wilson, J. Bousquet, S. Rak, and R. A. Pauwels 156. Genetic Models in Cardiorespiratory Biology, edited by G. G. Haddad and T. Xu 157. Respiratory-Circulatory Interactions in Health and Disease, edited by S. M. Scharf, M. R. Pinsky, and S. Magder 158. Ventilator Management Strategies for Critical Care, edited by N. S. Hill and M. M. Levy 159. Severe Asthma: Pathogenesis and Clinical Management, Second Edition, Revised and Expanded, edited by S. J. Szefler and D. Y. M. Leung 160. Gravity and the Lung: Lessons from Microgravity, edited by G. K. Prisk, M. Paiva, and J. B. West 161. High Altitude: An Exploration of Human Adaptation, edited by T. F. Hornbein and R. B. Schoene 162. Drug Delivery to the Lung, edited by H. Bisgaard, C. O’Callaghan, and G. C. Smaldone 163. Inhaled Steroids in Asthma: Optimizing Effects in the Airways, edited by R. P. Schleimer, P. M. O’Byrne, S. J. Szefler, and R. Brattsand 164. IgE and Anti-IgE Therapy in Asthma and Allergic Disease, edited by R. B. Fick, Jr., and P. M. Jardieu 165. Clinical Management of Chronic Obstructive Pulmonary Disease, edited by T. Similowski, W. A. Whitelaw, and J.-P. Derenne 166. Sleep Apnea: Pathogenesis, Diagnosis, and Treatment, edited by A. I. Pack 167. Biotherapeutic Approaches to Asthma, edited by J. Agosti and A. L. Sheffer 168. Proteoglycans in Lung Disease, edited by H. G. Garg, P. J. Roughley, and C. A. Hales 169. Gene Therapy in Lung Disease, edited by S. M. Albelda 170. Disease Markers in Exhaled Breath, edited by N. Marczin, S. A. Kharitonov, M. H. Yacoub, and P. J. Barnes 171. Sleep-Related Breathing Disorders: Experimental Models and Therapeutic Potential, edited by D. W. Carley and M. Radulovacki 172. Chemokines in the Lung, edited by R. M. Strieter, S. L. Kunkel, and T. J. Standiford 173. Respiratory Control and Disorders in the Newborn, edited by O. P. Mathew 174. The Immunological Basis of Asthma, edited by B. N. Lambrecht, H. C. Hoogsteden, and Z. Diamant 175. Oxygen Sensing: Responses and Adaptation to Hypoxia, edited by S. Lahiri, G. L. Semenza, and N. R. Prabhakar

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176. Non-Neoplastic Advanced Lung Disease, edited by J. R. Maurer 177. Therapeutic Targets in Airway Inflammation, edited by N. T. Eissa and D. P. Huston 178. Respiratory Infections in Allergy and Asthma, edited by S. L. Johnston and N. G. Papadopoulos 179. Acute Respiratory Distress Syndrome, edited by M. A. Matthay 180. Venous Thromboembolism, edited by J. E. Dalen 181. Upper and Lower Respiratory Disease, edited by J. Corren, A. Togias, and J. Bousquet 182. Pharmacotherapy in Chronic Obstructive Pulmonary Disease, edited by B. R. Celli 183. Acute Exacerbations of Chronic Obstructive Pulmonary Disease, edited by N. M. Siafakas, N. R. Anthonisen, and D. Georgopoulos 184. Lung Volume Reduction Surgery for Emphysema, edited by H. E. Fessler, J. J. Reilly, Jr., and D. J. Sugarbaker 185. Idiopathic Pulmonary Fibrosis, edited by J. P. Lynch III 186. Pleural Disease, edited by D. Bouros 187. Oxygen/Nitrogen Radicals: Lung Injury and Disease, edited by V. Vallyathan, V. Castranova, and X. Shi 188. Therapy for Mucus-Clearance Disorders, edited by B. K. Rubin and C. P. van der Schans 189. Interventional Pulmonary Medicine, edited by J. F. Beamis, Jr., P. N. Mathur, and A. C. Mehta 190. Lung Development and Regeneration, edited by D. J. Massaro, G. Massaro, and P. Chambon 191. Long-Term Intervention in Chronic Obstructive Pulmonary Disease, edited by R. Pauwels, D. S. Postma, and S. T. Weiss 192. Sleep Deprivation: Basic Science, Physiology, and Behavior, edited by Clete A. Kushida 193. Sleep Deprivation: Clinical Issues, Pharmacology, and Sleep Loss Effects, edited by Clete A. Kushida 194. Pneumocystis Pneumonia: Third Edition, Revised and Expanded, edited by P. D. Walzer and M. Cushion 195. Asthma Prevention, edited by William W. Busse and Robert F. Lemanske, Jr. 196. Lung Injury: Mechanisms, Pathophysiology, and Therapy, edited by Robert H. Notter, Jacob Finkelstein, and Bruce Holm 197. Ion Channels in the Pulmonary Vasculature, edited by Jason X.-J. Yuan 198. Chronic Obstuctive Pulmonary Disease: Cellular and Molecular Mechanisms, edited by Peter J. Barnes 199. Pediatric Nasal and Sinus Disorders, edited by Tania Sih and Peter A. R. Clement 200. Functional Lung Imaging, edited by David Lipson and Edwin van Beek 201. Lung Surfactant Function and Disorder, edited by Kaushik Nag

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202. Pharmacology and Pathophysiology of the Control of Breathing, edited by Denham S. Ward, Albert Dahan and Luc J. Teppema 203. Molecular Imaging of the Lungs, edited by Daniel Schuster and Timothy Blackwell 204. Air Pollutants and the Respiratory Tract: Second Edition, edited by W. Michael Foster and Daniel L. Costa 205. Acute and Chronic Cough, edited by Anthony E. Redington and Alyn H. Morice 206. Severe Pneumonia, edited by Michael S. Niederman 207. Monitoring Asthma, edited by Peter G. Gibson 208. Dyspnea: Mechanisms, Measurement, and Management, Second Edition, edited by Donald A. Mahler and Denis E. O'Donnell 209. Childhood Asthma, edited by Stanley J. Szefler and Søren Pedersen 210. Sarcoidosis, edited by Robert Baughman 211. Tropical Lung Disease, Second Edition, edited by Om Sharma 212. Pharmacotherapy of Asthma, edited by James T. Li 213. Practical Pulmonary and Critical Care Medicine: Respiratory Failure, edited by Zab Mosenifar and Guy W. Soo Hoo 214. Practical Pulmonary and Critical Care Medicine: Disease Management, edited by Zab Mosenifar and Guy W. Soo Hoo 215. Ventilator-Induced Lung Injury, edited by Didier Dreyfuss, Georges Saumon, and Rolf D. Hubmayr 216. Bronchial Vascular Remodeling In Asthma and COPD, edited by Aili Lazaar 217. Lung and Heart–Lung Transplantation, edited by Joseph P. Lynch III and David J. Ross 218. Genetics of Asthma and Chronic Obstructive Pulmonary Disease, edited by Dirkje S. Postma and Scott T. Weiss 219. Reichman and Hershfield's Tuberculosis: A Comprehensive, International Approach, Third Edition (in two parts), edited by Mario C. Raviglione 220. Narcolepsy and Hypersomnia, edited by Claudio Bassetti, Michel Billiard, and Emmanuel Mignot 221. Inhalation Aerosols: Physical and Biological Basis for Therapy, Second Edition, edited by Anthony J. Hickey The opinions expressed in these volumes do not necessarily represent the views of the National Institutes of Health.

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Inhalation Aerosols Physical and Biological Basis for Therapy

Second Edition

Edited by

Anthony J. Hickey

University of North Carolina Chapel Hill, North Carolina, U.S.A.

New York London

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Informa Healthcare USA, Inc. 270 Madison Avenue New York, NY 10016 © 2007 by Informa Healthcare USA, Inc. Informa Healthcare is an Informa business No claim to original U.S. Government works Printed in the United States of America on acid‑free paper 10 9 8 7 6 5 4 3 2 1 International Standard Book Number‑10: 0‑8493‑4160‑4 (Hardcover) International Standard Book Number‑13: 978‑0‑8493‑4160‑1 (Hardcover) This book contains information obtained from authentic and highly regarded sources. Reprinted material is quoted with permission, and sources are indicated. A wide variety of references are listed. Reasonable efforts have been made to publish reliable data and information, but the author and the publisher cannot assume responsibility for the validity of all materials or for the consequences of their use. No part of this book may be reprinted, reproduced, transmitted, or utilized in any form by any electronic, mechanical, or other means, now known or hereafter invented, including photocopying, microfilming, and recording, or in any information storage or retrieval system, without written permission from the publishers. For permission to photocopy or use material electronically from this work, please access www.copyright. com (http://www.copyright.com/) or contact the Copyright Clearance Center, Inc. (CCC) 222 Rosewood Drive, Danvers, MA 01923, 978‑750‑8400. CCC is a not‑for‑profit organization that provides licenses and registration for a variety of users. For organizations that have been granted a photocopy license by the CCC, a separate system of payment has been arranged. Trademark Notice: Product or corporate names may be trademarks or registered trademarks, and are used only for identification and explanation without intent to infringe. Visit the Informa Web site at www.informa.com and the Informa Healthcare Web site at www.informahealthcare.com

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Dedicated to

Richard Michael Evans and David Leslie Swift

Introduction

A field of medicine that has made astonishing advances during the last decade is that of inhalation aerosols, as will certainly be seen by the readership of this second edition of Inhalation Aerosols: Physical and Biological Basis for Therapy. The first edition, which appeared in 1996, ended with the very short section titled “Future Developments,” which predicted what could be coming from new developments in metered dose inhalers, dry powder inhalers, and nebulizers. Then, dry powder inhalers were in development—today, they are widely used. And so it goes for most of the topics discussed in the 1996 edition. The chapter titles and the contributors are the same for both volumes, but the texts are as new as the field! It is interesting to explore the several reasons why the field of inhalation aerosols progressed so quickly in relatively little time. First, the researchers and engineers focusing on the methodology and its application showed a great deal of foresight in anticipating what would work and in leading the effort toward better and more effective applications. When the field of inhalation aerosols began, it was almost all about the treatment of asthma. Today, it is used not only for a wide range of pulmonary diseases, but also for other systemic diseases, such as diabetes. At the same time, the pharmaceutical industry coupled its efforts with those of the engineers and the strong commitment of both disciplines led to the range of therapeutic aerosols that are now routinely used. But another element was needed—acceptance by the patients. Sure enough, patients were eager to accept new modes of drug administration, and the ability of these new technologies to deliver the medications to improve patient care stimulated the progress that we have witnessed. This volume may look the same as the 1996 edition, but it is not. It is the report of a medical journey paved with technological success that leads to better patient care. Dr. Anthony J. Hickey and his contributors are the architects of the progress that is reported.

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The series of monographs Lung Biology in Health and Disease is proud to present 10 years of most significant progress in the field of inhalation aerosols, and I am grateful to the editor and contributors to give the series the opportunity to present such advances. Claude Lenfant, MD Gaithersburg, Maryland, U.S.A.

Preface

A decade has passed since the original volume of this title was published. It has been the most eventful period in the history of inhaled drug products. During the 1990s, a large number of technologies were developed, which in turn resulted in the growth of specialized drug delivery companies. The international control of propellants and the growing biotechnology industry occurred in the shadow of the human genome project—the most significant scientific endeavor of this generation of scientists. The desire to improve drug delivery to and via the lungs was served by basic and applied science and engineering in a manner unsurpassed previously. Aerosol delivery of drugs, which was historically restricted to the treatment of asthma, has been extended to treat cystic fibrosis, chronic obstructive pulmonary disease, pulmonary infectious diseases, diabetes, and other diseases as yet under evaluation. It is clear that the available technology is capable of achieving controlled and targeted drug delivery for a range of diseases, but in many ways the technology has outstripped our fundamental understanding of the diseases we hope to treat. Which receptors are we targeting, where are they located, how and where are molecules absorbed, and what is the desired locally therapeutic dose? The precedent-setting Food and Drug Administration approval of the insulin product Exubera® as a dry powder aerosol dosage form occurring at about the same time that the agency announced its strategy for elimination of chlorofluorocarbon propellant–based products is a historical development predicted in the earlier volume but from which major new developments in the field can be anticipated. This volume was originally intended to link the mechanical and physicochemical properties of aerosols to the biology of their disposition. As an understanding of the lung biology in the context of drug delivery is of increasing importance, the need for a revised and updated version of this text is clear. Most of the chapters in the book cover areas that have seen significant developments in the last decade. Additional chapters have been added to address particular developments in specialized aerosol dosage forms.

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In the last decade, two of our most distinguished and well-respected colleagues, and contributors to the first edition of this book, have passed away. Dr. David Swift died very soon after publication of the first edition, and more recently the news of Dr. Richard Evans’ passing shocked the pharmaceutical aerosol sciences community. Both were remarkably young people and had much to look forward to, which makes their loss all the more poignant and leaves a large gap that would have been filled by their future scientific and technical contributions. Of the remaining original authors, most have continued to contribute to the field, and several leading researchers have been added to the list of contributors. I hope that our efforts have resulted in a volume that continues to serve as an introduction to those entering the field and a reference text for our colleagues and peers. Anthony J. Hickey

Contributors

Akwete L. Adjei

Kos Pharmaceuticals, Cranbury, New Jersey, U.S.A.

Ralph J. Altiere University of Colorado Health Sciences Center, Denver, Colorado, U.S.A. Hetal Amin University of Cincinnati College of Medicine, Cincinnati, Ohio, U.S.A. Bahman Asgharian CIIT Centers for Health Research, Research Triangle Park, North Carolina, U.S.A. Jonathan A. Bernstein University of Cincinnati College of Medicine, Cincinnati, Ohio, U.S.A. Meenakshi Bhat

Eli Lilly and Company, Indianapolis, Indiana, U.S.A.

Georges Caillibotte Air Liquide Research Center Claude-Delorme, Jouy en Josas, France Udayan Chokshi Chemical Engineering and Materials Sciences, Wayne State University, Detroit, Michigan, U.S.A. Timothy M. Crowder Oriel Therapeutics, Inc., Research Triangle Park, North Carolina, U.S.A. Sandro R. P. da Rocha Chemical Engineering and Materials Sciences, Wayne State University, Detroit, Michigan, U.S.A. Richard N. Dalby University of Maryland at Baltimore, Baltimore, Maryland, U.S.A. Richard M. Effros California, U.S.A.

Harbor–UCLA Medical Center, Torrance,

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Contributors

Richard M. Evans Inspire Pharmaceuticals, Durham, North Carolina, U.S.A. Bernard J. Greenspan California, U.S.A. Pramod K. Gupta

Versus Pharmaceuticals, Inc., San Diego,

Bausch & Lomb, Rochester, New York, U.S.A.

Anthony J. Hickey School of Pharmacy and Medicine, University of North Carolina, Chapel Hill, North Carolina, U.S.A. Keith A. Johnson Adroit Pharmaceutical Development, LLC, Durham, North Carolina, U.S.A. Ira Katz Air Liquide Research Center Claude-Delorme, Jouy en Josas, France, and Lafayette College, Easton, Pennsylvania, U.S.A. Julia S. Kimbell CIIT Centers for Health Research, Research Triangle Park, North Carolina, U.S.A. Joseph K. H. Ma West Virginia University, Morgantown, West Virginia, U.S.A. Ted Martonen CyberMedicine, Laguna Beach, California, and University of North Carolina, Chapel Hill, North Carolina, U.S.A. Ralph W. Niven

Innoven, Half Moon Bay, California, U.S.A.

Robson P. S. Peguin Chemical Engineering and Materials Sciences, Wayne State University, Detroit, Michigan, U.S.A. Yihong Qiu

Abbott Laboratories, North Chicago, Illinois, U.S.A.

Yongyut Rojanasakul West Virginia University, Morgantown, West Virginia, U.S.A. Mark Sacchetti GlaxoSmithKline, Research Triangle Park, North Carolina, U.S.A. Gabriela Sbirlea-Apiou Jouy en Josas, France

Air Liquide Research Center Claude-Delorme,

Parthiban Selvam Chemical Engineering and Materials Sciences, Wayne State University, Detroit, Michigan, U.S.A.

Contributors

xi

Steven J. Smith American Medical Association, Chicago, and Chicago Medical School, North Chicago, Illinois, U.S.A. Hugh D. C. Smyth College of Pharmacy, University of New Mexico, Albuquerque, New Mexico, U.S.A. David Swift

The Johns Hopkins University, Baltimore, Maryland, U.S.A.

David C. Thompson University of Colorado Health Sciences Center, Denver, Colorado, U.S.A. Schering-Plough, Kenilworth, New Jersey, U.S.A.

Susan L. Tiano

Michiel M. Van Oort GlaxoSmithKline, Research Triangle Park, North Carolina, U.S.A. Ronald K. Wolff

Nektar Therapeutics, San Carlos, California, U.S.A.

Libo Wu Chemical Engineering and Materials Sciences, Wayne State University, Detroit, Michigan, U.S.A. Yadong Yang

Crawford Communications, Inc., Atlanta, Georgia, U.S.A.

Contents

Introduction Claude Lenfant . . . . v Preface . . . . vii Contributors . . . . ix PART I: AERODYNAMIC BEHAVIOR 1. Deposition Mechanics of Pharmaceutical Particles in Human Airways . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1 Gabriela Sbirlea-Apiou, Ira Katz, Georges Caillibotte, Ted Martonen, and Yadong Yang I. Introduction . . . . 1 II. Deposition Model for Aerosolized Drugs . . . . 2 III. Results and Discussion . . . . 12 References . . . . 22 2. Thermodynamics of Inhaled Hygroscopic Drugs . . . . . . . . . . . . 31 Ira Katz, Gabriela Sbirlea-Apiou, and Ted Martonen I. Introduction . . . . 31 II. Methods to Study Aerosol Thermodynamics . . . . 33 III. Discussion of Aerosol Drug Hygroscopicity Within Human Airways . . . . 43 IV. Summary . . . . 46 References . . . . 51 3. Use of Mathematical Aerosol Deposition Models in Predicting the Distribution of Inhaled Therapeutic Aerosols . . . . . . . . . . . . 55 David Swift, Bahman Asgharian, and Julia S. Kimbell I. Introduction . . . . 55 II. Historical Review of Respiratory Tract Deposition Models . . . . 56 III. NCRP Lung-Deposition Model . . . . 65 xiii

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Contents

IV. V. VI. VII.

ICRP Lung-Deposition Model . . . . 67 General Features of Deposition Models . . . . 68 A Multiple-Path Particle Dosimetry Model . . . . 70 Aerosol Sampling Conventions Related to Respiratory Deposition . . . . 73 VIII. Special Features of Therapeutic Aerosols . . . . 76 IX. Future Developments in Deposition Modeling: A Model for Therapeutic Aerosols? . . . . 78 X. Summary and Conclusions . . . . 79 References . . . . 80 PART II: BIOLOGICAL CONSIDERATIONS 4. Physiology and Pharmacology of the Airways . . . . . . . . . . . . . . . 83 Ralph J. Altiere and David C. Thompson I. Introduction . . . . 83 II. Physiology of the Lung . . . . 84 III. Pharmacology of the Airways . . . . 107 References . . . . 118 5. Solute Transport Following Aerosol Deposition in the Lungs . . . 127 Richard M. Effros I. Introduction . . . . 127 II. Radioaerosols . . . . 128 III. Factors Affecting Clearance of Solutes from the Lungs . . . . 129 IV. Lipid Solubility . . . . 132 V. Electrolyte Transport . . . . 132 VI. Strong Acids . . . . 133 VII. Water Transport . . . . 134 VIII. Redox Indicators . . . . 135 IX. Macromolecules . . . . 135 X. Exhaled Breath Condensates . . . . 137 XI. Conclusions . . . . 140 References . . . . 140 6. Drug Metabolism and Enzyme Kinetics in the Lung . . . . . . . . . 147 Meenakshi Bhat, Joseph K. H. Ma, Yongyut Rojanasakul, and Ronald K. Wolff I. Introduction . . . . 147 II. Enzymatic Systems and Their Distribution in the Respiratory Tract . . . . 149

Contents

xv

III. Gross Anatomical Considerations in Aerosol Drug Delivery . . . . 161 IV. Experimental Models in Lung Metabolism Studies . . . . 164 V. Current Application of Aerosols in Drug Therapy . . . . 169 VI. Conclusions . . . . 173 References . . . . 175 7. Bioavailability and Pharmacokinetics of Inhaled Drugs . . . . . . 187 Akwete L. Adjei, Yihong Qiu, and Pramod K. Gupta I. Introduction . . . . 187 II. Formulation Effects and Bioavailability . . . . 189 III. Factors Affecting Bioavailability of Inhaled Drugs . . . . 194 IV. Pharmacokinetics of Inhaled Drugs . . . . 200 V. Clinical and Regulatory Issues . . . . 211 VI. Epilogue . . . . 215 VII. Conclusions . . . . 215 References . . . . 216 8. Therapeutic Uses of Lung Aerosol . . . . . . . . . . . . . . . . . . . . . . . . 219 Jonathan A. Bernstein, Hetal Amin, and Steven J. Smith I. Introduction . . . . 219 II. Systemic Routes of Medication Delivery for Asthma and Other Lower Respiratory Disorders . . . . 220 III. Aerosolized Use of Drugs in Asthma Therapy . . . . 228 IV. Cystic Fibrosis . . . . 244 V. Gene Therapy . . . . 247 References . . . . 248 PART III: PHARMACEUTICS AND PHARMACEUTICAL TECHNOLOGY 9. Atomization and Nebulizers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 253 Ralph W. Niven and Anthony J. Hickey I. Introduction . . . . 253 II. Atomization . . . . 254 III. Nebulizers . . . . 263 IV. Formulation . . . . 275

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Contents

V. Summary and Conclusions . . . . 278 References . . . . 279 10. Ultrasonic and Electrohydrodynamic Methods for Aerosol Generation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 285 Bernard J. Greenspan I. Introduction . . . . 285 II. Ultrasonic Nebulizers . . . . 285 III. Electrohydrodynamic Atomization . . . . 296 IV. Conclusions . . . . 302 References . . . . 302 11. Spray-Drying and Supercritical Fluid Particle Generation Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 307 Michiel M. Van Oort and Mark Sacchetti I. Background . . . . 307 II. Spray Drying . . . . 308 III. Particle Generation Using Supercritical Fluids . . . . 329 IV. Properties of Spray-Dried and SCF Powders and Methods of Characterization . . . . 336 References . . . . 342 12. Interfacial Phenomena and Phase Behavior in Metered Dose Inhaler Formulations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 347 Keith A. Johnson I. Nomenclature . . . . 347 II. Introduction . . . . 348 III. Estimating Suspension Stability . . . . 349 IV. Propellant Physical Properties and Phase Behavior . . . . 356 V. Experimental Design and Methods . . . . 364 VI. Summary and Conclusions . . . . 368 References . . . . 369 13. Molecular Scale Behavior in Alternative Propellant-Based Inhaler Formulations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 373 Libo Wu, Robson P. S. Peguin, Parthiban Selvam, Udayan Chokshi, and Sandro R. P. da Rocha I. Introduction . . . . 373 II. Alternative pMDIs . . . . 374 III. Microscopic Tools for the Development of Alternative pMDIs . . . . 380

Contents

xvii

IV. Conclusions and Outlook . . . . 392 References . . . . 393 14. Aerosol Generation from Propellant-Driven

Metered Dose Inhalers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 399 Hugh D. C. Smyth, Richard M. Evans, and Anthony J. Hickey I. Introduction . . . . 399 II. Observed Droplet Formation and Dispersion . . . . 403 III. Fundamentals of Droplet Formation . . . . 406 IV. Theoretical Framework Describing Droplet Formation . . . . 409 V. Semiempirical Model of Droplet Formation . . . . 412 VI. Conclusion . . . . 413 References . . . . 414 15. Medical Devices for the Delivery of Therapeutic Aerosols to the Lungs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 417 Richard N. Dalby, Susan L. Tiano, and Anthony J. Hickey I. Introduction . . . . 417 II. Pressurized Metered-Dose Inhalers . . . . 418 III. Dry Powder Inhalers . . . . 425 IV. Nebulizers . . . . 432 V. Selection of a Delivery System . . . . 437 VI. Future Developments . . . . 438 References . . . . 439 16. Next Generation Dry Powder Inhalation Delivery Systems . . . . 445 Anthony J. Hickey and Timothy M. Crowder I. Introduction . . . . 445 II. Formulation . . . . 448 III. Metering Systems . . . . 451 IV. Dispersion Mechanisms . . . . 452 V. Inhaler Designs (Patent Review) . . . . 454 VI. Characterization of Dry-Powder Inhaler Products . . . . 454 VII. Conclusions . . . . 456 VIII. General Conclusion . . . . 457 References . . . . 457 Index . . . . . . 461

PART I:

AERODYNAMIC BEHAVIOR

1 Deposition Mechanics of Pharmaceutical Particles in Human Airways GABRIELA SBIRLEA-APIOU

IRA KATZ

Air Liquide Research Center Claude-Delorme, Jouy en Josas, France

Air Liquide Research Center Claude-Delorme, Jouy en Josas, France, and Lafayette College, Easton, Pennsylvania, U.S.A.

GEORGES CAILLIBOTTE

TED MARTONEN

Air Liquide Research Center Claude-Delorme, Jouy en Josas, France

CyberMedicine, Laguna Beach, California, and University of North Carolina, Chapel Hill, North Carolina, U.S.A.

YADONG YANG Crawford Communications, Inc., Atlanta, Georgia, U.S.A.

I.

Introduction

The straightforward objective of this chapter is to demonstrate how aerosolized drugs can be targeted to relatively well-defined regions within the human respiratory tract by merely understanding the relative roles of the various factors affecting the airborne motion and subsequent deposition of inhaled particles. The deposition patterns of inhaled particles may be expressed as functions of three classes of variables: aerosol characteristics, ventilatory parameters, and respiratory tract morphologies. The efficiencies of the different deposition mechanisms of inertial impaction, sedimentation, and diffusion can be, in turn, formulated in terms of these variables. Therefore, by understanding the relative roles of the respective deposition processes, clinical personnel can effectively regulate particle deposition patterns within the lung.

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Sbirlea-Apiou et al.

For instance, in the upper tracheobronchial (TB) tree the deposition of large (1 m) particles can be primarily attributed to inertial impaction whereas in the more peripheral airways it may be ascribed to sedimentation. By recognizing the effects of (for example) ventilatory parameters, deposition due to inertial impaction can be markedly enhanced in the upper lung by increasing the inspiratory flow rate, or, conversely, deposition in the lower lung can be promoted by increasing the duration of a postinspiratory pause (i.e., breath-holding time). There are substantive reasons to selectively deposit drugs within the human respiratory tract. The most obvious would include the following: (i) An airway disease may not be manifested uniformly within the lung. That is, the disease will have a predilection for specific locations. Let us consider three specific examples of different scales of the heterogeneity of disease distribution among lung airways: on a compartmental level, cystic fibrosis may be dispersed throughout the TB tree; on a regional level, asthma has been clinically observed in the large upper, central, or small lower TB airways; and, on a localized level, bronchogenic carcinomas have been found mostly within bifurcations (1). (ii) Receptors and nerve endings are not homogeneously distributed among airways, but may be localized (2). Indeed, in some cases the anatomical locations are very focal (e.g., at carinal ridges within bifurcations). There would be, therefore, distinct benefits if inhaled drugs could be selectively deposited. Let us identify some of the most obvious positive features. (i) In clinical practice it is often necessary to deliver massive doses to the whole lung in order to get a required quantity of prescribed drug to airway cells at desired sites and thereby elicit a therapeutic response. (ii) The commensurate adverse side effects often experienced by patients following such high doses could be eliminated if drugs were targeted to appropriate sites. (iii) Aerosol therapy would become more cost effective. If drugs were targeted, the waste (i.e., lung overdose to get a requisite amount to a desired location) could be minimized, if not eliminated. Overall, the demands placed on physicians, nurses, and technicians engaged in aerosol therapy would become, more defined if predetermined quantities of drugs could be knowingly delivered to desired sites (e.g., to receptors in the treatment of asthma and malignant tumors in aerosol chemotherapy). Experimental and theoretical studies of the deposition of pharmaceutical aerosols have evolved since the 1st edition of Inhalation Aerosols. A discussion of key advances is given in the Summary. II.

Deposition Model for Aerosolized Drugs

A.

Elements of Biology and Physiology

Head and Throat Airways

When considering the delivery of drugs to the lung, the filtering efficiencies of the upper respiratory tract must be addressed. If the inhaled aerosol mass is Mi, the quantity that penetrates to the trachea, M, may be written as M Mi[1 p(m)][1 p(l)]

(1)

Deposition Mechanics of Pharmaceutical Particles in Human Airways

3

where p(m) and p(l) are the particle deposition efficiencies within the oropharyngeal region and larynx, respectively. Empirical formulas for such losses under passive breathing conditions have been presented by Martonen (3). The formulas may be appropriate for drugs delivered by metered-dose inhalers (MDIs) when used with spacers, dry powder inhalers (DPIs), and nebulizers. The deposition of inhaled particles in the airways of the human head and throat can be considerable even during breathing under normal, ambient conditions (4). The situation is exacerbated during the administration of medicinal aerosols, especially while using MDIs that employ propellants. The deposition of aerosolized pharmacologic drugs in the mouth and oropharyngeal region varies considerably with technique of application, but losses using the pressurized devices are routinely greater than 70% and can exceed 90%. The situation is somewhat better (i.e., particle losses are not as great) when using DPIs that are activated by the patient’s inhalation and nebulizers. Nevertheless, particle losses that occur proximal to the lung are a long-documented problem that continues to compromise the effectiveness of current aerosol therapy protocols (5–10). To reduce particle deposition during inhalation, improvements have continued to evolve in both design of hardware and techniques of delivery. For instance, spacer devices have been developed to be used with MDIs (11–14), new kinds of mechanical instrumentation have been constructed (15,16), and protocols of administration have been examined (17). It is the express objective of this work, therefore, to examine factors that affect the behavior and deposition of pharmaceuticals within the lung. Accordingly, in the balance of this text we shall focus on aerosolized drugs that do penetrate the mouth and oropharyngeal region. That is to say, particle deposition in the lung will be normalized to the dose that enters the trachea. If the reader chooses, alternatively, to normalize to the quantity inhaled, particle deposition in the upstream (i.e., to the lung) passages can be estimated using empirical expressions derived from the previously cited investigations. Lung Morphology

Descriptions of human lungs as symmetric networks are frequently dismissed outright as being unrealistic. We do not agree with such assessments. Symmetric morphologies are appropriate for certain applications and continue to evolve (18). We have used various symmetric (18–20) and asymmetric (21,22) morphologies to describe human lungs in aerosol therapy studies (23–28). Predicted particle deposition patterns were compared with experimental data from, inhalation tests with human subjects. Symmetric morphologies were demonstrated to be suitable for particle deposition modelling within lungs of children and adults over a wide range of conditions (particle sizes, breathing patterns, etc.). In our laboratory, we have made relica casts of human airways (29,30), constructed surrogate airway systems (31,32), and inspected the lung via the fiberoptic bronchoscope (33,34). We submit that although individual human

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lungs contain elements of asymmetry, symmetric morphologies may be suitable representations of average human lungs for a population. To express it differently, symmetric descriptions may well simulate the intersubject variabilities in airway dimensions and branching patterns observed among patients by physicians. A review of various symmetric and asymmetric morphologies for human lungs has been presented by McBride (35). It is perhaps worth emphasizing at this point that human lungs do not exhibit the marked degree of asymmetry documented in lungs of experimental animals (36–38). Laboratory surrogates such as dogs, rats, and guinea pigs possess very monopodial structures. Based on studies emphasizing regular features of the human lung, a morphology has been proposed by Weibel (18,19). The Model A is a symmetric, dichotomously branching system and will be used in this work. The TB airways are numbered in sequential order from the trachea (generation I 0) to the terminal bronchioles (I 16). To relate the terminology to conditions in situ, airway generations I 1, 2, 3, and 4 correspond to the main, lobar, segmental, and first subsegmental bronchi, respectively. The pulmonary (P) region, consisting of generations I 17–23, inclusive, is subdivided into three generations of partially alveolated respiratory bronchioles, three generations of alveolar ducts, and final alveolar sacs. The bifurcation and gravity angles among the airway network will be assumed to be 70° (39) and 45° (40), respectively. There are 2I identical airways in each generation. The dimensions of each generation of the lung are presented in Figure 1. Respiratory Parameters

To evaluate effects of intersubject variability in aerosol therapy regimens, an appropriate range of ventilatory conditions must be addressed. The parameters

Figure 1

Dimensions of airways within the adult human lung.

Deposition Mechanics of Pharmaceutical Particles in Human Airways Table 1

5

Breathing Patterns for Human Subjects at Prescribed Levels of Activity Physical state

Ventilatory parameters Tidal volume (mL) Breathing frequency (/min) Minute volume (mL) Flow rate (mL/sec)

Sedentary

Low

Light

Heavy

Maximal

500 14 7000 233.3

793 12.6 9992 333.1

1291 15.5 20,011 667.0

2449 24.5 60,001 2000.0

3050 40 122,000 4066.7

associated with prescribed respiratory intensities are given in Table 1. The values were recommended by Hofmann et al. (41). The corresponding average velocities in each airway are described in Figure 2. B.

Elements of Physics and Mathematics

Nomenclature

Dg G T k L(I) D(I) U(I) (I)

(I) Re(I) M C(Dg) V d t(I)

Particle geometric diameter, cm Particle density, g/cm3 Mean free path of air 7.0 106 cm Gravitational constant 980 cm/sec2 Absolute temperature 293 K Boltzmann constant 1.38 1016 g cm2/sec Length of generation I airway, cm Diameter of generation I airway, cm Air kinematic viscosity 1.5 × 101 cm2/sec Air absolute viscosity 1.84 × 104 g/cm sec Mean air velocity in generation I airway, cm/sec Inclination of generation I airway with respect to horizontal, degrees Angle of bend of generation I, degrees Airflow Reynolds number in generation I airway D(I)U(I)/•, dimensionless Particle mass, g Particle slip correction factor 1 + A(2/Dg), dimensionless where A 1.257 + 0.4 exp {1.1Dg/2)} Particle relaxation time mC(Dg)/(3 Dg), seconds Particle Stokes terminal settling speed G, cm/sec Particle diffusion coefficient kT/m, cm2/sec L(I ) U ( I ) ± V sin φ( I ) residence time of particles in generation I airways, seconds. [Note: +, or , denotes downhill, or uphill, flow in an airway when V has a component with, or against, fluid motion (31)]

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Figure 2

Air velocity profiles within the lung as a function of subject activity levels.

Primary Deposition Mechanisms

The major deposition processes are depicted in Figure 3, where particle behavior at a bronchial branching site is illustrated. The mechanisms are defined below. Inertial Impaction

Particles of sufficient momentum (product of mass and velocity) will be affected by the considerable centrifugal forces generated where the airway network, and thus convective fluid motion, changes direction abruptly. Sedimentation

Particles of sufficient mass may be deposited by the action of gravity when residence times within airways are large. Diffusion

Particle deposition can be a consequence of random Brownian motion. The efficiencies of these mechanisms will obviously depend on local respiratory tract geometries, particle parameters, and airstream characteristics. Therefore, the situation in inhalation therapy can be rather constrained, because, for a given patient (i.e., fixed morphology) and drug (i.e., prescribed aerosol), breathing is the only parameter that can be regulated.

Deposition Mechanics of Pharmaceutical Particles in Human Airways

Figure 3

7

Description of particle deposition mechanisms at an airway branching site.

Because particles are entrained and transported by an airstream, their trajectories are naturally affected by the magnitude and character of the airstream. The velocity distribution of air within the lung is determined by the tidal volume and breathing frequency parameters. Hence, deposition probability formulas must be indicative of the character of convective air movement. Although the respective mechanisms will be formulated below, a brief subjective description may be in order to put matters into perspective. Diffusion is most effective for submicrometer-sized particles. For larger particles, the inertial impaction deposition mechanism is dominant in upper TB airways where velocities are the maximum, and the effect of gravitational settling is most pronounced in distal airways where velocities are at a minimum. The fluid dynamics environment of the upper TB tree is characterized by flow instabilities initiated at the vibrating glottic aperture of the larynx. An airstream will enter the trachea as a laryngeal jet (42). Photographs of replica

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casts of human larynges that demonstrate the effects of inspiratory flow rates have been presented by Martonen and Lowe (43). Additional airstream instabilities may be attributed to the branching nature of the lung and neighborhood airway surface irregularities (44–46). The unstable motion will be dampened with progression to distal airways. Within such bifurcations localized airflow patterns have been studied (47–50). Laminar but asymmetric primary flow patterns were observed. In other experimental studies (51,52), secondary (double vortex) currents have been observed. To study aerosol deposition in the lung, the assumed airflow pattern must endeavor to be physiologically realistic. Accordingly, the following fluid dynamics pattern will be assumed in this work: turbulent airflow in bronchial generation I 0–3 (inclusive of the trachea and main, lobar, and segmental bronchi); and laminar flow with a plug velocity profile in TB generation I 4–16 (inclusive of the subsegmental bronchi to terminal bronchioles) and generation I 17–23 (the entire P compartment). Alternatively, laminar flow with a parabolic velocity profile may be assumed for the P compartment. However, under usual breathing conditions the length-to-diameter ratios of the airways are too small for flows to become fully developed. A system of particle deposition equations derived for such physiologically realistic flow patterns as described above has been derived by Martonen (53). The respective formulas, used herein, are presented in Table 2. We note that the trachea is a special case because the larynx affects inhaled particle trajectories (54). Particles entrained in the laryngeal jet will be deposited immediately downstream of the larynx, producing a “hot spot” described by the empirical equation of Chan et al. (55) p(t) 2.536Stk1.231 where Stk mean velocity in the trachea.

(2)

D2g U(0)/9D(0),

D(0) is the tracheal diameter and U(0) is the

Miscellaneous Deposition Processes

Other forces may become significant under special circumstances. The most prominent of these that are appropriate to particles within the human respiratory tract are represented below, albeit briefly. Table 2

Particle Deposition Formulas for Human Airways Deposition mechanisms

Flow condition

Inertial impaction P(I)

Laminar

2  2 1/ 2 + arcsin (e) ; e 1− e  π where e = θ( I )IU ( I ) D( I )

(

Turbulent

)

 −4e  1 − exp   π 

a

Sedimentation P(S)

(

)

Diffusion P(D)

2  2 + arcsin ( f ) ;  f 1− f  π where f = t ( I )v cos(θ( I )) / D( I )

where k = 4dL ( I ) / U ( I ) D( I )2

 −8τGt ( I ) cos (φ ( I ))  1 − exp   πD ( I )  

 −0.088d 3 / 4 Re ( I )7 / 8 L ( I )  1 − exp   2   U (I ) D (I )

1/ 2

4 (K / π)

1/ 2

− K;

Deposition Mechanics of Pharmaceutical Particles in Human Airways

9

Cloud Motion

This mechanism occurs when an array of particles settles as an entity with airborne characteristics different from its composite particles. In an experimental study using human replica casts, it has been demonstrated (1) that cloud motion is the mechanism that can dominate the deposition of cigarette smoke. Martonen (1) has identified morphological features and flow conditions that provide real opportunities for the initiation and propagation of particle clouds within the human lung. A theory of cloud motion was developed employing the concept of Smoluchowski (56). Cloud motion will take place when the Stokes settling velocity of an array, Uc, exceeds that of an isolated particle, Up, or Uc Up; conversely, individual particle motion occurs when Up Uc. The situation may be mathematically described as 2

 Dg  2/3 1 ρc( D ) ρc g   (6 + Re c ) 6  Dc 

(3)

where the upper inequality depicts the condition necessary for single particle settling to occur and the lower inequality signifies cloud motion. Dc and Rec denote the cloud diameter and Reynolds number within the human lung. Interception

This is likely to be the most effective deposition mechanism for aggregates and fibers. For such shapes, deposition can occur when a particle contacts an airway wall although its center of mass remains on a fluid streamline. Particle orientation is a critical factor in describing interception. Mathematical models of inhaled fibers have become ever more detailed, considering deposition within the whole lung (57), regional distributions (58,59), and localized deposition patterns (60–62). Empirical formulas for fiber deposition based on data from deposition experiments with simulated human airways and replica casts have been presented by Myojo (63) and Sussman et al. (64,65). Such theoretical and empirical findings, suitable for describing the airborne behavior of pharmacologic drugs as fibrous aerosols, should be integrated into improved aerosol therapy protocols (66–68). Electric Charge Effects

During the mechanical generation of aerosols electric charges may be produced on the particles. Electric charge effects have been studied using surrogate airway systems and human subjects (55,69–71). The data indicate that the effects will be most pronounced for submicron particles that possess correspondingly greater mobilities than larger (i.e., >1 m) ones. Electric charge effects within the human respiratory tract can be manifest in two ways, (i) image charge fields, and (ii) space charge forces. Regarding (i), it is the response between charges of opposite sign on an airway surface and a particle that create attraction and subsequent deposition. Regarding (ii), it is the repulsion between like-charged

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inhaled particles that may direct their motion toward airway walls resulting in deposition. Either effect may result in deposition being enhanced relative to an uncharged particle. It has been suggested (72–74) that space charge forces are of less significance than image forces. It should be emphasized that to reduce effects on deposition during controlled human subject exposures, it is standard operating procedure for aerosols to pass through an electric charge deionizer and thereby attain a Boltzmann equilibrium (i.e., no net electric charge) prior to inhalation. Simulation of Inhaled Aerosols

To describe particle deposition in the lung, we shall follow an inhaled bolus of aerosol throughout the TB and P airways. The method is depicted in Figure 4. The bolus designates an incremental volume of inhaled aerosol. Particles are continuously removed from the bolus (i.e., deposited on airway surfaces) as it travels throughout the lung during a cycle of breathing. This simulates popular experimental protocols for human exposures (75,76). The superposition principle advocated by Landahl (77) may be used to determine cumulative deposition P(C) P(I) + P(S) + P(D) + P(I)P(S)P(D) P(I)P(S) P(I)P(D) P(S)P(D) (4) P(C) considers interactive effects between deposition processes that have been simulated as being independent in the derivations of the equations in Table 2. Alternative ways to account for interactive effects between deposition processes have been presented as various investigators have attempted to isolate conditions where only selected mechanisms need to be addressed. For instance, let us consider two mechanisms, sedimentation and diffusion, which have received considerable attention in the literature. Simple superposition would indicate that P(C) P(D) + P(S) P(D)P(S)

(5)

For long, straight, smooth-walled tubes of circular cross-section, Heyder et al. (78) have presented an empirical equation P(C) P(D) + P(S) P(D)P(S)/[P(D) + P(S)]

(6)

For idealized tubes, Chen and Yu (79) have predicted that P(C) (P(S)2 + P(D)2 [P(S)P(D)]2)1/2

(7)

The latter formulas were proposed for the peripheral regions of the lung where diffusion and sedimentation are relatively effective. It must be emphasized, however, that those mechanisms are most relevant to particles of vastly different sizes. Moreover, empirical expressions reflect, inherently, their laboratory conditions that may not pertain to the lung. Likewise, theories for idealized tubes may not be applicable to the lung, where airways have, in fact, pronounced anatomical features that markedly affect fluid dynamics (33,34).

Deposition Mechanics of Pharmaceutical Particles in Human Airways

11

So it remains a question whether such alternative formulations have practical applications in drug delivery. The formulas presented in Table 2 and Eqs. (2) and (4) have been incorporated into a mathematical model describing the behavior of inhaled pharmacologic drugs (24). The model has been successfully employed to systematically study deposition patterns in human subjects as a function of aerosol polydispersity (25), ventilation (26), and lung morphology (27). In the cited investigations, the model was validated for use in the clinical environment. Calculated distributions of inhaled aerosols were in excellent agreement with data from volunteer test subjects and patients. In recent interdisciplinary efforts, the mathematical model has been introduced into the medical arena. Martonen et al. (80) have used it to interpret planar gamma camera images. Martonen et al. (81) have applied it to cystic fibrosis; specifically, the model was used to determine effects of diseaseinduced changes in airway dimensions upon the dispersion of inhaled aerosols.

Figure 4

Motion of a discrete volume of aerosolized drug.

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Results and Discussion

In panels A to C of Figures 5 to 7, the effects of breathing conditions and particle sizes on deposition are systematically studied. For a given figure (i.e., fixed

Deposition patterns of inhaled particles of designated sizes as a function of location within the human lung for sedentary breathing conditions. The deposition mechanisms are inertial impaction (), sedimentation (), and diffusion ().

Figure 5

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Deposition patterns of inhaled particles of designated sizes as a function of location within the human lung for light-activity breathing conditions. The deposition mechanisms are inertial impaction (), sedimentation (), and diffusion ().

Figure 6

respiratory intensity), the respective panels illustrate the effect of size. Conversely, for a given panel (i.e., constant particle size), the respective figures illustrate the effect of ventilatory parameters. Let us consider some examples.

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Deposition patterns of inhaled particles of designated sizes as a function of location within the human lung for heavy-activity breathing conditions. The deposition mechanisms are inertial impaction (), sedimentation (), and diffusion ().

Figure 7

In Figure 5 the influences of the deposition mechanisms of inertial impaction, sedimentation, and diffusion on particles of different sizes are addressed for sedentary breathing conditions. Specifically, the effects of the

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respective deposition processes on the spatial distributions of deposited particles are examined. Clearly, the deposition of submicron particles (panel A) is pronounced in the alveolated airways due to the action of diffusion. Likewise, the larger (3 m) particles are preferentially deposited in that region but for a completely different reason — namely, the effectiveness of the sedimentation mechanism (panel C). The deposition values in the peripheral airways are zero simply because the tidal volume for sedentary breathing is not sufficient to carry particles that far into the lung. The results for the 1 m particles (panel C) are intermediate to the other particle sizes. The effect of the increased inspiratory flow rate associated with the light activity level of breathing (Fig. 6) can be observed by comparing its C panel with Figure 5. The influence of inertial impaction can be discerned in the upper airways of the TB tree. Also, the larger tidal volume transports particles deeper into the lung so that particles are deposited by diffusion and sedimentation. In Figure 7, the role of inertial impaction becomes more evident. Indeed, it is the dominant deposition mechanism for the 3 m particles (panel C) that are selectively deposited in the large TB airways). In the more distal airways of the lung, deposition by diffusion (panels A and B) and sedimentation (panels B and C) are markedly enhanced by the tidal volume effect. The deposition curves of Figures 5 to 7 are of fundamental importance for the targeted delivery of inhaled drugs. The curves demonstrate, unambiguously, that individual particle deposition mechanisms are dominant in different regions of the lung. Moreover, the curves clearly show that effects of particle size and respiratory intensity, as formulated in the equations in Table 2, are distinctive and therefore can be integrated into aerosol therapy protocols. In the next step of the analysis (Fig. 8), we shall examine factors affecting the deposition efficiencies of the respective mechanisms that are dominant at prescribed locations within the lung. In each panel of Figure 8, deposition is plotted as a function of particle size for designated flow conditions. The illustrations are almost self-explanatory, so we shall make only a few observations to orient the reader. To begin, consider the upper airways. In panel A deposition in a generation I 4 airway, corresponding to the first generation of subsegmental bronchi, is simulated. For any selected particle size (i.e., the abscissa), the deposition efficiency increases monotonically with the magnitude of the inspiratory flow rate. However, the effect of flow rate is reversed in panels B (I 16) and C (I 20), where sedimentation and diffusion are addressed. In those downstream airways of the lung, deposition decreases as the level of respiratory intensity increases. This is simply because the effect of an increase in flow rate will be to reduce the residence time of a particle in an airway, so that influence of gravity and Brownian motion are commensurately reduced. Figures 9 to 11 are three-dimensional (3D) representations of particle deposition patterns. The distributions are expressed in terms of (i) inhaled particle size and (ii) location within the lung. These two parameters were deemed to be most important for implementation in the medical arena. For instance, if a clinician desires to target the delivery of a prescribed pharmacologic drug, it is important to know deposition as function of (i) the particle size

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Figure 8 Deposition efficiencies of specific mechanisms in prescribed airways. The breathing conditions are sedentary (), light activity (), and heavy activity ().

produced by a given generator (MDI, DPI, and nebulizer), and (ii) where the drug is to be deposited. To put matters into perspective, let us briefly consider drug delivery protocols in the clinical environment as envisioned at the time of the 1st edition. A

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Three-dimensional display of particle distribution within the human lung for sedentary breathing conditions. Figure 9

scenario could be this: a patient has a known airway disease (e.g., asthma) that determines the drug a physician selects to treat it (e.g., containing isoproterenol hydrochloride as the active bronchodilator agent) and that aerosolized drug is only available in certain particle sizes from different commercial devices. Let it be assumed that the physician wishes to deposit inhaled pharmaceuticals selectively within the upper, central, or lower airways within lungs. We suggest that the physician can use Figures 9 to 11 to determine the necessary breathing conditions and aerosol generators (i.e., particle sizes) to target the drugs to appropriate sites. Now, via a review of the literature we shall consider how recent works might affect the aforementioned vision. Research regarding particle deposition within the human respiratory system, especially as it pertains to inhalation therapy, has a worldwide scope. However, three groups have distinguished themselves by the extent and significance of their contributions. The group led by Martonen [(82–89) and other works cited later] has used a deterministic model to study a wide range of problems from general questions addressing the very strengths and limitations of modelling (85) to specific topics such as deposition in children’s lungs (86). The group led by Hofmann has used a stochastic model (90–96) to investigate, among other topics, the therapeutic effects of particle mass distributions (93) and cave aerosols (95). The group led by Finlay has used empirical modelling (97–107) to examine isotonic aerosols (100) and liposome-encapsulated ciproflaxin (103). Another widely used empirical model developed for regulatory compliance is called the International Commission on Radiological Protection (ICRP) model (108–112). Focused studies regarding electric charge effects (113,114) and fiber deposition (115,116) are examples of other subject matter investigated regarding deposition (117–126).

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Three-dimensional display of particle distribution within the human lung for light-activity breathing conditions.

Figure 10

A natural progression in the above research that is driven by limitations of the existing models and is fostered by advances in hardware (e.g., memory, power) and computational tools (i.e., software), is the application of computational fluid dynamics (CFD). To date, this has been most apparent regarding the calculation of particle deposition in the extrathoracic (ET) and TB regions of the respiratory tract where complex inertial effects are prevalent. The groups of Martonen [(127–129) with other publications referenced in following text],

Figure 11 Three-dimensional display of particle distribution within the human lung for heavy-activity breathing conditions.

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Streamlines from computational fluid dynamics simulations of steady inspiratory flow (0.5 L/sec) through the extrathoracic region: (A) front view and (B) side view.

Figure 12

Hofmann (130,131), and Finlay (132–135) have all contributed CFD papers to the inhalation therapy literature. These and other CFD studies (136–144) have focused on either idealized morphological models or ones based on single individuals. The next step, or logical progression, for this work would be to incorporate the great advances in imaging technologies that will permit the customization of CFD models for individual patients. Studies by our team indicate that unique aspects of an individual’s respiratory tract morphology might have a profound impact on the distribution of administered drugs (145–149) (Vial L, de Rochefort L, Fodil R, et al. In vitro validation of CFD simulation in human proximal airways reconstructed from medical images with hyperpolarized helium-3 MRI phase contrast velocimetry. J Appl Physiol 2006. In review). For example, Figure 12 presents different views of streamlines in the ET region as determined using laminar CFD simulations. It bears emphasis that the streamlines through the nasal cavity for this healthy subject are far from symmetric, and that flow through the larynx is quite complex. The morphology was obtained via high-resolution (0.4 mm 0.4 mm 0.6 mm voxels) computed tomography (CT) images. The CT data were incorporated into the commercial CFD package FLUENT (from FLUENT Inc.) in our study. A second simulation of flow through the TB morphology of a second healthy subject is shown in Figure 13. The asymmetry of the branching morphology at the first (main) bronchi creates vortices that may substantially affect particle deposition. Such secondary flow patterns may not be as pronounced in idealized morphological models. It is probable that airstream profiles such as those

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Streamlines from computational fluid dynamics simulations of steady inspiratory flow (0.5 L/sec) through the tracheobronchial region.

Figure 13

illustrated above may contribute to the intersubject variabilities observed with in vivo particle deposition studies. An important goal of our Air Liquide–CyberMedicine collaboration is to understand the complex phenomena associated with aerosol deposition processes so that models can be optimally simplified, thereby resulting in a software package that is a real clinical tool and can be used practically in the medical arena especially to formulate inhalation therapy regimens. At this juncture we shall recognize theoretical studies that have appeared in the peer-reviewed literature and have direct applications to aerosol therapy. The series of publications of Martonen and colleagues (127–129) show 3DCFD simulations and 3D trajectories of entrained particles within an anatomically realistic model of the human respiratory system. To be specific, the 3D morphology consisted of the complete ET region (i.e., nasal, oral, pharyngeal, and laryngeal passages) and part of the thoracic region (i.e., tracheal and large bronchial airways). The importance of the above studies is twofold. First of all, they simulate the transport and deposition of particles inhaled via the nose. This has obvious implication to the administration of drugs (e.g., flu vaccines) via the nasal route for rapid absorption and distribution via the circulatory system. Secondly, the results are controls with respect to which the effects of hygroscopicity can be objectively quantitated. That is, the results can function as baseline, reference data for materials presented in Chapter 2 of this 2nd edition. At the 36th Respiratory Care Journal conference titled “MDIs and DPIs in Aerosol Therapy” Martonen et al. (150) presented a 3D color illustration of

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the complete, contiguous respiratory system. Moreover, 3D simulations of aerosolized drug delivery via the mouth were given. The study is a direct application of mathematical modelling and computer simulations to inhalation therapy which is, of course, the subject of this textbook. In this 2nd edition of Inhalation Aerosols we shall emphasize that the deposition patterns presented in Figures 9 to 11 from the 1st edition are still valid. This is simply because the theoretically predicted deposition curves are within the error bars of reported human exposure data, especially when intersubject variabilities are recognized. The in silico aerosol dosimetry model used in this work has been extensively tested via comparisons of predicted deposition values with data from human subject studies, and has been validated (86,150–152). We shall note that the term “in silico” model has been employed to be compatible with in vivo and in vitro data from aerosol exposure studies with human subjects and replica airway casts, respectively. The term in silico is meant to indicate that theoretical formulations per se have been integrated with computer technology (software and hardware). The salient point to be made regarding theory-versus-experiment comparisons is that theoretical analyses can be performed to any desired degree of resolution; for instance, the computational end point could be the prediction of aerosol deposition in units of (gm/cm2) (i.e., aerosol mass deposited per unit airway surface area). However, in vivo deposition patterns cannot be measured to that extent using current technologies and protocols (86). That is why, of course, in vitro experiments with surrogate airway systems are so valuable (84). Therefore, regarding in vivo data, one is limited to making comparisons with the format in which those results have been published in the literature [e.g., total lung deposition values, compartmental deposition fractions in TB, and pulmonary (P) compartments]. Another factor that must be considered is the redistribution of deposited aerosols by natural clearance processes. Clearance from the TB tree is acknowledged to occur between 24 and 30 hours (e.g., about one day). Within the bronchial network localized effects may occur, that is, clearance may be reduced. Indeed, ciliostasis may occur within bifurcations leading to greatly increased exposures of underlying cells. In the alveolated region, particle removal may require a matter of days, perhaps weeks. Such redistributions have been included in our modelling efforts (153). In our Air Liquide and CyberMedicine laboratories we have placed an emphases on making the in silico dosimetry models increasingly realistic from biological and technical perspectives. To begin, we have endeavored to develop morphological descriptions of the human respiratory system that are as correct as currently possible to formulate and simulate. For instance, we have generated a model for the entire human respiratory system inclusive of the ET region and lung (TB and P) airways (i.e., from the lips and nostrils to the alveoli). A colored illustration of the 3D system has been presented by Martonen et al. (150). Once the geometries of the described airways (e.g., surface roughness) were defined, we addressed issues of related spatial orientation (e.g., asymmetry) of branching networks. For details of the evolution of the architecture of

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the system we refer the reader to Refs. 82, 127, and 154–165. Following, the definition of respiratory system morphology, matters associated with airstream patterns and trajectories of entrained particles could be considered. We refer the reader to Refs. 128, 129, and 166–178 to see how these scientific topics were handled in our analyses. The status of our in silico model integrating the above features is presented in publications (150–152). Of special relevance to the medical community is the application of our modelling and simulation efforts to airway diseases such as cystic fibrosis (81,179), chronic obstructive pulmonary disease (COPD) (180), cancer (87,181, 182), and asthma (89,183). We have also applied the aerosol dosimetry code to problems related to the systemic delivery of inhaled drugs (i.e., using lungs as the portal of entry into the body); specifically, diabetes has been addressed (184), although the general problem of protein delivery per se has been recognized (185). Mathematical modelling and computer simulation studies may have a special role in pediatric medicine simply because experimental data are so difficult to obtain for legal and ethical reasons. Accordingly, we have developed dosimetry codes to predict the behavior of inhaled drugs in children (85,186–189). In our continuing efforts to facilitate the use of mathematical modelling and computer simulations in the medical arena a technique has been developed to portray 3D respiratory system structures (190). Special glasses were supplied with the journal Cell Biochemistry and Biophysics to allow the viewing of the branching networks in 3D.

Acknowledgments Yadong Yang, one of the authors of the 1st edition of this chapter, was supported by funds provided by the U.S. Environmental Protection Agency under Collaborative Agreement CR817643 on Health Effects of Exposure to Air Pollutants through the Center for Environmental Medicine and Biology, University of North Carolina, Chapel Hill, NC, U.S.A.

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2 Thermodynamics of Inhaled Hygroscopic Drugs

IRA KATZ

GABRIELA SBIRLEA-APIOU

Air Liquide Research Center Claude-Delorme, Jouy en Josas, France, and Lafayette College, Easton, Pennsylvania, U.S.A.

Air Liquide Research Center Claude-Delorme, Jouy en Josas, France

TED MARTONEN CyberMedicine, Laguna Beach, California, and University of North Carolina, Chapel Hill, North Carolina, U.S.A.

I.

Introduction

The environment of the human respiratory tract can have a profound impact on the behavior of inhaled hygroscopic pharmaceuticals used for the prophylaxis and treatment of airway diseases, as detailed herein. This fact should be actively integrated into current drug-delivery protocols and, indeed, become a seminal element of future aerosol therapy programs. The critical parameters describing the atmospheres are temperature (T) and relative humidity (H). Although the interior of the lung is obviously very warm and humid, the values of the parameters in situ are difficult to quantitate and, therefore, the airways have not been systematically mapped. The actual T and H values depend on the route of breathing (i.e., via the nose or mouth). Regarding the extrathoracic region, the nasopharyngeal passages are much more effective in conditioning inhaled air than oropharyngeal passages (1–3). Spatial distributions of T and H profiles have been presented by Martonen and Zhang (4). Because of their chemical composition, many pharmaceuticals administered via inhalation are hygroscopic in character (5); that is, the drugs have affinities for water vapor, which is ubiquitous in the respiratory tract. Upon the uptake of water vapor, drug particles may change in size, shape, and density during transit within human airways. Therefore, their deposition efficiencies can be markedly different from identical preinspired particles (6–9).

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The subject of hygroscopicity has been reviewed by Morrow (10), with emphasis on an analysis of factors affecting the airborne behavior of inhaled particles. Subsequently, the applications of hygroscopic growth to aerosol therapy have been addressed in the review of Hiller (11). These two manuscripts afford a thorough compilation of relevant literature to which we shall, for the sake of brevity, refer the reader. In more recent times, the role of hygroscopicity in the delivery of pharmacologic drugs has received increasing and deserved attention from the perspectives of research and applications. Regarding pure research or the examination of factors affecting the physical chemistry of hygroscopicity, the associated thermodynamics problems have been addressed by Martonen (12), Eisner et al. (13), Ferron and Soderhohn (14), Soderholin and Ferron (15), and Hickey and Martonen (5). Regarding applications, analyses of the effects of hygroscopicity upon particle deposition have been modeled by Martonen et al. (9), Ferron et al. (8), Soderholm et al. (16), Ferron et al. (17), and Martonen et al. (3). The implications of hygroscopicity to drug delivery are elementary in concept and straightforward in practice. Indeed, the role of hygroscopicity is so fundamental to the administration of inhaled pharmaceuticals, especially for the success of targeted drug delivery, that it is difficult to comprehend the extent to which it has been neglected in the literature. For instance, in the medical arena, it must be acknowledged that the deposition patterns of inhaled hygroscopic drugs may not be related to particle sizes and densities as measured at sites of aerosol generation. Specifically, the aerosol characteristics of mass median aerodynamic diameter and geometric standard deviation determined at a nebulizer, metered dose inhaled, or dry powder inhaler are not relevant parameters in estimating the dose delivered to the lung unless hygroscopic growth following inhalation is considered. In this text, we shall address two avenues of research pertinent to the study of hygroscopicity: theoretical models and laboratory surrogates. However, we shall emphasize the former for two reasons. First of all, knowledge of theory is essential for understanding the kinetics of inhaled drugs via the correct identification and elucidation of the fundamentals of thermodynamics pertaining to hygroscopicity. Secondly, a theory of hygroscopicity suitable for inclusion into a validated particle deposition model [such as that developed expressly for drugs by Martonen (18)] and thereby capable of describing commensurate effects of hygroscopicity on drug deposition, may have a tremendous impact on improving the efficacies of inhaled drugs via their targeted delivery. To describe aerosol hygroscopicity in human airways, it is imperative that the thermodynamics of particle behavior be formulated in a dynamic environment; that is, it may not be appropriate to extrapolate data about hygroscopicity obtained under the static conditions more commonly addressed in the literature. The lung has unique temperature, relative humidity, and fluid-dynamic conditions that make it difficult (perhaps inappropriate) to use analyses derived for other problems involving engineering or biological systems. Accordingly, we shall organize our text in the following manner. We shall begin by discussing the

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33

theory widely used in current studies of aerosol hygroscopicity proposed by Ferron and colleagues (6,8,14–17). This will be done in sufficient detail to disclose its limited mechanisms of operation. Then, we shall discuss real fluiddynamic conditions within lung airways, specifically, the role of convection versus molecular diffusion, which affect certain tenets of their theory and may negate its applicability for the respiratory system. Finally, employing principles of engineering, we shall present a new theoretical approach to the analysis of thermodynamics suitable for the study of aerosol hygroscopicity, particularly pharmacologic drugs. Since the 1st edition of Inhalation Aerosols appeared, the study of thermodynamics and hygroscopic effects on aerosol deposition has continued. A discussion of the implications of this recent work is given in Summary. II.

Methods to Study Aerosol Thermodynamics

Upon the inspiration of hygroscopic aerosols, a complex set of thermodynamic and fluid-dynamic interactions begins between individual droplets and the entraining airstream. In a recent review, Sirignano (19) has described the general interactions of sprays with airstreams: (i) If there is any relative motion between a droplet and the air, fluid-dynamic phenomena such as boundary layers, wakes, etc. will appear. The increased transport of mass or energy beyond molecular diffusion that results from this relative motion is termed convection. Convective transport is highly coupled to the structure of a flow field. (ii) The relative motion between air and a droplet can also cause sheardriven circulation within it. Again, fluid motion will increase transport beyond that expected from molecular diffusion. (iii) Another convective process that may be present is a consequence of Stefan flow—that is, flow caused by the movement of the liquid–gas (i.e., droplet air) interface, which is hygroscopic growth per se. (iv) The flow field about the droplet may be disturbed owing to the presence of others (i.e., hydrodynamic interactions), the presence of vortical structures (e.g., separation regions), and turbulence (e.g., as formed by the laryngeal jet within the trachea). The shape of a droplet may be altered to such an extent by the flow field that it breaks apart. All of these fluid-dynamic phenomena (i.e., points i–iv above) may affect, to varying degrees, the transport of mass and energy between a droplet and a surrounding airstream. In fact, there can be a strong coupling between the larger-scale structure of the flow field and the smaller-droplet-scale heat- and mass-transfer processes. To be succinct, the thermodynamics of hygroscopic aerosols is a very challenging problem. Accordingly, we shall begin by considering the model(s) of hygroscopicity currently popular and subsequently identify potential improvements to theoretical approaches. A.

Present Theoretical Model(s)

The theory originally formulated by Ferron (6) is frequently used for predicting the sizes of hygroscopic aerosols used for pharmacological purposes. The latest permutations of that theory have been presented in the recent works of Ferron

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and colleagues (8,14–16). The analysis is based on the assumption that a particle reaches an equilibrium size a relatively short time after entering the lung (equilibrium being defined as the condition when the temperature of the ambient air within the lung and that of the particle’s surface are the same, and their watervapor concentrations are equal). Subsequently, there will be no mass transfer between them, and hygroscopic particle growth will have been completed. Let us consider some critical elements of the theory. It used a variation of Raoult’s law to determine the water-vapor concentration at the surface of a salt solution based on the mass and molecular weights of the salt and water, written as ms nw + jns mw = N= mg nw + (i + j )ns 1 + (i + j ) mw 1+ j

Mw Mg Mw Mg

(1)

where N is the ratio of the water-vapor concentration at the surface of the solution to the vapor pressure of pure water under identical thermodynamic conditions, i is the number of ions into which a dissolved salt molecule is dissociated in water, j is the number of water molecules per salt molecule, and nw and ns are the number of water and solute molecules in the solution, respectively. Mw and Ms are the molecular weights of water and the solute, while mw and mg are the mass of water and the solid (i.e., solute) particle, respectively. The equation proposed by Kelvin is used to account for the difference in watervapor concentration at the surface of a droplet (diameter deq) to that of a flat surface. Rw is the corresponding ratio  R1   4 σM w  exp = Rw = exp      ρdeq   RT ρdeq 

(2)

where ␴ is the surface tension of water, R is the gas constant, T is the absolute temperature, and ␳ is the density of water (liquid). The equilibrium condition requires that the water-vapor concentration at the surface of a droplet be equal to the water-vapor concentration in the surrounding air, c. Thus, the relative humidity can be related to the water-vapor concentration given in Eq. (1) via the expression H=

c = NRw csat (T )

(3)

where H is the relative humidity and csat (T) is the saturated water-vapor concentration at temperature T. Using Eq. (1) with (3), the equilibrium masses of water and solute can be determined for any given relative humidity. This

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total mass can be compared to the dry mass of the original solid (i.e., solute) particle and in turn, assuming a spherical shape for the resultant particle, a diameter ratio can be derived deq,1 deq,s

 ρ  = s   ρ1gs 

1/ 3

1/ 3

  M w  H (i + j )− jRw    ρ  =  s 1 + R −H  ρ1    M s  w    

(4)

where deq,1 and deq,s are the equivalent diameters of the particle in the air and the initial dry particle, respectively, ␳1 and ␳s are their respective densities, and the solid (i.e., solute) weight concentration in the particle is gs. The assumption of equilibrium is critical in using Eq. (4) with models describing the deposition of inhaled particles. One must estimate how long after a droplet enters the lung it will reach thermodynamic equilibrium with the ambient conditions. If this time is short enough, the hygroscopic growth model and the deposition model can be uncoupled, that is, the complex interactions between mass transfer on the droplet scale and the complex fluid-dynamic conditions within airways can be ignored. Ferron (6) addressed the issue in a most prudent manner. He assumed all thermodynamic processes between a droplet and the surrounding air to be due to molecular diffusion alone. Such an analysis afforded, inherently, the longest probable timescale for the heat- and mass-transfer processes to occur. Any of the modes of convection discussed earlier (i–iv) would have the effect of decreasing the time required to reach equilibrium. Ferron (6) also assumed that the heat capacity term could be neglected from the energy budget because it would be much smaller than the heat of condensation. The transfer of heat between a droplet and surrounding air is thus equal to the energy required for condensation, or LI ⫹ ␸⫽0

(5)

where L is the heat of condensation, I is the mass of water transferred to the droplet, and ␸ is the heat transfer. The governing equations for molecular diffusion of mass and heat in air are Fick’s law and Fourier’s law. They are, respectively  2 ∂c ∂ 2 c  ∂c = D +  ∂t  r ∂r ∂r 2 

(6)

∂T κ  2 ∂T ∂ 2T  = +   ∂t c p ρa  r ∂r ∂r 2 

(7)

and

where c is the concentration of water vapor in the air, t is the time, Dw is the diffusion constant of water-vapor molecules in the air, and r is the spatial variable

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in a spherical coordinate system whose origin is at the center of the droplet in Fick’s law. In Fourier’s law, T is the temperature, ␬ is the thermal conductivity of the air, cp is the specific heat of air at constant pressure, and ␳a is the density of the air. Ferron (6) utilized a solution to Eqs. (6) and (7) for constant boundary conditions of concentration and temperature at a droplet’s surface (c0, T0) and in the air distant from the droplet (c⬁, T⬁). The solution yielded functions for c(r, t) and T(r, t) in the air, which consequently yielded I and ␸ from the concentration and temperature gradients at the droplet’s surface, respectively. To determine the change in size of a droplet, Ferron (6) used the quasistationary mass-transfer equation I ⫽ 2␲deqDw(c0 ⫺ c⬁).

(8)

A similar result was obtained for Eq. (7) ␸ ⫽ 2␲deq␬(T0 ⫺ T⬁).

(9)

Eqs. (8) and (9) are quasistationary because properties of the droplet are assumed to be constant (i.e., uniform composition) with a form of equilibrium attained. In other words, this is the steady-state solution for the molecular diffusion problem described by Eqs. (6) and (7), with constant boundary conditions. However, it was presumed that the time for molecular diffusion to reach equilibrium would be much shorter than the time necessary for the droplet to reach its equilibrium size in the lung. Therefore, c0 and T0 could be updated in time to give the time-dependent mass transfer (and thus growth) for the droplet using Eqs. (5), (8), and (9). A numerical scheme was developed to solve this set of equations because c0 is a function of the temperature and diameter of the droplet. The conclusion of the above analysis that is pertinent to inhaled drugs is that, “particles with an aerodynamic diameter smaller than 0.2 g have reached the equilibrium size in the lung within 0.1 seconds. Furthermore, 1.0 ␮m particles obtain their equilibrium size within two seconds, while larger particles do not reach their equilibrium size during one respiration” (6). These conclusions were drawn after numerous assumptions were made pertaining to ambient conditions (T and H) in the lung and ventilatory parameters. However, it must be emphasized that, more often than not, the highly complex interactions between lung flow fields and particle growth (factors i–iv outlined at the beginning of Methods to Study Aerosol Thermodynamics section) must be considered to adequately model deposition. This subject is therefore addressed in the next section. A salient point being: as particles grow larger, the roles of inertial impaction and sedimentation upon deposition in the upper and lower airways, respectively, increase markedly (20). This is especially true for particles greater than 1.0 ␮m. B.

New Theoretical Model(s)

Recently, much research in the engineering community has been directed toward numerically determining mass-transfer rates between droplets and ambient flow

Thermodynamics of Inhaled Hygroscopic Drugs

37

fields (19). A great amount of the work has focused on evaporation—in practical effect, the reverse of the problem previously addressed herein. Unfortunately, at this time, detailed flow field–mass-transfer interactions exist only for uniform flow fields. For hygroscopic growth in pharmaceutical applications, a significant period occurs during passage through the oronasopharyngeal passages, the larynx, and the trachea. The flow fields in these regions are typically very complex and unstable (21). For example, consider the turbulent flow field in the trachea. Turbulent flows are characterized by a large range of length scales (or eddy sizes). The largest scale for the trachea is its diameter D. The smallest scale can be estimated by calculating the Kohnogorov microscales (22). First determine the dimension of the energy dissipation rate, ␧, which is established from the largescale dynamics

ε∼

U3 D

(10)

where U is the average velocity of air in the trachea. The dimension of the smallest scale, ␩, is related to the eddy size that dissipates energy viscously

 υ3  η∼   ε 

1/ 4

(11)

where ␯ ⫽ 0.153 cm2/sec is the kinematic viscosity of air at 37° C. For heavy breathing, consider an inspiratory flow rate of Q ⫽ 60 L/min and D ⫽ 1.8 cm. The average velocity is related to Q by U=

4Q πD 2

(12)

For the situation in the trachea described above, ␩ ~ 32 ␮m, is of the same order of magnitude as the individual particles within the flow field and six orders of magnitude less than the largest length scale, D ⫽ 1.8 cm. Unfortunately, as regards deposition modeling, Sirignano (19) has reported that there has been relatively little research performed on the interaction of turbulent eddies with individual particles. The problem of hygroscopic growth of aerosol particles in the lung must address complex, time-dependent mass-transfer processes, especially the turbulent nature of the flow field. At this juncture, probably any mode of analysis put forth by an author would seem to others to have severe limitations; however, this situation is not uncommon in engineering practice, and a standard approach has been to apply empirical results to predict future phenomena. In the balance of this text, the phenomenological approach to be taken to study the time-dependent nature of hygroscopic growth will be a lumped capacitance model for heat and mass transfer. The lumped capacitance model has been widely accepted as a simple yet powerful model for understanding many

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complex engineering systems. A heat–mass–transfer analogy will be used to explain how the lumped capacitance model might be used for the thermodynamic transfer processes involved in hygroscopic growth. Lumped Capacitance Heat-Transfer Model

Consider a particle of volume, V, surface area, A, internal energy, E, density, ␳, and specific heat, c, at temperature, T, and immersed in a fluid at temperature T⬁. The key assumption to be employed is a uniform temperature distribution in the particle. It will be valid if diffusional transport within the particle is much greater than convectional transport with the surrounding fluid. Our analysis will begin with the first law of thermodynamics relating the time rate of change of internal energy to the time rate of change of the temperature of the particle dE dT = ρcV dt dt

(13)

For this general system, we are not considering effects of phase changes because Eq. (13) is used only to introduce the heat–mass–transfer analogy, not to perform calculations for heat transfer for hygroscopic particles in the lung. For a system with no heat sources, a heat balance indicates that all net energy changes within the particle must be the result of a flux through its surface, or

dE = − hA (T − T∞ ) dt

(14)

where h is the average convection heat-transfer coefficient about the particle. Regarding modeling, the role of h is to reduce the complex fluid-dynamic and heat-transfer conditions over the surface of the body into a single value. This simplification is permissible because of the stated assumption of large diffusional transport within the particle. Eqs. (13) and (14), therefore, can be combined to form this simple, first-order, ordinary differential equation dT hA =− (T − T∞ ) dt ρcV

(15)

Using the initial condition T ⫽ T0 at t ⫽ t0, the solution to Eq. (15) is  hA  T − T∞ = exp  − t − t 0 ) ( T0 − T∞  ρcV 

(16)

The leading term in the exponent is the inverse time constant. The time constant is the critical parameter in determining the transient response of a lumped capacitance system and allows one to compare different systems easily. The lumped capacitance assumption can be readily checked by calculating the Biot number (Bi), which is a ratio of convective transport at the

Thermodynamics of Inhaled Hygroscopic Drugs

39

surface to diffusional transport within the particle. For a sphere, the lumped capacitance model is considered valid for Bi =

hr < 0.1 3k

(17)

where r is the radius of the sphere and k is the thermal conductivity of the material. A typical use of the lumped capacitance model would be to curve-fit the exponential function, Eq. (16), to experimental data. The fit would yield the value of the time constant from which h could be calculated. An alternate approach would use an empirical correlation for h in Eq. (16) to calculate the temperature history of the particle considered. For example, for heat transfer between a sphere of diameter D and a surrounding stream, Whitaker (23) proposed the following correlation:

(

)

µ  2/3 NuD = 2 + 0.4 Re1D/ 2 + 0.06 Re D Pr 0.4  ∞   µ

1.4

(18)

where NuD ⫽ h D/k is the average Nusselt number, Pr ⫽ v/␣ is the Prandtl number, v is the kinematic viscosity, ␣ is the thermal diffusivity, ␮⬁ is the absolute viscosity of the free stream, and ␮ is the absolute viscosity at the surface. Lumped Capacitance Mass-Transfer Model

In principle, a similar analysis can be performed for mass transfer by applying the heat–mass–transfer analogy (24). The analogy is based on the similarity of the general heat- and mass-transport equations, implying that a solution to Eq. (6) is a solution to Eq. (7). In this case, the lumped capacitance is a droplet of saline solution; the analysis does not hold for a dry salt particle. Instead of energy, it is the mass of water moving from the air to the droplet that will be simulated. In the mass-transfer system, the concentration of water vapor at the surface of the droplet, c (moles H2O/cm3 air) is analogous to T in the heat-transfer system. The mass balance of water for the droplet can be written [analogous to Eq. (15)] as d (cV (t )) dt

= − h m A (t )(c − c ∞ )

(19)

where hm is the average mass-transfer coefficient and c⬁ is the concentration of water vapor in the free stream distant from the particle. The right-hand side of Eq. (19) is the mass flux of moles of H2O across the boundary layer (Fig. 1). As previously noted, the lumped capacitance model assumes a uniform concentration within the droplet; thus, c is also the concentration within the droplet, and cV represents the total number of moles of H2O that constitute the droplet. Therefore, the left-hand side is the time rate of change in the total number of moles of H2O within the droplet.

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Figure 1

Description of the droplet–air model defined in Eq. (19).

Eq. (19) does not have a simple exponential solution because V(t) and A(t) are functions of time. It is beneficial to change the dependent variable from c to the droplet’s radius, r, where c is related to r in Eqs. (1) to (3). C.

Application of New Theoretical Model(s)

Now, let us consider an explicit example, the hygroscopic growths of saline droplets. For NaCl, the values of i and j in Eq. (1) are 2 and 0, respectively. The droplets will be of sufficient size that the Kelvin effect is negligible—that is, Rw ⫽ 1. Using these values, Eqs. (1) to (3) yield csat c = csat N = (20) m Mw 1+ 2 s mw M s The mass of salt, ms, in solution is a constant. The initial state is a saturated salt solution, so

ms

=

4 3 πr S ρ 3 0

(21)

where r0 is the initial radius, S is the solubility of the salt in water, and ␳ is the density of water (liquid). Similarly mw is mw =

4 3 πr ρ 3

(22)

where r is the current radius. Substituting Eqs. (21) and (22) into Eq. (20) yields c as a function of r csat c c= = sat (23) 3 K Sr M 1 + 2 03 w 1 = 3 r r Ms where K is a constant for a particular salt and particle size. The ambient concentration following Eq. (3) is c ∞ = Hcsat

(24)

Thermodynamics of Inhaled Hygroscopic Drugs

41

Eq. (19) can be transformed into a function of r using Eqs. (23) and (24) and the geometric formulas for the volume and surface area of a sphere. The function is

   4 3 csat  d  πr K   3 1+ 3    r  2  csat = −h m 4 πr  − Hcsat  K dt  1+   r3

(25)

Separation of variables may be employed on this ordinary differential equation and integration performed to obtain an analytical solution for r. Unfortunately, the solution to Eq. (25) is extremely cumbersome; however, it can also be determined numerically. We integrated Eq. (25) with the commercial software Polymath for NaCl (Mw ⫽ 58.44, S ⫽ 0.37), H ⫽ 0.988, and h m having the values 100, 300, and 1000. These solutions are plotted in Figure 2. The effect of increasing h m is to increase the rate at which a droplet approaches equilibrium size. The equilibrium size is determined by mw and H, hm does not affect it. Figure 2 also shows the experimental results of Martonen et al. (25) and Martonen (12). A useful aspect of the heat–mass–transfer analogy is that the empirical heat-transfer correlations apply to the mass-transfer case with a simple

Figure 2

Comparison of predicted particle growth rates with laboratory data.

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exchange of the appropriate terms. Thus, the correlation for the mass transfer between a sphere and the free stream is

(

)

ShD = 2 + 0.4 Re D + 0.06 Re D Sc 1/ 2

2/3

0.4  µ ∞ 

1/ 4

 µ 

(26)

which is Eq. (18) with ShD = hmD/Dw, the Sherwood number, replacing NuD and Sc ⫽ ␯/Dw, the Schmidt number, replacing Pr. D.

Surrogate Laboratory Systems

The bedrock question regarding the role of hygroscopicity in aerosol theory may be expressed thus: How can the hygroscopic properties of a drug be characterized so as to be applicable in a meaningful manner to the human respiratory tract? By hygroscopic properties we mean changes in the parameters of size and density. [Note: The issue of particle shape is, relatively, not too important (except for extreme forms like fibers), because dynamic shape factors can be used to estimate the effects of shape on the motion of airborne particles (26), and the effect of water-vapor uptake will be to create a liquid particle (or droplet).] The two most obvious answers to the question are: (i) develop more physiologically realistic theoretical models (this alternative has been discussed previously), and (ii) design more physiologically realistic laboratory surrogates. In the balance of this text, we shall address the latter subject. To accurately characterize the hygroscopic behavior of inhaled drugs, what is needed is a surrogate lung system that mimics the environment of the human respiratory tract. Then, aerosolized drugs could be introduced into the surrogate system at a prescribed flow rate and the hygroscopic growth rate of a particle measured as a function of its residence time spent inside the surrogate system. Of course, a series of curves would ultimately need to be generated to encompass a range of flow rates corresponding to different ventilatory parameters for different patients. To accomplish this goal, an engineering prototype has been designed and constructed (12). It is a branching network of airways (Fig. 3). The core of the surrogate lung is composed of contiguous sintered stainless-steel tubes. The tubes are naturally porous. An aerosol to be studied is introduced at the top of the core. The core is surrounded by an annular plastic jacket that is divided into many discrete sections. Thermodynamics of Inhaled Hygroscopic Drugs

A detailed view of an individual component is presented in Figure 4. The flow of water (liquid) into each section is independently regulated. The water entering each section is of a selected temperature, different from that of its neighbor. This difference will produce a gradient within the surrogate. The crucial operating point of the engineering lung is that where water vapor can

Thermodynamics of Inhaled Hygroscopic Drugs

43

Physical simulator of human lung airways intended for studies of aerosol hygroscopicity. Abbreviations: T, temperature; RH, relative humidity.

Figure 3

readily diffuse through the porous walls of the sintered stainless tubing, the water (liquid) circulating through the various sections of the jacket cannot. Therefore, the T and H profiles within the surrogate can be controlled by monitoring the temperature of water entering the sections, and regulating the flow rate and temperature of air/aerosol entering the core. The T and H profiles can be chosen to mimic desired conditions in the human lung such as, for example, described by the data of Table 1. Hence, when a hygroscopic aerosol is sampled, it will have passed through a system in which the environment is well defined, stable, and physiologically realistic.

III.

Discussion of Aerosol Drug Hygroscopicity Within Human Airways

In the simulations of hygroscopic particle behavior within airways, as depicted in Figure 5, two distinct conceptual approaches have evolved. For the sake of

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Figure 4 Construction and operating details of the lung simulator. Abbreviations: T, temperature; RH, relative humidity.

discussion, they may be classified simply as temporal and spatial. They are outlined below. A.

Temporal

Particle growth curves describing changes in sizes as functions of time are used. These growth curves may be determined by either abstract (see following case 1) or empirical (see following case 2) techniques. Case 1: The growth of inhaled particles may be described by theoretical formulas. This technique has been advocated, for example, by Ferron (6) and Ferron et al. (17). Case 2: The growth of inhaled particles may be described by experimental data. This technique has been used, for example, by Martonen et al. (9,25) and Martonen and Graham (27,28). Regardless of how the respective curves are determined, the salient point is that particle properties are associated with time, and not, for instance, with position in the respiratory tract. The common drawback of such curves is that both techniques of determining them have incorporated fixed T and H environments. That is, Ferron and colleagues (8,14–16) have assumed constant T and H values in the air surrounding a particle in their mathematical models. Likewise, Martonen and colleagues (5,25,29) have constructed surrogate

Thermodynamics of Inhaled Hygroscopic Drugs Table 1

45

Temperature (T ) and Relative Humidity (H) Values Measured in Human Subjects Respiratory tract atmosphere

Anatomical location Larynx I ⫽ 0 (trachea) I⫽0 I⫽0 I⫽0 I⫽0 I⫽0 I⫽0 I⫽0 I⫽0 I ⫽ 1 (main) I ⫽ 2 (lobar) I⫽3 (segmental) I⫽4 (subsegmental) I ⫽ 4–5 I ⫽10–11

Tin

Tex

Hin

Hex

30.6 ± 0.8 33.0–34.0 34.5 32.9 26.7 31.4–31.9 31.2 32.0 32.2 30.6 32.2 33.0 33.1

36.2 ± 0.3 36.0–37.0 35.8 34.4 – 33.2–33.4 32.6 33.0 33.4 – 33.7 34.0 –

90.0 – – – 82.7 73.6–80.4 – – – 85.8 87.0 – 91.3

99.0 – – – – – – – – – – – –

33.9

–

94.6

–

33.9 34.6

35.0 36.0

– –

– –

The subscripts “in” and “ex” identify the inspiratory and expiratory phases of a breath.

lungs in which the T and H values were prescribed. Therefore, when the respective theoretical (i.e., Ferron and colleagues) and empirical (i.e., Martonen and colleagues) growth curves were applied to the lung to calculate the effects of growth upon deposition, the implicit assumption was that the T and H environment within the lung was fixed. As might be anticipated, the two approaches have separate advantages and disadvantages. For case 1, a key advantage is that once the thermodynamics

Figure 5

Simulation of hygroscopic growth in the lung.

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Katz et al.

of hygroscopicity is formulated in terms of general material properties (e.g., heat conductivity and vapor pressure), the theory may subsequently be applied to any substance. A major disadvantage is that, to achieve solutions to the equations describing changes in particle sizes and densities as functions of surrounding T and H environments, it has proved necessary to make so many assumptions that the results may be of limited or questionable applicability to the human respiratory tract. For example, a critical assumption in the early theory of Ferron (6) was that coupling effects between particles and gases could be ignored. Eisner et al. (13) have clearly demonstrated that the assumption is not valid. However, it has continued to be employed (15,17). Regarding case 2, perhaps its main advantage is that experimental data measuring changes in particle sizes and densities are real (i.e., not abstract, as derived from mathematical models). A disadvantage is that the experiments, to be applicable to the respiratory tract, should be performed with physiologically realistic surrogates of airways that may be quite difficult to design and construct (5,25,29). As shown in Figures 3 and 4, appropriate surrogates have evolved. B.

Spatial

To date, the T and H environment of the human respiratory tract has not been systematically mapped. However, some information is available from measurements within oro- and nasopharyngeal passages and the lung to demonstrate the variations of T and H values to be anticipated within the respiratory tract. The data have been compiled by Martonen and Zhang (4) in a format suitable for studying aerosol hygroscopicity. The information is presented in Table 1. When such an environmental template is superimposed on the respiratory tract, deposition calculations can be performed assuming that particles have sizes and densities corresponding to the T and H values prescribed at specific airway sites, that is, hygroscopicity is related to position, not time. However, to complete such deposition calculations, data describing growth as functions of different T and H environments are needed, and precisely such information has been obtained by Martonen and colleagues in the aforementioned studies. For, although the T and H values were fixed in a given experiment, the T and H parameters were indeed different in different tests. Consequently, the effects of T and H on hygroscopic growth were in fact determined, albeit not in a single test but in a related set of tests. Using this protocol, aerosol deposition computations have been conducted by Martonen and Zhang (4). The differences between the two approaches outlined above (i.e., temporal and spatial) have been discussed by Martonen et al. (9).

IV.

Summary

Ferron and colleagues (6,8,14–17) have developed mathematical models describing processes of hygroscopic growth. The models have two rather severe limitations that compromise applicability to human airways. First of all,

Thermodynamics of Inhaled Hygroscopic Drugs

47

they ignored coupling effects (i.e., interactions) between inhaled particles and surrounding air. Eisner et al. (13), however, clearly demonstrated that the very presence of particles could indeed have a profound influence on thermodynamic conditions in the lung, especially affecting aerosol hygroscopicity. Secondly, Ferron and colleagues assumed that heat- and mass-transfer processes are controlled by molecular diffusion. But, as explained herein, convection is likely the salient mechanism affecting heat- and mass-transfer processes in airways, and may therefore be the more important factor (i.e., rather than molecular diffusion) regulating the behavior of inhaled hygroscopic pharmacologic drugs. A mathematical model describing factors affecting the behavior and fate of inhaled pharmacologic drugs has been defined (18). In a subsequent series of manuscripts, the model has been systematically tested. Specifically, particle deposition patterns have been examined as functions of: (i) patient ventilatory parameters (30); (ii) aerosol polydispersity (20); and (iii) airway morphologies (31). The model has been validated by comparisons of predicted deposition patterns with data from inhalation exposures with healthy, volunteer human subjects (18,26) and clinical data from patients with respiratory tract diseases (20,31). Most recently, the model has been employed in an investigation of the effects of disease-induced changes in airway morphologies upon the targeted delivery of inhaled drugs used in their treatment (32). It should be recognized that the atmospheres within the passages of the head, throat, and lung could exert a great influence upon the deposition of pharmacologic drugs. However, acknowledgment in the medical and pharmaceutical literature of the importance of drug hygroscopicity has been conspicuous in its absence. The objective of this manuscript was to take the next logical step in the analysis of factors affecting the successful administration of inhaled pharmaceuticals, namely, to study the effects of the material properties of a drug on its deposition patterns within the human respiratory tract. The salient question was: How do the physicochemical characteristics of a drug (that is, its size and composition) affect its airborne behavior and resultant sites of distribution? Previous theoretical and experimental studies performed in our laboratory, cited above, relevant to the issues of hygroscopicity and drug delivery, have established that the effects of hygroscopicity are quite complex. For instance, in the medical community, it is widely assumed that the effects of water-vapor uptake are to increase the aerodynamic diameters of inhaled particles. However, for bronchodilators, it has been clearly shown (33) that may not be the case. To be specific, the aerodynamic sizes of inhaled drugs may actually decrease upon the absorption of water vapor, depending on the chemical formulation of the drug and the initial sizes of inhaled particles. Moreover, the extent to which hygroscopic growth can occur within an airway depends on its residence time therein. Therefore, drug hygroscopicity is a direct function of localized fluid-dynamic profiles. In a recent sequence of simulations using the Cray Y-MP supercomputer, the kinetics of air within the human lung has been systematically investigated. The effects of naturally occurring anatomical features, ignored in other drug-deposition models in the current literature, have been quantitated. In order that they are experienced by

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an inhaled bolus of drug, the effects of the laryngeal jet (21), cartilaginous rings (34), and carinal ridge shapes (35) have been formulated. The subject of air motion has been recently reviewed by Martonen et al. (36). Therefore, it became apparent to our laboratory that it would be prudent to develop alternative theoretical and experimental protocols with which to study particle thermodynamics. By identifying the factors that affect hygroscopicity within the human lung, perhaps particle growth can be controlled to target delivery of inhaled pharmacologic drugs. As a direct consequence, the efficacies of the drugs would be enhanced. Since this chapter appeared in the 1st edition of Inhalation Aerosols, work regarding thermodynamics and hygroscopicity has covered pharmaceuticals (37–41), biological aerosols (42,43), and environmental pollutants such as tobacco smoke (44–49). Moreover, studies have addressed fundamental issues affecting atmospheric science (50–53), imaging techniques to measure deposition (54,55), and general efforts at modeling hygroscopic particles (56–60). For the purposes of this textbook, it is useful to categorize these works as they apply to the simulation, and thus prediction, of aerosol deposition within the human respiratory tract. Hygroscopic effects are a secondary, although potentially significant, deposition mechanism such as cloud motion, charge effects, and coagulation. The relative influence of various deposition processes affecting inhaled particles has been addressed by Isaacs et al. (61). Our computational results using hygroscopic growth measurements (7,12) of aerosolized saline solutions with the particle deposition code developed by Martonen and colleagues (7,18,62,63) indicate that hygroscopic effects can be significant. But, they are not a straightforward matter to assess. For example, Figure 6 shows the predicted deposition fractions as a function of lung generations for prescribed particle sizes. The simulations were performed using a Weibel (64) lung morphology, an inspiratory flow rate of 15 L/min, and a tidal volume of 500 mL. In the previous work, it has been clearly shown that a symmetric branching network is an acceptable depiction of human lungs, especially when a population (i.e., not an individual) is being addressed (18). There are two plots in each panel of Figure 6. The simulations assume that particles enter lungs with a specified size D0, and continue on their paths while experiencing water-vapor uptake and physical growth. As a control case, the particles do not experience hygroscopic growth. For 0.1 ␮m particles, hygroscopicity produces a marked decrease in deposition whereas for 2.0 ␮m particles, there is a clear increase in deposition due to hygroscopic growth. But for 4.0 ␮m particles, there is no significant change in deposition. It is important to note that these observations can be explained in terms of the efficiencies of active deposition processes. For instance, submicron particles are deposited primarily by diffusion, and the kinetics of Brownian motion are inversely related to particle size (61). Therefore, when 0.1 ␮m particles increase in size due to the uptake of water vapor, their deposition efficiency decreases. However, larger particles (e.g., 2 ␮m) are deposited most effectively by inertial impaction in the tracheobronchial tree and by sedimentation in the pulmonary compartment, and both mechanisms become more efficient as particles increase in size. For each of the

Thermodynamics of Inhaled Hygroscopic Drugs

49

Figure 6 Deposition fraction as a function of lung generation for monodisperse aerosols under hygroscopic and nonhygroscopic conditions. Original aerodynamic diameters (D0) of 0.1, 2.0, and 4.0 ␮m are considered in panels (A), (B), and (C), respectively.

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Katz et al.

particle sizes considered, the spatial distribution patterns among lung airways do not change as a consequence of hygroscopicity. As the citations (7,18,62,63) indicate, the hygroscopic aerosol deposition code has evolved over the period 1982 to 2005. It has become evermore biologically realistic, especially in terms of airway structures (e.g., cartilaginous rings), branching networks (e.g., asymmetries), flow conditions (e.g., velocity profiles), and clinical implications (e.g., single-photon emission-computed tomography analyses). Our Air Liquide–CyberMedicine research team is creating a generalized simulation tool to be used to optimize the design of inhalation methods for the targeted delivery of pharmaceuticals. Fundamental questions remain regarding hygroscopicity and its influence on aerosolized drugs. Some salient issues are identified below, which are of theoretical and clinical (i.e., practical) interest. First of all, is hygroscopic growth important for a prescribed aerosol under particular flow conditions? Finlay (39) has provided excellent guidance regarding this question using two nondimensional parameters. The first parameter is a guide for the growth of the particle per se. The second parameter is used to predict the need to consider transport to the gas carrying the aerosol such that a coupled analysis is necessary as first studied by Eisner et al. (13) and incorporated in our in silico dosimetry code (see Chapter 1). It would be of value to incorporate such a prediction technique (39) within a simulation tool. Secondly, can traditional mass diffusion models as developed by Ferron and colleagues (14–17) be employed or is it necessary to consider convective transport? In the respiratory tract, this will usually be the case when the flow carrying the particles is turbulent or transitional. But treating turbulence itself is difficult. Stapleton et al. (65) have concluded that better turbulence models are necessary to model flow in the respiratory tract, specifically in the mouth and throat. Martonen et al. (66) have performed original analyses of laminar, transitional, and turbulent flow conditions in extrathoracic airways for nasal breathing. This would be applicable to vaccines administered via the nose. The effects on particle deposition were a sensitive function of volumetric flow rates and particle sizes. Over a wide range of realistic situations, the transitional flow field yielded computational results in best agreement with experimental data describing particle deposition in the nasal region. The work presented in this chapter, namely a methodology for treating hygroscopicity with turbulence models as part of computational fluid-dynamics studies, is important because it is at the forefront of current modeling approaches. As reported herein, there is a wealth of information regarding hygroscopicity in the literature, although this material is not always consistent. Thus, there is a third issue to consider in developing a comprehensive tool, namely, how does one evaluate the reliability of results previously published? For example, some of the new citations used in this 2nd edition have addressed cigarette smoke, which is of great current interest because of its importance as an environmental pollutant of serious health concern. The models presented by Robinson and Yu (45–47) postulate hygroscopicity and coagulation as the causes of enhanced deposition (i.e., hotspots) of cigarette smoke. However, in a systematic series of studies previously published, Martonen and colleagues conducted extensive experimental

Thermodynamics of Inhaled Hygroscopic Drugs

51

and theoretical analyses of cigarette smoke deposition patterns in human airways (48,67–71). In the manuscripts describing the aforementioned laboratory tests, color illustrations were presented, which showed the heterogeneous deposition patterns of cigarette smoke within surrogate airway systems at inspiratory flow rates varying from 15 L/min (71), 30 L/min (67) to 60 L/min (48). These experimental results and related theoretical analyses conclusively determined the importance of cloud motion in the deposition of cigarette smoke over a range of flow rates and, therefore, indicated that there are fundamental errors in the assumptions that are the bases for the models presented by Robinson and Yu (45–47). The work of Broday and Robinson (49) is an example of a different problem in that cloud dynamics was considered but theoretical predictions were not consistent with the experiments and analyses of Martonen and colleagues previously published (48,67–71). In conclusion, the straightforward and seminal point of this chapter is this: the human respiratory system is a warm, humid environment and inhaled, hygroscopic particles will have affinities for the uptake of water vapor. As a direct result, the sizes and densities of aerosolized drugs will change following inhalation. Therefore, the deposition sites of aerosols will not be known by clinicians unless the physicochemical characteristics of these drugs are known. However, if the hygroscopic growth rates of drugs are measured in the manner of Martonen and colleagues (25) the data can be integrated into a validated dosimetry code (63) to predict spatial distribution patterns within human lungs. By accounting for such thermodynamic effects, the efficacies of inhaled drugs may be enhanced by directing them to sites desired by clinicians.

References 1. 2. 3. 4. 5. 6. 7. 8. 9.

10.

Swift DL, Proctor DF. Access of air to the respiratory tract. In: Brain JD, Proctor DF, Reid LM, eds. Respiratory Defense Mechanisms, Part I. New York: Marcel Dekker, 1977:63. Proctor DF, Swift DL. Temperature and water vapor adjustment. In: Brain JD, Proctor DF, Reid LM, eds. Respiratory Defense Mechanisms, Part 1. New York: Marcel Dekker, 1977:95. Martonen TB, Yang Y, Hwang D. Hygroscopic behavior of secondary cigarette smoke in human nasal passages. STP Pharm Sci 1994; 4:69. Martonen TB, Zhang Z. Deposition of sulfate acid aerosols in the developing human lung. Inhal Toxicol 1993; 5:165. Hickey AJ, Martonen TB. Behavior of hygroscopic pharmaceutical aerosols and the influence of hydrophobic additives. Pharm Res 1993; 10:1. Ferron GA. The size of soluble aerosol particles as a function of the humidity of the air. J Aerosol Sci 1977; 8:251. Martonen TB. Analytical model of hygroscopic particle behavior in human airways. Bull Math Biol 1982; 44:425. Ferron GA, Oberdorster G, Henneberg R. Estimation of the deposition of aerosolized drugs in the human respiratory tract due to hygroscopic growth. J Aerosol Med 1989; 2:271. Martonen TB, Menache MG, Hofmann W, Eisner AD. The role of particle hygroscopicity in aerosol therapy and inhalation toxicology. In: Crapo JD, Smolko ED, Miller FJ, Graham JA, Hayes AW, eds. Extrapolation Modeling of Inhaled Particles and Gases: Lung Dosimetry. New York: Academic Press, 1989:303. Morrow PE. Factors determining hygroscopic aerosol deposition in airways. Physiol Rev 1986; 66:330.

52 11. 12. 13. 14. 15. 16. 17. 18. 19. 20. 21. 22. 23. 24. 25.

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Katz et al. Hiller FC. Health implications of hygroscopic particle growth in the human respiratory tract. J Aerosol Med 1991; 4:1. Martonen TB. Development of surrogate lung systems with controlled thermodynamic environments to study hygroscopic particles: air pollutants and pharmacologic drugs. Part Sci Tech 1990; 8:1. Eisner AD, Graham R, Martonen TB. Coupled mass and energy transport phenomena in aerosol/vapor-laden gases. J Aerosol Sci 1990; 21:883. Ferron GA, Soderholm SC. Estimation of the times for evaporation of pure water droplets and for stabilization of salt solution. J Aerosol Sci 1990; 21:415. Soderhohn SC, Ferron GA. Estimating effects of evaporation and condensation on volatile aerosols during inhalation exposures. J Aerosol Sci 1992; 23:257. Soderhohn SC, Anderson DA, Utell MJ, Ferron GA. Method of measuring the total deposition efficiency of volatile aerosols in humans. J Aerosol Sci 1991; 22:917. Ferron GA, Karg E, Peter JE. Estimation of deposition of polydisperse hygroscopic aerosols in the human respiratory tract. J Aerosol Sci 1993; 24:655. Martonen TB. Mathematical model for the selective deposition of inhaled pharmaceuticals. J Pharm Sci 1993; 82:1191. Sirignano WA. Fluid dynamics of sprays - 1992 Freeman Scholar Lecture. J Fluids Eng 1993; 115:345. Martonen TB, Katz IM. Effects of aerosol polydispersity on deposition patterns within human lungs. J Aerosol Med 1993; 6:251. Martonen TB, Zhang Z, Lessmann R. Fluid dynamics of the human larynx and upper tracheobronchial airways. Aerosol Sci Technol 1993; 19:133. Tennekes H, Lumley JL. A First Course in Turbulence. Cambridge. MA: MIT Press, 1972:20. Whitaker S. Forced convection heat transfer for flow in pipes, past flat plates, single cylinders, single spheres, and flow in packed beds and tube bundles. AIChE J 1972; 18:361. Bejan AB. Heat Transfer. New York: Wiley, 1993:576. Martonen TB, Bell KA, Phalen RF, Ho A, Wilson AF. Growth rate measurements and deposition modelling of hygroscopic aerosols in human tracheobronchial models. In: Walton W, ed. Inhaled Particles V. Oxford: Pergamon Press, 1982:93. Stober W. Dynamic shape factors of nonspherical aerosol particles. In: Mercer TT, Morrow PE, Stober W, eds. Assessment of Airborne Particles. Springfield, IL: Charles C Thomas, 1972:249. Martonen TB, Graham RC. Hygroscopic growth: its effect on aerosol therapy and inhalation toxicology. In: Deposition and Clearance of Aerosols in the Human Respiratory Tract. Vienna: Facultas Universitalverlag Ges. m.b.H., 1987:200. Martonen TB, Graham RC. The effect of hygroscopic growth upon the regional dispersion of particulates in human airways. International Society for Aerosols in Medicine, Proceedings of the 6th Congress, Librairie Lavoisier, Paris, 1987:221. Martonen TB, Lowe JE. Measurements of hygroscopic growth rates of medicinal aerosols. In: Liu BYH, Pui DYH, Fissan HJ, eds. Aerosols. New York: Elsevier, 1984:1003. Martonen TB, Katz IM. Deposition patterns of aerosolized drugs within human lungs: effects of ventilatory parameters. Pharm Res 1993; 10:871. Martonen TB, Katz IM. Inter-related effects of morphology and ventilation on drug deposition patterns. J Pharm Sci 1994; 4:11. Martonen TB, Katz IM, Cress W. Aerosol drug deposition as a function of airway disease: cystic fibrosis. Pharm Sci 1995; 12:96. Martonen TB, Wilson AF. The influence of hygroscopic growth upon the deposition of bronchodilator aerosols in upper human airways. J Aerosol Sci 1983; 14:208. Martonen TB, Yang Y, Xue Z. Influences of cartilaginous rings on tracheobronchial fluid dynamics. Inhal Toxicol 1994; 6:185. Martonen TB, Yang Y, Xue X. Effects of carinal ridge shapes on lung air streams. Aerosol Sci Technol 1994; 21:119. Martonen TB, Yang Y, Xue Z, Zhang Z. Motion of air within the human tracheobronchial tree. Part Sci Technol 1994; 12:175. Finlay WH, Stapelton KW. The effect of regional lung deposition of coupled heat and mass transfer between hygroscopic droplets and their surrounding phase. J Aerosol Sci 1995; 26:655.

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59. 60. 61. 62.

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Finlay WH, Stapelton KW, Zuberbuhler P. Errors in regional lung deposition predictions of nebulized salbutamol sulphate due to neglect or partial inclusion of hygroscopic effects. Int J Pharmaceut 1997; 149:63. Finlay WH. Estimating the type of hygroscopic behaviour exhibited by aqueous droplets. J Aerosol Med 1998; 11:221. Johnson DL, Wenger EN, Polikandritou-Lambros M. Aerosolization and hygroscopic growth evaluation of lyophilized liposome aerosols under controlled temperature and relative humidity conditions. Aerosol Sci Technol 1996; 25:22. Asgharian B. A model of deposition of hygroscopic particles in the human lung. Aerosol Sci Technol 2004; 38:938. Reponen T, Willeke K, Ulevicius V, Reponen A, Grinshpun SA. Effect of relative humidity on the aerodynamic diameter and respiratory deposition of fungal spores. Atmos Env 1996; 30:3967. Johnson DL, Pearce TA, Esmen NA. The effect of phosphate buffer on aerosol size distribution of nebulized Bacillus subtilis and Pseudomonas fluorescens bacteria. Aerosol Sci Technol 1999; 30:202. Hofmann W, Morawska L, Bergmann R. Environmental tobacco smoke deposition in the human respiratory tract: differences between experimental and theoretical approaches. J Aerosol Med 2001; 14:317. Robinson RJ, Yu CP. Theoretical analysis of hygroscopic growth rate of mainstream and sidestream cigarette smoke particles in the human respiratory tract. Aerosol Sci Technol 1998; 28:21. Robinson RJ, Yu CP. Coagulation of cigarette smoke particles. J Aerosol Sci 1999; 30:533. Robinson RJ, Yu CP. Deposition of cigarette smoke particles in the human respiratory tract. Aerosol Sci Technol 2001; 34:202. Martonen TB, Musante CJ. Importance of cloud motion on cigarette deposition in lung airways. Inhal Toxicol 2000; 12:261–280. Broday DM, Robinson R. Application of cloud dynamics to dosimetry of cigarette smoke particles in the lungs. Aerosol Sci Technol 2003; 37:510. Berg OH, Swietlicki E, Frank G, et al. Comparison of observed and modelled hygroscopic behaviour of atmospheric particles. Contr Atmos Phys 1998; 71:47. Brechtel FJ, Kreidenweis SM. Predicting particle critical supersaturation from hygroscopic growth measurements in the humidified TDMA. Part I: Theory and sensitivity studies. J Atmos Sci 2000; 57:1854. Brechtel FJ, Kreidenweis SM. Prediciting particle critical supersaturation from hygroscopic growth measurements in the humidified TDMA. Part II: Laboratory and ambient studies. J Atmos Sci 2000; 57:1872. Lehtinen KEJ, Kulmala M, Ctyroky P, Futschek T, Hitzenberger R. Effect of electrolyte diffusion on the growth of NaCl particles by water vapour condensation. J Phys Chem A 2003; 107:346. Eberl S, Chan H, Daviskas E, Constable C, Young I. Aerosol deposition and clearance measurement: a novel technique using dynamic SPET. Eur J Nucl Med 2001; 28:1365. Chan H, Eberl S, Daviskas E, Constable C, Young I. Changes is lung deposition of aerosols due to hygroscopic growth: a fast SPECT study. J Aerosol Med 2002; 15:307. Sarangapani R, Wexler AS. Modeling particle deposition in extrathoracic airways. Aerosol Sci Technol 2000; 32:72. Broday DM, Georgopoulos PG. Growth and deposition of hygroscopic particulate matter in the human lungs. Aerosol Sci Technol 2001; 34:144. Schroeter JD, Musante CJ, Hwang D, Burton R, Guilmette R, Martonen TB. Hygroscopic growth and deposition of inhaled secondary cigarette smoke in human nasal pathways. Aerosol Sci Technol 2001; 34:137. Zhang Z, Kleinstreuer C. Computational thermodynamics analysis of vaporizing fuel droplets in the human upper airways. JSME Int J 2003; 46:563. Zhang Z, Kleinstreuer C, Kim CS. Water vapor transport and its effects on the deposition of hygroscopic droplets in a human airway model. Aerosol Sci Technol 2006; 40:52. Isaacs KK, Rosati JA, Martonen TB. Mechanisms of particle deposition. In: Aerosols Handbook: Measurement, Dosimetry, and Health Effects. Boca Raton, FL: CRC Press, 2005:75–99. Martonen TB, Musante CJ, Segal RA, et al. Lung models: strength and limitations. Respir Care 2000; 45:712–736.

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Katz et al. Martonen TB, Rosati JA, Isaccs KK. Modeling deposition of inhaled particles. In: Aerosols Handbook: Measurement, Dosimetry, and Health Effects. Boca Raton, FL: CRC Press, 2005:113–155. Weibel E. Morphometry of the Human Lung. Berlin: Springer-Verlag, 1963. Stapleton KW, Guentsch E, Hoskinson MK, Finlay WH. On the suitability of k?␧ turbulence modelling for aerosol deposition in the mouth and throat: a comparison with experiment. J Aerosol Sci 2000; 31:739. Martonen TB, Zhang Z, Yue G, Musante CJ. Fine particle deposition within human nasal airways. Inhal Toxicol 2003; 15:283–303. Martonen T, Hofmann W. Factors to be considered in a dosimetry model for risk assessment of inhaled particles. Rad Protect 1986; 15:225. Martonen TB, Horfmann W, Lowe J. Cigarette smoke and lung cancer. Health Phys 1987; 52:213. Martonen TB, Hofmann W, Balásházy I. Physical factors affecting lung deposition of cigarette smoke (with syncarcinogenic radon progeny effects) and mineral fiber particulate matter. In: Wehner AP, ed. Biological Interaction of Inhaled Mineral Fibers and Cigarette Smoke. Columbus, OH: Battelle Press, 1989:159–181. Martonen TB, Hofmann W. Dosimetry of localized accumulations of cigarette smoke and radon progeny at bifurcations. Rad Prot Dosim 1991; 38:81–89. Martonen TB. The behaviour of cigarette smoke in human airways. Am Ind Hyg Assoc J 1992; 53:6.

3 Use of Mathematical Aerosol Deposition Models in Predicting the Distribution of Inhaled Therapeutic Aerosols

DAVID SWIFT† The Johns Hopkins University, Baltimore, Maryland, U.S.A.

I.

BAHMAN ASGHARIAN and JULIA S. KIMBELL CIIT Centers for Health Research, Research Triangle Park, North Carolina, U.S.A.

Introduction

The purposes of this chapter are (i) to review the mathematical models used to predict the deposition of inhaled aerosol particles in the human respiratory tract, (ii) to consider how appropriate it is to use these models for therapeutic aerosol deposition, and (iii) to propose what features may need to be incorporated into future deposition models to allow for their use in predicting therapeutic aerosol deposition. The calculation of particle removal is a function not only of particle size but also of the breathing characteristics and the shape of the respiratory airways, as discussed in the previous chapter. There is a significant history of respiratory modeling with respect to aerosol deposition, stretching back many years, as detailed below. In general, the early models were rather simple in formulation, becoming more detailed as they were improved and elaborated in subsequent years. The models were driven by certain very specific needs, and modelers were quite explicit about stating the purpose of their models and their limitations. The early aerosol-deposition models were generally confined to aerosols that commonly occur in certain industrial settings, such as mining and mineral processing. These aerosols, primarily composed of mineral compounds, had †Deceased.

55

56

Swift et al.

certain properties that made their predicted behavior relatively simple with respect to aerosols that occur in other occupational, environmental, and therapeutic settings. It is thus necessary in applying such models to these other aerosols to take account of their properties. The major models for predicting respiratory deposition will be described, including their scope, purpose, and limitations. They will then be compared with respect to their predicted deposition features, the degree to which these predictions have experimental foundations, the level of detail of the predicted deposition, and the range of conditions addressed. In general, therapeutic aerosols have some unique features that must be considered in applying the models to these aerosols. These features cover a range of special properties, some rather simple in nature and others so “nonideal” as to make simplifying assumptions tenuous and predictive deposition equations practically impossible to derive from theory. The degree to which respiratory deposition of such aerosols can be predicted from the major models will be considered, giving attention to the validity of simplifying assumptions that are often made. Future developments in respiratory modeling that can address the behavior of therapeutic aerosols will be considered. At present, no model exists that is specifically intended to include all major types of therapeutic aerosols. Because the use of therapeutic aerosols has grown significantly, treating both localized and systemic conditions, the motivation to develop such a model has also grown. Modelers are an intrepid group in general and are willing to take on any challenge that is worth the effort. It is likely that there will be models developed, whose specific purpose is to predict the general behavior of therapeutic aerosols. These models will require some experimental verification to be performed to win general acceptance. Thus, improvements will continue to be made in deposition models, and the field of depositions modeling will be an evolutionary interaction among theoretical predictions, experimental measurements, and empirical equations.

II.

Historical Review of Respiratory Tract Deposition Models

A.

Background

Deposition models for particles in the human respiratory tract vary considerably in their degree of detail, range of conditions covered, foundational data (theoretical calculations or experimental measurements), and respiratory zones of consideration. Models that depend wholly or partially on theoretical calculations must have an assumed respiratory tract morphological structure, a model of airflow and particle transport through this structure with predictive equations for the deposition of particles for the mechanisms assumed operative, and an assumed set of breathing conditions. Depending on the purpose of the deposition model, the range of these defining parameters varies significantly. Generally speaking, the early deposition models assumed relatively simple lung morphologies, a small number of breathing conditions (based on conditions in occupational settings), and a limited range of particle sizes.

Mathematical Aerosol Deposition Models

57

Models for a specific respiratory zone, such as the nasal-pharyngeal-laryngeal-tracheal zone or the tracheobronchial (TB) zone, were often presented rather than models for the entire respiratory tract. With improvement in techniques for determining morphology of the respiratory tract, measuring breathing conditions, and calculating particle transport based on fluid flow in the airways, the models have become more sophisticated. Simultaneously, techniques for measuring both overall and regional deposition of aerosols have improved so that experimental verification of models is more readily accomplished. B.

Early Models (1935–1966)

Although experimental measurements of respiratory deposition of aerosols were reported by Drinker et al. (1) and Brown (2), the first mathematical model was presented by Findeisen (3). He divided the respiratory tract into nine generations, beginning with the trachea, progressing through three orders of bronchi and two orders of bronchioles, and terminating with alveolar ducts and sacs. As seen in Table 1, Findeisen assumed branching factors, dimensions, flow speeds, and transit times for each generation. The latter values were based on his assumed breathing pattern of two seconds inspiration and two seconds expiration with a tidal volume (TV) of 400 cm3, resulting in a constant flow rate of 200 cm3 sec1. Findeisen assumed simple expressions for the deposition of particles in each generation resulting from three mechanisms: Brownian motion, sedimentation, and impaction. He also assumed that the particles were spherical in shape and their density was 1 g cm3. He calculated deposition in each generation for seven particle diameters: 0.03, 0.1, 0.3, 1.0, 3.0, 10, and 30 m. For the three smallest sizes, the total deposition fractions in the respiratory tract (percent of inhaled aerosol mass) were 68%, 35%, and 34%, respectively, the deposition essentially being Table 1

Schematic Representation of the Respiratory Tract

Generation

Branching factor Number

Trachea 1 Main bronchi 2 1st bronchi 6 8 2nd bronchi 3rd bronchi 8 Terminal bronchioles 70 Respiratory bronchioles 2 Alveolar ducts 240 Alveolar sacs 2

Diameter (cm)

CrossResidence sectional Velocity time area (cm2) (cm° sec1)a (sec)a

1

1.3

11.0

1.3

150

0.07

2 12

0.75 0.4

6.5 3.0

1.1 1.5

180 130

0.04 0.02

100

0.2

1.5

3.1

65

0.02

770

0.15

0.5

14

14

0.04

54,000

0.06

0.3

150

1.3

0.22

1.1 × 105 2.6 × 107 5.2 × 107

0.05 0.02 0.03

0.15 0.02 0.03

220 8200 1.47 × 105

0.9 0.025 ~0

0.17 0.82 1.2

ventilatory flow rate of 200 cm3 sec1. Source: From Ref. 3.

aFor

Length (cm)

58

Swift et al.

confined to the last two generations. For 1.0 m-diameter particles, he calculated 97.4% total deposition, while for the two largest diameters the deposition was 100%. For 1 m particles, deposition was still primarily in the last two generations, but as particle diameter increased, the site of deposition moved proximally, the 30 m-diameter particles all being deposited in the trachea and first bifurcation (carina). A major shortcoming of the Findeisen model is that it neglected the airways above the trachea and in the nasal and oral passages. Findeisen also assumed a very simple anatomical generation scheme, perhaps based on the best anatomical information at that time. The calculation was limited to a single breathing pattern that was not physiologically reasonable; with the deposition mechanisms assumed by Findeisen, he could have extended the calculation to other, more realistic patterns of breathing. However, as a pioneering work, the Findeisen model established the “ground rules” of most future model calculations, and deserves an important place in the history of particle modeling. The next important model was that of Landahl (4), who modified the features of Findeisen’s model in several important respects. He added two upper-airway compartments, the mouth and the pharynx, and an additional alveolar duct generation [which he later removed in a revised version of his model (5)], and recognized that not all of the inspired air reached the most distal generation. Landahl’s calculations included four breathing conditions: (i) 450 cm3 TV, 15 breaths min1 (bpm), constant 350 cm3 sec1 flow rate; (ii) 900 cm 3 TV, 7.5 bpm, 300 cm 3 sec 1 flow rate; (iii) 1350 cm 3 TV, 5 bpm, 300 cm3 sec1 flow rate; and (iv) 1500 cm3 TV, 15 bpm, 1000 cm3 sec1 flow rate. Five unit-density (p 1 g cm3) spherical-particle diameters were considered—0.2, 0.6, 2.0, 6.0, and 20 m. Quantitatively, the results of Landahl’s calculations were not greatly different from those of Findeisen, with some exceptions. For the largest particles, deposition is still predominantly in the upper airways, but because the mouth and pharynx are included, a significant fraction of these particles is deposited in these airways as well as in the first few bronchial generations. As with Findeisen’s model, deposition shifted more distally as the particle diameter decreased from 20 to 0.2 m for all breathing conditions. For the largest diameter, 20 m, no particles penetrated beyond the eighth generation, the terminal bronchioles (Table 2). Landahl’s model was further refined by the addition of the nasal passage as an alternative upper-airway path (6). Because the nasal passage was assumed to be more efficient than the oral passage as a particle filter, aerosols that passed through the nasal passage had less penetration than those that passed through the oral passage for similar conditions otherwise. Landahl calculated that percent removal of particle diameters ranging from 1 to 40 m for three constant flow rates—72, 300, and 1200 cm3 sec1. Calculated values at certain conditions were compared to measured nasal deposition made by Landahl and his coworkers (7,8) and were found to be in reasonable agreement. Landhal did not publish a full respiratory deposition model including the nasal passage. Using the 12-generation anatomical respiratory schematic model of Landahl, Beekmans (9) presented a deposition model that assumed three

Mathematical Aerosol Deposition Models

59

Calculated Respiratory Tract Deposition Percent by Landahl

Table 2

2 m

0.6 m

0.2 m

tidal volume 0 0 1 2 4 9 7 19 11 25 5 83

0 0 0 0 1 2 2 6 5 25 0 41

0 0 0 0 0 0 1 4 3 8 0 16

0 0 0 0 0 0 1 6 4 11 0 22

300 cm3 sec1 flow, 8 sec cycle, 900 cm3 tidal volume Mouth 14 1 Pharynx 8 1 Trachea 11 1 1st bronchi 13 2 2nd bronchi 17 4 3rd bronchi 20 9 4th bronchi 8 7 Terminal bronchioles 6 24 Respiratory bronchioles 0 10 Alveolar ducts 0 27 Alveolar sacs 0 5 Total deposition (%) 97 91

0 0 0 0 1 2 1 7 7 44 4 66

0 0 0 0 0 0 1 4 6 17 2 30

0 0 0 0 0 1 1 6 6 23 3 40

300 cm3 sec1 flow, 12 sec cycle, 1350 cm3 tidal volume Mouth 14 1 Pharynx 8 1 Trachea 11 1 1st bronchi 13 1 2nd bronchi 18 5 3rd bronchi 21 10 4th bronchi 8 7 Terminal bronchioles 6 24 Respiratory bronchioles 0 12 Alveolar ducts 0 27 Alveolar sacs 0 5 Total deposition (%) 99 94

0 0 0 0 1 2 1 8 11 48 11 82

0 0 0 0 0 0 0 4 3 22 9 38

0 0 0 0 0 0 1 6 5 25 10 47

1000 cm3 sec1 flow, 4 sec cycle, 1500 cm3 tidal volume Mouth 18 1 Pharynx 10 1 Trachea 19 3 1st bronchi 20 5 2nd bronchi 21 12 3rd bronchi 9 20 4th bronchi 1 10 Terminal bronchioles 1 9 Respiratory bronchioles 0 3 Alveolar ducts 0 13 Alveolar sacs 0 18 Total deposition (%) 99 95

0 0 0 1 2 5 3 3 2 26 17 59

0 0 0 0 0 0 1 2 2 10 6 21

0 0 0 0 0 0 1 4 4 13 7 29

Generation cm3

sec1

20 m

300 flow, 4 sec cycle, 450 Mouth 15 Pharynx 8 Trachea 10 1st bronchi 12 2nd bronchi 19 3rd bronchi 17 4th bronchi 6 Terminal bronchioles 6 Respiratory bronchioles 0 Alveolar ducts 0 Alveolar sacs 0 Total deposition (%) 93

Source: From Ref. 5.

6 m cm3

60

Swift et al.

mechanisms of particle deposition—Brownian diffusion, sedimentation, and impaction. Beekmans attempted to correct the airway dimensions for expansion during inspiration, and allowed the volumes of the alveolar ducts and sacs (generations 10, 11, and 12) to change during breathing as the volume of the lung changed. He also recognized the role of mixing of tidal and residual air in the distal three generations, as shown in the experimental work of Altshuler et al. (10). Beekmans assumed equal inspiratory and expiratory times and a pause after each phase was allowed, during which time deposition by sedimentation and diffusion took place. For unit-density spherical-particle diameters ranging from 0.05 to 6.0 m, TVs from 450 to 3000 cm3 and breathing rates from 5 to 15 bpm, he presented calculated results for the total and alveolar (generations 10–12) deposition fractions. For total respiratory deposition, a minimum percent deposition of 15% to 30% occurred at a diameter of 0.3 m, depending on the TV and breathing rate. Total deposition approached 100% at 6 m (as a result of sedimentation and impaction), and increased below 0.3 m owing to diffusional deposition. For alveolar deposition, a minimum in the deposition versus particle-diameter curve occurred at 0.3 m, but above 1 m, deposition fraction increased to a maximum at 2.5 m, then decreased at larger diameters because sedimentation and impaction deposition occurred more proximally (upper and bronchial airways). These results illustrated an important principle, the difference between (i) the deposition efficiency, defined as the ratio of particles deposited in a particular zone to those entering that zone and (ii) the deposition fraction, defined as the ratio of particles deposited in a particular zone to those entering the respiratory tract. Large particles may have high deposition efficiency in the alveolar zones, but because few of them reach these zones, the deposition fraction is small. Calculational models of particle behavior in the respiratory tract invariably demonstrate this tendency, as do experimental measurements. A more detailed anatomical scheme of the human respiratory tract was presented by Davies (11), containing 15 generations, beginning with the mouth and ending at the alveolar sacs. However, there were no depositional models published employing this anatomical arrangement. The lung anatomical schemes published by Weibel (12) were used in many subsequent lung-deposition models, especially the symmetric model A shown in Table 3. It and the nonsymmetrical model (more difficult to use in deposition models) were based on very detailed anatomical examination of several excised normal adult human lungs inflated to approximately 75% of their vital capacity. During the period discussed above, there were a number of experimental studies of lung deposition, beginning with the aforementioned work of Drinker et al. (1) and. Brown (2), and including studies of Van Wijk and Patterson (13), Wilson and LaMer (14), Brown et al. (15), Landahl et al. (16), Altshuler et al. (10), and Dennis (17). The models described above were compared with these experimental studies and, broadly speaking, were in reasonable agreement with most of them. Several different particle-composition aerosols were used in these studies, including NaCl particles, CaCO3, triphenyl phosphate, ZnO, and

Mathematical Aerosol Deposition Models

61

Dimensions of Dichotomous Human Airway Model A by Weibel (12) (4800 cm3 Lung Volume, ¾ Maximal Inflation) Table 3

Name of airway

Generation no.

Trachea 0 Main bronchus 1 Lobar bronchus 2 Lobar bronchus 3 Segmental bronchus 4 Segmental bronchus 5 Bronchus 6 Bronchus 7 Bronchus 8 Bronchus 9 Bronchus 10 Terminal bronchus 11 Terminal bronchus 12 Bronchiole 13 Bronchiole 14 Bronchiole 15 Terminal bronchiole 16 Respiratory bronchiole 171 Respiratory bronchiole 18 Respiratory bronchiole 19 Alveolar duct 20 Alveolar duct 21 Alveolar duct 22 Alveolar sac 23

Number per generation

Diameter (cm)

1

1.8

2

1.22

4

Length (cm)

12.0

Total crosssection (cm2)

Total volume (cm3)

Accumulated volume (cm3)

2.54

30.5

30.5

4.76

2.33

11.25

41.8

0.83

1.90

2.13

3.97

45.8

8

0.56

0.76

2.00

1.52

47.2

16

0.45

1.27

2.48

3.46

50.7

32

0.35

1.07

3.11

3.30

54.0

64 128 256 512 1024

0.28 0.23 0.186 0.154 0.130

0.90 0.76 0.64 0.54 0.46

3.96 5.10 6.95 9.65 13.4

3.53 3.85 4.45 5.17 6.31

57.5 61.4 65.8 71.0 77.2

2048

0.109

0.39

19.6

7.56

84.8

4096 8192 16,384 32,768

0.095 0.082 0.074 0.066

0.33 0.27 0.23 0.20

28.8 44.5 69.4 113.0

9.82 12.45 16.40 21.70

94.6 106.0 123.4 145.1

65,536

0.060

0.165

180.0

29.70

174.8

31 × 105

0.054

0.141

300.

041.80

216.6

2.62 × 105

0.050

0.117

534.0

61.10

277.7

5.24 × 105 1.49 × 106 2.10 × 106 4.19 × 106 8.39 × 106

0.047 0.045 0.043 0.041 0.041

0.099 0.083 0.070 0.059 0.050

944.0 1600 3220 5880 11,800

93.20 139.50 224.30 350.00 591.00

370.9 510.4 734.7 1084.7 1675.0

stearic acid. Most of the studies measured only total deposition (by measuring the aerosol entering and leaving the whole respiratory tract) under several breathing conditions, with the exception of Wilson and LaMer’s studies with radiolabeled NaCl. A model of lung deposition that has been rather widely used is contained in the report of the Task Group on Lung Dynamics to the International Commission on Radiation Protection (ICRP), and later published (18). This model employed a slightly modified version of the Findeisen anatomical scheme, to which was added an additional compartment, the nasopharyngeal

62

Swift et al.

(N-P) airway. For the purpose of calculating the deposition of aerosols in this airway, the empirical equation of Pattle (19) for the inspiratory nasal deposition fraction, NI, was employed: NI 1.2 0.475 log (dP2Q)

(1)

(g cm3),

where is particle density dp is particle diameter (m), and Q is flow rate (cm3 sec1). For the conditions of breathing in this and other calculations, Eq. (1) represents nasal deposition by impaction and is important for particle diameter above 1 m. The quantity dp2 is the aerodynamic equivalent diameter (sometimes written dAED), and the product dp2Q is an impactive parameter that appears in deposition-efficiency equations for the nasal passage, the oral passage, and the bronchial airways, where velocities are relatively large. Calculations were performed for particle diameters ranging from 0.01 to 100 m for three TVs (750, 1450, and 2150 cm3) and a breathing rate of 15bpm. It was assumed that the aerosols were breathed through the nasal airways. The deposition fractions were grouped into three zones—the N-P, TB, and pulmonary (P). The results of the calculations are shown in Figure 1 (18). Note that nasal deposition cannot be characterized by Eq. (1) for particles less than approximately 1 m in diameter. As with previous deposition models, the total deposition fraction demonstrated a minimum at a particle diameter of approximately 0.3 m, and increased monotonically as particle diameter increased or decreased from this minimum. In this deposition model, it was also assumed that the particles were spherical, nonhygroscopic, nonevaporating, nonreactive, and that they behaved inertially as isolated particles (no “cloud effect”). The Task Group model accepted the principle that essentially spherical particles above 1 m in diameter could be assumed to behave according to their aerodynamic equivalent diameter. Furthermore, the calculations were performed for polydisperse aerosols whose mass distribution could be described by a logarithmic-normal distribution, and it was demonstrated that the regional deposition of such an

Figure 1 Regional deposition-fraction curves of the Task Group on lung dynamics as a function of particle diameter. Abbreviations: TV, tidal volume; RPM, respiration per minute. Source: From Ref. 18.

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aerosol would be expressed in terms of the aerosol mass median aerodynamic diameter, so long as the geometric standard deviation, g, did not exceed 3.5, a value rarely exceeded in practice. Although the Task Group lung deposition model was intended for radiation protection purposes, it was widely applied in other situations, including outdoor pollutant particles and occupationally related aerosols. In the same manner, the earlier deposition models were mainly intended to predict the behavior of mineral dusts and other occupational aerosols that normally met the assumptions of the models, such as particle shape, nonhygroscopicity, nonevaporation, and nonreactivity. In some cases, the model was inappropriately applied to situations that did not meet the assumptions, as when Beekmans (9) applied his model to the deposition of cigarette smoke and found significant lack of agreement. It should be noted that Beekmans recognized the hygroscopic nature of some cigarette-smoke particles, and believed that this would lead to disagreement. C.

More Recent Deposition Models (1964–1991)

Following the Task Group model and prior to the more recent U.S. National Council on Radiation Protection and Measurements (NCRP) and ICRP models, almost all the mathematical models that covered the entire respiratory tract have employed the Weibel symmetrical model A. This morphological lung structure was described by Weibel as having 23 generations: generation 0 denoted for the trachea, generations 1 to 16 for the bronchial airways (large bronchi, small bronchi, and bronchioles), and generations 17 to 23 for the gasexchange region (respiratory bronchioles, alveolar ducts, and alveolar sacs). The dimensions of each airway generation were given, allowing for calculation of deposition for each generation for the major mechanisms. The calculation of aerosol deposition on a generation-by-generation basis, known also as the Findeisen-Landahl-Beekmans compartmental model approach, has been used by several other investigators to predict the local deposition for each generation (20,21). There are several assumptions made in these computational models that should be pointed out, such as that the concentration profiles at the entrance of each generation are uniform. An alternative approach was presented by Taulbee and Yu (22) and later by Egan and Nixon (23), in which the human respiratory tract is treated as a continuously expanding duct (“trumpet model”), as one moves more distally into the lungs. In this approach, the phenomena of mixing between the tidal and residual air in the lung can be modeled, and the fact that aerosol deposition is a nonsteady process at the onset of aerosol exposure can be described. The validity of these models cannot be easily determined experimentally, especially for the more distal airways, because the spatial resolution of external detectors (e.g., gamma scintillation cameras) is inadequate for this purpose and small airways overlap larger ones. Thus, it is more common to group the generations into major airway regions, such as the large bronchi, the small bronchi, the bronchioles, and the P region (nonciliated airways and spaces).

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Regional or Local Deposition Models

An alternative approach to the “global” view of the respiratory tract for deposition modeling is a focus on a particular region or subregion such as the nasal passage, the large bronchi, or the alveolar spaces. The long-term objective of these modeling efforts is to piece the regions together to achieve an overall model of aerosol deposition. Investigators who have employed this approach are attempting to model a particular respiratory region over a range of particle diameters and breathing conditions, seeking to obtain a deposition-efficiency expression that can be used in a larger model. As an example, we consider the deposition models for the nasal passages, beginning with the work of Landahl and Black (7), whose experiments consisted of drawing aerosols of several materials into the nasal passages and out the mouth at four constant flow rates. The inlet and outlet size distributions were measured, from which deposition efficiency as a function of particle diameter was calculated. As a follow-up to this study, Landahl (4) developed a theoretical expression for particles in the diameter range from 1 to 40 m, where the mechanisms of deposition for particles were assumed to be impaction and sedimentation. The nasal passage was assumed to consist of four zones, which were modeled as simple geometric shapes. These zones were combined to give a total nasal removal efficiency curve. Empirical equations, such as Eq. (1) presented by Pattle (19) and used by the Task Group, have been extended to include wider particle-diameter ranges and flow rates, and improved based on more subjects studied. Experimental data from five studies were summarized by Heyder and Rudolf (24), who presented separate empirical equations for inspiration and expiration. Using more recent data from nasal replicate models, Cheng et al. (25) extended the diameter range through the ultrafine particle range to 1 nm diameter, so that an empirical equation covering five orders of magnitude of particle diameter can be employed for predicting nasal inspiratory and expiratory deposition efficiency. Their expression for nasal inspiratory deposition efficiency, NI, was NI = 1 exp(0.00168dar2 Q 12.8 D 0.5 Q 0.125)

(2)

where dar is aerodynamic particle diameter (m), D is particle diffusivity (cm2 sec1), and Q is flow rate (L min1). A similar expression was used to describe expiratory deposition efficiency with different constants. The exponential term containing dar2Q represents the impactive mechanism while the term with D0.5Q0.125 represents the diffusional mechanism. The terms are multiplicative, but each term is of negligible effect where the other term is significant. More recently, computational fluid dynamics (CFD) approaches have been used to predict particle deposition in three-dimensional (3D) reconstructions of the upper respiratory tract (URT). Some of the earliest CFD studies of particle transport in the URT include a laryngeal model constructed by Martonen and colleagues (26), an integrated nasal, oral, pharyngeal, laryngeal, and tracheal model of Yu and coworkers (27), and a nasal, pharyngeal, and laryngeal model by Sarangapani and Wexler (28). Like the early lung models, the

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primary focus of most of these early models was toxicological (27,28), although the advantages of particle-deposition prediction for pharmaceutical purposes were also recognized early on (26). The studies by Yu and colleagues were later extended to deposition calculations in nasal, laryngeal, or tracheal regions (29) and Schroeter and coworkers (30) looked at deposition in the nasal valve, turbinate, and olfactory regions within the nasal passages. To date, very few, if any, 3D models using CFD approaches have been directly applied to the transport of therapeutic aerosols in the URT. III.

NCRP Lung-Deposition Model

In view of the new theoretical and experimental information available and the interval since the publication of the ICRP Task Group Report, the NCRP appointed a task group (TG 2) to prepare a new lung deposition and clearance model for radiation protection purposes (31,32). The model consists of a description of the respiratory tract regions (Fig. 2), which contain the clearance pathways, a morphometric model of the adult human nasal and oral airways, a morphometric model of the TB and P regions [the Yeh and Schum (21) model], and a regional deposition curve (Fig. 3) based on equations for sedimentation, impaction, and diffusion of particles onto airway surfaces. The model was incorporated into a computer program that could be employed on a personal computer (PC). The computed regional deposition efficiency curves for the three breathing conditions originally used by the Task

Figure 2 Functional diagram of deposition region and clearance pathway definitions for the U.S. National Council on Radiation Protection and Measurements respiratory tract model. Abbreviations: LN, lymph nodes; A(t), absorption; M(t), mechanical clearance. Source: From Ref. 32.

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Figure 3 Regional deposition-fraction curves of the U.S. National Council on Radiation Protection and Measurements model as a function of particle diameter. Abbreviations: TB, tracheobronchial; P, pulmonary; NOPL, naso-oro-pharyngo-laryngeal. Source: From Ref. 32.

Group (15 breaths min1; 750, 1450, and 2150 cm3 TV) were derived from default values in the program, but these values could be changed to obtain deposition predictions for other conditions. For example, the article contains adopted breathing frequencies and TVs for different individuals of different ages, ranging from infancy to 18 years (adult), obtained by Phelan et al. (33). The report also contains morphometric corrections for the TB and P regions as a function of age that can be incorporated into the computer model as alternative parameters. The program allowed other values to be included into the deposition-calculation model based on future improvements of measurements or new conditions, such as individuals with altered airway morphology due to chronic disease. The NCRP model represented a philosophical break with past models in that it was not confined to healthy adult workers exposed to radioactive aerosols in occupational settings, but recognized the need to provide a flexible tool for a variety of conditions and ages. It allowed for the fact that environmental scientists, toxicologists, industrial hygienists, and pharmaceutical specialists were likely to use the model as well as health physicists. It attempted to use the most current and accurate information on aerosol deposition and clearance in constructing a “user-friendly” computer model designed for practitioners as well as researchers. The model contained, as stated above, default parameters that could be adopted in the absence of more appropriate values, or could be replaced by new values for computational and comparison purposes. A description of the software designed for the model was presented by Chang et al. (34).

Mathematical Aerosol Deposition Models IV.

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ICRP Lung-Deposition Model

The ICRP, likewise, undertook an effort to draw on experimental data that had been published since the Task Group model was published and to develop a new model of aerosol and vapor deposition and clearance. The Task Group on Human Respiratory Tract Models for Radiological Protection was assembled with expertise in the biological, physiological, chemical, radiological, and health physics aspects of inhaled radioactive aerosols and vapors. An overview of the ICRP respiratory tract model was presented by Bair (35), noting that specific deficiencies of the Task Group model were addressed in the development of the new model (36), such as: (i) the Task Group model omitted the dose calculation to the nasal and oral passages; (ii) radioactive gases were omitted; (iii) ultrafine particles were not included (dp 0.01 m); and (iv) the clearance of particles was simplified into three categories—clearance half-times in days, weeks, or years. As with the NCRP model, the ICRP model was intended to take advantage of new developments in lung macro- and micromorphology, new information in respiratory physiology, and progress in computer technology, making complex computation possible on PCs. There were several significant differences between the ICRP and NCRP models. The ICRP model was intended to provide annual limits of intake for radionuclides. It adopted a somewhat more complicated lung morphology, the dimensional aspects of which were taken from the Weibel dichotomous model. The regional divisions of the respiratory tract in the ICRP formalism include two extrathoracic zones (anterior nasal passage and nasopharynx-mouthpharynx-larynx), trachea and large bronchi, small bronchi-bronchioles, and alveolar ducts and sacs. An important reason for the additional subdivision in the ICRP model was the assumed slow-clearance compartments of the large bronchial and small bronchial-bronchiolar regions, having a half-time of approximately 20 days, in contrast to the “normal” fast clearance of the TB region, half-time approximately 100 minutes. This slow-clearance compartment of the TB region was based on the human data of Stahlhofen et al. (37) and animal data of Gore and Patrick (38). As with the NCRP model, the ICRP deposition model included corrections for age, mode of breathing (nasal or oronasal), ethnic factors, and sex. The lung was modeled as a series of particle filters (FindeisenLandahl-Beekmans approach) for the purpose of deposition-efficiency calculations for each region. The calculated fractional depositions for the extrathoracic, bronchial, bronchiolar, and alveolar-interstitial regions are shown in Figures 4 to 7 for nasal-breathing individuals of age ranging from three months to adult (18 years), including an adult woman. Corrections for hygroscopic particles were based on the model approach of Ferron (39), in which the hygroscopic growth or evaporation was assumed to take place entirely within the extrathoracic region of the respiratory tract. Specific examples of growth or evaporation were not given in the model, but the procedure for correction was outlined.

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Fractional deposition curves for the extrathoracic region (excluding anterior nares) for nasal breathing according to the International Commission on Radiation Protection model. Abbreviations: AMAD, activity median aerodynamic diameter; AMTD, activity median thermodynamic diameter; ET, extra-thoracic. Source: From Ref. 54.

Figure 4

V.

General Features of Deposition Models

Deposition models generally recognize the intersubject variability of experimental deposition measurements and adopt the view that the deposition model represents the “average” behavior in the population considered. However, it should be noted that deposition measurements have been made almost exclusively on adult, healthy, male, nonsmoking, Caucasian subjects. Little information exists on subjects outside this rather narrow slice of the population, except for a small number of studies on subjects who have acute and chronic respiratory disease (40,41). The results of these studies indicate both irregular local patterns of deposition and greater overall deposition than in normal subjects. For total deposition in the respiratory tract, the dependence on particle diameter predicts a minimum percent at a particle diameter of 0.4 to 0.8 m.

Figure 5 Fractional deposition curves for the bronchial region for nasal breathing according to the International Commission on Radiation Protection model. Abbreviations: AMAD, activity median aerodynamic diameter; AMTD, activity median thermodynamic diameter; BB, bronchial. Source: From Ref. 54.

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Fractional deposition curves for the bronchiolar region for nasal breathing according to the International Commission on Radiation Protection model. Abbreviation: bb, bronchiolar. Source: From Ref. 54.

Figure 6

At larger particle diameters (expressed as dAED, the aerodynamic equivalent diameter), the percent deposition increases due to (i) impactive deposition in the nasal passage and large bronchi and (ii) sedimentation deposition in the small bronchi, bronchioles, and P spaces. Depending on the breathing rate and TV, the total deposition reaches 100% at a particle diameter of about 10 m. At this diameter, almost the entire deposition takes place in the nasal passage for nose breathing, and this remains so for larger particles. For particles smaller than 0.4 m, the mechanism of particle diffusion becomes important, and the overall deposition increases with decreasing particle diameter. Diffusion does not depend on particle density, so the concept of aerodynamic diameter is not appropriate. The total deposition approaches 100% as the particle diameter reaches 3 nm. At a particle diameter of 1 nm, the total deposition is also 100%, and almost all the deposition takes place in the nasal passage, for nasal breathing. Thus, there is a symmetry between very

Fractional deposition curves for the alveolar-interstitial region for nasal breathing according to the International Commission on Radiation Protection model. Abbreviation: AI, alveolar-interstitial. Source: From Ref. 54.

Figure 7

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large and very small particles, except for the fact that the flow dependence of deposition for very small particles (1–3 nm diameter) is less than that for the large particles (8–10 m diameter). All models include calculations for the bronchial generations between the larynx and the alveolar spaces. Here, the general tendency is for particles in the diameter range 0.5 to 5.0 m to have greater deposition percent in the small airways than in the large bronchi. Particles of diameter 5 to 10 m, which penetrate the nasal or oral passage, have higher deposition in the medium and large bronchi than in the distal airways, especially at high breathing flow rates. Deposition in the P (alveolar) spaces is predicted to have two maxima, one at 3 to 4 m diameter and the other at 0.02 to 0.05 m, due, respectively, to sedimentation and diffusion. Decreases in P deposition percent at diameters above 4 m and below 0.02 m are due to the removal of particles in the more proximal airways by the same mechanisms. All deposition models present equations for the deposition efficiency of the nose and mouth for the mechanisms of impaction and diffusion. They are similar in form to Eq. (2) above. For particles in the diameter range 2 to 10 m and 1 to 20 nm, these deposition efficiencies are important in determining how much aerosol is able to penetrate the upper-airway zone and reach the more distal airways. This is an important feature for aerosols that are intended for delivery to the intrathoracic airways, such as many therapeutic aerosols.

VI.

A Multiple-Path Particle Dosimetry Model

One disadvantage of the ICRP and NCRP models is that they are semiempirically based and, as a result, cannot be used outside the measurement range. In addition, as mentioned before, the models use the Weibel symmetric lung geometry model. Thus, no information regarding the distribution of deposited particles among various airways can be found. This makes the use of these models limited in pharmaceutical applications because lung injury and diseased are most likely related to local deposition rather than regional deposition. A comprehensive deposition model utilizing a trumpet analogy (42) was developed based on the physics of airflow and particle transport in lung airways (43). Particle deposition could be calculated in any geometric structure of the lung. The new deposition model was named multiple-path particle dosimetry (MPPD) in that particle deposition would be calculated in every airway of the lung during a single breathing cycle. The deposition fractions were calculated following several computation steps. First, airflow velocities in all airways of the lung were calculated. Second, aerosol transport equations in an airway and distribution between daughter branches were obtained during a single breath. Third, the combined deposition efficiency of particles in each airway by various deposition mechanisms was calculated. Finally, a mass balance on the traveling particles in each airway was performed, and deposition fraction per airway, generation, region, and lobe could be calculated. Various assumptions were built into the deposition model to facilitate the calculation of particle-deposition fractions. First, the airflow velocity at a

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location within an airway was assumed uniform and proportional to the lung volume distal to that location. Consequently, the airflow rates decrease distally in an airway and within the lung. At the same time, lung airways expanded and contracted uniformly during a breathing cycle that includes inhalation, pause, and exhalation. Particle deposition was calculated at a lung volume midway between the lung at rest (functional residual capacity) and end of inhalation. Second, axial diffusion and dispersion of particles were neglected. Third, particles were insoluble, uncharged, monodisperse spheres, and were uniformly distributed in the tidal air. The assumptions discussed above did not compromise the accuracy of predictions. While the uniform-airflow assumption was used for the distribution of the aerosol among various airways due to convective transport, particle-deposition calculations were based on parabolic velocity profile, which occurs in lung airways a few generations into the lung. Axial diffusion is small for most particle sizes in the breathing range except for nanosized particles, which coagulate quickly to form particles of larger sizes. Cases of other types of particles can be readily added to the existing model. The MPPD model allows investigators to study the effect of lung heterogeneity on particle deposition. It is a valuable tool for predicting the required exposure dose of pharmaceutical aerosols and deposition site in the lung. Regional, lobar, and local deposition can be predicted accurately with this model. The model was also extended to children at different age groups (44). Prediction of regional deposition fraction by MPPD differs from NCRP and ICRP mainly due to differences in head [naso-oro-pharyngo-laryngeal, (NOPL)] deposition (Fig. 8). Only MPPD includes losses by diffusion in the NOPL region. Consequently, TB- and P-deposition predictions by MPPD are lower than that by the NCRP model (Fig. 3) due to filtering in NOPL. Variation in deposition with TV is similar between the two models. Deposition fraction increases with increasing of the TV. The trend is less apparent in the P

Figure 8 Predicted regional deposition-fraction curves of particles by multiple-path particle dosimetry as a function of particle size. Abbreviations: P, pulmonary; TB, tracheobronchial; NOPL, naso-oro-pharyngo-laryngeal.

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region and in the TB region for coarse-size particles because of the filtering of particles in the TB and NOPL regions. MPPD allows deposition calculations of particles per airway, generation, and region in each lobe of the lung. This information is useful for targeting pharmaceutical drugs to specific locations within the lung. TB deposition in each lobe of the lung for tidal breathing of 750 mL and 15 bpm is shown in Figure 9. Predicted deposition shows that deposition fraction in the apical and basal regions of the left and right lobes nearly match. Deposition fraction is highest in the basal lobes followed by that in the apical lobes. Deposition fraction is least in the right middle lobe. Inspections of deposition values indicate that deposition fraction at a location (e.g., region, lobe, etc.) is directly linked to the lung volume for that location. In fact, all deposition curves collapse to a single curve when deposition fraction is normalized with lung volume (44). Similar behavior is observed for lobar deposition of particles in the P region (Fig. 10). The shape of deposition fraction versus particle-size curve is similar in all lobes differing only in magnitude. The two deposition peaks in submicron (ultrafine) and coarse sizes are due to the filtering of particles in the NOPL and TB regions. Lung geometry models for children were developed based on airway measurements of the Utah Biomedical Test Laboratories (45) and published information on adults (12). The geometry models are included in the MPPD model for the prediction of particle deposition in the lungs of children at various age groups. One notable difference between MPPD predictions (Fig. 11) and the calculation results of ICRP (Figs. 6 and 7) is the trend of deposition with age. Lung geometries for children in the ICRP model are estimated by growth curves derived from morphometric measurements. Consequently, there is a well-defined relationship between age and deposition fraction. The geometry models in MPPD were constructed from morphometric measurements for specific ages that included intra-age variability associated with differences in

Figure 9 Tracheobronchial deposition fraction in different lobes of the lung predicted by multiple-path particle dosimetry. Abbreviations: LL, left lower; RL, right lower; LU, left upper; RU, right upper; RM, right middle.

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Pulmonary deposition fraction in different lobes of the lung predicted by multiplepath particle dosimetry. Abbreviations: RL, right lower; LL, left lower; RU, right upper; LU, left upper; RM, right middle. Figure 10

individual sizes. The lung geometries of typical-path (symmetric) and five-lobe symmetric structures were constructed (44). The deposition-fraction curves did not show a pattern with age. VII.

Aerosol Sampling Conventions Related to Respiratory Deposition

While respiratory deposition models were being developed, aerosol sampling conventions were being adopted based on existing deposition and clearance data and other factors related to the practical considerations of sampling. The first of these conventions was the “respirable” sampling definition, based on the concept that certain inhaled aerosol particles could reach the

Figure 11 Predicted lung deposition fraction by multiple-path particle dosimetry with oronasal breathing.

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nonciliated P spaces and potentially produce chronic lung diseases such as silicosis. Conversely, particles that were not respirable could not reach the deep lung, and were thus assumed to be incapable of producing the above health effects. The first respirable sampling definition was that of the British Medical Research Council, and is therefore known as the BMRC definition. It is defined by the equation (3) C/C0 1 (f / fc) where C is respirable concentration of particles having terminal velocity f, C0 is total concentration, and fc is twice the terminal velocity of a 5 m-diameter unit-density sphere. Thus, for a particle diameter dAED having a terminal settling velocity of f, the respirable fraction is C/C0. The definition was also adopted at the Pneumoconiosis Conference of 1960 and published in its proceedings (46). The respirable fraction as a function of aerodynamic diameter is shown in Table 4, where it is seen that the 50% respirable diameter is 5 m. This convention was adopted based on early observations of retained dust in the lungs of miners that did not exceed 5 m. The specific curve was designed to be in conformity with the properties of a particular sampling device, the horizontal elutriator, which separates particles according to their settling velocity under gravitational forces. Shortly thereafter, the U.S. Atomic Energy Commission adopted a separate definition of respirability, intending to establish sampling procedures that would

British Medical Research Council and U.S. Atomic Energy Commission Respirability Definitions

Table 4

Aerodynamic diameter (m)

Respirable percent

British Medical Research Council Definition (36) 0.0 100 2.2 90 3.9 70 4.5 60 5.0 50 5.9 30 6.9 10 7.1 0 U.S. Atomic Energy Commission Definition (37) 2.0 100a 2.5 75 3.5 50 5.0 25 10.0 0 a 90% respirable for 2.0 m adopted by American Conference of Governmental Industrial Hygienists. Source: From Refs. 36–37.

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separate particles that deposited in the P spaces and could produce high local radiation dose because of their insolubility and long retention. The definition adopted was based primarily on the human deposition data of Brown et al. (15). The curve was defined for five aerodynamic particle diameters (also shown in Table 4) with a 50% value at 3.5 m. This sampling convention was found to be met with an inertial cyclone sampler, one version of which could be employed as a personal sampler for occupational exposure estimates. Subsequently, the American Conference of Governmental Industrial Hygienists (ACGIH) adopted the definition into their threshold limit value (TLV) standards (47), with a modified value of 90% respirability at 2 m diameter rather than 100%. The U.S. Environmental Protection Agency (EPA), recognizing the sizesdependent deposition of ambient aerosols, sought to define the size-fractionating characteristics of an ambient aerosol sampler, based on existing deposition and health effects data. The first recommendation of EPA’s Health Effects Research Laboratory was for a 15 m 50% cut point, defining those particles below that sampling curve as “inhalable” dust. Later, in concert with the International Standards Organization (ISO), they adopted a 50% cut point of 10 m, denoting particles within this sampling definition as PM10. Following discussion between ISO and ACGIH, three definitions of particle-size-selective TLVs were proposed (48): (i) inspirable particles, those particles that can enter any part of the respiratory tract; (ii) thoracic particles, those that can penetrate the head airways (nose or mouth) and enter the thoracic region; and (iii) respirable particles, those that can penetrate the TB tree and enter the P-alveolar region. The third definition, it can be seen, is identical to the original definition of respirability, but is not limited to insoluble particles. For the thoracic and respirable definitions, the adopted acceptance curves are cumulative log-normal functions having a geometric standard deviation, g, of 1.5 and median values, respectively, of 10 and 3.5 m diameter. Two definitions were adopted for inspirable particles. The ACGIH curve is defined by the following: E 50[l exp(0.06dAED)]

(4)

where E is percent entering the respiratory tract. At small particle diameters, E approaches 100%, while for large diameters, E approaches 50%. Conversely, the ISO curve is defined by the following: E* 100 15[log10(dAED 1)]2 10 log10(dAED 1)

(5)

for 0 dAED 185 m. This curve is similar to the ACGIH curve at small diameters, but decreases steadily, reaching E 0 at 185 m diameter. Although there are three definitions above for respirable particles, they all refer to the particles that can penetrate to the P-alveolar region. In some instances, the term “respirable” is used to mean something else, such as that fraction that can enter the thoracic region. One must be careful when the concept “respirability” is employed, to make sure how the term is being defined and used.

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Special Features of Therapeutic Aerosols

Therapeutic aerosols are generated by several different means and have several physical forms. The nature of these generation methods and the physical forms of the aerosols are such that models of aerosol deposition discussed above are not appropriate to predict the behavior of these aerosols in the human respiratory tract. The generation of therapeutic aerosols may occur by pneumatic atomization, ultrasonic atomization, dry powder dispersion into air, hydraulic atomization, dispersion in a flash-evaporative propellant, free jet shear of a thin liquid film, or electrostatic atomization. Each of these generation processes has particular aspects that must be considered when attempting to predict the respiratory deposition of aerosols from devices employing these methods. Additionally, the physical form of the particles in the aerosol may be liquid drops, solid particles, solutions, or suspensions. Many such aerosols are inherently unstable physically because they are subject to evaporation, growth, reaction, or mutual repulsion, especially because they pass from the environment of the generation device and the delivery system to the respiratory tract with the subsequent changes of temperature and humidity. Although some environmental aerosols do exhibit some of these features, the models described above employ a set of assumptions (sometimes not explicitly stated), which are not met by all therapeutic aerosols. First, it is implicitly assumed that the aerosol particles in the air surrounding the breathing individual are essentially at rest with respect to their air environment, not traveling at significant velocity with respect to the air being breathed. This means that in the initial inspiratory phase, the particles travel convectively with the inspired air, which can be taken as the initial condition of their trajectory in the respiratory tract. It is also assumed that the particles outside the respiratory tract are spatially uniform so that their incoming concentration is not spatially variable. Furthermore, it is assumed that the separation of the particles is such that each particle behaves inertially, gravitationally, and diffusionally as a “lone” particle in a gas medium. This assumption is equivalent to the absence of a “cloud effect,” whereby neighboring particles have strong fluid-flow interactions that influence their paths and deposition. Particle concentrations, in which the average separation is over 25 times the particle diameter, fulfill this “no cloud effect” condition, but some therapeutic aerosols may not meet this criterion. In addition to the particle velocity assumption, the models are based on the assumption that the gas velocity outside the respiratory tract is not high with respect to the velocity imposed on the inspired air by the act of inspiration. Although breathing in convective wind conditions is “allowed” in the models, the inspiratory flow rate and upper-airway velocity are assumed to be determined by the tidal breathing conditions, so that the velocity “match” between the inspiratory flow and the convective conditions outside the head is reasonable. In most “normal” environment conditions, the velocity of air within the initial portions of the airways is similar to what it would be in “still air.”

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Deposition models generally assume that the particles do not change diameter or shape during their transit through the respiratory tract. There have been a number of models presented to account for the growth or evaporation of particles as they enter the respiratory tract (44,49). Furthermore, the models of evaporation or growth assume a uniform state of temperature and humidity at the respiratory airway surfaces, and that equilibrium is achieved in the entering airstream rapidly with respect to the airway surfaces. Although some therapeutic aerosols meet all these assumptions, many aerosols intended to deliver therapeutic substances do not meet these criteria, and their behavior must be considered in this light. Jet- and ultrasonic-nebulized aerosols are mostly produced from aqueous solutions or suspensions, leading to liquid droplet particles. It has been observed (50) that the evaporative phenomena that occur in the nebulizers before the aerosols leaving the device cause a significant drop in gas temperature, so that the aerosol entering the respiratory tract is saturated with water vapor at the temperature in the nebulizer and must come to temperature equilibrium within the airways eventually with attending equilibrium of the droplets. The final size of the droplets depends on the solute concentration of the droplet solution and the vaporpressure effect associated with small-particle curvature (Kelvin effect). Probably the most difficult aerosol to model with respect to the above assumptions is the metered dose inhaler (MDI) aerosol. In this aerosol generation device, the generation occurs explosively because the propellant containing the therapeutic substance disintegrates while passing out of the metering device at high velocity. As the propellant flash evaporates, the particles undergo very rapid decrease in diameter, from approximately 50 to 75 m diameter to the final size of the suspended particle, or the “dried solute.” The propellant discharge velocity is imposed on the evaporating particles, so that they move through the air downstream of the valve at high velocity, up to 70 m sec1 (51). This particle velocity is markedly higher than the inspiratory velocity of air during any reasonable level of breathing. This high velocity of particles from the MDI may result in inertial deposition on the surfaces of the oral passage before the particles mix with the inspired air surrounding the MDI jet. This assumes that the MDI device is inserted into the oral passage directly, rather than into a “spacer,” where the particles may undergo deposition before they enter the oral airway. Not only are the particles rapidly evaporating at high velocity as they pass through the entrance to the respiratory tract, but also the high-velocity gas jet, consisting of entrained air and evaporated propellant, is mixing with the inspired air, reducing the velocity of the jet and increasing the velocity of the inspired air. This mixing phenomenon will influence particle deposition in the oral airway and thus make it less likely that the deposition models discussed above can predict such aerosol behavior. For example, for particles in the 1 to 10 m diameter range, the models predict that oral deposition will increase with increasing inspiratory flow rate. With MDIs, it has been shown (52) that for inspiratory flow rates from 3 to 50 L min1, the percent deposition decreases with increasing flow rate, assumed to be related to the mixing between the jet and the inspired air.

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Aerosols produced by pneumatic atomization (jet nebulizers), hydraulic atomization, ultrasonic nebulization, and free jet shear atomization (Babington principle) (53) of aqueous solutions or suspensions are all likely to undergo evaporation or growth as they pass into and through the respiratory tract. Such diameter changes result in changes in the settling velocity of the particles and their inertial behavior, resulting in changes in the site and magnitude of deposition. Most such therapeutic aerosol-delivery systems involve such aqueous solutions or suspensions. Similarly, aerosols produced by electrostatic atomization undergo changes in their trajectories due to electrostatic repulsion and image force attraction to the airway surfaces (presuming that the particles retain electrical charge). Although some therapeutic calculations have been presented to predict how such phenomena will influence deposition, few experimental studies have been performed to verify these predictions.

IX.

Future Developments in Deposition Modeling: A Model for Therapeutic Aerosols?

Models of aerosol deposition recently developed have improved in their consideration of some of the more complicated features of the air-flow distribution in the human respiratory tract, and have, likewise, treated some special cases such as age changes of airway dimension and configuration, irregularshaped particles, and wider particle-diameter ranges. Little serious effort has been directed to the case of respiratory-disease changes because it has been difficult to generalize the anatomical and functional changes that occur in the myriad disease states. The only case of a “nonclassical” aerosol that has received significant attention has been the hygroscopic particle aerosol; most other situations in which the aerosol does not fit the assumptions listed above have received little attention. Even the hygroscopic aerosol situation has generally been limited to aerosols of common hygroscopic character such as NaCl aerosols, or solutions of such solutes. Despite the improvements in the deposition models, it must be accepted that the present description of fluid flow in the airways is somewhat simplified. Considerable improvement in this situation is likely in the future as CFD methods are more often combined with improved anatomical descriptions of the respiratory airways, ranging from the head airways to the P-alveolar spaces. What effects are these improvements likely to have on the prediction of aerosol deposition in the respiratory tract? It is not easy to say based on present knowledge. However, it should be noted that experimental measurements of deposition in “normal” human subjects have yielded results that show considerable biological variability, and this variability cannot be accounted for at present using functional or anatomical measurements. Thus, the improvements might get “lost” in the biological variability. On the other hand, some deposition effects might be quite significant when the phenomena discussed above are allowed for.

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What is the likelihood of being able to develop a specific model for predicting the deposition of therapeutic aerosols? If it were possible, would it be of value? There is some motivation for such a development in view of the growing need to design aerosol generation and delivery systems that can better predict regional and local deposition. This is particularly true in view of the increased specificity and likely cost of new therapeutic substances that promise great advances in the prevention and treatment of a number of respiratory diseases. The efforts to improve the prediction of aerosol deposition in environmental situations using new theoretical and experimental techniques ought to be complemented by similar efforts for therapeutic aerosols. Even though the special conditions for therapeutic aerosols present special challenges, some improvement can be realized by efforts to treat these situations. A model that covers all situations in a general fashion is not likely to have the same “simplicity” as does the present model for environmental aerosols. Especially in the area of predicting aerosol behavior in individuals with respiratory diseases leading to marked changes in lung morphology and flow distribution, the likelihood of a general model is remote. To date, there has been considerable effort in theoretical modeling for some cases without accompanying experimental verification. As with environmental aerosols, there needs to be a continual interaction between theory and experiment, and improvements in experimental techniques should be encouraged.

X.

Summary and Conclusions

We have sought in this chapter to review the development of respiratory-deposition models, to consider how applicable they are for therapeutic aerosols, and to suggest how improvements can be made to make existing models more applicable to the special conditions that are found in many therapeutic aerosol situations. We have also considered whether it is feasible to modify or remake existing models into a new model (or models) that is uniquely appropriate to therapeutic aerosols with their special characteristics. Respiratory deposition models have undergone considerable expansion and sophistication in the years since the first models were presented by Findeisen (3) and Landahl (4). This is due to better knowledge of respiratory anatomy and physiology and to increasing knowledge of airflow and deposition mechanisms that occur in the airways. Improvements in deposition prediction have been incorporated into theoretical models and confirmed by experimental studies. Nevertheless, considerable biological variability remains that cannot be strictly accounted for, and it is unlikely that better flow models will significantly improve this situation. However, changes in deposition due to breathing parameters and changing particle characteristics can be better accounted for. Existing models rest on a set of assumptions that must be made to arrive at simple predictive equations, but these assumptions are not met for many therapeutic aerosols. It can therefore be concluded that such aerosols require

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new models that take into account the failure to meet these assumptions. It is not obvious that general predictive deposition equations can be developed to cover most of the situations with therapeutic aerosols, but efforts to cover some of the less complicated situations should be undertaken. It is concluded that efforts to improve the predictive models for environmental aerosols should be matched by similar efforts for therapeutic substances in aerosol form, in view of the considerable payoff such efforts could realize. Thus, the challenge is for deposition modelers and pharmaceutical scientists to collaborate in the development of models that will both predict behavior of existing aerosol systems and lead to improved design of new systems that can better target specific drugs in known quantity to desired sites in the respiratory tract.

Acknowledgment Revision of this chapter was funded by the National Institutes of Health (NIH/NHLBI 1 R01 HL073598).

References 1.

2. 3. 4. 5. 6. 7. 8. 9. 10. 11. 12. 13. 14. 15.

Drinker P, Thompson RM, Finn JL. Quantitative measurements of the inhalation, retention and exhalation of dust and fumes by man. I. Concentration of 50 to 450 mg per cubic meter. J Ind Hyg Toxicol 1928; 10:13–25. Brown CE. Quantitative measurements of the inhalation, retention, and exhalation of dusts and fumes by man: II. Concentration below 50 mg per cubic meter. J Ind Hyg Toxicol 1931; 13:285–291. Findeisen W. Uber das Absetzen kleiner, in der Luft suspendierter Teilchen in der menschlichen Lunge bei der Atmung. Arch Ges Physiol 1935; 236:367–379. Landahl HD. On the removal of airborne droplets by the human respiratory tract. I. Lung Bull Math Biophys 1950; 12:43–56. Landahl HD. Schematic representation of respiratory tract (Table 3.2). In: Hatch TF, Gross P, eds. Pulmonary Deposition and Retention of Inhaled Aerosols. New York: Academic Press, 1964. Landahl HD. On the removal of airborne droplets by the human respiratory tract. 11. The nasal passages. Bull Math Biophys 1950; 42:161–169. Landahl HD, Black S. Penetration of airborne particulates through the human nose. J Ind Hyg Toxicol 1947; 29:269–277. Landahl HD, Tracewell T. Penetration of airborne particulates through the human nose. II. J Ind Hyg Toxicol 1949; 31:55–59. Beekmans JM. The deposition of aerosols in the respiratory tract. Can J Physiol Pharmacol 1965; 43:157–172. Altshuler B, Yarmus L, Palmes ED, et al. Aerosols deposition in the human respiratory tract. Arch Environ Health 1957; 15:292–303. Davies CN. A formalized anatomy of the human respiratory tract. In: Davies CN, ed. Inhaled Particles and Vapours. Oxford: Pergamon Press, 1961:82–87. Weibel ER. Morphometry of the Human Lung. Berlin: Springer-Verlag, 1963. Van Wijk AM, Patterson HS. The percentage of particles of different sizes removed from dust–laden air by breathing. J Ind Hyg Toxicol 1940; 22:31–35. Wilson IB, LaMer VK. The retention of aerosol particles in the human respiratory tract as a function of particle radius. J Ind Hyg Toxicol 1948; 30:265–280. Brown JH, Cook KM, Ney FG, et al. The influence of particle size upon the retention of particulate matter in the human lung. Am J Public Health 1950; 40:450–458.

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Landahl HD, Tracewell TN, Lassen WH. On the retention of airborne particulates in the human lung. Arch Ind Hyg Occup Med 1951; 3:359–366. Dennis WL. Discussion of paper by C. N. Davies. In: Davies CN, ed. Inhaled Particles and Vapours. Oxford: Pergamon Press, 1961:88–90. Task Group on Lung Dynamics. Deposition and retention models for internal dosimetry of the human respiratory tract. Health Phys 1966; 12:173–208. Pattle RE. The retention of gases and particles in the human nose. In: Davies CN, ed. Inhaled Particles and Vapours. Oxford: Pergamon Press, 1961:301–311. Gerrity TR, Lee PS, Haas FJ, et al. Calculated deposition of inhaled particles in the airway generations of normal subjects. J Appl Physiol Respir Environ Exercise Physiol 1979; 47:867–873. Yeh H-C, Schum GM. Models of human lung airways and their application to inhaled particle deposition. Bull Math Biol 1980; 42:461–480. Taulbee DB, Yu C-P. A theory of aerosol deposition in the human respiratory tract. J Appl Physiol 1975; 38:77–85. Egan MJ, Nixon W. A modelling study of regional deposition of inspired aerosols with reference to dosimetric assessments. In: Dodgson J, McCallum RI, Bailey MR, Fisher DR, eds. Inhaled Particles VI. Oxford: Pergamon Press, 1988:909–918. Heyder J, Rudolf G. Deposition of aerosol particles in the human nose. In: Walton WH, ed. Inhaled Particles IV. Oxford: Pergamon Press, 1977:107–125. Cheng Y-S, Yeh H-C, Swift DL. Aerosol deposition in human nasal airway for particles 1 nm to 20 m: a model study. Radiat Prot Dosim 1991; 38:41–47. Martonen TB, Zhang Z, Lessmann RC. Fluid dynamics of the human larynx and upper tracheobronchial airways. Aerosol Sci Technol 1993; 19:133–156. Yu G, Zhang Z, Lessmann R. Fluid flow and particle diffusion in the human upper respiratory system. Aerosol Sci Technol 1998; 28:146–158. Sarangapani R, Wexler AS. Modeling particle deposition in extrathoracic airways. Aerosol Sci Technol 2000; 32:72–89. Martonen TB, Zhang Z, Yue G, et al. 3-D particle transport within the human upper respiratory tract. J. Aerosol Sci 2002; 33:1095–1110. Schroeter, J.D., Kimbell, J.S., Asgharian, B. Analysis of particle deposition in the turbinate and olfactory regions using a human nasal computational fluid dynamics model. J. Aerosol Med. 2006, In press. National Council on Radiological Protection and Measurements (NACRE). Deposition, Retention and Dosimetry of Inhaled Radioactive Substances. Report 125, Bethesda, MD, 1997. Phelan RF, Cuddihy RG, Fisher GL, et al. Main features of the proposed NCRP respiratory tract model. Radiat Prot Dosim 1991; 38:179–184. Phelan RF, Oldham MJ, Kleinman MT, et al. Tracheobronchial deposition predictions for infants, children and adolescents. In: Dodgson J, McCallum RI, Bailey MR, Fisher DR, eds. Inhaled Particles VI. Oxford: Pergamon Press, 1988:11–21. Chang IY, Griffith WC, Shyr LJ, et al. Software for the draft NCRP respiratory tract dosimetry model. Radiat Prot Dosim 1991; 38:193–199. Bair WJ. Overview of ICRP respiratory tract model. Radiat Prot Dosim 1991; 38:147–152. International Commission on Radiological Protection. Human respiratory tract model for radiological protection. Annals of the ICRP, 1994. (ICRP Publication 66, Vol. 24, Nos. 1–3.) Stahlhofen W, Scheuch G, Bailey MR. Measurement of the tracheobronchial clearance after aerosol bolus inhalation. In: Dodgson J, McCallum RI, eds. Inhaled Particles VII. Oxford: Pergamon Press, 1994:189–196. Gore DJ, Patrick G. A quantitative study of the penetration of insoluble particles into the tissues of the conducting airways. In: Walton WH, ed. Inhaled Particles V. Oxford: Pergamon Press, 1982:149–161. Ferron GA. The size of soluble aerosol particles as a function of the humidity of the air. Application to the human respiratory tract. J Aerosol Sci 1977; 8:251–267. Laube BL, Swift DL, Wagner HN, et al. The effect of bronchial obstruction on central airway deposition of a saline aerosol in patients with asthma. Am Rev Respir Dis 1986; 133:740–743. Dolovich MB, Sanchis J, Rossman C, et al. Aerosol penetrance: a sensitive index of peripheral airways obstruction. J Appl Physiol 1976; 40:468–471.

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Swift et al. Yu CP. Exact analysis of aerosol deposition during steady state breathing. Powder Technol 1978; 21:55–62. Asgharian B, Hofmann W, Bergmann R. Particle deposition in a multiple-path model of the human lung. Aerosol Sci Technol 2001; 34:332–339. Asgharian B, Ménache MG, Miller FJ. Modeling age-related particle deposition in humans. J Aerosol Med 2004; 17(3):213–224. Mortensen JD, Schaap RN, Bagley B, et al. Final report: a study of age specific human respiratory morphometry. Technical Report TR 01525-010, University of Utah Research Institute, UBTL Division, 1983. Orenstein AJ, ed. Proceedings of the Pneumoconiosis Conference, Johannesburg, 1959. London: J. and A. Churchill, 1960:620. American Conference of Governmental Industrial Hygienists. Threshold Limit Values of Airborne Contaminants for 1968. Cincinnati: ACGIH, 1968. Air Sampling Procedures Committee. Particle Size-Selective Sampling in the Workplace. Cincinnati: ACGIH, 1985. Robinson RJ, Yu CP. Theoretical analysis of hygroscopic growth rate of mainstream and sidestream cigarette smoke particles in the human respiratory tract. Aerosol Sci Technol 1998; 28:21–32. Mercer TT, Tillery MI, Chow HY. Operating characteristics of some compressed-air nebulizers. Am Ind Hyg Assoc J 1968; 29:66–78. Clark AR. Medical aerosol inhalers: past, present, and future. J Aerosol Med 1993; 6:224. Swift DL. Apparatus and method for measuring regional distribution of therapeutic aerosols and comparing delivery systems. J Aerosol Sci 1992; 23(suppl 1):S495–S498. Litt M, Swift DL. The Babington nebulizer: a new principle for generation of therapeutic aerosols. Am Rev Respir Dis 1972; 105:308–310. James AC, Stahlhofen W, Rudolf G, et al. The respiratory tract deposition model proposed by the ICRP Task Group. Rad Prot Dos 1991; 38:159–165.

PART II:

BIOLOGICAL CONSIDERATIONS

4 Physiology and Pharmacology of the Airways

RALPH J. ALTIERE AND DAVID C. THOMPSON University of Colorado Health Sciences Center, Denver, Colorado, U.S.A.

I.

Introduction

Understanding inhalation aerosol therapy requires a knowledge of lung function, particularly as it relates to the mechanical properties of the lung during the process of ventilation. This chapter is intended to supply basic information on lower airway physiology and pharmacology. The upper airways, consisting of the nasopharynx region, have not been included, even though the mouth and pharynx are critical regions to consider in inhalation aerosol therapy because of the potential for significant impaction and loss of aerosolized substances. Although emphasis has been placed on lung function under normal conditions, alterations in lung function caused by pathophysiologic states have been incorporated into the text where appropriate. Numerous monographs, books, and articles have been written on the subjects of both pulmonary physiology and pharmacology. It is not the intention of this chapter to provide an extensive discussion of these subject matters but rather to present an overview of each discipline with references to more comprehensive written materials. The chapter is divided into two major sections: pulmonary physiology and pulmonary pharmacology. In the first section, emphasis is placed on physiological aspects of lung function, including lung structure and function, pulmonary ventilation, mechanical properties of the lungs, pulmonary function tests, and altered lung function in several prominent pulmonary disease conditions. The section on pulmonary pharmacology considers basic pharmacodynamics and 83

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explores briefly the pharmacology of agents affecting airway smooth muscle and mucous secretion. Several topics related to airway pharmacology, such as transport mechanisms, drug metabolism, and pharmacokinetics are discussed in chapters X, Y, and Z of this volume. II.

Physiology of the Lung

A.

Structure and Function of the Lung

General Features

Lung function is intimately linked with the structural features of the airways and lung parenchyma. Knowledge of the structural characteristics provides a basis for understanding lung function in normal and abnormal physiological conditions and for understanding the opportunities and difficulties associated with aerosol inhalation delivery of drugs. The airways of the lungs provide a pathway of normally low resistance to the bulk flow of air into and out of the lung periphery where alveoli perform the essential function of gas exchange. Based on this concept, the lungs have been divided into two general compartments or zones—the conducting zone and the respiratory zone as depicted in Figure 1 (1,2). The conducting zone consists of the first 16 generations of airways comprising the trachea (generation 0), which bifurcates into the two mainstem bronchi that further subdivide into bronchi that enter two left and three right lung lobes. The intrapulmonary bronchi continue to subdivide into progressively smaller diameter bronchi and bronchioles. The conducting zone ends with terminal bronchioles that are devoid of alveoli. Accordingly, the function of the conducting zone is to move the air by bulk flow into and out of the lungs during each breath. The respiratory zone consists of all structures that participate in gas exchange and begins with respiratory bronchioles containing alveoli. These bronchioles subdivide into additional respiratory bronchioles, eventually giving rise to alveolar ducts and finally to alveolar sacs. The acinus is defined as the unit comprising a primary respiratory bronchiole, alveolar ducts, and alveolar sacs. Several models of airway branching have been developed (2–5). Models of branching airways are shown schematically at the bottom of Figure 1. Two characteristic features of branching airways strongly influence lung function: decreasing airway caliber and increasing airway surface area. As the airways branch, they become smaller in diameter and length but greater in number and cross-sectional area. In keeping with the patterns of branching observed in human lungs, models have been constructed that are consistent with development of the airways, namely dichotomous branching where each branch bifurcates to form two new branches. The two daughter airways from the same parent may be similar (regular dichotomy) or may differ (irregular dichotomy) in diameter and length, but, in general, the diameter of the daughter airways is smaller than that of the parent. The branching ratio is 2 and, accordingly, the number (N) of branches in each generation (z) is N(z) ⫽ 2z. This pattern of branching and decreasing size holds true for the conducting airways, i.e., through generation 16. The decrease in airway caliber (from 1.8 to 0.06 cm) along with a relatively small increase in total

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Model of airway morphology: (Upper) Conducting (trachea—terminal bronchioles) and respiratory (respiratory bronchioles—alveolar sacs) zones of the airways. (Lower) Regular dichotomy (A) and irregular dichotomy (B) of airway branching. Source: From Refs. 1, 2.

Figure 1

cross-sectional area from trachea to terminal bronchioles (2.5–180 cm2) ensures optimal conditions for bulk flow of air through the larger airways down to the terminal bronchioles (6). In the adult, the total volume of conducting airways is approximately 150 mL (designated as the anatomic dead space) that is exchanged by bulk airflow when breathing 500 mL of air at rest.

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In the terminal airways that comprise the respiratory zone and acini, airways continue dichotomous branching, but diameters of respiratory bronchioles and alveolar ducts change very little with each new generation (from 0.06 to 0.024 cm). Thus, the total airway cross-sectional area nearly doubles with each generation beyond generation 16 (from 180 to 10,000 cm2). The total alveolar surface area approaches 140 to 160 m2 in the adult human (7). As a result of the enormous increase in surface area, bulk flow of air decreases rapidly within the respiratory zone until movement of air within alveoli occurs entirely by diffusion (Fig. 2) (5,8). The presence or absence of cartilage is another structural feature of the airways that undergoes significant change along the tracheobronchial tree and influences airway function (9,10). The trachea consists of a single tube structure composed of C-shaped cartilage rings connected posteriorly by smooth muscle and connective tissue. It provides the path of least resistance to airflow due to its large diameter, the relatively rigid structure provided by the cartilage rings and the limited range over which it can restrict its internal diameter through contraction of the smooth muscle. Inside the thoracic cavity, the trachea bifurcates into two main stem bronchi. These bronchi also contain cartilage rings and offer little resistance to airflow due to their large diameter and rigid structure. As the airways penetrate the lung parenchyma and continue dichotomous branching, cartilage continues to decrease in size and amount, forming irregular shaped plates rather than rings around the airways. The reduced cartilage content renders the bronchi progressively less rigid and more compliant as well as collapsible. Bronchioles are distinguished from bronchi by the lack of cartilage altogether, making them far less rigid and more distensible and collapsible than bronchi.

Generation-dependent changes in airway cross-sectional area and flow of air. As the generation of the airways increases (i.e., diameter decreases but number of airways increases), the velocity of airflow (as measured by mass flow velocity) decreases and airway cross-sectional area increases. The respiratory zone is identified by the shaded area. Source: From Ref. 1.

Figure 2

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Under normal conditions, the bronchi and bronchioles remain patent, allowing the free flow of air into and out of the respiratory zone. In disease states, these small airways may become occluded or may collapse, thereby preventing air from entering or leaving the distal respiratory areas supplied by the obstructed airway. These and other pathophysiological events that adversely affect lung function will be addressed in more detail in separate sections below. Another structural feature of the distal airways within the respiratory zone is the presence of channels for collateral flow between acini and between alveoli (11). Interalveolar pathways are known as pores of Kohn. Although they were considered artifacts of fixation for many years, they now are recognized to be present in the lungs of all mammalian species studied, although to varying extents. The pores vary in size from 3 to 13 ␮m and in shape from cylindrical to hourglass. The relative physiological importance of these pores in collateral ventilation between adjacent alveoli is not resolved. In addition to pores between alveoli, there also exist other channels for collateral ventilation. Martin (12) demonstrated the presence of interbronchiolar or interductal channels in dog lungs. Short tubular interacinar ducts subsequently were confirmed to exist in human lungs (13). An alternate pathway was described by Lambert (14), in which channels connect respiratory bronchioles and alveolar ducts and sacs in adjacent acini and have since been found in human lungs (15,16). Together, these three types of collateral channels provide alternate pathways for airflow. The relative contribution of each type of channel depends upon the specific condition of the lungs, e.g., interbronchiolar channels would supply air to an adjacent acinus deprived of air by occlusion of its primary bronchiole, whereas bronchiolar–alveolar channels would provide collateral flow to an area peripheral to an obstructed alveolar duct. However, neither of these alternative pathways would be useful if obstruction occurred in a larger, more central bronchus. In this circumstance, interalveolar pores may provide collateral flow between adjacent units of the lung supplied by different bronchi. Other morphological and histological features of the airways change along the tracheobronchial tree and can have important influences on lung function under normal and pathophysiological conditions. Numerous cell types contribute to the characteristic features of different regions of the airways, such as epithelial cells that line the lumen of the airways, glands that secrete mucus, and smooth muscle that surrounds airways, nerves, blood vessels, and resident cells within the airways. Epithelium

The entire respiratory tract is lined with a continuous sheet of epithelial cells that vary in type and function throughout the tracheobronchial tree (Table 1) (17–19). The epithelial lining of the airway lumen serves principally to separate the external environment of inhaled air from the internal environment of the subepithelial airway structures. Protection of the internal airway structures afforded by the epithelial cells is accomplished by mucous secretions and by the specialized tight junctions between epithelial cells that limit the penetration

88 Table 1

Altiere and Thompson Airway Surface Epithelial Cells

Cell Airways Ciliated columnar

Location

Goblet

Trachea through respiratory bronchioles Trachea and bronchi

Basala

Trachea and bronchi

Clara

Bronchioles

Serousb Brushc Neuroendocrine (Kultschitsky; APUD) Alveoli Type I pneumocyte

Sparse Sparse Bronchi

Type II pneumocyte Type III pneumocytec (alveolar brush)

Principal function Mucociliary escalator; glycoprotein secretion Mucus secretion; progenitor for ciliate cells Progenitor cells; aid in attachment of columnar cells to basement membrane Glycoprotein secretion; progenitor for ciliated and Clara cells; xenobiotic metabolism Serous fluid secretion Transitional cell Chemoreceptor, paracrine functions Gas exchange surface; fluid transport Surfactant secretion; progenitor for type I cells Remains to be established

aFunction

of basal cells not resolved. in fetal lung and several nonhuman mammalian species cFound in nonhuman mammalian species; rarely in human. Source: From Refs. 17, 18, 22–24. bFound

of inhaled substances through intercellular channels into the subepithelial areas (5,20). Large molecules must then pass through epithelial cells and may be subject to metabolism (refer to chapters X and Y in this volume). In disease states, the epithelial cell lining may be damaged or disrupted, leading to increased penetration of inhaled substances (20,21). The most prominent epithelial cell is the ciliated columnar cell (20). These cells line the airways from trachea through terminal respiratory bronchioles and decrease in size in proportion to the airways. Its primary function is to move mucus upwards from the lower airways to the trachea and pharynx where it is swallowed or expectorated. Ciliated cells also secrete glycoproteins that contribute to the mucous layer that covers the epithelium. Goblet cells that secrete mucus are present throughout the larger airways down to the small bronchi but are not found in bronchioles (25). A third prominent type of epithelial cell found in the trachea and bronchi is the basal cell. The basal cell has the potential to serve as a progenitor cell for goblet and ciliated cells. Experiments in laboratory animals suggest, however, that goblet cells may be progenitor cells for ciliated cells (26,27). Clara cells, rather than goblet cells, are

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found in bronchioles of human lungs (25,28). Like goblet cells, these cells are not ciliated and secrete a mucus-like substance. They also serve as stem cells for ciliated and mucus-secreting cells (29–31). Clara cells also participate in the metabolism of xenobiotic substances. Alveoli are lined primarily by two epithelial cell types (32). Type I alveolar epithelial cells comprise the most prominent cell type (17,33). It is through type I cells that gases diffuse to allow for oxygen and carbon dioxide exchange with pulmonary capillary blood. Type II cells produce a surfactant that lines the luminal surface of the alveoli and also serve as progenitor cells for type I cells (17,34). Alveolar brush cells have been identified in alveoli of the nonhuman mammals and tentatively termed the “type III pneumocytes” (22). Their function remains to be elucidated, although chemoreceptor and/or fluid transport functions have been speculated. Gland Cells

Submucosal glands lie beneath the surface epithelium and are found in the trachea and in cartilaginous airways but are absent from noncartilaginous bronchioles (Fig. 3). These glands consist of four regions from distal to proximal ends: serous tubules lined with serous-secreting cells (these tubules form acini or blind-ended tubules); mucous tubules lined with mucus-secreting cells; collecting ducts composed of columnar epithelial cells; and a ciliated excretory duct. Serous fluid formed in the distal tubules migrates into mucous tubules and carries mucus with it into collecting ducts and finally into ciliated excretory ducts that expel the mucus onto the luminal surface of the airway epithelium (35–39). Mucociliary Escalator

The mucus produced by surface epithelial cells and submucosal glands contains a complex mixture of mucous glycoproteins, immunoglobulins, and

Figure 3

Histological characteristics of the bronchus and bronchiole. Source: From Ref. 40.

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numerous other substances. The mucous layer is biphasic and consists of a lower viscosity sol layer in contact with the ciliated surface epithelium and an upper more viscoelastic mucus or gel layer. This biphasic mucous layer serves several important functions. It protects the epithelium from dehydration, helps to humidify inhaled air to saturation, and provides a protective barrier through trapping of inhaled particles and through antimicrobial and metabolic properties (39). The mucous layer and substances trapped in it (inhaled particles, bacteria, viruses, drugs, etc.) are removed from the lungs by the upward movement of mucus in the tracheobronchial tree caused by the coordinated beating of cilia on surface epithelial cells (17,20,41–43). This movement is referred to as the mucociliary escalator. The rate of mucous movement varies with the airway region and is determined by the number of ciliated cells and beat frequency of cilia. Movement is slower in small airways than in larger airways where there is a relatively greater number of ciliated cells having a greater ciliary beat frequency. More rapid movement of mucus in large airways is necessary to accommodate the inflow of mucus from the larger number of smaller airways. Airway mucus eventually reaches the pharynx where it is either swallowed or expectorated. The rate of mucous clearance can be an important consideration in drug delivery to the lungs, especially when prolonged duration of action within the airways is desirable. The drug may become trapped in and removed by the mucociliary escalator. Mucous production and clearance are affected by many factors (42,44,45). Coughing, for example, greatly increases mucous clearance by providing a large propulsive force. Conversely, ciliary dysfunction or hypersecretion of thick, viscous mucus greatly slows or prevents mucous clearance, as might occur in severe asthma, chronic bronchitis, or cystic fibrosis. In these situations, mucous plugging can obstruct airways, prevent movement of air into and out of distal airways and alveoli, and provide an environment for recurring infections. Under these conditions, the delivery of drugs to the airways, as well as their therapeutic efficacy, can be adversely affected by the production of a thicker mucous layer through which the drugs penetrate more slowly and by retarding or preventing drug penetration to the site of action due to mucous plugging of the airway lumen. Smooth Muscle

Smooth muscle is found throughout the tracheobronchial tree (40). It is separated from the epithelium by the lamina propria, a region containing connective tissue, blood vessels, and nerves. In the trachea, smooth muscle forms a layer on the posterior surface joining the ends of the C-shaped cartilage. In contrast, smooth muscle completely surrounds the central bronchi and bronchioles, acting like a sphincter to constrict the airways when activated to contract. Smooth muscle surrounding the airways is found down to the level of terminal bronchioles. Strands of smooth muscle can be found in respiratory bronchioles and as far as alveolar ducts, although greatly diminished in amount. As will be described, numerous substances including endogenous neurotransmitters, hormones, and mediators, as well as numerous drug entities, act on airway

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smooth muscle to either activate the muscle to contract or to inhibit activation and cause the muscle to relax. When activated sufficiently, smooth muscle can cause extensive narrowing of the airways or bronchoconstriction, especially in terminal bronchi and bronchioles, causing a rise in resistance to airflow as occurs in asthma. Conversely, relaxation of airway smooth muscle leads to maintenance of an open airway with little resistance to airflow. In disease states, airway smooth muscle can also be stimulated to increase in size (hypertrophy) or cell number (hyperplasia or proliferation) such that greater contractility and/or impingement on the airway lumen can give rise to physiological and structural reduction of luminal diameter and increased resistance to airflow. Resident Cells

The two most prominent cells that reside within the airways are mast cells and alveolar macrophages. Alveolar macrophages are migrating mononuclear cells found in the interstitium and lumen of alveoli (41). These cells phagocytize foreign substances, including particulates and microorganisms, and remain within the alveolus or migrate to the mucociliary escalator or lymph tissue. When activated, these macrophages release numerous biologically active mediators and enzymes (48) that can affect airway function. Mast cells found in the walls and lumen of central and peripheral airways (49) contain various biologically active mediators (both preformed and newly generated) that directly alter airway function including smooth muscle, mucous-secreting cells, and vascular cells (50,51). Under pathophysiological conditions, other cells such as neutrophils, eosinophils, and lymphocytes can be found within the airways. These cells release numerous biologically active substances that can affect vascular permeability, airway smooth muscle, mucous secretion, and nerves either directly or indirectly through actions on endogenous neurotransmitters and other mediators. Discussion of the many effects of these cells is beyond the scope of this chapter. Innervation

The airways are innervated by afferent sensory nerves and efferent motor nerves (Table 2) (52–54). Sensory nerves respond to various stimuli in both normal and potentially adverse conditions and usually evoke a response in the airways through a central nervous system—mediated reflex arc (52). Slowly adapting receptors or pulmonary stretch receptors are located in smooth muscle of central airways and respond to airway stretch during breathing patterns. They are thought to participate in reflex control of respiratory drive. Rapidly adapting receptors or irritant receptors are located within epithelium of central airways and are sensitive to various stimuli such as chemical or irritant stimuli (e.g., inhaled particulates or inflammatory mediators released locally within the airways), mechanical stimulation, and pulmonary microembolism. When activated, these receptors trigger a reflex bronchoconstriction through activation of efferent cholinergic motor nerves. Afferent C-fibers are located in the epithelium and between smooth muscle cells (55). These nerves

92 Table 2

Altiere and Thompson Principal Innervation of the Airways

Nerves Afferent Slowly adapting receptor (pulmonary stretch receptor) Rapidly adapting receptor (irritant receptors) Unmyelinated C-fibers Efferent Cholinergic Adrenergic Nonadrenergic noncholinergic inhibitory

Proposed function Breuer-Hering inflation reflex Sensation of inhaled irritants, mechanical stimuli, pulmonary emboli, or edema Sensation of inhaled irritants, mechanical stimuli Bronchoconstriction, mucous secretion Vasoconstriction Bronchodilation, mucous secretion

contain peptide transmitters such as tachykinins (substance P and neurokinin A), calcitonin gene-related peptide, and others. Afferent C-fibers are activated by chemical, particulate, or mechanical stimuli that evoke a reflex bronchoconstriction through efferent cholinergic nerves (56), or, under specific experimental conditions, may elicit reflex bronchodilation through the activation of noncholinergic nonadrenergic inhibitory nerves (57–59). Stimulation of afferent C-fibers, e.g., by inhaled irritants, also can evoke antidromic release of tachykinins from the nerve terminals within the airways. The tachykinins released directly into the airways have been shown to have numerous effects, including airway smooth muscle contraction, neuromodulation, and, most importantly, proinflammatory effects such as increased vascular permeability and mucosal edema (60,61). Efferent nerves to human airways consist of autonomic excitatory cholinergic nerves and inhibitory nonadrenergic noncholinergic nerves, both of which are carried in the vagal nerve trunk (62,63). Cholinergic nerves innervate airway smooth muscle and submucosal glands. When activated, these nerves release acetylcholine (ACh) from their nerve terminals and cause contraction of airway smooth muscle and bronchoconstriction (64), as well as an increase in mucous and serous secretion (65,66). Nonadrenergic noncholinergic nerves provide the sole inhibitory neural input to human airways and, when activated, cause relaxation of airway smooth muscle and bronchodilation (62,63). Although the neurotransmitter of these nerves has not been identified with certainty, nitric oxide may play a critical role in mediating nonadrenergic noncholinergic responses in human airways (67–69). These nerves also may influence mucous secretion (70). Human airways have no functional adrenergic innervation to smooth muscle, but adrenergic nerves contribute to regulation of mucous secretion (44,65) and bronchial and pulmonary blood flow. Despite the lack of a functional adrenergic innervation, airways contain ␤-adrenergic receptors that respond to circulating epinephrine and mediate bronchodilatory effects, as evidenced by the powerful smooth muscle relaxant actions of numerous ␤-adrenergic receptor bronchodilators used to treat bronchospasm associated with asthma and similar conditions.

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Vasculature

Blood supply to the lungs is divided between the pulmonary and systemic circulations (71–76). The pulmonary circulation consists of the vascular bed that leaves the right heart via the pulmonary artery, branches into a dense, sheet-like pulmonary capillary bed that surrounds the alveoli, and finally coalesces to form pulmonary veins that drain into the left heart. Thus, 100% of the cardiac output flows through the pulmonary circulation. The principal function of the pulmonary circulation is the gas exchange of oxygen and carbon dioxide with air in the alveoli. Deoxygenated blood enters the pulmonary circulation from the right heart, and oxygenated blood returns to the left heart. To accommodate the entire cardiac output, the pulmonary circulation operates as a low-pressure, low-resistance vascular bed. Under normal conditions, pulmonary blood vessels are dilated with little or no intrinsic arterial smooth muscle tone (in contrast to the systemic circulation), giving rise to a mean pulmonary artery pressure of 15 mmHg (vs. systemic mean pressure of 100 mmHg). An important property of pulmonary blood vessels is that they have relatively thin walls comprising less connective tissue and smooth muscle than systemic vessels and, therefore, they are distensible and collapsible. The combination of low-pressure, distensible, collapsible vessels and interstitial and alveolar pressures determined by factors other than gravity (see section below on pulmonary mechanics) gives rise to asymmetrical distribution of blood flow in the lungs that is sensitive to gravity. Three major flow conditions in the upright position can result from these factors: zone 1 in the upper portion of the lungs where alveolar pressure ⬎ arterial pressure ⬎ venous pressure, thereby collapsing pulmonary blood vessels and stopping flow; zone 2 in the midportion of the lungs where arterial pressure ⬎ alveolar pressure ⬎ venous pressure such that blood flow is regulated by compression of microvessels at the venous outflow from alveolar walls (expanded alveoli can compress these structures); and zone 3 in the lower portion of the lungs where arterial pressure ⬎ venous pressure ⬎ alveolar pressure (due to hydrostatic pressure effects on pulmonary blood pressure but not on alveolar pressure) such that blood flow is independent of alveolar pressure and dependent on cardiovascular parameters (driving pressure from right heart, vascular geometry, and vascular smooth muscle tone). Under normal conditions, the pulmonary circulation operates in zones 2 and 3, except that far upper portions of the lungs may experience zone 1 conditions during diastole. Pulmonary blood flow is regulated by passive and active mechanisms. Passive mechanisms involve recruitment and distension of pulmonary microvessels. Such passive regulation occurs, for example, in exercise when cardiac output is increased. To accommodate the increased blood flow without large increases in pressure, pulmonary microvessels are distended and blood flow in them increases without a substantial rise in pressure. Thus, the lung operates more in zone 3 conditions. A second regulatory mechanism involves active processes. The most prominent active regulatory mechanism in the pulmonary circulation is the response to alveolar hypoxia known as hypoxic

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pulmonary vasoconstriction. This mechanism is closely linked to the gas exchange function of the pulmonary circulation. When alveoli are poorly ventilated, resulting in low oxygen tension, pulmonary arteries constrict, thereby reducing blood flow to these poorly ventilated regions of the lung and directing blood flow to better ventilated regions. In certain pathophysiological conditions, e.g., when large regions of the lungs are poorly ventilated, as might occur in severe cases of chronic bronchitis, pulmonary artery pressure rises, which places a greater load on the right heart and can result in right heart hypertrophy and failure. Pulmonary vascular pressure also can be affected by numerous other endogenous substances that can constrict or dilate pulmonary blood vessels (77). The pulmonary circulation also serves other important purposes related to metabolic functions, whereby endogenous substrates are converted to bioactive hormones (e.g., conversion of angiotensin I to angiotensin II by angiotensin-converting enzyme located on pulmonary vascular endothelial cells), bioactive substances are degraded (e.g., bradykinin, 5-hydroxytryptamine, and peptide hormones), or bioactive mediators are released into the circulation (e.g., prostacyclin) (78–83). Chapter X of this text provides more detailed descriptions of these functions. The lungs receive a second blood supply by way of the systemic circulation (71,76). This blood supply is commonly referred to as the bronchial circulation. It has a normal mean systemic blood pressure of 100 mmHg, and perfusion of lung structures by the bronchial arteries is therefore not affected by changes in body position or lung pressures as is the pulmonary circulation. The bronchial circulation supplies oxygenated blood to the trachea and intrapulmonary airways from central bronchi down to bronchioles. The larger bronchial arteries originate from the aorta and intercostal arteries. The bronchial circulation provides nutrients to all structures of the tracheobronchial tree, including the epithelium, smooth muscle, nerves, and glands, as well as pulmonary arteries and veins and the pleura surrounding the lungs. The bronchial circulation normally does not perfuse terminal respiratory units (respiratory bronchioles, alveolar ducts, and alveoli) that receive nutrients from the pulmonary circulation. Although venous drainage of the bronchial circulation varies among species, in humans, it appears that about half the bronchial circulation, i.e., that portion, which perfuses the trachea and extrapulmonary bronchi, drains into the systemic circulation and the right heart via bronchial veins. The remainder, i.e., the portion that perfuses intrapulmonary airways and other structures, appears to empty into the pulmonary veins and the left heart via bronchopulmonary anastomoses. Thus, drugs delivered to the lower airways can enter the systemic circulation through absorption into the bronchial circulation or into alveolar capillaries of the pulmonary vascular bed. The bronchial circulation subserves several important and vital functions. It participates in air conditioning by assisting in the humidification and warming of inspired air in the trachea and bronchi. It aids in gas exchange and nutrient supply to alveoli when the pulmonary circulation is obstructed (e.g., thrombosis or embolism) by dilating, enlarging, and forming connections with precapillary arteries in the pulmonary circulation to supply blood flow to alveolar capillaries.

AQ1

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It also plays a central role in inflammatory conditions of the lung by contributing to mucosal edema and delivery of inflammatory cells and mediators to the airways. The bronchial circulation is the only angiogenic portion of the lung circulation in the adult, i.e., all new blood vessel growth comes from the bronchial circulation and not from the pulmonary circulation. Thus, bronchial vessels vascularize all new tissue in the lungs, such as scar tissue or tumors. Lung Functions

Lung functions can be broadly subdivided into two principal categories— respiratory and nonrespiratory functions. Respiratory functions include the two principal respiratory activities of gas exchange and acid–base balance. Nonrespiratory functions include all other aspects of lung function, such as endocrine and metabolic activities. Nonrespiratory functions of the lung have been alluded to in the previous sections, and further details can be found elsewhere in this volume. A brief summary of respiratory functions follows. Gas Exchange

The airways can be subdivided into two main regions: the conducting zone and the respiratory zone (as described above in the subsection on general features of lung structure). Exchange of oxygen and carbon dioxide gases, which comprises the principal function of the lungs, occurs in the respiratory zone, specifically in respiratory units, comprising respiratory bronchioles, alveolar ducts, and alveoli. All airways in which gas exchange does not occur comprise the conducting zone. The conducting zone acts as a conduit for airflow between the environment and the respiratory zone. The volume of air within the conducting zone is not available for gas exchange and is termed the anatomic dead space. In addition to its function as a conduit for airflow, the conducting zone performs two other important functions, gas buffering and air conditioning. Dead space air volume is about 150 mL in the adult and comprises a large fraction (30%) of a normal tidal breath of 500 mL (although it comprises a small fraction (⬍5%) of total lung volume; vide infra). It contains a mixture of gases having a composition between that of inspired air and expired alveolar air after gas exchange has occurred, i.e., it contains relatively less oxygen and more carbon dioxide than inspired air. Thus, dead space air buffers the air entering the alveoli from the external environment. A critical function of the conducting zone is air conditioning. As already mentioned in the section on the vasculature, the bronchial circulation helps to maintain the water content and temperature of the large central airways needed to humidify and warm inspired air, such that air reaching alveoli is saturated with water and is at body temperature. The respiratory zone provides the structures needed to permit efficient gas exchange between the respiratory and cardiovascular systems. Exchange of oxygen and carbon dioxide occurs across the alveolar epithelium, capillary endothelium, and their respective basement membranes having a total distance of less than 0.5 ␮m. The total gas exchange area is approximately 120 to 160 m2 in the adult human lung. The combination of a large surface area and a short

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diffusion distance contributes to efficient gas exchange. In some regions of the alveolus, an interstitial space comprising connective tissue elements separates the alveolar epithelium and capillary endothelium basement membranes. These structures provide a longer distance through which gases must travel for exchange to take place and also appear to be the sites where water and solute exchange occur (84). Edema in these structures can hinder efficient gas exchange. Passage of solutes through these and other lung cells has been described (85) and will be discussed in chapter X of this volume. Gas exchange across the alveolar–capillary barrier occurs by diffusion. The direction and extent of gas diffusion is determined by the concentration or partial pressure gradient of gases between alveolar air and capillary blood. The partial pressure of oxygen in alveoli is normally 104 mmHg, whereas in deoxygenated pulmonary capillary blood, it is about 40 mmHg. Hence, the driving force is for oxygen to travel from the alveolus to capillary blood. Carbon dioxide has the opposite pressure gradient, normally having a greater partial pressure in pulmonary capillary blood (45 mmHg) versus alveolar air (40 mmHg), thereby providing a driving force for carbon dioxide to leave capillary blood and enter the alveolus. Techniques are available for measuring diffusion capacity, and the reader is referred to other sources for description of these methods (86,87). Under optimal conditions, all alveoli would be well ventilated, and all pulmonary capillaries would be well perfused, such that ventilation and perfusion are perfectly matched. However, such conditions often do not exist. When alveoli are well ventilated but associated capillaries are poorly perfused, e.g., due to thrombosis, embolism, etc., ventilation to these alveoli is wasted and is termed “alveolar dead space.” The sum of alveolar and anatomic dead space is termed “physiologic dead space.” In contrast, conditions may exist such that alveoli are not well ventilated (due to bronchoconstriction, mucous plug, atelectasis, etc.), but perfusion of these regions continues. In this situation, deoxygenated blood contributes to a physiological shunt that delivers inadequately oxygenated (i.e., deoxygenated) blood to the left heart and systemic circulation. Under such conditions, ventilation and perfusion must be matched to maintain optimal gas exchange. Mechanisms such as hypoxic pulmonary vasoconstriction (vide supra) aid in ventilation-perfusion matching by shunting blood flow away from poorly ventilated areas to alveoli with adequate ventilation. Acid–base balance is closely linked to lung function through the exchange of carbon dioxide (75,86,88). Carbon dioxide formed through cellular metabolism is in equilibrium with carbonic acid, bicarbonate, and hydrogen ion in blood. Control of carbon dioxide content in blood by gas exchange in alveoli greatly influences blood hydrogen ion concentration and, thereby, acid–base balance. Thus, ventilation of the lungs has a direct effect on blood pH as shown: ↓ ventilation → ↑ [CO2] in blood → ↑ blood [H+] → respiratory acidosis ↑ ventilation → ↑ [CO2] in blood → ↓ blood [H+] → respiratory alkalosis Accordingly, alterations in ventilation can influence blood pH. For example, impaired ventilation, which may result from airway obstruction or central

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nervous system depression and reduced ventilatory drive, would lead to respiratory acidosis. Conversely, hyperventilation, as may occur at high altitudes or by fever or emotional circumstances, would cause respiratory alkalosis. Alternatively, pulmonary ventilation acts to control acid–base disturbances caused by nonrespiratory factors such as altered kidney function, vomiting, diarrhea, etc. The change in pH triggers the appropriate alteration in respiration, such that changes in the rate and depth of ventilation can compensate for the acid–base imbalance and return pH back toward normal. This mechanism occurs through chemoreceptors sensitive to hydrogen ion concentrations that in turn alter central nervous system control of respiratory drive. B.

Evaluation of Pulmonary Function

Tests of pulmonary function provide objective, quantifiable measurements of lung function and are utilized for various purposes. They are used diagnostically to evaluate diseases that affect heart and lung function (although they usually cannot provide a definitive diagnosis of specific disease), to screen persons at risk for pulmonary disease, to assess prognosis, and to assess preoperative risks. Pulmonary function tests also are used to monitor the effectiveness of drug therapy for lung or heart disease, to monitor the course of disease or occupational exposure of injurious agents on lung function, and to detect adverse reactions to drugs with known pulmonary toxicity. Understanding pulmonary function tests and their usefulness require familiarity with basic concepts of lung physiology. A brief overview of these concepts will be provided below. For more detailed information on these topics, the reader is referred to the numerous publications that are available that describe in detail basic pulmonary physiology (86–90). Pulmonary Ventilation

One of the basic concepts of respiratory physiology involves lung volumes and capacities. Determination of lung volumes and capacities can provide important information on the status of lung function, as discussed below. Definitions and estimates of these values are shown in Table 3. The amount of air moving into or out of the lungs, viz. tidal volume, inspiratory and expiratory reserve volumes, and vital capacity, can be measured through spirometry. A diagrammatic representation of a spirometer tracing is shown in Figure 4. Estimates of volumes of air remaining in the lungs after expiration, e.g., residual volume and functional residual capacity, are made by gas dilution methods (for volume of air in open airways only) or body plethysmography methods (for determination of all gas within the lungs whether they are well ventilated, poorly ventilated, or not ventilated). Measurements of lung volumes generally are normalized to body size (height, weight, or surface area), age, and sex of the subject, thereby allowing comparison with standardized or predicted lung volumes. In this manner, lung pathophysiology can be detected using relatively simple procedures, although changes in lung volumes may not be diagnostic of specific pulmonary diseases. For example, vital capacity may be decreased by pulmonary factors such as

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Table 3

Lung Volumes and Capacities in Healthy Subjects Volumesa (mL)

Volume or capacity Tidal volume (VT) Inspiratory reserve volume (IRV) Expiratory reserve volume (ERV) Residual volume (RV) Inspiratory capacity (IC) Functional residual capacity (FRC) Vital capacity (VC)

Total lung capacity (TLC)

Definition Volume of air inspired or expired during a normal breath Maximal volume that can be inspired after a normal tidal inspiration Maximal volume that can be expired after a normal tidal expiration Volume of air remaining in the lungs after a maximal expiratory effort Maximal volume of air that can be inspired after a normal tidal expiration ⫽ VT + IRV Volume of air remaining in the lungs after a normal tidal expiration ⫽ ERV + RV Maximal volume of air that can be expired from the lungs after a maximal inspiration ⫽ IRV + VT + ERV Volume of air in the lungs after a maximal inspiratory effort ⫽ IRV + VT + ERV + RV

Female (20–30 yr)

Male (20–30 yr)

Male (50–60 yr)

400b

500b

500b

1800b

2700b

2100b

1000

1600

1000

1000

1600

2400

2200

3200

2600

2000

3200

3400

3200

4800

3600

4200

6400

6000

aValues bTidal

are approximate volumes in healthy seated subjects (86). volume is variable; IRV varies depending on VT

reduction in lung distensibility (as occurs in pneumonia or alveolar edema), increases in lung stiffness (e.g., due to interstitial edema and respiratory distress syndrome), and lung hyperinflation due to air trapping leading to increased residual volume and functional residual capacity (as may occur in emphysema or airway obstruction in asthma). Vital capacity also can be decreased by extrapulmonary factors such as limited thoracic expansion (as occurs in kyphoscoliosis or pleural fibrosis), limited descent of the diaphragm (e.g., in pregnancy or abdominal ascites), or nervous and muscular dysfunction affecting respiratory muscle activity. Measurement of lung volumes provides a static picture of lung function. The dynamic function of the lungs is the process of ventilation or the rate of movement of air into and out of the lungs. The most crucial aspect of this process is alveolar ventilation, i.e., the amount of air that ventilates the alveoli where gas exchange occurs. Alveolar ventilation must keep pace with the metabolic needs

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Figure 4 Example of a spirometric tracing demonstrating different measures of lung volumes and capacities.

of the body to supply oxygen and dispose of carbon dioxide, i.e., the rate of alveolar ventilation per minute is the important factor. Determination of the total volume of air that a person breathes each minute is a simple matter of measuring the expired volume of air for each breath (represents VT) and the number of breaths over a fixed period of time, e.g., 500 mL per breath and 15 breaths/min yield a total ventilation (also termed minute ventilation) of 7500 mL/min. This rate represents the total amount of air that moves into and out of all airways and alveoli. To determine alveolar ventilation, the amount of air present in the conducting zone of the airways where gas exchange does not occur must be subtracted out of the calculation. This volume is known as anatomic dead space and is approximately 150 mL in an adult male. Thus, alveolar ventilation is (500–150 mL) ⫻ 15 breaths/min ⫽ 5250 mL/min. It should be recognized that equivalent levels of total ventilation can yield dramatically different levels of alveolar ventilation. For example, rapid, shallow breathing of 250 mL VT at 30 breaths/min yields a total ventilation of 7500 mL/min but alveolar ventilation of only (250–150 mL) ⫻ 30 breaths/min yields a total ventilation of 3000 mL/min, which may be

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inadequate to meet metabolic needs. Conversely, slow, deep breathing can result in excessive alveolar ventilation, e.g., 750 mL VT at 10 breaths/min yields a total ventilation of 7500 mL/min and an alveolar ventilation rate of 6000 mL/min. Larger tidal volumes dilute anatomic dead space air to a greater extent, thereby providing more fresh air to alveoli. In addition to anatomic dead space, areas of the lung may be well ventilated but poorly perfused such that gas exchange is greatly diminished. As noted previously, the volume of air that ventilates alveoli but is not exchanged with pulmonary capillary blood is known as alveolar dead space. The combination of anatomic and alveolar dead space is the physiologic dead space. This dead space can be estimated by determining the difference in the partial pressure of carbon dioxide in arterial blood and in expired air (assuming that these quantities are equal when there is complete gas exchange in the lungs). In normal, healthy individuals, anatomic and physiologic dead space are nearly equivalent, i.e., there is little or no alveolar dead space. In pathophysiological conditions, however, physiologic dead space can greatly exceed anatomic dead space due to inequality of blood flow and alveolar ventilation. The end result is a lower effective alveolar ventilation rate in terms of gas exchange with pulmonary capillary blood. Mechanical Properties of the Lungs and Airways

Under normal conditions, there is little work of breathing, i.e., it takes very little energy to inhale and exhale. Two important properties of lung tissue and fluid determine the characteristic mechanical events that occur in the lung during breathing and affect the work of breathing. The principal properties of the lungs that permit such effortless inflation and deflation of the lungs during breathing are lung compliance and elasticity (i.e., ability to stretch or inflate and to return to its original size and shape) and low resistance to airflow through the airways. Thus, under normal conditions, air flows unobstructed into and out of the lungs that easily expand to accommodate the increase in volume and then recoil to their initial preinspiration volume to expel air. Pressures and Air Movement

Movement of air into and out of the lungs is driven by pressure differentials or gradients across the lungs. When respiratory muscles (diaphragm and intercostal muscles) contract to expand the thoracic cavity, a force is applied to the lung surface that causes expansion of the lungs. Lung expansion occurs because the lungs are compliant and distensible. By expanding, a negative pressure is created within the lungs, specifically in airways and alveoli. Air flows down its pressure gradient into the airways and alveoli that now have a lower pressure relative to the external atmospheric pressure. Changes in lung pressures relative to atmospheric pressure can be summarized as follows (Fig. 5). At the start of inspiration, alveolar pressure is arbitrarily set at zero (atmospheric). Intrapleural pressure is about ⫺5 cmH2O because elastic recoil of the lungs counteracts the forces of the chest wall to recoil outwards. Thus, a negative pressure is generated in the intrapleural space between the lungs and the chest

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Figure 5

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Ventilatory parameters during a single breath.

wall. Intrapleural pressure can be measured directly via a needle inserted into the space or, more commonly, via an esophageal balloon that provides a reasonably accurate measure of changes in intrapleural pressure. Upon inspiration, a greater negative intrapleural pressure is generated as the chest wall moves outward against the elastic recoil of the lungs, reaching a maximal value of about ⫺7 to ⫺8 cmH2O under normal conditions. The expansion of the lungs by the greater negative intrapleural pressure causes alveolar pressure to decrease (becomes negative relative to atmospheric pressure) until it reaches a maximum value of about ⫺1 cmH2O under normal conditions, providing the pressure gradient for air to flow into the airways and alveoli (depicted as negative flow in Fig. 5). In obstructive diseases of the airways, alveolar pressure becomes much more negative to overcome resistance to airflow (vide infra). As the alveoli fill with air, flow decreases and alveolar pressure returns toward zero. At end inspiration, intrapleural pressure is maximal but alveolar pressure is zero, eliminating the driving pressure for

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airflow that then ceases. The difference between intrapleural pressure and alveolar pressure is the transpulmonary pressure, which provides a measure of elastic lung recoil at each point of lung expansion. When inspiration is complete and the lungs are inflated, respiratory muscles relax and elastic recoil properties of the lung cause it to return back to its original state prior to inflation, thereby expelling the inspired air. Intrapleural pressure returns to ⫺5 cmH2O and alveolar pressure increases to about +1 cmH2O, thereby creating the pressure gradient to allow air to flow out of the lungs to the external environment (depicted as positive flow in Fig. 5). Throughout this cycle of a normal inspiration and expiration, airways remain open to allow air to flow into and out of the lungs with relative ease. Expiration may be forced, such as during a cough or when blowing up a balloon, by contracting respiratory muscles. Such forced expiratory maneuvers can increase intrapleural pressure to reach positive values, causing a large increase in alveolar pressure above atmospheric pressure and creating a large pressure gradient for air to flow out of the lungs at greater velocity. Lung Compliance

Lung compliance is determined by the relationship between pressure and volume changes in the lungs, specifically the slope of the pressure–volume curve (V/P). As pressure becomes more negative, the lungs expand or increase in volume. Under normal conditions, lungs are more compliant at low lung volumes than at high lung volumes, i.e., there is a greater increase in volume for a given change in pressure at low initial lung volumes. Normal human lungs have a compliance of approximately 200 mL/cmH2O, i.e., for every 1 cmH2O decrease in intrapleural pressure, the lungs expand by 200 mL of air. The pressure–volume curve for lung inflation is not identical to the curve for lung deflation. During inflation, the lung appears to be less compliant than during deflation, i.e., the deflation curve lies to the left of the inflation pressure–volume curve. This property is known as hysteresis (Fig. 6). In practice, compliance often is measured statically, whereby the lung is inflated to a specified volume and air is released in increments with breath holding to record changes in

Figure 6

Hysteresis associated with a normal breath.

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intrapleural pressure via an esophageal balloon. Methods also are available to measure compliance during continuous inspiration and expiration maneuvers. This measure is termed dynamic lung compliance. It can be altered not only by changes in lung tissue or surfactant that affects static compliance (vide infra) but also by small airways resistance when determined at high breathing frequency. Alveoli supplied by airways with higher resistance cannot fill sufficiently during the shorter periods of inspiration and are termed slow alveoli or alveoli with long time constants. As a result, these alveoli cannot contribute to overall lung compliance and drop out at higher breathing frequencies, thereby lowering dynamic compliance measurements. Therefore, dynamic compliance measurements are thought to reflect resistance of peripheral airways, as well as elasticity of peripheral airways and lung parenchyma. Two principal elements of the lungs account for most of its compliance and elastic recoil characteristics. Elastic connective tissue elements contribute to lung elasticity. During inflation, these fibers that are crisscrossed tend to stretch much like interwoven fabrics stretch and then recoil when released. A second very important component of lung pressure–volume behavior is the low surface tension of the fluid lining alveolar luminal surfaces. This low surface tension is produced by surfactant that is synthesized and secreted by type II alveolar cells. Intermolecular repulsive forces between surfactant molecules account for its ability to lower the surface tension. These forces change with the degree of lung (alveolar) inflation such that the reduction in surface tension is greater at low lung volumes when alveoli are partially deflated. It is this variable surface tension that accounts for the hysteresis observed in pressure–volume curves. Surfactant contributes to lung compliance in three important ways. First, low surface tension increases lung compliance by making it easier to expand alveoli and reducing the work of breathing. Second, surfactant stabilizes alveoli to keep them from collapsing at low lung volumes that would occur if they were lined with pure water having a relatively high surface tension. Alveolar stability also is maintained by interdependence mechanisms whereby adjacent alveoli support each other and tend to prevent collapse or overinflation. Third, surfactant acts to keep alveoli dry by reducing hydrostatic pressure outside capillaries, thereby reducing transudation of fluid. All three actions of surfactant result in a stable, compliant lung that is relatively easy to expand and that will not collapse during elastic recoil and lung deflation. Changes in lung compliance occur with age or may indicate dysfunction. For example, increased compliance occurs in emphysema and in old age due to alterations and/or loss of elastic tissue elements. In these conditions, it is easier to inflate the lungs but more difficult to deflate them. Conversely, restrictive lung diseases such as pleural or interstitial pulmonary fibrosis lead to decreased compliance that results from the fibrotic tissue that has far less distensibility than normal lung elastic tissue. Similarly, pulmonary edema leads to decreased compliance, in part due to alterations or loss of surfactant. Reduced compliance causes an increased work of breathing that results from the need to generate greater pressure gradients to inflate the less compliant lung.

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The second major factor involved in air movement in and out of the lungs is resistance to airflow in the airways. Airway resistance is determined by the pressure–flow relationship, specifically AP/flow, where AP is the difference between pressure at the airway opening (nose or mouth) and at the alveoli, i.e., the pressure gradient that drives the airflow. Thus, airway resistance must be determined under dynamic conditions, i.e., during airflow. Atmospheric pressure is measured directly, airflow is measured with a pneumotachograph, and alveolar pressure can be measured by plethysmography. The units of resistance are cmH2O/L/sec. Normal values for adults range from 0.6 to 2.4 cmH2O/L/sec at 0.5 L/sec airflow that comprise relatively low resistances to airflow. These measurements represent total airway resistance. A longitudinal partitioning of airway resistance exists within the lung, such that approximately 80% of total airway resistance occurs in the large airways (⬎2 mm diameter) and 20% occurs in small airways (⬍2 mm diameter). Small airway resistance is low because airflow is laminar and its velocity is low,; there is a much larger number of small airways with a greater total cross-sectional area than the larger airways, and the diameter of successive branches does not decrease dramatically within these small airways (vide supra). Therefore, determinations of airway resistance are thought to measure principally the caliber of the central airways, bronchi, and bronchioles. Airway conductance, the reciprocal of airway resistance, also is used as a dynamic measure of airway caliber. A number of factors influence airflow resistance, including density and viscosity of the gas and whether there is laminar or turbulent airflow. However, the principal determinant of airflow resistance is the diameter of the airways. Under normal conditions, the airways are dilated, and resistance to airflow is low. Several circumstances can lead to decreased airway diameter and increased airflow resistance. Contraction of airway smooth muscle reduces airway caliber particularly in small bronchi and bronchioles which that have little or no cartilage to add rigidity. Airway wall edema also can reduce airway diameter causing an increase in resistance. Obstruction by excessive mucus can reduce airway diameter or even occlude airways causing a significant rise in resistance. In each circumstance, airflow into or out of the lungs is impaired. Lung volume also influences airway resistance by affecting airway diameter. At high lung volumes, radial traction of surrounding lung tissue on the airways tends to increase airway caliber, thereby lowering resistance. Patients with high airway resistance tend to breathe at high lung volumes to help maintain open airways. Conversely, at low lung volumes, small airways can be compressed, leading to increased resistance to airflow. Dynamic compression of small airways during expiration at low lung volumes greatly limits airflow, especially during forced expiration when intrapleural pressures can exceed atmospheric pressure. In such conditions, closure of these small airways may occur, which leads to trapping of air distal to the compressed airway. Chronic air trapping, as might occur in lung diseases, such as moderate—to—severe asthma or emphysema, damages respiratory units to further compromise lung function. An important point relative to aerosol inhalation therapy is that regional differences in ventilation routinely occur in the lungs. There are several reasons

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for such regional variations in ventilation. One is topographical, whereby lower regions of the lungs (in an upright individual) are better ventilated that upper zones. Lower zones of the lung have a smaller resting volume and, as noted previously, the lungs are easier to inflate at low volumes than at high volumes. Thus, under normal conditions, lower regions of the lungs tend to be better ventilated because they operate on the steep slope of the pressure–volume curve. Other causes for uneven ventilation include regional alterations in lung distensibility or airway resistance. The pattern of ventilation inequality under these circumstances will depend upon the frequency of breathing. For example, an area of low compliance will fill rapidly, but total volume will be less than normal. Conversely, an area with increased airway resistance will fill slowly and incompletely during a typical breath relative to regions of normal compliance and resistance. However, slow breathing would allow alveoli in the latter condition to fill more completely. Measurements of Pulmonary Function

As noted in the preface to this section, pulmonary function tests are useful in a variety of settings, such as in the diagnosis of lung or cardiac disease, evaluation of the effectiveness of therapy for heart or lung diseases or monitoring the course of disease, or exposure to injurious agents. Often, tests consist of determining airway caliber during forced expirations to determine the presence of obstructive or restrictive lung disease. For more detailed information, the reader is referred to the numerous texts and other publications describing these tests of pulmonary function (86–95). Peak Flow Measurements

These measurements are perhaps the simplest tests of expiratory airflow to determine airway function. Subjects inhale completely [to total lung capacity (TLC)], then expire forcibly (both rapidly and completely) into a peak flowmeter, which records the maximal flow rate of expiration. These instruments are simple to operate and often are provided to patients with lung dysfunction, especially asthmatics, for self-evaluation and documentation of lung ventilatory function. Forced Expiratory Flow Measurements

A useful and simple test of pulmonary function is the measurement of a single forced expiratory maneuver. Spirometric methods are used to measure the time course of the forced expiratory volumes. The subject starts at TLC and exhales as rapidly and as completely as possible into the mouthpiece of a spirometer that records the volume of air expired over time as depicted in Figure 7. The volume of air exhaled in the first second is termed the “forced expiratory volume” (FEV1), and the total volume exhaled is the “forced vital capacity” (FVC). The FEV1 is normalized to body size, age, and sex for comparison with standardized values. Additionally, the FEV1/FVC ratio is useful in interpreting these pulmonary function tests. As shown in Figure 7, a normal FEV1/FVC ratio is 0.8. A shallow, prolonged expiratory curve is indicative of airway obstructive disease, because airway obstruction worsens airway compression during forced expiration, significantly slowing exhalation of air from the lungs.

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FEV1/FVC measurements under normal (norm), obstructive (obs), and restrictive (res) airway conditions. Abbreviations: FEV1, forced expiratory volume in one second; FVC, forced vital capacity. Source: From Ref. 96. Figure 7

In this situation, both FEV1 and FVC are decreased but the decrease in FEV1 is much greater leading to a decrease in the FEV1/FVC ratio. In contrast, a sharp curve with a low FEV1 and a low FVC indicates restrictive lung disease, i.e., there is little or no airway obstruction, but a decrease in vital capacity is present. In this instance, the FEV1/FVC ratio may be normal or increased. Forced Expiratory Flow Rate

A related measurement is the forced expiratory flow rate (FEF25–75%) that represents the average flow rate determined over the midportion of the expiration. Generally, FEF25–75% is closely related to FEV1. Flow–Volume Curves

Another useful measure of pulmonary function is the flow–volume curve. Figure 8 depicts flow–volume measurements for normal lung function and for obstructive and restrictive lung diseases. In normal function, flow rate upon forced expiration rapidly reaches a peak at which point airway compression occurs. From this point onward, expiration is effort independent, because greater effort simply compresses more airways and does not expel air at a greater rate. Thus, expiration depends upon lung elastic recoil and resistance of airways upstream from the collapsed points in the airways. Obstructive lung disease causes the curve after the maximal flow rate to assume a concave appearance or a reduced maximal flow and a curve shifted to higher lung volumes (as depicted in the figure). These changes occur because of early airway closure, which may result from enhanced airway constriction caused by smooth muscle spasm in asthma, excessive mucus narrowing the airway lumen, or reduced radial traction caused by loss of parenchyma in emphysema. Conversely, restrictive lung disease could result

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Flow–volume measurements under normal, obstructive, and restrictive airway conditions. Source: From Ref. 96.

Figure 8

in more rapid exhalation due to increased lung elastic recoil without obstructed airways. Maximal flow rate also is reduced in restrictive disease, and expiration occurs at lower lung volumes caused by decreased compliance and the inability to inflate the lungs normally. Lung Volumes

As discussed previously, measurement of lung volumes by spirometry, gas dilution, and plethysmography also can yield valuable information about lung function. Ventilation-Perfusion Ratios

More sophisticated tests can determine matching of ventilation-perfusion ratios and estimate physiologic dead space (vide supra). Airway Resistance and Compliance

Measurements of airway resistance and compliance have already been discussed (vide supra). These tests are more complex than simple expiratory flow measurements and could require invasive methods such as placement of an intraesophageal balloon. Nevertheless, they offer a more reliable measure of airway caliber and peripheral lung function.

III. A.

Pharmacology of the Airways Introduction

As described in the previous section on airway physiology, a primary function of the airways and lungs resides in facilitation of gas exchange between the environment and the blood. To this end, the airways have a continuous integument of epithelium to protect the internal environment of the organ, together

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with mechanisms to expurgate noxious substances from the lungs, such as the mucociliary escalator and alveolar macrophages. Further, caliber of the airways may be modulated to influence characteristics of gas flow, potentially to limit access of irritants to the lower airways or to enhance impaction of inhaled particles. Many (if not all) of these functions are continually subject to regulation by agents released from nerves and cells resident within the lung, by blood-borne compounds and by inhaled substances, including drugs. The following discussion will focus on the pharmacology of agents that influence principally airway caliber by regulating smooth muscle tone and mucous secretion. B.

Pharmacodynamics

Receptors are protein-based, macromolecular complexes existing within cell membranes that, upon interaction with specific agents, change conformation and lead to the triggering of a cellular response, such as smooth muscle cell contraction or relaxation. A receptor shows considerable selectivity in the nature of the agent that interacts with it. As evidence of this, the nomenclature of receptors generally reflects the name of the agent that activates it. For example, cholinergic receptors are stimulated by ACh. Muscarine activates muscarinic receptors and nicotine stimulates nicotinic receptors. ACh acts on both of the latter receptor types leading to the definition that these are subtypes of cholinergic receptors. Agonists interact with a receptor to produce a cellular response, the magnitude of which will be determined by the number of receptors occupied by the agonist and the intrinsic activity of the agonist (or the ability of the agonist to “activate” the receptor during occupation). Full agonists are those agents that can produce a maximal receptor-mediated tissue response. Agents with zero intrinsic activity bind to the receptor but do not produce a response. These agents, by definition, are receptor antagonists, specifically competitive antagonists. The fraction of total receptors (of a specific receptor type) required to be activated to produce a maximal tissue response varies from agonist to agonist. For example, full agonists may need only occupy all or only a small fraction of the total receptor population to produce a maximal tissue response. Partial agonists, on the other hand, have a lower intrinsic activity and require occupation of the total receptor population to produce a maximum effect, the magnitude of which is less than that induced by a full agonist (Fig. 9). As noted, antagonists are compounds that bind to (or occupy) the receptor but do not elicit a response. In the presence of an antagonist, agonist-induced responses are inhibited or prevented. There are three broad categories of antagonists. The first involves antagonists that compete with the agonist for the receptor. Competitive antagonists are those agents that bind reversibly to the receptor. Inhibition caused by these agents can be overcome by increasing agonist concentrations. Irreversible antagonists are agents that bind covalently to the receptor. This form of antagonism irreversibly prevents access of the agonist to the receptor and cannot be surmounted by increasing agonist concentration. As a consequence, the total receptor population available to an agonist is reduced, and a decrease in the maximal effect of the agonist is observed. The second category involves noncompetitive antagonists, which are agents

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Theoretical concentration-response curves for drugs with different intrinsic activities.

that bind to a site distinct from the agonist-binding site of the receptor but that hinders or prevents agonist-induced receptor activation. The final category relates to functional antagonists. These are compounds that inhibit agonistinduced responses by eliciting a counteracting tissue response. For example, in muscle, a drug that causes relaxation would be a functional antagonist of a drug that causes contraction. In vivo, an antagonist can appear to elicit a response. However, such effects are the result of the antagonist inhibiting an ongoing intrinsic agonistinduced response. For example, continual activation of cholinergic nerves innervating airway smooth muscle will result in maintained bronchoconstriction due to the continuously released ACh acting on muscarinic cholinergic receptors to contract the smooth muscle. Under these circumstances, administration of a muscarinic cholinergic competitive antagonist such as atropine will cause bronchodilation. The effect of atropine, therefore, results not from its ability to actively relax airway smooth muscle but rather from it preventing ACh from interacting with the cholinergic receptor to cause bronchoconstriction. Application of a ␤2-adrenoceptor agonist also would induce bronchodilation under the same circumstances. However, in this instance, activation of ␤2-adrenoceptors on the smooth muscle directly causes smooth muscle relaxation and, thereby, functionally antagonizes the bronchoconstriction elicited by ACh (Fig. 10). Structure–activity analyses evaluate the biological activity of chemically modified agonists or antagonists to delineate the structural requirements for binding of the agent to its receptor and for agonist activity. Such studies have led to the identification of receptor subtypes, i.e., isoforms of receptors of the same type but with different ligand (agonist or antagonist) activity profiles. For example, there are three known subtypes of ␤-adrenoceptors, identified as ␤1, ␤2, and ␤3 receptors, each subserving specific cellular functions. Molecular biological techniques have now permitted cloning and expression of receptor proteins and provided important insights into receptor structure and the expression of receptor subtypes. Furthermore, site-directed mutagenesis

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Figure 10 Comparison of competitive and functional antagonism. ACh binds to muscarinic receptors to elicit airway smooth muscle contraction (denoted as +++). The competitive muscarinic receptor antagonist prevents receptor activation by ACh and thereby prevents its contractile effects. The functional antagonist (e.g., ␤-adrenoceptor agonist) does not interact with the muscarinic receptor but prevents the contractile effects of ACh by promoting smooth muscle relaxation (denoted as ---). Abbreviation: ACh, acetylcholine.

studies (in which amino acids in the receptor protein are deleted or substituted) have elucidated amino acid sequences necessary for (i) agonist/antagonist binding, (ii ) location of the receptor in the cell membrane, or (iii ) linking the receptor with intracellular signal transduction pathways. Remarkably, many receptors with different agonist specificities share common structural features, exemplified best by the numerous endogenous transmitters and mediators that act on a variety of G-protein–coupled receptors (Table 4). C.

Signal Transduction Mechanisms

Interaction of an agonist with its receptor results in a conformational change in the receptor protein such that a signal transduction (or second

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Some Agents that Act on G-Protein–Coupled Receptors

Neurotransmitters

Inflammatory mediators

Acetylcholine Epinephrine Norepinephrine Substance P Vasoactive intestinal peptide Arachidonic acid metabolites Leukotrienes Prostacyclin Prostaglandins Thromboxane Bradykinin Histamine Formyl peptide met-lue-phe

Source: From Ref. 97.

messenger) pathway is activated. Several transduction pathways linking receptor activation to response have been and continue to be identified. Principal transduction pathways involved in control of airway smooth muscle tone are illustrated in Figure 11. Several of these will be considered briefly. The reader is directed to other more comprehensive treatises about signal transduction in the lung (98–100).

Examples of signal transduction pathways. Receptor (A) activation can lead to initiation of second messenger cascades through G-protein linkage to intracellular enzymes, such as adenylate cyclase (AC), phospholipase C (PLC), or to ion channels (each requiring different G-proteins). Activation of these systems can then elevate intracellular concentrations of cyclic nucleotides [e.g., cyclic adenosine monophosphate (cAMP)], lipid mediators [e.g., inositol 1,4,5-trisphosphate (IP3) and diacylglycerol (DAG)], or ions (e.g., Ca2+), which activate (or inactivate) other enzymes such as cAMP-dependent protein kinase A (PKA), protein kinase C (PKC), or calmodulin-dependent enzymes (CmDE). In addition, agents (such as nitric oxide) may directly activate intracellular enzymes [such as soluble guanylate cyclase (sGC)], or receptors (B) themselves may act as guanylate cyclase. In either case, elevations in cyclic guanosine monophosphate (cGMP) can activate cGMP-dependent protein kinases (PKG) to initiate a second messenger cascade. It is important to recognize that the illustrated pathways are simplified and represent only the first few steps of a complex transduction cascade.

Figure 11

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Guanine nucleotide–binding regulatory proteins (G-proteins) are the first step in linking many receptors with signal transduction cascades. These are proteins consisting of three different subunits, ␣, ␤, and ␥. The ␤ and ␥ subunits are closely associated with each other and may function to anchor the G-proteins to the cell membrane (101). The ␣ subunit possesses sites for (i) recognition of receptors and transduction molecules (102), (ii) hydrolysis of guanosine triphosphate (GTP) (GTPase activity) (97), and (iii) phosphorylation by other enzymes (104,105). It is primarily the ␣ subunit that interacts with the transduction pathways. Activation of a receptor linked to the G-protein results in dissociation of the ␣ subunit from the ␤ and ␥ subunit complex. The activated ␣ subunit is able to influence the activity of enzymes or ion channels that in turn contribute to the cellular response. The specific subtype of a subunit linked to a particular receptor will determine which enzyme or ion channels are affected. The actions of several different G-protein ␣ subunits are summarized in Table 5. Adenylate cyclase, an enzyme that catalyzes the conversion of adenosine triphosphate to cyclic adenosine monophosphate (cAMP), is subject to control by G-proteins in that it is stimulated by Gs and inhibited by Gi (100). As such, receptor activation of these G-proteins will have opposing effects on cAMPdependent enzymes further down the signal transduction cascade such as protein kinase A. Gq activates phospholipase C (109). Accordingly, receptors linked to these G-proteins elicit cellular responses by phospholipase C–dependent transduction pathways. Activation of phospholipase C results in the formation of inositol 1,4,5-trisphosphate (IP3) (one of many inositol phosphates produced) and diacylglycerol (DAG) from membrane phosphatidylinositides. IP3 acts on its receptor to promote the release of calcium ions from intracellular stores (110). Increases in cytosolic calcium lead to the activation of calcium/calmodulin-dependent protein kinases and subsequent protein phosphorylation. In smooth muscle, this series of events concludes in contraction. DAG activates protein kinase C, which, through phosphorylation of selected proteins, can also influence cellular function (111). Ion channel function also can be influenced by G-proteins. For example, Gs activates calcium ion (Ca2+) channels (112) to promote the influx of Ca2+ into the cell and consequent activation of Ca2+/calmodulin-dependent protein Table 5

Activities of ␣-Subunits of G-Proteins

G-protein GS Gi Go Gq G12/13

Activity Stimulates adenylate cyclase; activates Ca2⫹ ion channels Inhibits adenylate cyclase; inhibits/activates K⫹ ion channels Regulates Ca2⫹ ion channels Activates phospholipase C Activates guanosine nuclear exchange factors for Rho

Source: From Refs. 100, 106, 107.

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kinases. Gi activates potassium ion (K+) channels to influence cell membrane depolarization (113). Further, there is accumulating evidence that the ␤␥ subunit may activate potassium channels (106). G-protein function can be modulated by a recently discovered group of proteins called regulators of G-proteins signaling (RGS proteins). These intracellular molecules repress G-protein activity by increasing the rate at which ␣ subunits rebind ␤ and ␥ subunits or by preventing the ␣ subunits from interacting with their intracellular target molecules (100). D.

Bronchoactive Agents

This section focuses on the pharmacological actions of agents on airway smooth muscle and on mucous gland secretion. While there is a wealth of information related to pulmonary pharmacology in animals, the information provided relates to studies conducted in humans or in human tissue (except where noted otherwise). This section is provided only as a brief summary of airway pharmacology and readers are directed to other, more comprehensive reviews in previous volumes of this series. Neurotransmitters Cholinergic Innervation

Airway smooth muscle tone is subject to regulation by neural and humoral mediators. Human airway smooth muscle is innervated by cholinergic (constrictor) and nonadrenergic noncholinergic inhibitory (NANCI) (relaxant) nerves (115,116). Inhalation of an irritant can result in the activation of a central nervous reflex increase in cholinergic and NANCI nerve activity. ACh released from postganglionic cholinergic nerves acts upon muscarinic M3cholinoceptors on airway smooth muscle to cause constriction (Fig. 12) (117). Administration of muscarinic cholinoceptor agonists such as methacholine mimics the actions of neurally released ACh to induce bronchoconstriction through an action on M3 receptors. Such effects can be prevented or reversed by muscarinic cholinoceptor antagonists such as atropine or ipratropium. Muscarinic M2 cholinoceptors are located presynaptically on postganglionic cholinergic nerves and mediate inhibition of ACh release (Fig. 12, inset ) (117). As a result, ACh released from postganglionic nerves can inhibit its own release, i.e., via negative feedback inhibition. It has been hypothesized that the development of selective M3 antagonists may lead to more effective inhibitors of reflex bronchoconstriction by preventing the M2 antagonisminduced facilitation of ACh release (which normally would act to abrogate antagonism of the postsynaptic M2 receptor) (117). However, a large population of M 2 receptors have been associated with airway smooth muscle of several species (118) and appear to repress airway smooth muscle relaxation by inhibition of adenylate cyclase (119). Submucosal glands also receive cholinergic innervation (120). Mucous secretion is promoted by muscarinic cholinoceptor activation (possibly of the M3 subtype), be it a result of cholinergic nerve activity or application of a muscarinic agonist (65,117).

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Central reflex control of human bronchomotor tone. In addition to the central reflex arc shown, other reflex mechanisms may exist including a local sensory reflex (axon reflex) and a direct sensory nerve/motor nerve ganglia reflex. Abbreviations: ACh, acetylcholine; NANCI, nonadrenergic noncholinergic inhibitory; NO, nitric oxide; sGC, soluble guanylate cyclase; cGMP, cyclic guanosine monophosphate; GTP, guanosine triphosphate.

Figure 12

NANCI Innervation

Airway exposure to irritants can also result in activation of NANCI nerves (59,121). Under conditions of elevated bronchomotor tone (i.e., during bronchoconstriction), this reflex manifests as bronchodilation. As yet, no inhaled stimulus has been identified, which evokes solely a NANCI bronchodilator response. Rather, irritant exposure elicits a biphasic bronchomotor response comprising an initial cholinergic bronchoconstriction succeeded by a NANCI bronchodilator phase. The neurotransmitter mediating NANCI responses in the airways remains to be established with certainty. Vasoactive intestinal peptide (VIP) has been proposed (122) but the absence of selective antagonists has limited the definition of its precise role. VIP is an effective relaxant of isolated airway smooth muscle (122,123) but is only poorly effective as a bronchodilator in vivo (124). VIP also inhibits airway mucous secretion (125) in the ferret by a mechanism involving presynaptic inhibition of ACh release (126). An analog of VIP, Ro25-1553, is more potent than VIP as an airway smooth muscle relaxant (127) and effective as a bronchodilator when administered by inhalation to patients with asthma (128). Other evidence supports the participation of a gaseous molecule, nitric oxide, in mediating the NANCI response. Evidence in favor of nitric oxide as the NANCI transmitter centers on the ability of inhibitors of nitric oxide synthesis to attenuate or abolish NANCI responses in airways isolated from human lungs (68,69,129). Conversely, nitric oxide is a weak bronchodilator in humans (130), and NANCI bronchodilator responses

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in vivo, at least in the cat, are insensitive to nitric oxide synthase inhibition (131). Thus, the precise role of nitric oxide, VIP, or other potential neurotransmitters of airway NANCI nerves remains the subject of continued investigations. Adrenergic Innervation

Airways of several species such as the dog, guinea pig, and rabbit receive innervation by adrenergic nerves (116,132). Norepinephrine is released from postganglionic sympathetic nerves on ␤-adrenoceptors to cause relaxation of airway smooth muscle. In all species, including humans, that lack direct adrenergic innervation of airway smooth muscle (132), circulating epinephrine released from adrenal glands activates adrenoceptors of the ␤2 subtype to mediate bronchodilation (115,116,133). However, the responsiveness of the airways to circulating catecholamines or exogenously applied receptor agonists may not be uniform in the population, given that functional polymorphisms have been identified in the ␤2-adrenoceptor, which affect their responsiveness to activation (134,135). Alpha-adrenoceptor activation exerts little effect on human airway smooth muscle tone (136). Mucous secretion from submucosal glands may be influenced by adrenoceptor agonists such that ␣-adrenoceptor agonists, and agonists of ␤1 and ␤2 adrenoceptors stimulate secretion (137). Agonists of ␤2-adrenoceptors are the primary bronchodilator therapy for reversible obstructive airway diseases such as asthma. Accordingly, many of these selective agonists have been developed as effective bronchodilators with rapid onset of action (e.g., albuterol and pirbuterol) or for more sustained actions (e.g., formoterol and salmeterol) (138). Sensory Neuropeptides

Afferent nerves innervating the airways were traditionally thought to serve a sensory role in central nervous system reflexes. However, a subgroup of sensory nerves innervating the airways have been identified as containing neuropeptides such as the tachykinins, substance P, and neurokinin A (139). As noted, a variety of stimuli evoke the release of tachykinins from sensory nerves in the airway tissue (60). The tachykinins so released can elicit airway mucous secretion or smooth muscle contraction by activation of neurokinin NK1 and NK2 receptors, respectively (140). The ability of inflammatory mediators such as bradykinin, histamine, and prostaglandins to stimulate tachykinin-containing sensory nerves led to the notion that an axon reflex may contribute or mediate bronchoconstrictor responses caused by these agents (141). However, only meager neurally mediated tachykininergic contractile responses could be elicited in human isolated airways (142). The equivocal effects of potent and selective tachykinin receptor antagonists in asthmatic patients (143) bring into question the role of endogenous tachykinins as bronchoconstrictors. Inflammatory Mediators

Antigenic activation of inflammatory cells such as mast cells can result in the release of preformed mediators and/or synthesis of many nascent mediators

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with potent biological actions. Preformed mediators include histamine, cytokines, and proteases including chymase and tryptase. Nascent mediators include adenosine, bradykinin, platelet-activating factor, and the arachidonic acid metabolites, leukotrienes and prostaglandins. Histamine, released primarily from secretory granules of mast cells in the airways, induces bronchoconstriction by an action on histamine H1 receptors on airway smooth muscle cells (144). Activation of histamine H2 receptors promotes mucous secretion. Aside from the direct effect on smooth muscle, histamine is able to elicit central reflex cholinergic nerve-mediated bronchospasm by stimulation of H1 receptors on sensory nerve endings in the airways (144), i.e., act similarly to the irritant shown in Figure 12. Several cytokines are released by mast cells and they include tumor necrosis factor ␣ (TNF) and interleukins (IL) 3, 4, 5, 6, 8, 10, and 13 (145). The functions of these proteins are more associated with the regulation of the allergic response, inflammation, and the immune system rather than with airway smooth muscle directly. Of these, IL4, IL5, IL10, and TNF enhance the responsiveness of nonhuman airway smooth muscle to contractile stimuli (146). Inhaled TNF increases airway sensitivity to inhaled cholinergic agonist through a mechanism likely to be secondary to the induction of airway inflammation (147). Mast cell–derived tryptase exerts many of its pathophysiological effects through cleavage and activation of protease-activated receptor 2 (148). It is a weak contractile agonist of airway smooth muscle (149) but is able to enhance the contractility of histamine (149,150). In addition, tryptase stimulates airway smooth muscle proliferation (149,151). By contrast, chymase represses airway smooth muscle proliferation (152). It can promote airway serous gland secretion (153). The physiological role of these enzymes relative to other mast cell–derived mediators in airway function remains an area of active research, particularly given their potential for regulating the activity of biologically active peptides in the environment of the degranulating mast cell (154). The potential importance of tryptase is suggested by current interest in the development of tryptase inhibitors (155). The presence of mast cell–derived proteases in inflamed airways represents a potentially important consideration for the delivery of peptide-based drugs to the airways by inhalation. Kininogenase-like activity may be released from mast cells to elicit the formation of bradykinin from ␣2-globulins in the plasma (156). Bradykinin induces bronchoconstriction by stimulation of airway sensory nerves to induce central reflex cholinergic nerve-mediated bronchoconstriction while having very little direct airway smooth muscle contractile activity (157). The bronchoconstrictor effect is thought to be mediated via the bradykinin B2 receptor (158). The recent development of selective, nonpeptide antagonists of bradykinin receptors should allow a clearer definition of the role played by bradykinin in airways diseases (159). Adenosine also induces bronchoconstriction in atopic and asthmatic humans indirectly through the release of mast cell–derived mediators and activation of a central cholinergic reflex (160,161). The actions on mast

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cells appear to be mediated by the adenosine A2A and A2B receptors (162). In cultured human epithelial cells, adenosine acts on the adenosine A1 receptor to promote mucin production (163). The influx of calcium ions accompanying mast cell activation can lead to the activation of a variety of enzymes including phospholipase A2, an enzyme primarily responsible for the cleavage of arachidonic acid from membrane phospholipids. In the lung, arachidonic acid serves as a substrate for the enzymes cyclooxygenase (or prostaglandin synthase) or 5-lipoxygenase, leading to the formation of prostaglandins and leukotrienes, respectively. In contrast to several of the aforementioned mediators, arachidonic acid metabolites or eicosanoids are synthesized de novo upon cell activation and are not preformed and stored within the cell (164). Cyclooxygenase metabolites that affect airway caliber include prostaglandins (PG) D2, E2, and F2␣, thromboxane (TX) A2, and prostacyclin. Of these, PGD2, PGF2␣, and TXA2 exert direct contractile actions on the airway smooth muscle (165–167). PGE2 and prostacyclin can cause airway smooth muscle relaxation (166,168). However, PGE2 has also been reported to promote contraction of airway smooth muscle (166,169). This may manifest in vivo as bronchoconstriction in asthmatic subjects (170). The effects of PGE2 on mucous secretion are equivocal whereas PGF2␣ increases mucous secretion (171,172). In addition to direct actions on smooth muscle, PGD2 and PGF2␣ may induce bronchoconstriction by activation of a cholinergic reflex (170,173). The receptors mediating prostaglandin-induced contractile or relaxant responses on human airway smooth muscle are thought to be of the prostanoid TP and EP2 subtypes, respectively (174,175). Airway smooth muscle contraction induced by prostaglandin E 2 may be mediated by EP3 receptors (176). Although prostaglandins are able to elicit bronchospasm, cyclooxygenase inhibitors are ineffective in the treatment of asthma; rather, inhibitors of the cyclooxygenase-1 isozyme, such as aspirin, may induce bronchospasm in asthmatics (177,178). The other pathway of enzyme-catalyzed arachidonic acid metabolism leads to the generation of lipoxygenase metabolites. The 5-lipoxygenase products, leukotrienes C4, D4, and E4, are potent constrictors of airway smooth muscle (179,180) through an action on the cysLT1 receptor subtype, at least in human airways (181). Both leukotriene C4 and D4 stimulate submucosal gland secretion (182). Given the potency of leukotrienes as bronchoconstrictors, selective inhibitors of their synthesis (5-lipoxygenase inhibitors) or receptors have been developed and shown to be very effective in the treatment of asthma. Examples of such agents include the 5-lipoxygenase inhibitor, zileuton, and the leukotriene receptor antagonists, montelukast, zafirlukast, and pranlukast (183). Other prostaglandin-like molecules, isoprostanes, can be generated by the action of reactive oxygen species. Upon activation, mast cells can generate reactive oxygen species that interact nonenzymatically with membrane phospholipids (184). Several of these, including 8-iso-PGF2␣ and 8-iso-PGE2 induce contraction of human airway smooth muscle by an action on the prostanoid TP receptor (185).

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Xanthines such as theophylline and its salt, aminophylline, exert bronchodilator actions in humans. The exact mechanism by which these agents elicit this effect remains enigmatic and, consequently, several mechanisms have been proposed. A direct smooth muscle-relaxing effect of the xanthines may occur through their ability to inhibit phosphodiesterase, the enzyme responsible for the hydrolysis of cAMP (186,187). Resultant elevations in intracellular concentrations of cAMP (by inhibition of degradation) then activate similar relaxant transduction pathways, as would agonists that induce cAMP synthesis through activation of adenylate cyclase, such as ␤-adrenoceptor agonists or VIP. Another possible action involves antagonism of adenosine receptors, thereby preventing the airway smooth muscle contractile effects of adenosine (188). Structure–activity relationships for the xanthines indicate a correlation between xanthine-induced inhibition of lung cAMP phosphodiesterase and airway smooth muscle relaxant potency. No such relationship with relaxation is apparent in relation to adenosine antagonist activity of the xanthines (189). The predominant phosphodiesterase isozymes in human airway smooth muscle are types 3 and 4 and are selective for cAMP (190). This observation, together with the identification that inflammatory cells express, primarily, phosphodiesterase 4 (PDE4) (191), propelled the development of phosphodiesterase isozyme–selective inhibitors. Cilomilast and roflumilast are examples of PDE4 inhibitors currently under clinical investigation (192), although their therapeutic actions are targeted more as anti-inflammatory agents than as bronchodilators. The majority of therapeutic agents that directly relax airway smooth muscle do so by elevating intracellular cAMP. However, agents that act to increase intracellular cyclic guanosine monophosphate (cGMP) also can induce airway smooth muscle relaxation (193). These include natriuretic peptides such as atrial natriuretic factor (194,195), B-type natriuretic peptide (196), and urodilatin (197) and “nitro” compounds such as glyceryl trinitrate, isosorbide dinitrate, and sodium nitroprusside (198–200). The putative NANCI transmitter, nitric oxide, is a direct activator of intracellular soluble guanylate cyclase, one of the enzymes responsible for cGMP production (193). When inhaled, however, it is a weak bronchodilator (130,201). Nitric oxide has been suggested to have a stimulatory effect on mucous secretion (202). Exploration of the airway effects of the cGMP-mediated agents is commonly limited by the profound cardiovascular effects that attend the doses needed to induce bronchodilation. Nevertheless, cGMP-dependent bronchodilator agents will continue to be an active area of research, with pulmonary delivery and the development of cGMP phosphodiesterase isozyme-specific inhibitors being potentially fruitful avenues of endeavor.

References 1. 2.

Weibel ER. Morphometry of the Human Lung. Berlin: Springer Verlag, 1963:1. Weibel ER. Lung morphometry and models in respiratory physiology. In: Chang HK, Paiva M, eds. Respiratory Physiology: An Analytical Approach. New York: Marcel Dekker, 1989:1–56.

Physiology and Pharmacology of the Airways 3. 4.

5. 6. 7. 8. 9. 10. 11.

12. 13. 14. 15. 16. 17. 18. 19. 20. 21. 22. 23. 24. 25. 26. 27. 28. 29. 30.

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5 Solute Transport Following Aerosol Deposition in the Lungs

RICHARD M. EFFROS Harbor–UCLA Medical Center, Torrance, California, U.S.A.

I.

Introduction

Interest in the manner in which solutes cross the pulmonary epithelium has grown with the recognition that the lungs can be used as a convenient portal for administering medications that are currently injected because they cannot be absorbed from the gut. It has long been understood that gases (e.g., amyl nitrite and carbon monoxide) and lipid-soluble solutes (e.g., nicotine and alcohol vapor) can be readily absorbed through the lungs, but it was thought that absorption of nonvolatile and lipid-insoluble molecules, such as proteins, would be too slow to permit effective therapy. This problem has been overcome with modified pharmaceuticals and the use of new devices that can deliver larger doses of these medications (most notably insulin). However more information is needed regarding the factors that influence the rates at which these molecules traverse the barriers separating the gaseous and blood compartments of the lungs. The enormous surface area and attenuated membranes of the alveolar compartment of the lungs suggests that aerosols containing medications for extrapulmonary treatment should be designed to reach the alveolar compartment, and this review will be concerned with absorption of solutes from the distal regions of the lungs.

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Effros II.

Radioaerosols

Much of our understanding about the movement of solutes between compartments in the lungs in man has been derived from the clearance of radioaerosols. Aerosolized radionuclides were introduced 35 years ago as a method for determining the distribution of ventilation in the lungs (1). These radioactive indicators are usually delivered to the airways and airspaces in aqueous droplets smaller than 2 ␮m in diameter. Once the indicators become deposited on the lung surfaces, the regional distribution of radioactivity in the lungs stabilizes, permitting measurements of aerosol distribution over periods that may exceed 30 minutes. The relative stability and resolution of images obtained with radioaerosols compared to those observed after administration of gaseous indicators such as ¹³³Xe has contributed to the popularity of the radioaerosol approach. However, not long after the radioaerosol approach was introduced, it was noted that indicators such as 99mTc-DTPA and 99mTCO4– are lost quite rapidly from the lungs of many patients, suggesting that they had diffused across the pulmonary epithelial surfaces of the lungs and been carried away in the blood. This observation led us and others to utilize radioaerosols for detecting abnormalities in the permeability of the pulmonary epithelial barrier (2–5). Since then, the radioaerosol approach has found application in a wide variety of illnesses that affect the integrity of the pulmonary epithelium [see reviews in Refs. (6,7)]. Some recent observations made with 99mTc-DTPA are indicated in Table 1. An alternative agent, “pertechnegas,” has also been introduced for detecting increases in pulmonary epithelial permeability (23–26). This indicator was derived from “technegas,” which represents an ultrafine suspension of Table 1

Some Recent Studies of Epithelial Permeability to 99mTc-DTPA

Observations regarding clearance rates Increased with aging Increased with oxygenators in cardiopulmonary bypass Increased with mild unstable asthma Increased with radiation pneumonitis, methacholine Increased with ozone exposure Smoking increases correlated with associated with oxidant stress and inflammatory cells in lavaged lungs Not increased in stable asthma Surfactant attenuated increases after detergent lavage Increased with inverse ratio ventilation Increases not correlated with either X ray or pulmonary function abnormalities in interstitial lung disease Increased with exposure to cotton, hemp flax H2O2 increases moderated with adenyl cyclase activation Bronchial clearance of 99mTc- not affected by smoking Clearances of 113In-DTPA and 99mTc-DTPA similar Increased in systemic lupus erythematosis Increased with platelet-activating factor

Species

Reference

Mice Humans Humans Humans Dogs

(8) (9) (10) (11) (12)

Humans Humans Rabbits Rabbits

(13) (14) (15) (16)

Humans Guinea pigs Rabbits

(17) (18) (19)

Humans Humans Humans

(20) (21) (22)

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99m

Tc metal crystals that are completely surrounded by graphite hexagons. Although it is not a true gas, these virtually inert particles are purportedly much smaller (5–30 nm across and 3 nm thick) than droplets produced in nebulizers. They are not hygroscopic and behave much as a gas in the airways, permitting efficient penetration to the smallest airways. Some evidence suggests that technegas particles actually tend to agglutinate, and particle size, may therefore, be as large as 225 nm (26). Technegas is produced by heating pertechnetate (99mTcO4–) to extremely high temperatures in a graphite crucible in an environment of pure argon. Pertechnegas can be produced in the same apparatus if a low concentration of O2 is also present in the device. The 99mTc in this agent is not sealed from the environment and 99mTcO4⫺ is released when the particles come in contact with water. Clearance of pertechnegas is consequently more rapid than that of 99mTc-DTPA. However, like 99mTc-DTPA, pertechnegas clearance is accelerated among patients who smoke and those with pulmonary fibrosis. Because clearances of pertechnegas and 99mTcO4⫺ are much more rapid than that of 99mTc-DTPA, a large dose must be given in a relatively short interval, making these agents clinically less practical. Of interest, has been the observation that smokers have accelerated clearance of 99mTc-DTPA before significant changes in pulmonary function or X-ray appearance are detected. However rapid clearance is associated with bronchoalveolar lavage evidence for oxidant stress and airway inflammation (13). As indicated below, insulin is also cleared more rapidly among smokers, and smoking increases the permeability of the pulmonary epithelium to the bronchodilator, terbutaline (27). These observations suggest that significant increases in epithelial permeability occur with smoking and can be detected well before patients become symptomatic or experience any change in conventional pulmonary function tests. The aerosol procedure, therefore, appears to provide a sensitive method of detecting early lung injury. III.

Factors Affecting Clearance of Solutes from the Lungs

Because most of the surface area upon which droplets ⬍ 2.0 ␮m in diameter become deposited is in the gas-exchange area of the lungs, it is of particular importance to consider the manner in which solutes diffuse out of the alveoli. The alveoli contain two continuous layers of cells that separate blood flowing through the lungs from the air to which they are exposed: The epithelium and the endothelium. [See schematic diagram in Fig. 1 and detailed descriptions by Weibel (28,29).] There is good reason to believe that transport of small hydrophilic solutes and macromolecules from the alveoli to the capillaries is limited by the permeability of the epithelium, rather than that of the endothelium. Schneeberger and Karnovsky (30) and Pietra et al. (31) reported that when intravascular pressures were increased, horseradish peroxidase moves through the junctions of the endothelial cells but movement into the alveoli is blocked by the interepithelial junctions. They subsequently showed that the epithelial cells are joined by a complex network of sealing strands, whereas adjacent endothelial cells are connected by only one or two strands,

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Effros

Schematic diagram of the arrival of aerosol droplets in the airspaces. Most of the alveolar surface is covered by flat Type II pulmonary epithelial cells. Droplets become deposited on the epithelial lining fluid, which lies under a layer of surfactant. Once the droplets have merged with the ELF, hydrophilic molecules diffuse through the intercellular junctions of the epithelium to the interstitium or are carried by vesicles into the capillaries. Junctions between epithelial cells are tighter than those between endothelial cells. Surfactant within the myelin bodies of the type II pulmonary epithelial cells is secreted into the ELF in the form of tubular myelin, which may slow diffusion of solutes to the interepithelial junctions. Lipophilic solutes cross directly through lipid membranes.

Figure 1

which are frequently discontinuous (32). This morphologic evidence that the epithelium is less permeable than the endothelium is supported by physiologic observations, which suggest that the reflection coefficients of the epithelium to small solutes are greater than those of the endothelium (33), and tracer movement of labeled mannitol and urea out of the pulmonary vasculature appears to be slowed at the pulmonary epithelial border (34,35). Increases in the rates of diffusion of solutes out of the airspaces presumably reflect alterations in the epithelium rather than the endothelium. Before entering the tissues of the lungs, solutes must traverse a thin layer of fluid epithelial lining fluid (ELF). This fluid tends to collect at the corners of the alveoli and is covered by an attenuated layer of surfactant. Although it is generally believed that the ELF and surfactant cover the entire surface of the alveoli, this hypothesis has been challenged (36). It has been suggested that surfactant may slow solute diffusion out of the alveoli (37–40). Complex tubular myelin structures within the ELF (Fig. 1) may impair diffusion of hydrophilic solutes out of the airspaces. Once droplets containing radionuclides enter the fluid lining the airspaces “ELF”, the radionuclides probably diffuse to neighboring airway surfaces. The rate of indicator clearance out of the fluid lining the airspaces can be predicted with the equation: dQ/dt ⫽ ⫺PS∆c

(1)

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where Q is the quantity of indicator in the air space, t is time, P is the permeability of the blood–gas barrier, S is the surface area of the barrier, and ∆c is the concentration difference across the barrier. For the sake of simplicity, the ELF can be conceptualized as a thin planar film overlying the alveolar surface with a thickness of Ax. If plasma concentrations remain low, equation (1) becomes: Ca,t /Ca,0 ⫽ e⫺(P/∆x)t

(2)

where Ca,t is the concentration in the air space at time t, Ca,0 is the initial airspace concentration, and ∆x is the thickness of the ELF in which the indicator is deposited. This equation deserves some emphasis because it shows that it is the thickness of the fluid layer rather than its surface area that influences the fractional rate at which the concentrations of the indicators in the airspaces decrease. For example, if the same amount of indicator were deposited in two adjoining alveoli rather than one, but the depth of fluid remained unchanged in each, then the fractional rate of decrease would remain the same. On the other hand, if the aerosol droplets actually increased the thickness of fluid overlying the alveolar surface, then the rate of loss of the indicator from the airspaces would be slowed. Equation (2) predicts a monoexponential decline with a slope of k ⫽ ⫺ P/∆x. The presence of multiexponential decreases in the amount of indicator remaining in the lungs is frequently interpreted in terms of the presence of heterogeneity in the permeability of the epithelium. Alternatively, it could represent the presence of serial barriers through which the indicator passes before reaching the blood. Although the loss of solutes from the airspaces is frequently referred to as “clearance,” which has units of volume/time, this is a misnomer because the units of P/∆x are time⫺1 (e.g., the fractional decline per minute). Although the rate of decline is sometimes indicated in terms of the halftime of decline (which is equal to ⫺0.693/k), it is preferable for statistical reasons to use “k” because this limits the variability of the normal values of clearance. Mucociliary transport contributes to the clearance of indicators that become deposited in the airways (e.g., 40). The contribution of mucociliary transport to the clearance of radionuclides from the airspaces can be estimated by measuring the rates at which impermeant indicators, such as 99mTc-sulfur colloid, are lost from the lungs. Unless the rate of diffusion of the solute is very rapid or the rate of perfusion is extremely slow, clearances of diffusible, extracellular radionuclides are not significantly affected by lung perfusion. For example, the clearance of 99m Tc-DTPA from the lungs cannot be reliably used to detect pulmonary embolism because sufficient flow derived from the bronchial circulation persists to carry away the 99mTc-DTPA that crosses the pulmonary epithelium and reaches the pulmonary capillaries (41). Although administration of radioaerosol droplets that are less than 2 ␮m in diameter favors peripheral distribution, the exact site of deposition cannot be predicted. There is some reason for believing that the time required for indicators to cross bronchiolar and bronchial walls is greater than that through the

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Effros

walls of the alveoli (42), but this probably depends in part on the depth of the fluid overlying the epithelium of these barriers, as indicated in equation (2). IV.

Lipid Solubility

Since the early studies of Collander (43) on solute diffusion into the cells of the alga Nitella, it has been recognized that the single most important property that governs the rate of solute diffusion into cells is its lipid-to-water solubility coefficient. The close relationship between lipid solubility and the rate at which solutes are lost from the airspaces was demonstrated in the early studies of Chinard et al. (44). The low solubility of ions in lipid is responsible for the slow rate at which they diffuse across the cell membranes. On the other hand, the rapidity with which respiratory and other gases cross the alveolar–capillary barrier can be attributed to the fact that they readily dissolve in the cell membranes. In studies with labeled 14 CO2 and H14CO⫺ we found that the H14CO⫺ diffuses across the epithe3 3 lium at a rate that is at least 600 times slower than that of 14CO2 (45). Similarly, acetic acid crosses the epithelium more rapidly than acetate and NH3 diffuses across this barrier more rapidly than NH4+ (46). Differences between the solubility of the ionized and uncharged species of a weak acid–base pair could conceivably be used to define the pH of the ELF. The pH of the epithelial fluid is normally about 6.9 (47). For an acid, HA, which dissociates to form H+ and A⫺, it would be expected that as the ELF becomes more acidic, the indicator would cross the barrier more rapidly. On the other hand, for a base, R, which combines with H+ to form RH+, acidity would decrease the rate at which the solute crosses the alveolar–capillary barrier. The similarity in the venous outflow of labeled water, CO2, alcohols, and acetone after instillation into the airspaces (26,28) suggests that during the time it takes the perfusate to traverse the pulmonary capillaries, they all equilibrate between the airspaces and the perfusate (see below). This means that the loss of these indicators from the lungs is limited by the rate of perfusion rather than the permeability of the alveolar–capillary barrier. With cooling, the solubility of lipophilic indicators in the lipid membranes falls, and the rates at which solutes such as antipyrine, methanol, and acetone diffuse across the pulmonary epithelium also decrease (48). The effect of temperature changes on permeability can be described in terms of the Arrhenius equation: P ⫽ e⫺E/ a /RT

(3)

where P is the permeability, Ea, is the activation energy, R is the gas constant (1.987 cal mole⫺1 ° C⫺1), and T is the absolute temperature. V.

Electrolyte Transport

Sodium, chloride, and water are secreted into the airspaces in the fetus. At the time of birth, fluid transport is reversed and most of the fluid remaining in the

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lungs is reabsorbed. Vectorial reabsorption of fluid from the lungs is based upon movement of Na+ from the alveoli into the epithelial cells through epithelial sodium channels (ENaC) on the apical surfaces of alveolar epithelial cells (49,50). There is evidence that both the flat type I cells (which represent one-third of the epithelial cells but 97% of the alveolar surface) and the more spherical type II cells participate in this process (51,52). The type II cells also secrete surfactant and act as precursor cells for the Type I cells. The movement of sodium and anions into the cells results in an osmotic gradient that is responsible for associated water movement across the cell membranes. Transport through the Na+ channels is upregulated by c-AMP agonists such as the beta-adrenergic agents and is inhibited by amiloride. Apical transport of sodium is passive and is responsible for most of the resistance of the epithelial cells to the movement of fluid across the epithelial surface. Sodium is actively transported out of the epithelial cells across the basolateral membranes through Na+- K+-ATPase transporters, which are inhibited by ouabain but are enhanced by dopamine (53). Energy for most of the transport of electrolytes across cell membranes is provided by this pump. Additional information concerning electrolyte and H+ transport across the pulmonary epithelium can be found elsewhere (54,55). VI.

Strong Acids

It has been known for many years that following aspiration of strong acids, the pH of the inspired fluid is rapidly neutralized in the lung, but the factors responsible for this phenomenon have not been defined. Aspiration of gastric fluid, which has a pH as low as 1.0, is a common occurrence in patients with a wide variety of swallowing problems. Rapid neutralization of 0.01 HCI in saline (pH ⫽ 2.0) depends in part on the efficiency with which anions, specifically HCO3⫺ and Cl ⫺, are exchanged across the alveolar–capillary barrier (56). It is likely that much of the transport of these anions across the basolateral surface of alveolar epithelial cells takes place through a specific member of the anion exchanger (AE) family, the AEII protein (57). Like the AE proteins in the red cell membrane (formerly referred to as the band III protein), the pulmonary AE permits exchange of these ions down concentration gradients. Transport of HCO3⫺ and Cl ⫺ across the apical membranes probably also depends on a variety of other transporters, including the cystic fibrosis transmembrane conductance regulator (CFTR) (58,59). This has led us to speculate that some of the damage that occurs to the bronchi in cystic fibrosis may be due to abnormalities of CFTR, which slow the rates at which HCO3⫺ is transported into the airspaces and Cl⫺ is removed. Although it has been generally assumed that the bronchiectasis that occurs in children with cystic fibrosis is due to inspissation of the bronchial secretions, this does not explain why bronchiectasis is not seen in asthmatic patients, who also have a problem with thick secretions. On the other hand, bronchiectasis is characteristic of patients who chronically aspirate large volumes of acid (60). The tendency for bronchiectasis to occur in the upper lobes of patients with cystic fibrosis may reflect the fact that much of the damage occurs when they are infants and spend much of the day in a recumbent position.

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Effros VII.

Water Transport

The history of transcellular water transport in the lungs can be found in the previous edition of this chapter and will not be repeated here (see also Ref. 113). Research regarding the role of aquaporins in water transport in the lungs is reviewed in several important publications (61,62). Considerable progress has been made regarding the role of aquaporins in water transport across the pulmonary endothelium and epithelium since the last edition of this text, but much remains to be learned. More than 10 isoforms of aquaporin have been described in human tissues. These are generally divided into those channels that transport only water, and those that transport water and small molecules such as glycerol and urea. Water movement through these channels is symmetric and driven by osmotic gradients. In most organs other than the intestines, movement of water through aquaporins is much faster than water movement through the lipid bilayers, cotransporters, or intercellular junctions. Aquaporin 1 is present in both the pulmonary and bronchial endothelium, aquaporin 3 is found in the large airways, aquaporin 4 is present throughout the airways, and aquaporin 5 is expressed in both the type I cells and the submucosal glands of the large airways. Aquaporins 8 and 9 have also been reported in the lungs (63,64). Aquaporins have not yet been described on the basolateral surfaces of adult type I or type II epithelial cells (65). Utilizing transgenic mice, it has been possible to show that aquaporins 1 and 3 play important roles in facilitating flow of water in response to osmotic gradients. However it is by no means clear, whether they have a significant influence on the formation or reabsorption of airway edema. No effect on edema formation or reabsorption could be found in several experimental models (61). However, aquaporin 1 did seem to be of some importance in transvascular fluid movement. Edema formation along airway walls was more pronounced in normal subjects than those who lacked aquaporins 1 after fluid challenges (66). Extracellular acidification reduced the permeability of aquaporin 3 in the airways (67). Although aquaporins 3 and 4 also enhance osmotically driven flows of water across the epithelium, they do not seem to have an effect on the hydration of the airways or isosmolar fluid reabsorption (61). However, deletion of aquaporin 5 from the upper airway glands decreases secretions from these glands and increases protein concentrations by a factor of two (61). Aquaporin 5 expression is also increased by beta agonists (68). Aquaporin 1 appeared to be redistributed to the plasma membranes by exposure to lipopolysaccharides (69). Water transport through aquaporins in cellular membranes in response to hydrostatic pressure differences is probably slowed by the rate at which electrolytes are transported. Movement of water without electrolytes results in an osmotic gradient, which tends to slow further shifts in water. It is possible that transport of fluid with electrolytes from the lungs in response to hydrostatic pressure gradients is sufficiently slow that water transport through other structures (lipid bilayers, cotransporters, or intercellular junctions) suffices to permit osmotic equilibration as electrolytes leak or are transported between the airspace, cellular, and interstitial compartments. This would make it difficult to

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detect the effect of aquaporins. Other potential roles of the aquaporins not directly related to water transport are discussed in a recent review (70). VIII.

Redox Indicators

Methylene blue has been used to estimate the volume of fluid in the airspaces prior to lavage. We reported that the transport of this redox dye from the airspaces is greatly accelerated if it is instilled into the airspaces of isolated lungs with the reducing agent, Na2S2O4 (71). The oxidized form of methylene blue is converted to the reduced, colorless, nonionized form in the presence of this reductant. In contrast, K3Fe(CN)6, a strong oxidant, slowed losses of methylene blue from airspace. Both ascorbate oxidase and superoxide dismutase also slowed movement of methylene out of the airspaces, suggesting that both ascorbate and superoxide are responsible for reduction and absorption of methylene from the airspaces. Most of the reducing activity of the ELF of rodent lungs is based on the presence of extracellular reductants, especially ascorbate (72). In contrast, reduced glutathione is responsible for most of the reduction of redox indicators that occurs in primates, which do not synthesize ascorbate. The redox potential of the ELF could reflect inflammation in the lungs. Strategies for developing aerosolized radionuclides to detect local inflammation in the lungs could be based on the selection of indicators that, like methylene blue, become more lipid soluble with changes in redox potential, resulting in an increase in clearance from the lungs. IX.

Macromolecules

A.

Therapeutic Agents

The advent of inhaled insulin therapy has rekindled interest in aerosols containing proteins (73–76). The advantages of inhaling rather than injecting insulin were understood soon after insulin was isolated (74), but practical problems associated with its administration have frustrated this form of therapy until very recently, and a variety of preparations are now undergoing tests. As expected, the bioavailability of small proteins decreases with molecular weight: the bioavailability of leuprolide is greater than that of calcitonin, which is greater than that of insulin. One problem with inhaling insulin is variability in the rate at which it is absorbed. For example, it has been found that smokers are more apt to develop hypoglycemia after inhaling insulin, a phenomenon which has be attributed to increased permeability of the pulmonary epithelium. This observation confirms the earlier reports that smokers also have increased clearance rates of 99mTc-DTPA, petechnegas, and terbutaline (see above). It is likely that absorption of inhaled proteins will be greater among patients with other diseases associated with pulmonary inflammation because these accelerate clearance of 99mTc-DTPA. It can be anticipated that aerosols will be used to deliver other therapeutic proteins in the near future.

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Effros B.

Diffusion of Macromolecules Through “Pores”

Several comprehensive reviews have appeared, which address the manner in which proteins and other macromolecules are absorbed across epithelial and endothelial barriers (77–79). Both Ohtani et al (80) and Conhaim et al (81) found that the rate of absorption of macromolecules across the pulmonary epithelium is inversely related to molecular size, confirming the conclusions of Effros and Mason (82). Multiple additional studies in five species are consistent with the conclusion that the rate of transport of large molecules tends to be slower than that of smaller molecules (77). An inverse relationship between molecular size and weight would be consistent with diffusion in an aqueous medium in which the ratio of the permeability (P) of one molecule to another can be predicted from the moleuclar weight (M) of these molecules on the basis of the Graham equation: P1 = P2

M2 M1

(4)

As the size of the molecule approaches the diameter of the interendothelial junction, permeability of the molecule will fall disproportionately. On the basis of the transport of hetastarch, Conhaim et al (81) concluded that the epithelial membranes could be modeled with epithelial pores of two diameters, 5 nm and 17 nm. C.

Vesicular Transport of Macromolecules

Evidence has been reported that both the alveolar type I and type II cells contain albumin and immunoglobulins in intracellular vesicles (81,83–85). The presence of macromolecular molecules in intracellular vesicles does not prove that they are actually transported across cell membranes (“transcytosis”). For example, surfactant molecules are taken up from the airspaces in vesicles but are then metabolized and the constituents are returned to the airspaces. In contrast to passive diffusion through interepithelial junctions, movement of relatively inert molecules in vesicles can be relatively insensitive to molecular size. For example, Matsukawa (86) found that uptake of 70 kDa and 150 kDa dextran molecules from the airspaces are similar, suggesting bulk uptake by a vesicular mechanism. Two major categories of intracellular vesicles have been identified in eukaryotic cells. Caveolae are flask-shaped structures which are abundant on pulmonary endothelial and epithelial membranes, especially type I cells (87). They are characterized by the presence of a family of proteins referred to as caveolins that play a major role in the formation and behavior of caveoli. Caveoli can detach themselves from the surface of the cells and transport proteins and even viruses across cellular membranes. A second group of vesicles are surrounded by a cage of clathrin molecules, which assemble on the cell membrane and subsequently detach when the vesicle reaches its destination

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(88). These vesicles play an important role in sorting, recycling, and delivery of proteins to the cell membranes or to lysosomes, where they are metabolized. Some molecules bind to sites within vesicles, a process that increases concentrations in these structures above those in the extracellular medium, thereby accelerating clearance. As reviewed by Hastings et al, human growth hormone, vasoactive intestinal protein, albumin, and globulin—all move through the epithelium much more rapidly than would be expected on the basis of strictly diffusional processes (77). Furthermore, many of these molecules show evidence for saturation binding. Below a concentration of 0.3 g/dl, the movement of albumin displays saturation kinetics, whereas no evidence for saturation is found at higher concentrations. Inhibitors of endocytosis appear to decrease endocytic transport of albumin (89). Hastings et al estimate that at concentrations of albumin encountered in pulmonary edema, 98% of albumin reabsorption occurs through paracellular pathways (77). Loss of proteins from the airspaces may also reflect metabolic degradation. Macrophages have been shown to ingest 125I-albumin instilled into the airspaces (77). Furthermore a variety of catalytic enzymes are secreted by epithelial cells, which can catabolize many proteins (77). Immunologic sensitization with proteins has been shown to block protein absorption (90–92). This could potentially decrease the efficacy of inhaled insulin therapy and might be associated with the development of allergic reactions. Various strategies can be used to inhibit proteolysis of therapeutic agents. For example, adding polyethylene chains of various sizes to insulin may serve to slow protein degradation or immune reactions, and the delivery of insulin and other polypeptides to the systemic organs can be sustained. D.

Particulates

Movement of particulates out of the lungs appears to be very slow compared to that of molecules in solution (77). Many of these particles can be taken ingested by macrophages and if not metabolized, they may be stored for long periods in lymph nodes in patients with occupational exposures. Certainly asbestos particles make their way into distant sites (93), but the role of macrophages in this process remains unclear. In addition to the possibility of conventional ultrafine particles (⬍ 100 nm) passing through the pulmonary epithelium, the emergence of even smaller particles encountered in nanotechnology (1–4 nm diameter) may represent an even greater risk for both pulmonary and extrapulmonary toxicity to employees (94). X.

Exhaled Breath Condensates

Although the focus of this chapter and text has been on exogenous aerosols, it is now clear that the respiratory tract and adjoining oral cavity also produce endogenous aerosols that can contain solutes such as acids and inflammatory mediators. These can be produced in the lungs and upper aerodigestive tract during exhalation and can then be deposited in different portions of the respiratory tract during inhalation. It has long been assumed that the formation of

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droplets in the exhaled air is primarily related to cooling of the water vapor that is released by evaporation from the surfaces of the pulmonary epithelium. Humidification of the lungs is essential for gas exchange and the exhaled air is normally kept close to 100% saturated with water vapor. The lungs produce approximately 500 mL of water each day, which can be colleted by cooling the exhaled air in a variety of condensers. Because most of the water remains as a gas in the lungs, it cannot deliver solutes to other areas of the lungs. However, the presence of nonvolatile cytokines and electrolytes in condensates indicates that droplets of fluid have also been released from the airways and are then captured in the condensers (95). Of particular interest have been reports indicating that cytokine concentrations in the exhaled breath condensates (EBC) are increased in a variety of inflammatory lung diseases (96,97). It has been suggested that these increases can be used to detect pulmonary inflammation and its response to therapy. However, little attention was given to the possibility that such increases could also be attributed to an increase in the number and/or size of respiratory droplets that are incorporated in the condensate To estimate the contribution of respiratory droplets to the EBC and calculate the concentrations of cytokines, and other molecules in the fluid that line the airways (ELF), estimates must be made of the dilution of a reference indicator by water vapor (98,99). It is assumed that the concentrations of the reference indicators in ELF are the same as those in the plasma, in other words: Cplasma ⫽ CELF

(5)

The dilution, D, of respiratory droplets by water vapor can then be calculated from the ratio of plasma to EBC concentrations, C: D=

C plasma C EBF

(6)

Once D has been determined, then the concentrations of any nonvolatile solute (CS,EBC) in the ELF can be calculated by the simple equation: C S,ELF = D × C S,EBC

(7)

The likelihood that plasma and ELF concentrations are the same is greatest if the indicator has the following properties: (a) It readily diffuses between the plasma and ELF, (b) it is not significantly produced or destroyed within the lungs, and (c) it is not volatile. Urea has all of these properties and has been used to estimate dilution. For example, if plasma urea concentration is 5 mmol/L and that in the EBC is 0.5 ␮mol/L, then the dilution of respiratory droplets by water vapor is 1:10,000. We also have used the total nonvolatile cation concentrations for this purpose (Na+ + K+ + Ca++ + Mg++). Selection of total cations was based upon the assumption that the respiratory tract is isotonic and the fact that these cations represent most of the cations in the plasma and extracellular space. Alternatively, the conductivity of the plasma and EBC can be used for this purpose, but the EBC samples must first be lyophilized to remove volatile contaminants (primarily NH4+ and HCO3⫺), which usually represent the principal ions in the EBC (98–101).

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139

Measurements of D with these three techniques yielded similar results and averaged ⬃ 1:10,000 (98–100). In other words only 0.1␮L of respiratory droplets are present in each mL of EBC. Water vapor contributes ~99.99% of the volume of the EBC. Because of the extreme dilution of respiratory droplets by water vapor, assays for cytokines and other nonvolatile molecules must be correspondingly sensitive and many of the problems with EBC studies are associated with insufficient sensitivity of current assays. It must also be emphasized that this correction for dilution of ELF by water vapor is only appropriate for nonvolatile solutes and no simple technique is available for volatile solutes such as H2O2. One of the most intriguing observations made in EBC studies is the tendency for the EBC to be relatively acidic in patients with obstructive lung diseases such as asthma, chronic obstructive lung disease, and cystic fibrosis (102). This phenomenon was referred to as “acidopnea” and it was suggested that it reflected acidification of the airways. Furthermore, because NH4+ concentrations in the EBC of patients with acidopnea are also low, the investigators concluded that acidopnea could be attributed to a decreased rate of NH4+ synthesis in the lungs (103). Although this explanation seemed plausible, there is reason to believe that acidopnea reflects oral rather than airway acidification. NH4+ represents more than 90% of the cations in normal subjects. Because concentrations of NH4+ in the EBC can be removed by collecting these samples from an endotracheal or tracheostomy tube (98,99,101), it can be concluded that most of the NH4+ is derived from the mouth, where it can be produced by bacterial degradation of urea to NH3, rather than from the lungs. It has been known for many years that oral NH3 serves to neutralize acidic aerosols such as those containing SO2 from power plants and it may also neutralize other acids derived from the mouth or stomach (104). EBC samples must contain anions in concentrations that match those of NH4+ to ensure electroneutrality. We recently found that HCO3– is the corresponding anion in most individuals and HCO3– concentrations are closely correlated with those of NH4+ (105). The concentration of HCO3– also depends in part upon the PCO2 in the sample. It has been the practice of most investigators to vigorously flush the EBC samples with an inert gas such as argon prior to pH measurements (102). This is intended to minimize the effect of end-tidal CO2 on the samples and the pH of the samples rises as CO2 is removed. However, it is difficult to remove all of the CO2 in this manner and even ambient concentrations of CO2 (⬍ 0.1%) are sufficient to have a marked effect on pH. HCO3⫺ usually remains the principal anion (105). If all of the CO2 were removed, then the pH of the EBC would increase above 9.0 in most samples, and OH– would become the principal anion. Because pH values this high have never been reported, it indicates that not all of the CO2 has been removed. Using recommended degassing procedures, we also found that much of the CO2 remained in the samples, and there was a concomitant loss of other important constituents that play a role in setting the EBC pH, such as water and NH4+. To avoid these artifacts, we have chosen to expose the samples to room air over a 30-minute period. This sets the PCO2 to ambient levels. Concentrations of acetic acid appear to be increased in the EBC samples of some patients with asthma or COPD (105,106). Although it was suggested

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that this might reflect increased production in the lungs, no evidence for this has been reported. Acetic acid represents the most abundant organic anion in the saliva and it appears to reflect production by anaerobic organisms in the mouth (107). Similarly, serum acetate is related to anaerobic organisms in the gut: Serum concentrations increase from 70 ␮moles/L to 2000 ␮moles/L following ingestion of oat bran (108). We found that acidification of the saliva was associated with a decrease in EBC NH4+ concentrations (reflecting conversion from NH3 to NH4+ in the saliva), whereas concentrations of EBC acetate were increased (reflecting conversion of acetate to acetic acid) (105). It therefore, appears that acidopnea can reflect acidification of the saliva. This could be related to gastroesophageal reflux, which is common in patients with a variety of obstructive lung diseases and cough (109). It was possible to show that both NH4+ and acetate in the EBC arrived as the corresponding gases (NH3 and acetic acid) rather than in respiratory droplets, which are insufficient to have an appreciable effect on EBC pH. Like other volatile substances (e.g., NH3, H2O2 and NO2), much of the volatile acids and bases in the exhaled air are derived from the mouth, rather than the lungs and EBC pH cannot be considered a reliable index of changes in airway pH. Acetic acid aerosols have been used as a provocative agent to test for airway sensitivity in patients with excess coughing (110). Furthermore, acetic acid is a component of air pollution that appears to cause lung damage in children (111,112). To the extent that oral gases are inhaled, it is possible that the lungs can be injured by what amounts to an endogenous form of air pollution. XI.

Conclusions

More than five decades of research concerning pulmonary epithelial permeability to aerosolized solutes are now beginning to bear fruit. We now understand some of the basic properties of solutes that either facilitate or slow absorption from the airspace compartment. This may help us to engineer agents that are retained within the airspaces for specific periods of time, providing rapid or prolonged action of the medications. Aside from the fact that the discomfort and inconvenience of injections are avoided, inhalation of medication avoids delays associated with liver metabolism following ingestion and delays related to diffusion away from sites of injection. Studies with radionuclides indicate that consideration will have to be given to the effects that pulmonary disorders and smoking may have upon the rate at which drugs are absorbed from the lungs. It should also be possible to use aerosolized indicators to determine whether lung injury is occurring from underlying pulmonary disorders and to detect injuries caused by aerosolized medications at a relatively early stage, before they become more serious. References 1. 2.

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Harding SM. Recent clinical investigations examining the association of asthma and gastroesophageal reflux. Am J Med 2003; 115(Suppl 3A):137–143. Mitsuhashi M, Mochizuki H, Tokuyama K, et al. Hyperresponsiveness of cough receptors in patients with bronchial asthma. Pediatrics 1985; 75(5):855–858. Gauderman WJ, Gilliland GF, Vora H, et al. Association between air pollution and lung function growth in southern California children: results from a second cohort. Am J Respir Crit Care Med 2002; 166(1):76–84. Gauderman WJ, Avol E, Gilliland F, et al. J.The Effect of Air Pollution on Lung Development from 10 to 18 Years of Age N Engl J Med 2004; 351(11):1057–1067. Effros RM, Darin C, Jacobs ER, et al. Water transport and the distribution of Aquaporin 1 in the pulmonary airspaces of rats. J. Appl.Physiol 1997; 83(3):1002–1016.

6 Drug Metabolism and Enzyme Kinetics in the Lung

MEENAKSHI BHAT Eli Lilly and Company, Indianapolis, Indiana, U.S.A.

JOSEPH K. H. MA and YONGYUT ROJANASAKUL West Virginia University, Morgantown, West Virginia, U.S.A.

RONALD K. WOLFF Nektar Therapeutics, San Carlos, California, U.S.A.

I.

Introduction

A.

Role of the Lung in Drug Absorption and Metabolism

Inhalation as a noninvasive alternative to oral drug administration offers several advantages in drug absorption. The lung is the only organ, other than the heart, through which the entire cardiac output passes. Drug absorption is regulated by a thin alveolar–vascular permeability barrier, which, in parts of the alveolar region, shows an air–blood pathway of less than 0.5 ␮m (1). In the adult human lung, the number of alveoli in the deep airway ranges from 200 million to 600 million, resulting in an enormous epithelial surface area that has an estimated value of 140 m2. While these properties are uniquely adapted to promote gas exchange through passive diffusion, they also provide a mechanism for efficient transport of drugs from the circulation to lung tissues and from lung tissues to the bloodstream. Drugs that are administered intravenously may accumulate in high concentrations in the lung following the first passage. On the other hand, the pulmonary epithelium may allow systemic absorption of aerosolized drugs manifold faster than the gastrointestinal tract (2). Thus, there is a great potential that many drug therapies may be improved via inhalation drug delivery, pending further advances in aerosol technology. In addition to dosage-related problems associated with the administration of inhalation aerosols, drug metabolism is another area that must be evaluated in order to fully appreciate the lung as an alternative route for drug delivery. 147

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The lung’s role extends far beyond the exchange of gases. Its metabolic function can result in the conversion of compounds to reactive products or inactivation of bioactive agents. One example of the lung’s metabolic action on blood-borne chemicals is the induction of lung-specific toxicity of paraquat. The lung appears to exhibit selective uptake mechanisms for various amines and peptides, which, when coupled with metabolism, served to inactivate the circulating agents. In the case of paraquat, the drug is internalized in lung cells by an energy-dependent process that is inhibited by putrescine (3). The toxicity results from the biochemical alterations that follow the entry of paraquat into the cells. The drug is reduced by NADPH- and NADH-requiring enzymes, resulting in the production of superoxide anion and the depletion of cellular NADPH (4). These studies show that the lung’s uptake–metabolism dual capacity can lead to altered drug effect or toxicity. Nitrofurantoin also undergoes cyclic reduction–oxidation and directly injures lung tissue, specifically pulmonary endothelial cells, by generating toxic O2 radicals within cells (5). B.

Metabolism and Pulmonary Drug Delivery

The study of drug metabolism in the lung necessitates some special considerations as it is a complex organ comprised of over 40 different cell types, while in comparison, the liver consists of 80% hepatocytes. There is scant knowledge available about the drug-metabolizing activities of the lung. Most of these studies are focused on the role of pulmonary cytochrome P-450 (CYP-450) isozymes, a group of enzymes known to convert lipophilic agents into more polar, water-soluble metabolites. This knowledge is mostly applicable to the metabolic fate of lipophilic drugs. There is an increasing interest in the development of protein and peptide drugs in contemporary biomedical research. The use of these agents may replace current low-molecular weight drug therapy and/or introduce innovative forms of treatment by countering a disease at the molecular level. Proteins and peptides that are susceptible to a variety of enzymatic degradations, are good candidates for the development of aerosol delivery systems. In addition to the monolayer nature of the alveolar epithelium and its large surface area, which are more conducive for peptide absorption, the lung may also have a relatively less enzymatic activity toward protein and peptide drugs. Indeed, studies by Wang et al. (6) recently showed that the degradation of enkephalins by peptidases in the alveolar epithelium was significantly less than that observed in other tissues. Several diseases of genetic origin in the lung can be targeted through inhalation drug therapy. These include, among others, the ␣1-antitrypsin deficiency, an autosomal-recessive disorder characterized by the development of emphysema in patients between the ages of 20 years and 40 years. ␣1-antitrypsin is a glycoprotein normally produced by hepatocytes and mononuclear phagocytes. It serves as the major inhibitor of neutrophil elastase, an omnivorous protease that is capable of destroying elastin as well as certain parts of the protein components of connective tissue. The critical role of ␣1-antitrypsin is highlighted by the hereditary disease caused by its deficiency, which accounts for 2% of all cases of emphysema in the United States (7). Possible strategies

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for the treatment of this disease include aerosol delivery of the protein (8,9) or the administration of a leukoprotease inhibitor (10). In general, proteins and peptides are not metabolized by the CYP-450 enzymes, as they are usually hydrophilic. They are, however, susceptible to degradation by various proteases, which may very well be the controlling forces of the bioavailability of drugs administered by pulmonary inhalation.

II.

Enzymatic Systems and Their Distribution in the Respiratory Tract

The lung is exposed to foreign substances through either the airways or the bloodstream. It is not surprising that the lung possesses localized defense mechanisms to deal with the presence of chemicals. The development of aerosol drug delivery systems, thus, requires careful evaluation of the aerosol deposition pattern, the aerosol components, and their metabolic reactions. The chemical nature of the drug can also yield clues on cell-specific metabolism especially if it resembles endogenous substrates acted upon by the lung. All metabolizing enzymes found in the liver are also found, although in lesser contents, in the lung. The spectra of metabolizing enzymes identified include phase I-reactive CYP-450s, flavin-containing mono-oxygenase (FMO), monoamine oxidase (MAO), aldehyde dehydrogenase, NADPH-cytochrome P-450 reductase, esterases, and epoxide hydrolase; and phase II-conjugating enzymes glutathione S-transferase (GST), UDP-glucuronyltransferase (UDPGT), sulfotransferase, N-acetyltransferase, and methyltransferase. The phase I reactions involve substrate modification by attachment or alteration of reactive function groups such as -OH, -NH2, or -COOH, through oxidation, reduction, and hydrolysis. The phase II pathways involve reaction of phase I metabolites with endogenous molecules such as glucuronic acid, to yield highly polar, readily excretable conjugates. These enzyme systems are located at various subcellular sites, e.g., the cytosol, mitochondria, and smooth endoplasmic reticulum. Tremendous progress has been made in the characterization of extrahepatic tissue CYP expression, function, and regulation employing quantitative RT-PCR. The profile of extrahepatic enzymes is being given increased consideration to allow for greater insights into the involvement in lung cancers, etiology of asthma, and multiple chemical sensitivity. A.

Distribution of Metabolism Enzymes in the Lung

Upper Respiratory Tract

The upper respiratory tract may be taken as the airways down to the junction of the larynx and trachea. The nasal vestibule is lined by skin, whereas elsewhere the nasal epithelium is comprised of ciliated and nonciliated columnar, basal, and goblet cells. The olfactory epithelium is located in the upper part of the nasal cavity and possesses CYP-450 and NADPH cytochrome P-450 reductase activity (11). The CYP-450 levels in olfactory epithelium are lower than those found in the liver, but the specific enzyme activities, expressed as product/time/nmol

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P-450, are comparatively much higher (12). Fetal and adult human nasal mucosa have been found to contain CYP-450 genes for 2A6, 2A13, 2B6, 2C, 2J2, and 3A (13–17). The nasal cavity has been shown to exhibit epoxide hydrolase, carboxylesterase, and aldehyde dehydrogenase activities, and their specific activities were found to be higher than those in the respiratory epithelial homogenates (18–20). In addition to the phase I enzymes, the nasal tissue also contains phase II enzymes including GST, glucuronyltransferases, and sulfotransferases (21,22). The relatively high enzyme activity of the nasal cavity is of prime importance for consideration in nasal drug delivery. Human CYP 2A6, 2A13 (14), 2B6 (23), and 2S1 (24) have been detected in tracheal tissue. CYP2A13, 2F1, and 2S1 are preferentially found in the respiratory tract and hence should be given special consideration as they could be responsible for lung-specific toxicity (23). Lower Respiratory Tract

The lower respiratory tract may be divided into the tracheobronchial, bronchiolar, and alveolar regions. The tracheobronchial region is lined by pseudostratified ciliated columnar epithelium, beneath which are mucous glands, connective tissue, cartilage, smooth muscle, and nerves. Various cell types including basal, ciliated, mucus, serous, and Clara cells and migratory lymphocytes are present in this region. The CYP-450 enzymes are distributed throughout the region, although they may be concentrated in only a few cell types (25). FMO appears to be absent in the larynx and trachea, whereas in the carina and the conducting, pulmonary, and small airways, this enzyme activity is comparable to or exceeds that of the liver (26). The bronchiolar region is characterized by the presence of a columnar epithelium comprising mainly of two cell types: the ciliated cells and the Clara cells. The apical regions of the Clara cells have a high concentration of CYP450 (27–29). The alveolar epithelium consists of type I and II pneumocytes and brush alveolar or type III cells. In addition, this region contains alveolar macrophages (AMs). Both the Clara cells and the type II pneumocytes contain CYP-450 isozymes and NADPH cyctochrome P-450 reductase, as shown by immunochemical studies and immunocytochemical electron microscopy (29–31). High levels of non-CYP-450–dependent enzymes, including FMO, GST, UDP-GT, and epoxide hydrolase, have also been detected in Clara cells (29,32). In contrast, type II cells were shown to have much lower GST and UDPGT activities, and the epoxide hydrolase activity was undetectable (33). AMs do not appear to contain CYP-450 (34,35); however, rabbit AMs have been shown to exhibit rich NADPH cytochrome P-450 reductase and N-acetyltransferase activities (34,36). Type I cells exhibit little or no metabolic activity; thus, they are susceptible to drug toxicity. Butylated hydroxytoluene, for example, is selectively toxic to type I cells (37). Many microsomal CYP-450s have been detected in the human lung including 1A1, 2B6, 2E1, 2F1, 3A4, 3A5, 4B1 (38), 1A2 (39), 1B1 (40–41), 2A6 (42,43), 2A13 (15), 2C (42,44), 2D6 (45), 2J2 (46), and 2S1 (24). All functional CYP genes in a CYP2 gene cluster on chromosome 19, including CYP2A6, 2A13, 2B6, 2F1, and 2S1 (47) are expressed in the respiratory tract.

Drug Metabolism and Enzyme Kinetics in the Lung B.

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Enzyme Systems

Pulmonary Mono-oxygenase Systems

The CYP-450 mono-oxygenase systems in the lung are involved in the metabolism of endogenous compounds such as fatty acids and steroids and lipid-soluble xenobiotics that enter through the circulation or via inhalation. In general, xenobiotics are degraded to water-soluble, readily excretable metabolites. However, some metabolism can lead to the formation of bioactive or carcinogenic products. In fact, the lung is one of the most susceptible organs to benzo[a]pyrene (BP)-induced carcinogenesis (48). Pulmonary CYP-450 enzymes, similar to those of the hepatic system, are located in the endoplasmic reticulum and are comprised of various P-450 isozymes and NADPH cyctochrome P-450 reductase. The level of CYP-450 in the lung shows approximately a 10-fold (0.05–0.5 nmol/mg protein) species variation and is 5–20 times lower than that of hepatic preparations (33,49,50). While this may suggest lower enzyme capacity for the lung, substrates such as p-xylene are in fact more rapidly metabolized (per nmol P-450) in lung than in the hepatic preparations (51). Substrate Specificity of Pulmonary CYP-450

The characterization of CYP-450 isozymes in rabbit and rat lung has been reviewed by Devereux et al. (52). Two major isozymes that account for over 90% of the total CYP-450 in rabbit lungs have been identified and designated as P-4502B1 and P-45041B1 (53–56). A third isozyme was identified as P-4501AI, which accounts for 1–3% of the total CYP-450 (57). The same isozymes have been reported in rat lung tissue (56,58), with the P-4502B1 being the major isozyme (58). Isozyme P-4502BI is known for its O-deethylation activity on 7-ethoxycoumarin (7-EC) (57), whereas P-4504BI is active in the conversion of aromatic amines to mutagenic products (59). In comparison to the 7-EC-O-deethylation by P-4502BI, P-4501AI catalyzes the O-deethylation of 7-ethoxyresorufin (7-ER) (52,57). This isozyme is responsible for the conversion of BP to BP9, 10-diol and BP7, 8-diol in rabbit lungs (52,60). The latter is known to form a possible carcinogenic substance via further metabolism (48). Rats exposed to silica, a respirable dust known to produce oxygen radicals, have been shown to exhibit increased pulmonary P-450 content, especially that of the P-450IAl (61). P-4504B1 has also been shown to act on BP to form metabolites that covalently bind to DNA (62). Comparison of Pulmonary and Hepatic Enzymes

Immunochemical studies showed that the antibody to P-4502BI from rabbit lung cross-reacts with the major hepatic enzyme induced by phenobarbital and inhibits both enzymes in the same fashion (57,63), suggesting that these are the same enzymes. Indeed, the purified pulmonary and hepatic enzymes also give the same rates of 7-EC-O-deethylation (57) and have the same amino acid composition (64,65) and peptide patterns in response to partial protein hydrolysis (63), thus establishing the presence of isozyme 2B1 in both tissues. Studies by

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Guengerich and Mason (66) showed the same results in rat tissues, i.e., both the pulmonary and the hepatic CYP-450 contains P4502BI. CYP-4504BI appears to be a lung-specific isozyme, as it has been found in lung tissues of rodents and monkeys. Its presence in the liver was detected only in the hamster (56). In humans, the P-4504BI mRNA is present in the lung but not in the liver (67). Because of the localized nature of the CYP-450 system, it may mediate drug toxicity (68,69) in specific lung cells and play an important role in the design of inhalation drug therapy. Because the lung contains isozyme that is not present in the liver, drug delivery by inhalation may also lead to altered metabolism patterns. CYP2A13 is active toward many compounds such as 2⬘-methoxyacetophenone, 2,6-dichlorobenzonitrile, hexamethylphosphoramide, N,N-dimethlyaniline, N-nitrosodiethylamne, and N-nitrosomethylphenylamine. CYP2A13 also appears to be the most efficient CYP enzyme in the metabolic activation of the tobaccospecific carcinogen, 4-(methylnitrosamino)-1-(3-pyridyl)-1-butanone (NNK). Subjects carrying the CYP2A6*4C allele were found to have lower risk for tobacco-related lung cancer. 8-Methoxypsolaren, an inhibitor of CYP2A6 was found to efficiently prevent the occurrence of adenoma caused by NNK in A/J mice (70–71). NADPH Cytochrome P-450 Reductase

Pulmonary NADPH cytochrome P-450 reductase occurs in AM, Clara, and type II cells. In rats, this pulmonary reductase has been shown to be identical with the hepatic enzyme in molecular weight, and in enzymatic and immunochemical responses (72). The purified pulmonary NADPH cytochrome P-450 reductase from rabbit was shown to have a specific activity of about 50,000 U (nmols cytochrome c reduced/min/mg protein) and a monomeric weight of 72–79 kD (53,57). The chemical and immunochemical properties of the pulmonary enzyme, including peptide mapping, antibody reactions, and its inhibition of microsomal mono-oxygenase activity, are also identical to those of hepatic reductase (55,57,73). NADPH cytochrome P-450 reductase was shown to activate the toxicity of paraquat (4,69) and nitrofurantoin (74) to lung cells by enzyme reduction of the nitro group, yielding a free-radical oxygen to regenerate the parent drug and superoxide anion. The latter causes lipid peroxidation and a depletion of cellular NADPH (68). Flavin-Containing Mono-oxygenases

FMOs are found predominantly in Clara cells and, to a lesser extent, in type II cells (29). Substrate types acted upon by these enzymes include secondary and tertiary amines and sulfur-containing compounds. The CYP-450 systems are active in the N-oxidation of nonbasic amines but not in the metabolism of basic amines owing to their strong ligand binding. Hence, the N-oxidation of basic amines, such as cocaine (75) and the phenothiazines (76), is carried out by FMO. Weak bases such as aromatic amines are substrates for both systems. FMO exists in a number of isozymes, and their pattern of distribution may vary between the lung and the liver within the same species and among different species, thereby eliciting differences in the metabolism of drugs such as phenothiazines and other

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amines. In rabbits, e.g., prochlorperazine and trifluoperazine are substrates for FMO, but chlorpromazine and imipramine are not (76). In rats, however, all four drugs are metabolized (77). FMO2 is the major lung FMO expressed in most mammals and nonhuman primates (78–80). FMO2 exhibits a number of distinct properties of substrate specificity, thermostability, and resistance to detergent inhibition. FMO2 in humans expresses a very interesting genetic polymorphism in expression (79,81–83). All of the Caucasians and Asians genotyped to date have 2 alleles with a c.1414C → T transition replacing a Gln with a premature stop codon. This truncated protein is not found in human lung samples; expressed p.X472 does not incorporate FAD, is catalytically inactive, and is probably destroyed rapidly due to incorrect folding (81,84). Interestingly, 27% of individuals of African descent (79,81) and 5% of Hispanics possess at least 1 allele coding for the full length, enzymatically active protein (FMO2.1, p.Q472). Individuals genotyped as possessing at least 1 FMO2*1 allele express full length, catalytically active protein. Interestingly, the FMO2 gene is also truncated in laboratory but not wild strains of rats and the use of R. norvegicus as a model of human FMO2 gene polymorphism has been proposed (85). As a result of this finding, it is being proposed that individuals with the FMO2*1 allele metabolize drugs and xenobiotics differently than individuals with no functional FMO2 in the lung and can be particularly susceptible to the toxicity of particular Scontaining chemicals. FMOs bioactivate thioureas through S-oxygenation to toxic sulfenic and/or sulfinic acid metabolites (86,87). The sulfenic acid is capable of undergoing redox cycling following conjugation with glutathione, resulting in oxidative stress due to depleted glutathione and NADPH (87–91). This is thought to be an important mechanism of thiourea toxicity in the lung. In the case of thioether-containing organophosphate pesticides, the presence of FMO2*1 allele can have a protective effect. However, the relative contribution of individual FMO enzymes in the metabolism of a given drug is difficult to determine as, to date, there are few if any FMO-family–specific or even selective substrates. Thus, one cannot phenotype an individual for a given FMO, as is possible for many of the CYPs (92). Ethionamide is structurally related to thiobenzamide and is an antitubercular agent. Ethionamide is a prodrug that must be S-oxygenated by an FMO in the target organism, Mycobacterium tuberculosis, to exert its effect. The metabolic activation of ethionamide by the bacterial FMO is the same as the mammalian FMO activation of thiobenzamide to produce hepatotoxic sulfinic and sulfinic acid metabolites. Ehionamide has been found to be a substrate for FMO1, FMO2, and FMO3. Esterases

Nonspecific esterases are widely distributed in various tissues of the body and act on esters, amides, and thioesters. In the lungs, these enzymes are present in high levels in AM and, to a lesser extent, in alveolar type I and 11 cells (93). Hydrolysis of beclomethasone dipropionate (BDP) by lung esterases to its monopropionate and beclomethasone has been demonstrated (94). The presence of esterases in AM may allow for the delivery of prodrugs into the alveolar region. Esterases secreted by AMs can lead to the degradation of protein and peptide drugs that are

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contained in dry powder aerosols as a postphagocytic event. Indeed, the depletion of AM by liposome-encapsulated dichloromethylene diphosphonate was demonstrated to cause dramatic enhancement of the systemic absorption of IgG and human chorionic gonadotrophin after intratracheal instillation in rats (95). Macrophages have been found to contain aminopeptidase N (96–97), carboxypeptidase M (99), and DPP-IV (99–100). UDP-GT and Sulfotransferases

UDP-GT enzymes are membrane bound and are localized in the microsomal fraction of both liver and lung homogenates (101). Substrates for these enzymes include phenols, thiols, and amines. Human bronchial cells can metabolize 1-naphthol to sulfate and glucuronide conjugates (102). Rat and guinea pig AM contain UDP-GT, while rabbit lung seems to be completely devoid of such activity (103,104). In rats, three isozymes—p-nitrophenol UDP-GT, 3␣-hydroxysteroid UDP-GT, and 17p-hydroxysteroid UDP-GT— have been located in Clara, type II, bronchial, and ciliated bronchiolar epithelial cells (105). Different drugs may be acted upon by different isozymes. Pulmonary glucuronyltransferase is activated by surface-active agents or enzymes that perturb the microsomal membrane. Cetylpyridinium chloride, digitonin, and trypsin all increase enzyme activity with either p-nitrophenol or 4-methylumbelliferone as a substrate (101). Sulfuric acid esters can be formed by reaction of aromatic or aliphatic hydroxyl groups (phenol or alcohol) with 3⬘-phosphoadenosine-5⬘-phosphosulfate (PAPS). Enzymes carrying out this conjugation are localized in the soluble fractions of both liver and lung. Significant activities of phenol sulfotransferase on 4-methylumbelliferone and p-nitrophenol have been detected in rabbit lung. The human lung is reportedly a poor site for sulfation compared to the liver and the GI tract (106–107). Northern blot analysis demonstrated the presence of SULT2B1 mRNA species approximately 1.4 kb in length in human trachea (108). SULT2B1a and SULT2B1b were found to be capable of catalyzing the sulfation of the hydroxysteroid DHEA with apparent Km values of 8.8 ␮M and 11.6 ␮M, respectively. This finding may have implications for inhaled steroids classically employing for treating asthma. Although the various CYP enzymes have been implicated in carcinogenesis, it has only recently been appreciated that sulfotransferases, although classically thought of as a detoxifying enzyme, can lead to the formation of carcinogenic adducts. This can be explained by the fact that the sulfate group is electron withdrawing and may be cleaved off heterolytically in appropriate molecules, thus leading to the formation of a strongly electrophilic cation. Reactive sulfate conjugates however only show strong toxicity if they are generated within the indicator cell, due to their insufficient penetration of cell membranes (109). C.

Metabolism of Endogenous Substrates and Peptides by the Lung and Its Implication on Drug Delivery

Monoamine Oxidase Activity

Because the entire cardiac output perfuses through the lung, the pulmonary vascular bed can be regarded as a sluice gate for controlling the level of bioactive

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agents allowed to enter the blood circulation. Indeed, studies have shown that the lung selectively removes endogenous norepinephrine (NE), 5-hydroxytryptamine (5-HT), and ␤-phenylethylamine (PEA) from the blood and converts angiotensin I to angiotensin II. This implies that drugs with properties similar to the endogenous compounds may also be candidates for uptake and metabolism by the lung. Knowledge of the combined effect of uptake and metabolism may prove useful in the elucidation of the therapeutic/toxic outcome of bioactive agents and in the design of drugs to be targeted to the lung. Inactivation of Monoamines

The observation by Starling and Verney (110) that a serum vasoconstrictor substance, later identified as 5-HT (111), was detoxicated on passage through a heart–lung preparation provided the first demonstration of the “metabolic” function of the lung. It is now widely accepted that the lung actively removes and inactivates monoamines. The enzyme kinetics for NE and 5-HT showed a wide divergence in Km values measured using lung homogenates and isolated perfused lung (112,113). This was attributed to the involvement of an intracellular enzyme, the mitochondrial MAO. Because substrates have to cross the cell membrane, the wide divergence results from the varying accessibility of the enzyme to the substrates in the two techniques. The mitochondrial MAO from rat lung shows similar substrate specificities and pH optima to those of the liver and brain (114). It contains two subtypes: MAO-A that deaminates 5-HT and NE; and MAO-B that deaminates PEA, as confirmed by both substrate and inhibitor specificities (115). The MAO-A and MAO-B activities are inhibited by clorgyline and deprenyl, respectively. The enzyme kinetics of the two subtypes are different in several aspects. Nonhydroxylated substrates generally have a lower Km than hydroxylated substances. While the metabolism of NE and 5-HT is rate limited by cellular uptake, inactivation of PEA in lung homogenates did not differ significantly from that of the perfused lung. Studies using 14C-PEA showed that following a 3-min infusion through isolated rat lung, the uptake of drug was not saturable up to 200 ␮M, while deamination of PEA was saturable at a lower concentration (116). In comparison, the uptake of 5-HT and NE was shown to be saturable at 10 ␮M, and was not affected by MAO-A inhibitors (112). When the lung was treated with selective inhibitors of MAO-B, however, uptake of PEA was markedly reduced, indicating that PEA metabolism is not uptake limited but enzyme limited (112). Studies employing quantitative autoradiography using either radiolabeled MAO-specific inhibitors allowed for the determination of the regional distribution of MAO activities within an organ (117). PET imaging employing [11C]-labeled MAO inhibitors or deuterium-substituted [11C]-labeled inhibitors (118,119) provided a useful noninvasive way of a measure of MAO activity. Recently, immunohistochemical approaches were employed with MAO-A– or MAO-B–specific monoclonal antibodies. Alveolar pneumocytes were found to express MAO-A and MAO-B at approximately similar levels. Both enzymes were found to be present in the bronchiole/alveolar smooth muscle and bronchiolar epithelial cells in the lung (120). Studies combining in

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situ hybridization with cRNA probes and immuohistochemistry of adjacent tissue sections demonstrated comparable levels of MAO-A– and MAOB–specific mRNA and MAO-A and MAO-B (121). The physiological role of pulmonary MAO, i.e., deamination of locally released NE and/or circulatory dopamine, norepinephrine and epinephrine, is consistent with this finding. These enzymes can probably additionally act on circulatory amines from the diet that are not metabolized by the intestine. Pulmonary Uptake of Basic Amines

Various amino drugs such as diphenhydramine, chlorpromazine, tetracaine, propranolol, etc., have been shown to accumulate in the lung (122–124). Studies on the uptake mechanisms (125–127) indicated that the amines should exhibit some lipophilicity with pKas greater than 8.0, and that the steady-state uptake of amines into the lung consists of two components: a saturable process at low concentrations, and a diffusion process at high concentrations. The saturable component reflects intracellular uptake due to facilitated diffusion but not carrier-mediated, transport or metabolism. Efflux studies of imipramine and methadone indicate that the amine drugs are accumulated in different pools in the lung (128,129). A slowly effluxing pool accounted for 30% of pulmonary uptake of imipramine, indicating the persistence of basic amines in the pulmonary system. The accumulation of certain amphiphilic amines such as chlorphentermine results in drug-induced phospholipidosis. This may be attributed to the binding of the cationic amine with phospholipids, thereby preventing the calcium-dependent degradation of phospholipids by phospholipases (130–131). In rats treated with chlorphentermine, excessive lipids and lipid-bound chlorphentermine were found in AM and type II cells (132). Some bisbenzylisoquinoline alkaloids, e.g., tetrandrine, were also shown to accumulate in the lung and bind strongly to AMs at the plasma membrane (133,134). These alkaloidal amines were shown to mediate internal calcium rise and reduce the rates of mitochrondrial ATP synthesis in type II cells (135). The important binding sites in the uptake of basic amines are the lysosomes. Yokogawa and colleagues modeled the tissue distribution for basic drugs. Basic drugs bind to nonlipophilic and lipophilic structures in cells, although a significant portion of the drugs enter the lysosomes via a pH-dependent mechanism. The protonated drugs were found to sequester within lysosomes (136). Peptide Degradation Pulmonary Control of Endogenous Peptides

In addition to its action on basic amines, the lung also removes or inactivates certain physiological peptides from the circulation. Bradykinin was the first peptide shown to be inactivated, by the lung via enzyme-mediated hydrolysis (137). Studies showed that about 80% of the infused bradykinin was lost during a single passage through the heart and lung chambers, thus indicating a very rapid effect. The lung was also shown to cause rapid conversion of angiotensin I to angiotensin II, with a rate of 50% in 10 seconds (138–139), suggesting that the converting enzyme was easily accessible. Pulmonary endothelium seems to be

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impermeable to peptides for which there is no uptake mechanism. It is likely that the rapid action is due to the presence of membrane-bound enzymes in direct contact with the blood, although certain uptake mechanisms coupled with intracellular enzymes would also allow the lung to control the circulating peptides. Ryan et al. (140–141), using radiolabeled bradykinin or angiotensin I, showed that the radioactivity has the same transit time as blue dextran, suggesting that enzyme action probably occurs in the vascular space, at the endothelium. Indeed, both bradykinin and angiotensin I have been shown to undergo degradation in highly purified lung plasma membranes (142). These peptides are known substrates for the angiotensin-converting enzyme (ACE), a membrane-bound peptidase that has been localized morphologically to crypts in the endothelial cells. ACE functions as a carboxyl-terminal peptidyl dipeptidase. The enzyme is anchored to the membrane by covalent bonds that can be released by treatment with detergent or trypsin. ACE acts on both angiotensin I and bradykinin. The latter is in fact a more active substrate (Km ⫽ 4 ␮m) than angiotensin I (Km ⫽ 30 ␮m) (143). The bradykinin metabolism illustrates the lung’s capacity to control certain circulating peptides. Inactivation of bradykinin does not depend on blood enzymes, as the degradation is seen even during passage through the lung free of blood (140). Several homologues of bradykinin, all exhibiting the five peptide bonds of bradykinin that undergo enzyme-mediated cleavage, are metabolized at rates slower than that of bradykinin: Lys-bradykinin at one-half the rate; Met-Lys and phyllokinin at one-tenth the rate; polistes kinin, no metabolism (140,144,145). It appears that the working enzyme(s) is either highly specific for bradykinin or sequestered in an exclusive microenvironment where some substrates have access and others do not (145). Nevertheless, this enzyme process is remarkably selective. It readily degrades bradykinin but has no effect on other vasodepressor peptides, e.g., eledoisin, substance P, and physalaemin that are known to undergo hydrolysis in lung homogenates. Proteolytic and Antiproteolytic Balance

AMs are the predominant phagocytic cells for the pulmonary defense against inhaled microorganisms, particles, and toxic agents. Other cell types may be recruited by AMs through cytokine-mediated immune-inflammatory responses, thus maintaining the homeostasis of the lung. Common sources for proteases are AMs and other inflammatory cells such as neutrophils, eosinophils, basophils, T-lymphocytes, and mast cells. Neutrophil proteases include, among others, elastase, cathepsin G, collagenase, and plasminogen activator. Neutrophil elastase has been implicated in many lung disorders, especially emphysema. Cellular release of protease can occur via phagocytosis; exocytosis due to stimulation by chemicals, endotoxins, or particles; or leaching due to cell damage or death. These enzymes play a key role in the degradation of proteins or peptides administered into the respiratory tract. The lung also preserves a delicate balance between the proteolytic and the antiproteolytic activity. As by necessity, proteases capable of attacking extracellular components of the lung parenchyma must be controlled by

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antiproteases that interact with the catalytic sites of the proteases, thereby inhibiting the enzyme action. The known extracellular lung antiproteases include ␣1-antitrypsin, secretory leukoprotease inhibitor, ␣1-antichymotrypsin, ␣2-macroglobulin, and tissue inhibitor of metalloprotease. Peptide Metabolism in the Respiratory Tract

Studies have shown that insulin (Asu1,7)-Eel Calcitonin (ECT), and fluoresceinlabeled dextrans are absorbed from the lung, while they are poorly absorbed from the gastrointestinal (GI) tract (146–148). The intrinsic permeability as measured in membrane electric resistance for the lung tissues including tracheal, bronchial, and cultured alveolar epithelium are slightly lower but comparable to that of the intestinal tissues (6,149). The alveolar epithelium, however, exhibits much greater surface area and is thus more advantageous for absorption of macromolecules than the GI system. Pulmonary absorption of proteins such as ECT and insulin are lower than that of dextrans with similar molecular weights. This may be attributed to the enzymatic hydrolysis of the proteins in the lung. The absorption of intratracheally delivered ECT in rat lung, for example, exhibited a pharmacological availability, as measured hypocalcemic effect, of 2.7% of the intravenous dose when given alone, but approached 100% when given with 20 mM bacitracin, a potent aminopeptidase inhibitor that was shown to inhibit the enzyme degradation of ECT by 78% (147). Although some drugs such as interferon (150) may undergo little catabolism on passage through the lung, proteins and peptides are in general subjected to hydrolysis by proteases. The pulmonary tissue, as seen from enzyme kinetics studies of enkephalins, exhibits strong aminopeptidase and dipeptidyl carboxyl-peptidase activities. Taking Met-enkephalin (Tyr-Gly-GlyPhe-Met) as an example, the metabolites obtained from studies using isolated rat lung (151) or cultured alveolar epithelium (6) contain mainly Tyr and TyrGly-Gly, with Tyr-Gly as a minor product indicating some dipeptidyl peptidase activity. These results are consistent with the finding that [D-Ala2]Metenkephalinamide (Tyr-Ala-Gly-Phe-MetNH2) was not hydrolyzed by the enzymes at the alveolar epithelium (6). This peptide is by design resistant to aminopeptidase and dipeptidyl carboxyl peptidase due to the D-Ala2 isomeric configuration of the N-terminal peptide bond, and the change of C-terminal carboxyl group to amide group. Cleavage of Gly-Phe bond in Met-enkephalin implies the action by enkephalinase or ACE on the endothelial side. Gillespie et al. (151) showed that the formation of Tyr-Gly-Gly in the isolated perfused lung study was inhibited by captopril, and ACE inhibitor, but not by thiorphan, an enkephalinase inhibitor. The latter effect may reflect the poor absorption of thiorphan by the lung. Interestingly, bestatin, an aminopeptidase inhibitor, greatly diminished the formation of tyrosine while augmenting the production of Tyr-Gly-Gly (151). Various protease inhibitors have been shown to improve pulmonary peptide absorption. Table 1 summarizes the reported use of these inhibitors on peptides with positive results (147,151–155). Compounds such as sodium glycocholate and bacitracin not only inhibit protease activity but also enhance the transport of protein and peptide drugs (146–147). It appears that

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the peptide degradation problems in the lung can be controlled by inhibition of the protease activity and/or employing modified peptide analogs. Incubation of LHRH and [D-Ala6]-LHRH with isolated primary type II, type I, and A549 cells demonstrated comparable rates of degradation of LHRH with type II and A549 cells. The enzymes responsible for degradation of LHRH were found to be EP 24.11, EP 24.18, and ACE. The rate of degradation of LHRH was found to be ten-fold lesser than in type II or A549 cells. The degradation rate of [DAla6]-LHRH was found three- to eight-fold lesser compared to LHRH and was found to be resistant to EP 24.15 and EP 24.11 but was susceptible to ACEmediated cleavage. Differentiation of type II to type I alveolar epithelial cells in culture was accompanied with a dramatic decrease in the proteolytic activity of isolated type I cells (156). The data indicate that the A549 monolayer retains proteolytic activities comparable to that of type II alveolar epithelial cells and hence may not be truly representative of the barrier function of type I alveolar epithelial cells that comprise the key interface for pulmonary drug delivery. It has been found that natural mammalian peptides that are less than 30 aa in length, are degraded rapidly by ubiquitous peptidases and have very poor bioavailabilities. This may be due to the necessity of the lung to act on neuropeptides that modulate lung function and exert transient effects (157). Proteins with MW approximately between 6 kDa and 50 KDa, however, tend to be resistant to peptidases and exhibit good bioavailabilities on inhalation (157–158). Table 1

Transport and Enzymatic Degradation of Some Proteins and Peptides in the Lungs

Protein or peptide I. Enkephalins (6,151)a Leu-Enkephalin Met-Enkephalin Met-Enkephalin [D-Ala2]MetEnkephalinamide II. Interferons (IFN) (150) Natural human IFN-␣ & rHuIFN-␣2

III. Substance P (152) Substance P

Experimental model

Degradation product

Enzyme inhibitor

IPL-rat

Tyr-Gly-Gly Tyr Tyr-Gly-Gly Tyr

Captopril Bestatin Thiorphan –

CAE-rat

None

–

IPL-rabbit Intravascular circulation Intrabronchial

Isolated lung parenchymal strips (guinea pig)

Negligible in – Absorption into epithelium with degradation –

–

Phosphoramidon Thiorphan Captopril (at high drug levels)

(Continued)

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Table 1 Transport and Enzymatic Degradation of Some Proteins and Peptides in the Lungs (Continued) Protein or peptide IV. Vasopressin (153) [125I]Arginine vasopressin V. Insulin (146,154) Insulin (dimeric hexameric)

Experimental model CAE-rat

Degradation product Cys1-[125I]Tyr2

Enzyme inhibitor Camostat mesylate

des-Phe-Insulin B1 LH-rat and rabbit

BI-2

–

des-Phe-Val-Insulin (rabbit)

– Aprotinin

Insulin

In situ—rat intratracheal

– STIb Bactitracin Na-glycolate LMb

VI. Calcitonin (147,155) (Asu1,7)-Eel calcitonin

Salmon calcitonin

In situ—rat intratracheal

–

Rat lung homogenate

–

In situ—rat intratracheal

–

Na-glyco cholate Bacitracin Nafamostat mesilate Na-glycocholate Bacitracin Bacitracin Chymostatin pCPIb Phosphoramidon

VII. LHRH (156) LHRH; [D-Ala6]-LHRH

Alveolar Type IIType I monolayer A549 monolayer

LHRH 4-10, LHRH 6-10 LHRH 7-10

Catpril Thiorphan, EDTA

a Numbers in parentheses indicate references cited. Abbreviations: IPL, isolated perfused lung; CAE, cultured alveolar epithelium; LH, lung homogenates; STI, soybean trypsin inhibitor; LM, N-lauryl-␤-D-maltopyranoside (absorption enhancer); pCPI, potato carboxypeptidase inhibitor.

Insulin and antidiuretic hormone have been in use for several decades. With the advent of recombinant DNA technology, increasing numbers of proteins including growth hormone, interferons, interleukins, tumor necrosis factor, tissue plasminogen activator, colony-stimulating factors, and erythropoietin are being produced and tested for therapeutic effects. The endothelium selectively degrades

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certain peptides with membrane-bound proteases. Given that a large number of such enzymes of unknown physiological roles also exist, drug delivery through the pulmonary route may have the advantage of circumventing problems associated with the endothelial degradation of proteins and peptides, whether they are for local or systemic action. Because these agents cannot be delivered orally owing to degradation and low absorption in the GI tract, it appears prudent that the pulmonary route be tested as an alternative to injections. The pulmonary route would be ideal for the delivery of proteins such as cytokines, ␣1-antitrypsin, catalases, etc., intended for therapeutic actions in the lung. The presence of a balanced proteolytic activity in the normal lung implies that in disease states, this balance is probably not the same. Various proteases may be activated or released to result in increased peptide degradation. In addition, potential effect of the delivery system on the macrophage-orchestrated inflammatory sequelae should also be considered, especially in disease states where migratory inflammatory cells may be present. On the other hand, permeability may be enhanced and it is not clear what the overall effect on drug uptake might be. For instance, it has been shown that in cigarette smokers, who presumably have some lung disease, that blood levels of inhaled insulin are enhanced compared to normal (159). The delivery of protein and peptide drugs to diseased lungs needs more study. III. A.

Gross Anatomical Considerations in Aerosol Drug Delivery Central Region: Mucociliary Clearance

The tracheobronchial airways are protected by a mucociliary apparatus that serves to entrap and eliminate particles from the lung via a combined activity of mucous secretion from the submucosal glands and goblet cells, and the beating of cilia protruding from the luminal surface of the columnar epithelial cells (160–161). Mucin is stored in a dehydrated state but is spontaneously hydrated upon secretion to produce a gel that floats atop a layer of periciliary fluid. The complex but coordinated beating pattern of the cilia moves this mucous layer toward the pharynx, thereby providing a clearance mechanism for the lung. The submucosal glands are found in the trachea and bronchi, and the surface cells are found as far down as the terminal bronchioles, which are about 1mm in diameter. Thus, it is only the central region of the lung that is protected by the mucociliary apparatus. Any agent that bypasses this central region has the potential to be absorbed or to exert a local effect in the peripheral lung. Regulatory Control of MCC

Both cholinergic and adrenergic innervations of animal submucosal glands have been detected, and each cell type receives fibers from both types of nerve. There are, however, differential responses, as serous cells are more responsive to aadrenergic agonists, while mucous cells are more responsive to ␤-adrenergic agents. The administration of drugs that are inherently adrenergic in nature probably will have some effect on the mucociliary apparatus. MCC is under dual regulatory control, i.e., a neurohumoral control of the ciliary beating frequency (CBF) by a ␤-adrenergic mechanism mediated through the effect of cAMP on

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the ciliary axoneme and a control of the mechanosensitivity of ciliated cells by mechanisms mediated through intracellular calcium (162–163). The latter may result from a direct stimulation of the ciliated surface, or from nervous or hormonal stimulation of mucus secretion leading to increased CBF. Thus, a very broad spectrum of stimuli may increase the rate of MCC (159). MCC is impaired in lung diseases such as immotile cilia syndrome, bronchiectasis, CF, and asthma (164). Hence, the pre-existing disease and its contribution to the deposition and clearance of aerosolized drugs deserve special attention. Aerosol Deposition

All components of the aerosol system including particle size and its chemical content may interact with the mucociliary system. The deposition of the aerosol particulates in the lung depends largely on their mass median aerodynamic diameter (MMAD). Particles having MMAD of 3 ␮m or less exhibit good alveolar penetration (165). Between 3 ␮m and 5 ␮m, particles are uniformly deposited throughout the lung, and particles with a MMAD of more than 5 ␮m are deposited mainly in the conducting airways (166). Hence, even the most optimally formulated aerosol system is likely to come into contact with the mucociliary apparatus. The effect of aerosol deposition and MCC on systemic drug absorption was studied by Colthorpe et al. (167) using a gamma scintigraphic technique. Aerosolized insulin (2–3 ␮m MMAD) in rabbit showed a peripheral:central lung deposition ratio of 1.52. In comparison, an instilled dose gave a ratio of 0.32, indicating major deposition of drug in the central lung. The measured bioavailabilities for the aerosolized and instilled drug were 57.2% and 5.6%, respectively. This 10-fold difference was attributed to mucociliary clearance (MCC) of aerosol deposited in the central region of the lungs. Drug Effects

Although the literature contains many conflicting results due to the complexity of the mucociliary system, effects of some adrenergic and cholinergic drugs on MCC have been documented. Epinephrine and isoprenaline, administered to healthy subjects from an intermittent positive pressure breathing apparatus, showed stimulatory effect on MCC (168). Sympathomimetic drugs in general are known to stimulate mucus secretion (169,170), CBF, and mucociliary transport (162,171). The cilioexcitatory effects of the ␤2-agonists can be blocked by low concentrations of propranolol (10–7 M), which has been shown to have no acute effect on MCC when compared to the placebo effect in healthy subjects (172). The effects of cholinergic drugs on MCC have been well documented in both in vitro and in vivo animal experiments. Acetylcholine and pilocarpine were shown to increase CBF in vitro, while subcutaneous administration of bethanecol to healthy, nonsmoking subjects resulted in a significant increase in MCC from the proximal airways, which may be attributed to increased mucus secretions (173,174). The anticholinergic drugs showed varying effects on MCC. Oral administration of hyoscine resulted in a statistically significant

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retardation of MCC (172). However, administration of methylscopolamine IM did not alter clearance. Other drugs that are indicated in the relief of bronchospasm are the methylxanthines. Aminophylline, given orally, has been shown to promote lung clearance in patients with obstructive airways (175). Some drugs are known to be strongly ciliotoxic; these include propranolol (176) and a wide range of pharmaceutical preservatives (177). B.

Peripheral Lung: Alveolar Epithelium

The alveolar wall consists of a specialized epithelium and a closely apposed network of capillaries supported by a delicate interstitial matrix. Nearly 95% of the alveolar surface area is covered by a single layer of epithelial type I pneumocytes joined by tight junctions. These terminally differentiated cells are vulnerable to a variety of injuries, and have little if any regenerative or replicative potential. In contrast, type II cells are more resistant to oxidative injury (178) and are capable of proliferating and repairing the epithelial lesions (179). AMs are residents of the alveolar spaces and are primarily responsible for the host defense mechanism in the deep lung. Oxidant Antioxidant Balance

Gas exchange is a vital function of the lung, and the effect of an aerosol delivery system on this function should be minimized. The mechanism for gas exchange primarily relies on diffusion, owing to the lung’s high oxygen tension and the differences in Po2 and Pco2 between the blood and the alveoli. This suggests that the air sacs should be kept relatively fluid free so as to keep the diffusion distance to a minimum. Fluid balance in the alveoli can be altered owing to oxidative injury to the alveolar–capillary barrier. The level of oxidants in the lung may be increased by the direct effect of the administered drug or by its activation of the AM-orchestrated phagocytic defense mechanism. At the onset of phagocytosis, AMs show a marked increase in oxygen consumption, which is accompanied by an increase in glucose uptake. The accelerated uptake, known as the respiratory burst, results in the production of superoxide anion catalyzed by membrane-bound NADPH enzymes and the formation of hydrogen peroxide. Free ferrous ions may participate in the Fenton reaction, further producing hydroxyl radicals that are highly destructive to proteins, DNA, and lipid plasma membranes. Hence, it is not surprising that the normal functioning of the lung requires a delicate balance in the handling of oxygen in order to maintain the integrity of the epithelial barrier. Overexposure to oxidants can severely compromise the permeability of the alveolar epithelium (180). An in vitro system for the measurement of oxidative cell injury has been established in our laboratory that allows for fluoromicroscopic analysis of intracellular oxidation in cells exposed to various oxidants using 2⬘,7⬘-dichlorofluorescein diacetate as a marker (181). An increase in oxidative status of the lung can also result in increased metabolic activity, such as the sulfoxidation of a methionine residue in ␣1-antitrypsin, which renders the antiproteolytic enzyme inactive, and leads to the deleterious effects of neutrophil elastase on the lung parenchyma (182).

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Studies on the mechanisms of drug absorption have provided valuable information on the permeability of, and the potential drug toxicity to, the pulmonary epithelium (183). Hydrophilic drugs were shown to undergo slow, diffusion-limited transport. The rate of diffusion was found to be inversely proportional to the molecular weight, with pore sizes corresponding to drug diffusion through tight junctions between type I cells (184–189). Lipophilic drugs were shown to exhibit simple diffusion at rates determined by their partition coefficient via a classical lipoid pore membrane structure. Recent studies demonstrated the promise the pulmonary route may hold for small MW drugs especially those that are subject to intestinal efflux and those that present high polar surface areas (190). This may be due to the high local concentrations of drugs generated leading to a saturation of transporters at the site of absorption. Studies carried out in our laboratory have shown that the tight junction permeability is mediated by intracellular calcium rise and depolymerization of membrane microfilaments, a process that accompanies the action of many absorption enhancers (191,192). Interestingly, cellular damage and epithelial damage induced by oxygen radicals are also mediated by a rise in intracellular calcium (180,193). C.

Aerosol Interaction with Pulmonary Endothelium

The pulmonary endothelium is composed of metabolically active, functionally responsive cells that serve to monitor the circulating bioactive agents (194,195). In addition, the permeability of the endothelium for hydrophilic compounds is at least 10 times greater than that of the alveolar epithelium (192). Aerosolized drugs, especially those meant for systemic delivery, will come into contact with the endothelium. A number of drugs such as bleomycin and monocrotaline have been shown to damage the pulmonary vasculature either through their oxidative properties or through the generation of toxic metabolites (196–198). Other drugs known to cause interference in the localized metabolic activity include desmethylimipramine (DMI), a tricyclic antidepressant drug that has been shown to decrease pulmonary clearance of 5-HT and PEA in a dose-dependent manner (199). The PEA metabolism was diminished entirely owing to inhibition of pulmonary MAO by DMI, while the clearance for 5-HT was lowered due to decreases in both drug uptake and MAO activity (199). Hence, the possibility of drug interactions with the endothelial components cannot be overlooked. Preincubation of cultured endothelial cells with aspirin was found to offer protection from hydrogen peroxide-induced toxicity possibly due to binding or chelation to free cytosolic iron (200).

IV.

Experimental Models in Lung Metabolism Studies

Various methods that have been used to study enzyme activities in other tissues are also applicable to studies of drug metabolism in the lung. These include in vivo models involving the whole organism and in vitro preparations for the

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evaluation of mechanisms and cellular involvement. The interpretation of results from in vitro studies needs to be made with caution, as artifacts can arise depending on the type of method used, and the enzyme processes may not be the same in whole organs as in purified preparations. In intact organs, cells are arranged in close contact with other cell types, blood vessels, and extracellular fluid. There exists an intricate organization of metabolic compartments and diffusion barriers, which controls the access of enzymes to their substrates. A similar organization also exists at the intracellular level, involving coordinated arrangement of subcellular organelles for the functioning of the cell as a whole. Thus, various in vitro models, ranging from isolated perfused lung to tissue and cell preparations, may vary significantly in their enzyme activity and kinetics. A.

In Vivo Methods

The in vivo uptake and metabolism of drugs by an organ may be studied by determining the arteriovenous difference in drug concentration together with the blood flow across the whole organ. This method has been applied with success in organs such as the liver. One of the significant drawbacks of this approach in pulmonary studies is that the lung has a large blood flow, as it receives the entire cardiac output. Because of this large blood flow, even a substantial pulmonary effect may result in only a small change in the concentration gradient across the lung. The arteriovenous difference for glucose, for example, was calculated to be 0.02 ␮mol/mL based on an isolated lung experiment (201). The detection of such a small concentration difference would be difficult and not above the range of the experimental error. The absorption and metabolism of drugs delivered through the pulmonary route may be studied using an intratracheal rat model reported by Enna and Schanker (185). Intratracheal instillations may be a useful preliminary model to assess feasibility of the employment of the pulmonary route for systemic absorption (202). Blood samples can be obtained to determine drug concentrations and degradation products due to lung metabolism. This model has been used to evaluate the lung as a potential route to deliver insulin and calcitonin and its protease activities (146,147). The in vivo studies are of importance in the assessment of overall drug absorption across an organ, but show little information on the mechanisms of drug transport and metabolism. B.

In Vitro Methods

Isolated Perfused Lung

The isolated perfused lung, prepared from rats, guinea pigs, or rabbits, has been a versatile model for studying the metabolism and pulmonary uptake of bioactive agents through the endothelium (203). It offers several advantages as a lung model, including the absence of other organs that can mask the pulmonary effect and the evaluation of drug metabolism in conditions where the tissue structure and cellular organization remain intact. Several experimental parameters such as the perfusion rate, perfusion fluid, temperature, oxygen partial pressure, and the content of test material can also be manipulated to aid in the study design.

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The lung perfusion techniques, including surgical procedures and the perfusion apparatus, have been studied and described in detail by Smith and Bend (204). The key elements of the perfusion apparatus include a peristaltic pump for circulating perfusion fluid, pH meter, an artificial thorax with an upper reservoir (both are jacketed for temperature control), an animal ventilator, and a recirculating waterbath. The artificial thorax serves as a chamber for the isolated perfused lung that, through connectors built in a Plexiglas disk covering the chamber, is attached to the upper reservoir via the pulmonary artery cannula and to the animal ventilator via the tracheal cannula. A mixture of 5% CO2 and 95% O2 or warm, humidified air to which sufficient CO2 has been added to control the pH of the perfusion medium can be used as the ventilating gas. Krebs–Ringer bicarbonate buffer fortified with 4.5% bovine serum albumin and 5 mM glucose has been the most widely used perfusion medium. Albumin has been shown to increase drug solubility and prolong perfusion without developing edema. Application

Studies of drug uptake by the lung can be carried out in two ways. The lung can be perfused with a drug-containing medium that is recycled continuously for serial determination of drug concentrations. Or simultaneous analysis of arterial and venous concentrations across the organ may be made via a “once-through” perfusion. The potential that drug may be delivered through the trachea and monitored in the continuously recycled perfusion medium has not been fully explored using this model. The perfused lung model has a viability of about 4 hours, which is sufficient time for most drug uptake and metabolism studies but which may not be sufficient for studies of slow reactions involving conjugating enzymes (205,206) or for monitoring drugs delivered through the trachea. Application of this model requires skillful surgical preparations and careful monitoring of potential lung edema and atelectasis during perfusion. The lung does not fulfill all the requirements for isolated perfusion studies in that an ideal organ for this purpose would have a single arterial supply and venous drainage and consist of a single cell type. However, its applicability to both epithelial and endothelial studies makes the isolated perfused lung model invaluable. Lung Slices

The preparation of tissue slices in studies of drug metabolism is simple but is subject to variations. Lung slices present two cut surfaces that can lose intracellular contents such as enzymes, cofactors, and ions. The supply of substrates and the removal of metabolic products are unnatural, being through the cut surfaces rather than via the capillaries. The thickness of lung slice is also a variable that will affect gas and metabolite diffusion. Studies using rat lung and a McIlwain tissue slicer showed that the oxygen consumption was optimal when the slice thickness was between 0.6 mm and 1.0 mm (207). This may not reflect the rate in isolated perfused lung preparations, as Faridy and

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Naimark (208) have shown that the oxygen consumption for the intact lung is lower than that of lung slices. Studies have shown that lung slices are also sensitive to lowered temperature during sample preparation. This method has been used to demonstrate the uptake of paraquat by lung tissue (3,209). Tissue Minces and Homogenates

Although these are the simplest preparations to employ, they may exhibit inconsistency in enzyme activity. The oxygen uptake of minced lung tissue is very low because of leakage of intracellular metabolites. These methods can overestimate enzyme activity due to enhanced access of enzymes by the substrates and produce different metabolic pattern. The latter is seen from studies on chlorcyclizine and imipramine, which show that these drugs are oxidized by subcellular lung preparations but not by the intact lung when delivered via the vascular system (128,210,211). For the characterization of total enzyme capacity, lung tissue may be homogenized and prepared to yield microsomal enzymes and, soluble fraction containing peptidase activity. Isolated Cell Systems Cell Types

Isolated cell populations may be used to localize the pulmonary metabolic activities to specific cell types in the lung. The isolation of all the different cell types present in the lung poses a formidable challenge since it contains over 40 different types of cells; fortunately, not all of them are likely to play a significant role in xenobiotic drug metabolism. The two important cell types with major cytochrome P-450 contents are Clara cells and alveolar type II pneumocytes. The latter is of special interests because this cell type is metabolically active in the regulation of pulmonary surfactants and, as a progenitor of type I cells in the repairment of the alveolar epithelium. Other cell types such as alveolar macrophages, tracheal epithelial, type I, endothelial, and mesothelial cells may also be important candidates for study, especially in the evaluation of cell damage by inhaled chemicals and peptide degradation. Cell Isolation

Alveolar macrophages can be readily obtained by bronchoalveolar lavage (133). The interstitial cell types are usually obtained by digesting the lungs with proteolytic enzymes and purifying the cells by centrifugal elutriation (135). A typical procedure for cell isolation from rats may involve the following steps: Animals are anesthetized with sodium pentobarbital (150 mg/kg body wt). The heart and lung are quickly removed, and the lung is perfused with saline to remove blood cells. AMs are removed by tracheal lavage with ice-cold phosphate-buffered medium. The lung is then filled with a buffered medium containing DNase and elastase or other proteolytic enzymes and incubated for 30 minutes at 37°C to free lung cells. Following this enzyme digestion, the lung is minced and the digestion is arrested by incubating the minced tissue in phosphate-buffered medium containing 25% fetal calf serum and 0.006% DNase, to prevent cell clumping, at

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37°C for 10 minutes. The cell suspension is then strained through nylon mesh of successively smaller pore sizes. These cells are then ready for elutriation separation of different cell types. In the isolation of cells from rabbit, Devereux and Fouts (29) have shown that a centrifugal elutriation procedure consisting of four stepwise increases in flow rate with elutriator rotor speed of 2,000 rpm allows for the separation of type II cells and Clara cells. Identification of type II and Clara cells can be made by phosphine 3R fluorescence and by light microscopy using a nitroblue tetrazolium staining method, respectively. The use of different proteolytic enzymes has been shown to yield different proportions of pulmonary cells that may not represent the normal lung cell population. Alteration of enzyme activity may also occur during the isolation process. For example, protease I can drastically inhibit the mixed function oxidase activity if internalized, although it does not alter the enzyme activity of cells with an intact membrane. Alveolar Epithelial Monolayer

Because of the difficulty in establishing in vitro type I alveolar epithelial cell lines, the epithelial monolayer approach developed using type II cells is a very useful model for the elucidation of mechanisms on alveolar epithelial permeability and enzyme activity for proteins and peptides. Most studies on pulmonary peptide absorption have focused on the overall transport of peptides rather than transport mechanisms, using anatomically complex, intact lungs. The epithelial monolayer offers an alternative model that complements the whole-lung studies. The method of isolation and culture of alveolar type II cells under sterile conditions routinely carried out in our laboratory (6,192) is summarized in Table 2. Formation of the monolayers, grown in tissue-treated Nucleopore filters, can be monitored by daily measurement of electrical resistance, which reaches maximum (–1500 ⍀cm2) on or about day 6 of culture. Other studies have reported similar results (212,213). Figure 1 shows the electron micrographic evidence of the transformation of the cuboidal type II cells into squamous, type I-like monolayer and the development of tight junctions after 6 days in culture. These morphologic changes are similar to those observed in developing type I cells in vivo (179). Studies by Danto et al. (214) also show that type II cells grown in vitro develop phenotypic characteristics of type I cells. Alveolar type II cells of human origin have been cultured to generate type I-like morphology with a peak TEER of 2100 ⍀cm2 (215,216). In order to assess the Calu-3 cell line’s potential as a metabolic model, two CYP-450 isozymes (1A1 and 2B) were examined for functionality and inducibility. The cell line was found to contain other CYPs, however, did not demonstrate inducibility hence, bringing into question its utility especially for the evaluation of metabolism of lipophilic drugs (217). Similarly, A549 cell lines, comprising of a human lung adenocarcinoma cell line was found to exhibit very low aminopeptidase activity compared to primary alveolar epithelial cell monolayers. It however was found to express functional CYP1A1 and 2B6 isozymes (218,219). Hence, the utility of various cell lines needs to be carefully evaluated against the nature of the drug to be studied. The alveolar surface contains type II, type I,

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Table 2 Preparation of Rat Pulmonary Type II Cells and Culture for Alveolar Epithelial Monolayers Preparation

Methods and reagents

Intact lung

Perfusion (0.9% NaCl) Brochoalveolar lavage (Ca2+ and Mg2+ free, 145 mM NaCl, 5 mM KCl, 9.35 mM Na2HPO4, 1.9 mM NaH2PO4, 5 mM glucose, pH 7.4) Enzymatic digestion [phosphate buffer with elastase (40 U/mL) and DNAase (0.006%) Incubation of minced lung (phosphate buffer with 25% fetal bovine serum and 0.006% DNAase) Filtration and centrifugation Sterile Percoll density gradient (second band) Identification: phosphine 3R fluorescence; purity: >70% Culture medium: 1:1 F12 and Eagle’s-modified maximum essential medium supplemented with 5 mg/mL insulin, 0.1 mg/mL epidermal growth factor, 4 mg/mL transferring, 0.5 mg/mL hydrocortisone, 100 U/mL penicillin, and 10 mg/mL gentamicin Differential adherence (5% CO2, 37oC): Filter: 0.4 ␮m pore, 1.2 cm2 tissue-treated nucleopore filter; 2 ⫻ 106 cells/cm2 Culture: change medium every 48 hr; remove unattached cells Characterization (>90% type II cell purity): Electrical resistance Transepithelial potential Electron microscopy

Mixed cells

Type II cells Epithelial monolayer

and alveolar macrophages in close proximity with endothelial cells. Models employing coculture conditions have been employed to create conditions that are closer to the in vivo situation. Cocultures of human lung epithelial cell lines (A549 and NCI H441) and human primary microvascular endothelial cells (HPMEC) stimulated with dexamethasone. The model demonstrated a significant drop in TEER values after a 24-hour exposure to basolaterally applied TNF␣. Such models may be of utility to examine drug metabolism that occur due to a contribution of the various cells acting in concert possibly due to release of fibrogenic cytokines (220). A triple-cell culture system employing alveolar epithelial cells, endothelial cells, and macrophages has been described (221). Models employing viable cascade impactors may be of value in exploring subtle effects of particle deposition and possible stimulation of drug metabolism resulting from biomechanical stimuli (222,223). V.

Current Application of Aerosols in Drug Therapy

A.

Drugs Administered for Local Action

Asthma

Aerosols have been used primarily for the treatment of bronchospasm in asthmatic patients. Both new drugs and new therapeutic approaches against asthma have been reported with proven efficacies in recent literature. Formoterol is a

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Electron microscopic examination of the alveolar epithelial monolayer developed from pulmonary type II cells. (A) Transmission electron micrograph (TEM) of type II cell culture plated on tissue-treated Nucleopore filter at day 1. (B) TEM of developed type I cell-like monolayer and tight junction at day 6. (C) Scanning electron micrograph (SEM) of epithelial monolayer (same as B) with markings of tight junctions. Bar indicates I ␮m for TEM and 10 ␮m for SEM. Figure 1

new ␤2 agonist that has a rapid onset of action and long duration of drug effect. Doses of 12 ␮g and 24 ␮g formoterol administered as a dry powder inhaler were shown to exert a similar onset of action to 400 ␮g salbutamol, but displayed a considerably longer duration of action (224). Another drug, salmeterol, when administered as a metered dose inhaler, was also shown to have a greater efficacy and a longer dose-dependent duration of action (9–18 hours), although a slower onset of action than albuterol. The inclusion of long-acting drugs will avoid nocturnal awakenings for asthma control, improve compliance, and reduce disruption at work or school (225). Treatment of asthma may also be improved using innovative approaches in the administration of the therapeutic agent. Continuous aerosol administration has been suggested for the treatment of acute asthma to replace the frequent, intermittent administrations in an emergency situation. The feasibility of continuous administration of albuterol using an

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acorn-type nebulizer attached to a vented face mask with two intervening conduit adapters has been demonstrated (226). Anticholinergic drugs are effective in the treatment of asthma as catecholamine excretion and cholinergic stimulation are increased during the night in asthmatic patients. Studies on the combination therapy of oral terbutaline and inhaled ipratropium have indicated additive bronchodilating effects (227). The comparison of aerosolized BDP and sustained-release theophylline for the treatment of mild-to-moderate chronic asthma in children (age 6–16) was studied (228). Both drugs were effective, but aerosol BDP resulted in comparable symptom control with less bronchodilator use, fewer courses of systemic steroids, and fewer side effects than theophylline, although caution must be exercised in treating children with BDP owing to possible growth velocity suppression. Other aerosolized drugs such as furosemide and nedocromil, used either alone or in combination, were shown to have a protective effect against exercise-induced asthma in children (229). A bronchoselective calcium entry blocker, RWJ 22108, has also been shown to alleviate symptoms of asthma, and has little effect on blood pressure when given as an aerosol (230). This agent is of interest because calcium is involved in many responses in asthma that may be controlled by blockage of calcium entry in the airway smooth muscle. The effect of the lung on the metabolism of the above-mentioned drugs remains unclear. Early studies have shown that isoproterenol in man undergoes metabolism following two major routes. The orally administered drug is mainly converted to isoproterenol O-sulfate, while the inhaled drug is O-methylated by the catechol I-methyltransferase (COMT) and is excreted as O-methylisoproterenol sulfate in the urine (231,232). Accumulation of the O-methylated metabolite, which has ␤-adrenergic antagonist action, has been reported as responsible for the development of tolerance to the bronchodilating action of isoproterenol. The fate of some structurally related but more selective ␤2-adrenoceptor agonists is different. Terbutaline and salbutamol are not candidates for metabolism by COMT (233). These studies show that the metabolic activity in the lungs may play an important role in the drug’s pharmacological action. Respiratory Infections and Conditions

The application of the aerosol system in the treatment of respiratory tract infections offers the advantage of providing localized drug effect and lowered systemic drug toxicity. Aerosol therapy is potentially very useful in the treatment of infections arising from immunocompromised conditions. A recent study showed the use of amphotericin B aerosol as a prophylactic treatment to prevent invasive pulmonary aspergillosis in neutropenic patients who have undergone bone marrow transplantation (234). Inhalations were found to be easily administered and well tolerated, with minimal systemic absorption of the drug. Opportunistic infections are common in AIDS patients, especially the occurrence of Pneumocystis carinii pneumonia. Pentamidine given as aerosols has been used as a prophylactic measure (235).

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The efficacy of the aerosolized aminoglycoside, tobramycin, was evaluated in patients with clinically stable cystic fibrosis and Pseudomonas aeruginosa infections in a multicenter clinical study (236). Direct aerosol delivery of high doses of tobramycin decreased the density of P. aeruginosa in the sputum by a factor of 100 with no production of ototoxicity or nephrotoxicity. The forced expiratory volume in one second (FEV1) was also increased compared to placebo. These results showed that the pulmonary route is both efficacious and safe for the administration of the tobramycin and it is now an approved drug (237). Other antibiotics have also been useful to treat respiratory infections and this is an area worthy of additional investigation. Hall (238) reviewed the use of aerosolized antibiotics including tobramycin, gentamycin, vancomycin, and others. It was concluded that “aerosolized agents for the treatment of acute pulmonary infections offer the potential advantages of targeting the agent to the site of the infection and achieving relatively high levels of drug in the secretions with minimal systemic toxicity and side effects.” Additionally, the American Academy of Pediatrics (239) noted in a recent review the extensive use of inhaled antibiotics over several decades. Palmer et al. have demonstrated that this is a useful approach in the treatment of ventilator-acquired pneumonia (240). Aerosolized antibiotic delivery is a promising area that deserves further attention to improve treatment of respiratory infectious diseases. Other approaches are being investigated as well, such as in the treatment of cystic fibrosis, an aerosolized recombinant secretory leukoprotease inhibitor was shown to decrease the amount of active neutrophil elastase in the epithelial lining fluid from patients with cystic fibrosis and inhibit the neutrophil elastaseinduced damage to the respiratory epithelium (241). Lung Cancer

An increased understanding of molecular events particularly in the polyunsaturated fatty acid metabolic (particularly linoleic and arachidonic acid) pathways have led to the identification of novel therapeutic targets and biomarker strategies (e.g., elevated PGE2 levels in BAL fluid in patients with bronchogenic carcinoma (242) for the treatment of lung cancer). NSAIDs, including COX-2 inhibitors were found to be active in vitro in NSCLC. Cigarette smoke activates nuclear factor kappa B (NF-␬B) in NSCLC with induction of NF-␬B–regulated gene products COX-2, cyclin D1, and MMP-9. The effects were blocked by celecoxib (243,244). B.

Drugs Administered for Systemic Action

Volatile anesthetic gases are the one of the oldest classes of drugs delivered by inhalation. The total amount of drug given is relatively large, which corresponds to blood levels of 10–100 mg/L. Hence, it is likely that the metabolic activity of the lung is saturated. Agents such as cyclopropane, enflurane, and halothane undergo minimal metabolism. The exact contribution of the lung to the metabolism of general anesthetics is not well known. Studies on the aerosol delivery of proteins and peptides to humans have begun to appear in the literature with encouraging results. Laube et al. (245)

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studied the delivery of aerosolized insulin to diabetic patients and recorded a 55% decrease in fasting plasma glucose, indicating a significant drug transport from the airways to systemic circulation. The delivery of aerosolized leuprolide acetate to healthy subjects, reported by Adjei et al. (246), also resulted in a substantial bioavailability, 28% relative to subcutaneous injections, for the peptide formulated as suspension aerosols. The peptide bioavailability was found to be formulation dependent, with the suspension aerosol being fourfold more effective than the solution aerosol. The effect of lung metabolism on the bioavailability of peptide drugs in relation to the aerosol delivery system has not been characterized. Studies have shown, however, that the deposition of aerosolized drug in the lung is an important factor in regulating drug bioavailability. Using a gamma scintigraphic imaging technique, Colthrope et al. (167) demonstrated that deposition of aerosolized insulin in rabbit lung has a peripheral central lung ratio (penetration index) of 1.52, whereas the instilled dose has a penetration index of 0.32. There is a 10-fold decrease in bioavailability for the intratracheal insulin (5.6%) when compared to aerosolized drug (57.2%), which has been attributed to the MCC of aerosol in the central lung. The aerosol deposition, which shows substantial amounts in both the central and the peripheral regions, may also affect the enzyme degradation process in the lung, as different cell types exist in these regions with varying enzyme activity. It appears that both the mucociliary transport and the enzyme activity may have a strong influence on the absorption of aerosolized peptides. The rate of drug elimination from the lung may include the rates of drug transport from the airways to blood, MCC, and enzyme kinetics. Concurrent advances in particle and device engineering thereby allowing readily dispersible particles to be manufactured, which lead to better lung deposition have recently led to an expansion in the utility of the pulmonary delivery platform for the delivery of drugs for systemic use (247). It is now speculated that due to the larger mass that can be delivered in a more cost-efficient manner, possibly a wide spectrum of drugs such as proteins, peptides, oligonucleotides, and vaccines may be delivered systemically via the lung (248). As of March 2006, Exubera™ (human inhaled insulin) received almost simultaneous approval in Europe and the United States for the treatment of diabetes.

VI.

Conclusions

Studies on the metabolic functions of the lung have provided a wealth of information on the way in which the lung handles either endogenous substrates exhibiting lung uptake affinity or exogenous chemicals that may be activated to cause lung-specific toxicity or carcinogenesis. These investigations have helped in establishing the various enzyme systems and their distribution in the lung. Study of the effect of lung metabolism on drug bioavailability is a more recent focus stemming from the growing awareness that the respiratory tract may be an effective route to deliver therapeutic agents, especially proteins and peptides that are degraded or not well absorbed through the GI tract. This is a direct consequence of the combined advances in biotechnology, which allows

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for economic production of bioactive macromolecules, and in aerosol technology that has provided marked improvements in inhaler devices and aerosol formulation. Protein and peptide drugs are unlikely candidates for metabolism by the cytochrome P-450 systems, but are subjected to degradation by proteolytic enzymes in the lungs. Current studies on the enzyme degradation process are carried out based on the use of various protease inhibitors and, in some studies, measurement of degradation products. The alveolar region has been shown to contain aminopeptidases, serine proteases, dipeptidyl carboxypeptidases, trypsin, and endothelial enzymes such as ACE. It has been well established by inhibitor action that the proteases play a key role in determining the bioavailability of peptide drugs. Conclusions on the enzymatic process made by using inhibitors alone could yield confusing results regarding the mechanism and the location of the enzymatic degradation. Studies have shown that one inhibitor may have an effect on one protein degradation process but not on other systems, suggesting that each drug may be responsive to a set of enzymes. It should be emphasized that the inhibitory activities of the enzyme inhibitors are not always specific (249), hence do not show direct evidence of enzyme activity. Certain inhibitors, such as aprotinin, leupeptin, STI, etc. (serine protease inhibitors), also exhibit preferential activity toward either membrane-bound or cytosol enzymes (154). In addition, the location of membrane enzymes relative to the apical and basolateral sides of the epithelium is also different (153). These results suggest that various lung models used in the study may lead to different enzymatic actions. That the lung is anatomically complex and contains more than 40 different cell types also contributes to the difficulty in assessing the lung’s enzymatic capacity. A welldefined lung model coupled with the characterization of the degradation products should be very useful in elucidating the mechanism of enzyme action and the specific inhibitor effect. It appears that the design of a satisfactory protocol to assess the metabolic activity of the lung toward a specific peptide drug remains a challenge. Alveolar epithelial monolayers have been demonstrated in several studies to be a valuable model for the evaluation of the membrane enzymes of the alveolar epithelium (6,153,250). With the use of protease inhibitors and the measurement of degradation products, this model allows one to determine the specific enzyme activity of the pulmonary epithelium toward a peptide drug and its effect on drug permeability. Aerosol formulation, which has been shown to have effects on peptide absorption (246), may also affect the enzymatic degradation of the drug. Aerosol formulation and other factors can also influence the aerodynamic particle size, which in turn is a primary determinant of deposition pattern in the lung. Because the host defense mechanisms are primarily MCC and alveolar macrophage phagocytosis in the distal airways, different patterns of metabolic degradation may result from different aerosol formulations. This suggests that integrated experimental approaches are needed and may involve the use of lung homogenates, isolated perfused lung, and isolated cells working in concert, allowing the elucidation of the mechanism of enzyme degradation.

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Despite the fact that alveolar macrophages are a key cell type in pulmonary defense against inhaled substances, surprisingly few studies have been devoted to the effects of AMs on peptide permeability and enzyme degradation. Given that the alveolar type I-like monolayer is a useful model for the characterization of the epithelial protease activity and drug transport, a combined use of AMs and the monolayer system may further contribute to the understanding of the enzyme activities in the distal lung. In summary, many studies have demonstrated that various protease inhibitors and absorption enhancers can be used to promote pulmonary absorption of peptides for systemic drug effect. Aerosolized insulin and other drugs have already been studied in humans (241,245,246). Although aerosol delivery is associated with variability in the fraction of the administered dose deposited, this may not be an obstacle in the future, pending further improvement in inhaler devices and aerosol formulation. The capacity of the lung to metabolize drugs is not as high as that of the liver. This, in conjunction with the vast surface and the thin nature of the alveolar epithelial barrier, makes the lung highly feasible route for the delivery of aerosolized drugs.

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7 Bioavailability and Pharmacokinetics of Inhaled Drugs

AKWETE L. ADJEI

YIHONG QIU

Kos Pharmaceuticals, Cranbury, New Jersey, U.S.A.

Abbott Laboratories, North Chicago, Illinois, U.S.A.

PRAMOD K. GUPTA Bausch & Lomb, Rochester, New York, U.S.A.

I.

Introduction

The evolution of lung drug-delivery technology has undergone significant advances over the last few decades, so that today it is one of the most effective and rapid modalities for treating pulmonary diseases. Many of these diseases are localized in the lung, examples being asthma, allergic airway disease (AAD), chronic obstructive pulmonary disease, and adult respiratory distress syndrome (ARDS) or emphysema. Of the series of drugs used in these conditions, the most common therapeutic entities include ␤-agonists, steroids, mucolytics, and cholinergics. However, a large number of compounds have recently been aimed at for systemic delivery via the airways, especially those that are challenging or infeasible for oral delivery, such as peptides and proteins. If systemic concentrations achieved by lung delivery could exceed minimum effective therapeutic requirements, the lung might become an effective route for drugs intended for treating certain systemic diseases. Insulin, leuprolide, and a variety of other therapeutic peptides and proteins that are in various phases of clinical development might be effective drug therapies tomorrow. In fact, Exubera, inhaled insulin using a proprietary inhalation device and powdered insulin formulation developed by Nektar Therapeutics, was recently approved for marketing by the 187

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Food and Drug Administration (FDA), marking the first success of many development programs of pulmonary delivery of insulin. The lung is a complex organ. Starting with the trachea, a hollow tube that branches asymmetrically into daughter tubes, the lung concludes at the finest of airway structures, called alveoli. Significant enzymatic reactivity of the lung coupled with this systemic subdivision of the airways creates a geometrically complex pathway and biophysically unfriendly environment for inhaled particles. This makes quantitation of lung drug bioavailability difficult. Further, lung deposition of pharmaceutical aerosols is generally less than 100% of the nominal dose. This is due often to complex biophysical factors associated with filtration mechanisms of the respiratory system (1). For locally acting drugs, this filtration effect may not carry a significant pharmacologic or therapeutics risk, because absorption is not a prerequisite for therapeutic effect. However, for drugs intended for systemic delivery, lung filtration effects are crucial for efficacy, because this influences the extent of deposition as well as throughput of drug to the body. Beyond filtration of inhaled drug particles, inertial impaction and device losses (i.e., drug retained by the mouth adapter and actuator) contribute to incomplete delivery of pharmaceutical aerosols to the lung (1–3). Inefficient delivery of drugs to the lung may be the single largest cause of low drug absorption from the human lung. In addition, physicochemical characteristics of the drug—namely, solubility, partition coefficient, permeability, stability to metabolizing enzymes, molecular weight, and stimulation of the alveolar macrophage clearance mechanism (4)—may play a significant role in the extent of deposition as well as throughput of drug to the body. These factors introduce a major problem in inhalation drug delivery relative to bioavailability assessment. First is the issue of absolute bioavailability based on intrinsic properties of the drug, such as charge, size, solubility, dissolution rate, permeability, and aggregation. Second is extrinsic bioavailability, wherein in vivo and biophysical factors having an effect on throughput of drug to the systemic circulation are analyzed. Examples include membrane filtration or size exclusion, enzymatic deactivation, tissue extraction, and mucociliary and alveolar macrophage clearance mechanisms. This chapter describes a paradigm for assessing fractional bioavailabilities of lung-delivered compounds. The discussion includes examples relating pharmacokinetic (PK) end points with aerosol formulation variables. Representative formulation variables are summarized in Table 1. Two mechanistic approaches for estimating lung bioavailability are also presented. Discussions of these approaches Table 1 Physicochemical and Biophysical Factors Affecting Bioavailability of Pharmaceutical Aerosols Surface energetics Surface charge Electrophoretic mobility Aggregation propensity

Dispersion phenomena

Biophysical properties

Solubility of drug Dispersibility of surfactant Partition coefficient of drug Moisture-sorption potential Inertial impaction

Vapor pressure of propellant Nonvolatile solvents Aerodynamic diameter Particle velocity Particle density

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include compartmental and noncompartmental PK modeling with examples, which include lung absorption of lipophilic and nonlipophilic drugs. II.

Formulation Effects and Bioavailability

Because of its large surface area and low enzymatic reactivity, and circumvention of hepatic first-pass effect, drug absorption from the lung is generally rapid and efficient compared to the nasal and oral routes (3,5). With most drugs, there is almost quantitative absorption—i.e., essentially 100% bioavailability (6) of the fraction of drug that actually deposits in the lower airways. In this sense, lung drug delivery could be considered effective and efficient compared to parenteral routes of administration. Yet, formulation effects and biophysical parameters related to device and patient could have a significant impact on bioavailability of inhaled drugs. Three types of inhalation drug-delivery systems are used in therapeutics today. These include metered dose inhalers (MDIs), dry-powder inhalers (DPIs), and nebulizers. Mechanistically, deposition of aerosolized drugs from these delivery systems involves inspiratory flow and breathing maneuvers of the user as well as dynamics of the aerosolized drug. Of these, aerodynamic particle size, particle velocity or inertial impaction, sedimentation, and diffusion in the airways are most important. Particles greater than about 10 ␮m deposit predominantly in the upper airways (i.e., throat and trachea) by impaction. Particles less than or equal to 3 ␮m generally deposit in the lower airways (alveoli and acini) by sedimentation. Generally, though, many pharmaceutical aerosol products contain drug particles 5 ␮m or less in diameter and thus deposit mostly in the transitional respiratory zones of the lung. Throughput is a term used to describe the amount of drug deliverable from aerosol devices. Respirable fraction (RF) describes the fraction of aerosolized dose-surviving filtration and impaction mechanisms of the nasopharynx. Throughput and drug-delivery efficiency from MDIs, DPIs, and nebulizers vary greatly. Dosimetry from MDIs is generally believed to be more efficient and, from patient-convenience standpoint, more reproducible than DPIs and nebulizers. This distinction is not clear-cut because dosimetry generally is linked to physicochemical properties of the drug. Factors causing aberrances in drug deposition in the airways from MDI, DPI, or nebulizer can thus have a significant impact on bioavailability of the drug. These factors are common to all forms of inhalation aerosol formulations, with the exception of vehicle effects such as is the case with MDIs and DPIs. These common factors are briefly described in the following section. A.

Metered Dose Inhalers

MDIs are by far the most popular method for pulmonary drug delivery, although DPIs have gained universal market acceptance in recent years because of stratospheric ozone depletion of chlorofluorocarbon (CFC) propellants. These systems are tamperproof and capable of delivering very accurate and reproducible doses of aerosolized drugs to the lung. Typical MDI formulations

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contain an active ingredient, generally in solution or suspension, along with inactive excipients (e.g., surfactants, suspending agents, and protective agents), propellants, and solvents. Drug Characteristics

Solubility and concentration are two variables that can be altered to improve deposition and absorption characteristics of drugs from the lung. Generally, lipophilic compounds are more readily absorbed from the lungs than their hydrophilic counterparts (7). In dilute suspensions, gravitational phase separation of the dispersed phase is proportional to the square of particle diameter. Thus, coarse particles will form suspensions that are inherently less stable physically and fraught with potentially significant content uniformity problems compared with dispersions created with fine particles. But a dispersion that consists of finely milled particles could also have content uniformity issues as well, if interparticulate charge interactions are excessive to the extent that aggregation and caking phenomena start to occur. For suspensions, increasing drug concentration to optimize throughput generally also improves suspension quality by viscosity increases, but this often reduces RF of the emitted dose. Thus, particulate interactions, micromeritics, polarity, charge, and morphology may be specific drug characteristics that can have a significant impact on bioavailability of inhaled aerosols. Surfactants

Surfactants generally aid formulation processing and minimize particle aggregation, thus improving physical stability and dose uniformity in pharmaceutical aerosols. However, interaction of surfactants with container/closure components and their potential effects on extractables can have an impact on drug stability. Surfactants that either are immiscible with propellants or are nonvolatile could decrease the vapor pressure of the aerosol, thus lowering the energy required for “drying” the drug particles. This results in the formation of coarse aerosols, thus decreasing the extent of lower-airway deposition (8). Conversely, increasing surfactant concentrations in propellant systems have been demonstrated to decrease vapor pressure and droplet-volatilization rates in MDI products (9–12). The impact of this on formulation performance is increased upper-airway deposition, particularly losses of drug to tracheobronchial milieu and a resulting decrease in extrinsic bioavailability (13). Solvents

Solvents in pharmaceutical MDIs are usually inert and help solubilize drug and surfactant. Solvents in MDIs serve also as carriers for delivering medicament to the airways. Ethanol, propylene glycol, polyethylene glycol, and glycerin are the most commonly used solvents in pharmaceutical MDIs. However, like liquid surfactants, these solvents often alter vapor pressure and stability of aerosol formulations. For example, 20% v/v ethanol in CFC propellant blends has been shown to cause a significant drop in vapor pressure with a subsequent increase in droplet size of the aerosol (8). An increase in “wetness” of the product due to

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increased proportions of solvent often leads to reduced lower-airway deposition, thus altering extrinsic bioavailability of the aerosolized drug. Propellants

In general, there are two primary types of propellants, liquids and gases, both providing enough force to aerosolize liquid mixtures. Liquefied propellants are gases that exist as liquids under pressure. Because the aerosol is under pressure, the propellant exists mainly as a liquid, but it will also be in the headspace as a vapor. As the product is used up, some of the liquid propellant turns to vapor and keeps the head space full of vapor. In this way, the pressure in the container remains essentially constant and spray performance is maintained throughout the life of the aerosol. When the liquid propellant emerges from the device, and if intimately mixed with product droplets, these will become vaporized into still smaller droplets dependent upon the pressure of propellant in the aerosol. Compressed gas propellants really only occupy the headspace above the liquid in the container. When the aerosol valve is opened, the gas “pushes” the liquid out, the amount of gas in the headspace remaining literally constant but occupying more space. Thus, the pressure will drop during the life of the container until completely empty. Unlike liquefied propellants, there is no liquid to instantly vaporize when product emerges from the actuator; the only means of creating the droplets required to form an acceptable spray is by mechanical action as it passes through both the valve and the actuator. The choice of these components is critical, and fortunately the suppliers of both valves and actuators are able to offer specifications, which meet the necessary requirements. There are a variety of materials that are used as conventional propellants pharmaceutical and personal care products. These are described below. Liquefied Petroleum Gas

Aerosol propellant–grade liquefied propellant gas (LPG) consists of highpurity hydrocarbons derived directly from oil wells, and as a by-product from the petroleum industry. They consist of a mixture of propane, isobutane, and n-butane. These propellants are used in most aerosols today, and have been used for many years in household aerosol products. These gases are flammable, and this is reflected in the classification of aerosols, which contain them. Dimethyl Ether

This is an alternative liquefied propellant, and is more common in personal care products, and some air fresheners. CFCs and Hydrofluroalkanes

These LPGs used to be very common prior to the discovery that they either affect the ozone layer or act as greenhouse gases. The CFCs, trichlorofluoromethane (CFC-11), dichlorodifluoromethane (CFC-12), and dichlorotetrafluoromethane (CFC-114) have significant ozone-depleting potential, and are no longer used in consumer aerosols in the western world.

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They are, however, permitted in inhalation and nasal aerosols, as used in the treatment of asthma around the World, until hydrofluroalkanes (HFCs) are used to reformulate these products. The HFCs, tetrafluoroethane (HFC-134a) and heptafluoropropane (HFC-227) are greenhouse gases but not to the extent of carbon dioxide, and are being considered alternatives to CFC propellants. Nonsoluble Compressed Gasses (e.g., Compressed Air and Nitrogen)

These are sometimes seen in consumer products, and are an environmental alternative to LPG. Soluble Compressed Gasses (e.g., Carbon Dioxide)

This is another alternative to LPG, but has limited use, mainly with alcoholic systems, such as air-treatment products, deodorants, and personal care products. Some of the key properties of CFC and HFC propellants are summarized in Table 2. As can be seen, HFC-134a has a much higher vapor pressure and a correspondingly lower boiling point than most CFCs or CFC blends, typically used in pharmaceutical MDI products. A higher vapor pressure increases particle velocity and subsequent upper-airway deposition as a result of inertial impaction. This decreases RF and bioavailability of the aerosolized drug. Alternatively, a fine aerosol product formulated with HFC-134a would be quite dry by the time the aerosolized drug particles exit the mouth adapter. Such particles could survive filtration and impaction losses, thus depositing in the respiratory zones of the lung and presenting a significantly higher extrinsic bioavailability. Other Excipients

Surfactants, solvents, and propellants are the main category of excipients used in pharmaceutical MDI products. Other excipients such as flavoring agents (e.g., menthol in Aerobid®), complexing agents, and antimicrobial Physical Properties of Aerosol Propellants Used for Pharmaceutical Metered Dose Inhaler Aerosols Table 2

Formula Molecular weight Boiling point (°F) Boiling point (°C) Vapor pressure at 25°C (77°F) (psi) Vapor pressure at 25°C (77°F) (bar) Liquid density at 25°C (77°F) (lb/ft3) Liquid density at 25°C (77°F) (g/cm3)

CFC-11

CFC-114

HFC-227ea

CFC-12

HFC-134a

CCl3F 137.4 74.9 23.8 0.6

CClF2CClF2 170.9 38.8 3.8 16.3

CF3CHFCF3 170.0 2.5 ⫺16.4 51.3

CCl2F2 120.9 ⫺21.6 ⫺29.8 79.8

CH2FCF3 102.0 ⫺14.9 ⫺26.1 81.9

1.1

3.5

5.5

5.7

90.9

86.5

81.8

75.3

0.04 92.1 1.48

1.46

Abbreviations: CFC, chlorofluorocarbon; HFC, hydrofluroalkane.

1.39

1.31

1.21

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preservatives (e.g., cetylpyridinium HC1 in Duo-Medihaler®) (14) may influence physicochemical properties of the drug, thus having an impact on the extent of absorption from the airways. In recent studies, a variety of compounds have been shown to be absorption and bioavailability modulators of pulmonary-delivered drugs (15–17). This approach is particularly useful for drugs that are inefficiently absorbed from the lungs due to reasons such as molecular size, charge, and liability toward local enzymes. One such study (16) revealed that N-lauryl-␤-D-maltopyranoside and bacitracin significantly enhanced the pulmonary absorption of insulin. These excipients increased the pharmacologic bioavailability of insulin from approximately 12% to 80% (16). B.

Dry-Powder Inhalers

Dry-powder systems utilize drug blends in a suitable carrier for delivery to the lungs. Lactose is the single, most commonly used drug carrier for DPIs (e.g., Ventolin Rotacaps) (14). Jet milling and spray drying are the two most common techniques to produce fine particles suitable for use in DPIs. Using ␤-galactosidase as an example, Broadhead et al. (18) showed that spray drying could significantly reduce the activity of this enzyme. However, the inclusion of additives such as sucrose and trehalose produced respirable particles (geometric mean diameter ⫽ 4.3 ␮m) without compromising its enzymatic activity. Spray-dried particles presented a narrower size distribution in formulations than a mechanically micronized batch, but a high surface roughness of the former resulted in low RF. Recrystallized lactose has been proposed as a superior alternative to spray-dried lactose for generating respirable particles (19). C.

Nebulizers

Nebulized systems involve conversion of a liquid formulation into vapor with the aid of energy (e.g., ultrasonication), which is then inhaled by the patient. The formulation may contain cosolvents and other pharmaceutical aids to ensure satisfactory physical and chemical stability of the drug. For example, several marketed products for nebulization contain preservatives (e.g., benzalkonium chloride in Proventil, Ventolin, and Alupent), stabilizers (e.g., disodium edetate in Airet, Alupent, and Isoetharine), sweetener (e.g., sodium saccharin in Isuprel HCI), and salts and acids to adjust tonicity and pH (14). Effects of these excipients on aerosolization efficiency, lung deposition, and bioavailability of nebulized drugs needs characterization. The excipient could have an impact on drug absorption if it causes configurational changes in the drug molecule or if it affects the integrity of pulmonary epithelia. For example, a decrease in pH of insulin solution from 7.0 to 3.0 improves pulmonary bioavailability of this drug from 13.1% to 41.6%. This was attributed to the existence of insulin in the monomeric form at the lower pH 3.0 (15). It is emphasized, however, that a large fraction of the nebulized drug, often greater than or equal to 80% of the nominal dose, is entrained in the apparatus. Of the fraction that is inhaled, up to 80% or greater is deposited in the lungs (20). Hence, the fraction of nebulized dose absorbed from nonpulmonary routes [e.g., the gastrointestinal tract (GIT)] is

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very small. This is quite different from MDIs and DPIs, where a significant fraction of the inhaled drug is swallowed and may contribute toward the overall absorption and bioavailability of the drug.

III.

Factors Affecting Bioavailability of Inhaled Drugs

Tissues within the body use oxygen to drive the various reactions and functions necessary to maintain homeostasis. The respiratory system supplies oxygen to the body and removes carbon dioxide from venous blood. It also removes atmospheric contaminants and particulate matter from the inspired air. If these pollutants are not removed, they may deposit deep in the peripheral compartments of the lung, thus accumulating or getting transported to more remote organs and tissues within the body to cause injury and disease. Many atmospheric contaminants penetrate deep into the respiratory system to produce ill effects (21). Gaseous contaminants such as carbon monoxide, insecticide mists, benzene, carbon tetrachloride, and sulfur dioxide cause respiratory inflammation and blood poisoning if inhaled in large quantities (22). Particulate allergens such as pollen, airborne microorganisms such as bacteria and viruses, industrial fumes, and environmental dusts of sorts can cause a variety of respiratory diseases either in the nasopharyngeal tree or in the deeper compartments of the lung (23). Because the lung serves as a reservoir for atmospheric contaminants, it is an excellent route for administering drugs into the respiratory system for the treatment of various localized diseases such as AAD and ARDS. Because of the effect of endocytic and phagocytic transport mechanisms for particles reaching the alveoli, the lung also provides a noninvasive route for systemic administration of drugs that are otherwise not well absorbed via conventional routes of administration. Drugs administered via the airways, therefore, can have local as well as systemic effects. However, the extent of drug absorption from the lung depends on deposition and distribution of the aerosolized drug in the airways. Pharmacologic effects arising from lung deposition of these drugs, like those after oral or parenteral administration, are usually dose dependent. Therefore, regardless of the site of action, guidelines for lung drug delivery must address and answer three key issues from bioavailability standpoint: 1. 2.

3.

What in vitro tests might one use to measure the extent of lung deposition of the drug? What physical models are appropriate for characterizing deposition and clearance of inhaled drug particles in the airways in both the presence and the absence of disease? How well do in vitro functional performance data correlate with in vivo PK and pharmacodynamic (PD) data to provide insights on the effect of age, gender, and disease on biophysical performance of aerosolized drugs?

A brief review of these issues is given in the following section of this chapter.

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In Vitro Variables (or Factors)

Particle Size

A large body of data on lung deposition of particles are available from industrial and occupational hygiene. Data published by the Task Group of the International Radiological Protection Commission confirmed that the extent of particle deposition at different sites within the respiratory system was dependent on aerodynamic particle size (24). The study identified three areas of particle deposition within the respiratory system. Particles of 8 ␮m or greater deposited in the nasal regions pf the airway axis, whereas a sizable fraction of particles less than 3 ␮m deposited in the pulmonary zones of the lung. Particles between 3 and 8 ␮m largely deposited in the nasal and tracheobronchial zone, namely the conducting and transitional zones of the lung. In vitro measurement of particle size is, therefore, crucial in characterizing aerosol formulations with a subsequent goal of enabling the prediction of lung deposition (25). Moisture

Several distinct differences between environmental and pharmaceutical aerosols must be recognized. In contrast with environmental pollutants or dust particles of interest to occupational hygienists, aerosolized drug particles undergo significant morphological changes as they move down the airways. Pharmaceutical aerosols are biphasic systems that usually contain a solid or liquid phase and a single, continuous, gaseous phase, usually air. For most MDIs, the continuous phase is a liquefied halocarbon or hydrocarbon propellant blend. After its exit from the spray jet of an MDI, the continuous phase rapidly flashes off, leaving the dispersed phase exposed to conditions within the pulmonary system. The loss of solvent initially causes a lowering in particle diameter. This size reduction is rapidly reversed for hygroscopic particles as they begin to pick up moisture from the respiratory tract (26–29). Being mostly diffusion controlled, the loss of propellant and simultaneous sorption of moisture onto these particles occur in a few milliseconds. This increase in moisture-uptake rate facilitates upper-airway deposition, thus increasing drug removal by mucociliary clearance with a subsequent decrease in pulmonary bioavailability of the drug. Solution aerosols, usually because they contain nonvolatile cosolvents, impact largely in the upper airways, thus sequestering a large fraction of the drug in mucus, which is then lost by swallowing into the stomach. Ideally, to obtain a realistic estimate of bioavailability, the fraction of drug absorbed in the GIT must be blocked. For some drugs, this may be done, for example, with concurrent oral administration of activated charcoal (30). Particle Velocity

Usually in vitro product qualification tests are conducted under controlled room-temperature conditions. Storage temperature and humidity conditions often used to monitor product stability differ from conditions in the

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respiratory tract. The implications of this difference on inhaled drug particles must be kept in mind. Velocity of aerosolized drug particles often decreases dramatically as a result of inertial forces and drag by the humid air in the lung. This facilitates upper-airway deposition and subsequent mucociliary clearance of the aerosolized drug. Particle velocity in aerosols usually is proportional to droplet or particle size. Thus, increasing particle velocity leads to impaction losses in aerosol devices such as the mouth adapter and also losses by impaction in the nasopharynx. The resulting decrease in lower-airway deposition also decreases pulmonary absorption of the aerosolized drug. Respirable Fraction

The term RF represents a dose fraction of aerosolized drug particles small enough in diameter (=5 ␮m) to escape the filtration machinery of the airways. Particles in this size range are capable of depositing in the lungs. Impaction and light-scattering methods have been developed for use in assessing RF and, therefore, the potential for pulmonary deposition of inhalation products (2). There is frequently a good relationship between RF and bioavailability of inhaled drugs in humans (31). B.

In Vivo Variables (or Factors)

A basic understanding of biopharmaceutics and PKs is essential to the design and evaluation of formulations with desirable efficacy and control. For inhalation drugs, rate-limiting mechanisms, such as dissolution kinetics, membrane permeability, diffusion, and clearance by various cells and neurotransmitters in the lung, can have a profound effect on the extent of uptake and distribution of compounds administered to the lung parenchyma. Drug absorption and bioavailability assessment are, therefore, often confounded, particularly in cases where these drugs are administered for their systemic rather than local effects in the lungs. Likewise, locally acting drugs in the airways may render their optimal PD effects via the direct effect they might have on lability, i.e., compliance of the airways such as occur during an asthma attack or a seasonal allergic rhinitis. Plasma as well as urine drug analysis can be used, however, to estimate the PKs of pulmonary-delivered drugs (32). Variables that can have a significant impact on in vivo performance of aerosolized drugs, examples upper-airway deposition and swallowing, are therefore crucial because these can have significant regulatory and clinical consequences during Phase I and II clinical studies. Some of these issues are summarized in the following section. Deposition

Absorption rates of lung-delivered compounds is generally lower after intratracheal instillation than after aerosol delivery. Figure 1 displays the correlation between absorption rate constant following instillation and aerosol delivery of 12 model compounds to rabbits (33). A good correlation (r2 ⫽ 0.980) with a slope of 1.62 suggests that, on average, the absorption rate constant of these compounds is about 60% higher after aerosol delivery than

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Correlation between absorption rate constant following intratracheal and inhalation delivery of 12 compounds in mouse (❍) and rabbits (●).

Figure 1

that after instillation as a solution. Similar results were obtained in a mouse model, where aerosol delivery generally resulted in 80% higher drug absorption rate than after instillation as a solution. This may be due to a higher fraction of drug depositing in the peripheral zone of the lung following aerosolization compared to instillation. The large surface area coupled with a greatly reduced thickness of the absorption barrier at the alveolar region facilitates drug diffusion into the systemic circulation at a relatively faster rate than drug absorption from the tracheobronchial region. Particle deposition in the lung is sensitive to several model-dependent factors, and it is therefore important to identify these factors when extrapolating data from one animal species to another. There are also significant differences in branching of the lung, when a comparison between the animal species and the human is made. These differences will affect airflow and particle impaction within the respiratory system. The physical branching obviously cannot be changed, but one should keep in mind its potential effects, when comparing data from different species. Lung Function Parameters

Forced vital capacity, forced expiratory volume, inspiratory flow rate, and tidal volume are some of the key lung function parameters used to measure the extent of lung health or lung disease. These lung function parameters are also crucial factors that can affect particle deposition characteristics in in vivo studies. Lung function parameters are sensitive to a variety of disease states. When airway resistance is compromised, such as in asthma and ARDS, particle deposition characteristics are altered (34,35). The net result is a shift in the

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absorption profile of the aerosolized drug. Normal subjects with indirect or direct exposure to dust from mines, or tobacco smoke can also have confounding effects on lung health and extent of lung uptake. The Endocrinologic and Metabolic Drugs Advisory Committee of the FDA (panel 47779, EMDAC) reviewed and adjudged inhaled insulin absorption from lung to be affected by smoking, and that this can have a measurable impact on therapeutics (36). Therefore, it is important to consider the status of the patient when developing an aerosol for clinical use. Lung deposition studies of an aerosolized drug should be carefully designed so that effects of model-dependent variables, i.e., lung physiology and respiration rate, are kept at a minimum. Lung Physiology, Cells, and Disease

The lung contains three basic components—air, blood, and tissue. The architectural arrangement of these three basic components provides optimal conditions for gas exchange and efficient resistance to the movement of air and blood. This subject is discussed in depth in most lung physiology texts; a good review may be obtained from Kilburn (37). The tracheobronchial tree also provides for efficient removal of particulate matter in inspired air by a highly specialized transport mechanism called mucociliary clearance. Details of the anatomical arrangement within the lung and cell line differentiation of lung tissue as a basis for drug delivery may be found elsewhere (4). It is noteworthy that after entry into the upper airways of the lung, air passes through a variety of regions (generations or branches) before gas exchange into blood occurs. These regions consist of 23 dichotomous branchings of the airways beginning from the trachea, with each subsequent pair of branches having a smaller diameter than the parent. The geometry of this architecture is fairly constant in humans but can be altered significantly by smoking, various diseases, and environmental pollutants. For example, cigarette smoking has been suggested to increase alveolar permeability (38), but excessive and prolonged smoking can induce physiologic changes in the lung that can affect lung function as well as bioavailability of inhaled drugs. C.

Biophysical Issues

It is generally known that inhaled particles do not enter alveolar air spaces. Still, after escaping the filtration mechanisms of the lung, cellular debris, degraded myelinate surfactant materials, microorganisms, and fine particulate drug matter may enter the alveolar spaces, depending on their size and deposition characteristics. Once such foreign particles enter the alveoli, they become targets of and are engulfed by alveolar macrophages (4). These cells, usually up to four per alveolus, migrate around the alveoli beneath the surfactant layer and ultimately enter the acinus via the terminal bronchiolar lumen. Inhaled drug and foreign particles of size lesser than or equal to 3 ␮m may be absorbed from the lung primarily by alveolar macrophages (39). Most drug particles also may transport into the body fluids by passive diffusion across epithelial cells (40). For drug particles to get to their sites of absorption within the lung, critical factors that could affect the dynamics of the aerosolized particle must be identified. A brief review of these factors are given below.

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Airway Subdivision

There is significant asymmetry in branching within the lungs of animal species when compared to human lung. This systematic subdivision of the lung furthermore complicates the dynamics of inhaled particles due primarily to tortuosity of the pathway traversed by the particle and the potential for impaction high up in the lung network. Thus, for efficient delivery of a drug to the respiratory zones of the lung, formulation factors should attempt to minimize the impact of filtration mechanisms of the lung. Where animal models are used to estimate formulation performance in vivo, efforts must be made to ensure that lung drug deposition characteristics in such animal models are consistent and can be qualitatively extrapolated to humans. Airway Geometry

The extent to which a drug formulator can effectively use any physical or biological system to model lung deposition and clearance of drug particles depends largely on an appreciation of the complex architecture of the human lung. Several morphologic models now exist that clearly describe the geometry of the human lung. The most widely used model for describing the morphologic structures within the lung was given initially by Weibel (41). A diagrammatic representation of the Weibel model is shown in Figure 2 with data from Bouhuys (42). Using a scheme based on that proposed by Weibel, the branched airways of the lung have been characterized to fall into three distinct compartments. These are the following. Conducting Zone

This region consists of the first 16 generations. The airways of the conducting zone are described as rigid tubes that consist primarily of cartilage in the walls

Figure 2

Ref. 41.

A schematic representation of airway branching in the human lung. Source: From

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and that symmetrically divide or bifurcate beginning with the trachea and ending with the terminal bronchioles. These are incapable of providing gas exchange with venous blood. The conducting zone is an anatomically dead space, whose purpose is to provide little resistance to the movement of air to distal compartments of the lung. It is noteworthy that drug absorption in this zone is low. Transitional Zone

This consists of generations 17 through 19. The walls of these airways, the respiratory bronchioles, consist of few alveoli lined with alveolar–capillary tissue for limited gas exchange. Little drug absorption occurs in this region, but the extent of absorption depends on intrinsic properties of the drug, i.e., solubility, permeability, and diffusivity across the absorption barrier. Respiratory Zone

Generations 20 through 22 contain vascularized structures that extend to the most distal parts of the lung, where gas exchange occurs across single cell lines. These cells generally form a continuum, namely alveoli sacs, each of which contains alveolar ducts, venules, and arterioles. The respiratory zone is highly vascularized and consists of lobules that have three to five terminal bronchioles. Each lobule stems from one segmental bronchiole from the transitional zone. The terminal bronchioles supply air to the smallest structural units of the lung, each called an acinus. The acini are the structures within which gas exchange occurs, primarily by molecular diffusion (4). Significant drug absorption occurs in this region of the lung, and it especially should be the primary target for drugs intended for systemic delivery such as peptides.

IV.

Pharmacokinetics of Inhaled Drugs

PKs is an interdisciplinary science, which encompasses quantitative studies of the kinetics of absorption, distribution, metabolism, and excretion of therapeutic agents in the body. A basic understanding of PKs is essential to the design and evaluation of inhalation formulations, especially in situations where bioavailability and PDs are clinical end points for product selection. For lungdelivered drugs, once a given drug enters the body, its in vivo disposition can be characterized by the PK relationships historically used for intravenous (IV) and orally administered drugs based on blood or urine data. However, there are certain aspects of drug absorption unique to lung delivery. Either rate or extent of drug absorption from the lungs may be determined by model-dependent or model-independent methods. Compartmental models, linear system analysis, and observational methods are currently being used for evaluation of drug absorption (43). Compartmental models assume that absorption occurs by a specific rate process. On the basis of the superposition principle, linear system analysis generates amount or percentage of drug absorbed as a function of time. Thus, absorption profiles for different formulations can be

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directly compared (44). The extent of absorption can also be estimated based on plateau values of the absorption profiles. However, it requires IV data to characterize drug disposition. The observational method is the most common method for assessing rate and extent of absorption. The rates of absorption are compared using peak drug concentration (Cmax) and the time for peak drug concentration to occur (tmax) The extent of absorption is determined by area under the plasma concentration–time curve from time zero to time infinity (AUC8) The following section discusses some methods most useful to characterize systemic absorption of inhaled drugs based on blood concentration data. A.

Compartmental Models

One-Compartment Open Model with First-Order Absorption

Scheme 1 represents a single-compartment model for a drug absorbed and eliminated from the body by apparent first-order processes after a single dose

Scheme 1

administration to the lung. Following lung deposition, the drugs are typically absorbed into the bloodstream and/or eliminated (or irreversibly bound) by processes such as mucociliary clearance and/or chemical and enzymatic degradations, etc. In certain cases, these processes can be approximated by first-order kinetics. For example, the kinetics of mucocilliary activity was found to follow monoexponential model (45). According to this scheme, the absorption of the inhaled drugs is primarily governed by two first-order processes, where Xa is the amount of drug at the absorption site, Xb is the amount of drug in the body, Xe is the amount of drug eliminated, and Xc the amount of drug removed by the nonabsorptive clearance mechanisms. The parameters Ki, K, and Kc, are the first-order absorption, elimination, and local clearance rate constants, respectively. The equation describing Xb, as a function of time, is as follows: X b (t ) =

K i X a (0) − Kt (e − e − K at ) (K a − K )

(1)

In most cases, only a fraction (F) of the total administered dose (D) is absorbed, which can be represented as Xa(0). Therefore, Eq. (1) can be expressed in concentration terms as C (t ) =

K i FD (e − Kt − e − K at ) (K a − K )V

(2)

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where V is the apparent volume of distribution. The parameters of the model can be estimated by fitting data to Eq. (2) using programs based on nonlinear least-squares regression. It should be noted that the apparent absorption rate constant (Ka) is the summation of true absorption (Ki) and local clearance (Kc) rate constants. If the drug is orally bioavailable, a fraction of the dose inhaled may be swallowed and absorbed through the GIT. Scheme 2 shows this possible scenario, where Ks represents the first-order rate constant of GI

Scheme 2

absorption. The parameters F1 and F2 are fractions of the total dose absorbed from lung and GIT, respectively. Thus, the profile for drug concentration in the body is the result of absorption from lung and GIT:

 Ki( Ks − Ka )+ KsKc -Kat KsKc - K st   (Ka − K )(Ka − K ) e + (Ka − Ks)(K − Ks) e +  FD   C(t ) =  V  Ki(Ks − K )+ KsKc -Kt   (Ka − K )(Ks − K ) e  

(3)

where F ⫽ F1 + F2. In order to prevent absorption of inhaled drug from the GIT, concurrent oral administration of charcoal for adsorption of swallowed drug has been suggested (30). However, the usefulness of this approach is drug dependent and often unsuccessful for drugs with high oral bioavailability. It should be noted that Ki in the model includes absorption rate constants from central and peripheral lungs. Although there might be differences between the two regions in the rate of absorption, no attempt has been made to model drug absorption from the two parallel routes in the lung. Two-Compartment Open Model with First-Order Absorption and Central Compartment Elimination

The two-compartment model, with input and elimination occurring from the central compartment, is represented by Scheme 3.

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Scheme 3

The concentration in the central compartment as a function of time can be obtained by solving a system of differential equations of the model:  ( K − α ) e − αt ( K 21 − β)e −βt   21 + + Ki FD  ( K a − α )(β − α ) ( K a − β)(α − β)  C (t ) =   −K t Vc   ( K 21 − K a )e a    (α − K a )(β − K a ) 

(4)

where Vc is the volume of the central compartment, K12 and K21 are first-order distribution rate constants, and ␣ and ␤ are hybrid rate constants. If data are fitted to Eq. (4), the model parameters can be derived from the coefficients and exponents (43). Once again, Ka, in this model, represents the summation rate constants for absorption and nonabsorptive clearance. It should be noted that for some multicompartment drugs with prolonged absorption it is difficult to define distribution rate constants unless IV data are available. B.

Noncompartmental Models

Drug Disposition and Bioavailability

The basic system for noncompartmental analysis is shown in Scheme 4.

Scheme 4

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In this system, there is a central homogenous kinetic space (sampling compartment), where input and elimination occur, and a peripheral heterogeneous kinetic space (44). For a linear system, the elimination rate constant, K10, is first order, such that

dXe = K 10 Xc = ClC dt

(5)

where Xe and Xc are amounts of drug eliminated and in the sampling kinetic space at time t, respectively; and C is the drug concentration in the sampling kinetic space at time t. Thus, the total systemic clearance, Cl, is given by

Cl =

FD AUC

(6)

AUC is normally determined by linear trapezoidal summation or its combination with logarithmic trapezoidal summation. The accuracy of the extrapolated area of total AUC may be influenced by the sampling and assay limitations. In many PK applications, disposition of drugs is described by polyexponentials: n

C (t ) = ∑ Aie − α i t i =1

(7)

where C(t) is the plasma drug concentration, and Ai and ␣i are the coefficient and the exponent of the ith exponential term, respectively. The clearance can be obtained by

Cl =

FD n

∑ Ai / αi i =1

(8)

Therefore, in a bioavailability study with each subject as his or her own control, the fraction (F) of the total dose (D) absorbed or bioavailability can be estimated from the AUC based on Eq. (6) or (8): F=

Cl lung × AUClung × DIV Cl IV × AUCIV × D lung

(9)

The use of AUC from two different dosing periods for the evaluation of extent of absorption assumes that clearance is unchanged during each period so that Cllung and ClIV cancel out in the equation. Rate of Absorption

Parameters for assessing rate of absorption include Cmax and tmax obtained directly from the data without interpolation. However, Cmax depends on both rate and extent of absorption, and it is quite insensitive to changes in rate (41). The tmax variable can only have discrete values depending on the sampling time.

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This parameter often has reliability problems because its frequency is not normally distributed. Other alternative measures of rate of absorption have been proposed, which include partial AUC computed up to tmax (AUCp), and Cmax/AUCp, Cmax/tmax, and Cmax/AUCtmax. The latter three parameters attempt to normalize Cmax with a factor representing the extent of absorption. Additional techniques, i.e., feathered slope (SLf) and feathered AUC (AUCf) derived from peeling, have been used to subtract an estimated elimination component from the concentration curve to isolate drug absorption. Bois et al. (46) tested these methods using simulations and observed some degree of insensitivity to changes in rate of absorption. In general, each rate measure has advantages and limitations that depend on the kinetic properties of the drug and its formulation (46). Mean Residence Time

Mean time parameters, such as mean residence time (MRT), have been widely applied in PKs. They are useful in studying specific models as well as less differentiated, more general system models (44). Many important concepts, definitions, and computations on this subject have been thoroughly discussed by Veng-Pedersen (47). By definition, MRT is the average total time the drug molecule spends in the introduced kinetic space. It depends on the site of input and the site of elimination. When the elimination of the molecule follows first-order kinetics, its MRT can be expressed by (47) ∞

∫ t × C (t )dt = AUMC MRT = 0 ∞ AUC ∫0 C (t )dt

(10)

where AUMC is area under the moment curve. Estimates for MRT can be calculated by fitting C(t) to polyexponential equation followed by integration, or by using trapezoidal rules. Mean Arrival Time

For noninstantaneous input into a kinetic space, such as pulmonary delivery, the MRT estimated from extravascular data includes a contribution of the mean transit time for input, known as mean arrival time (MAT, or mean input time or mean absorption time) (47). The MAT of drug molecules represents the average time taken to arrive in that space (48), and it can be estimated as ∞

∫ t × fin(t )dt = AUMC MRT = 0 ∞ AUC ∫0 fin(t )dt

(11)

where fin(t) denotes an arbitrary rate of input into the kinetic space. For lung delivery, the MAT can be determined according to the equation MATlung = MRTlung − MRTIV

(12)

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The term MAT lung thus obtained represents the mean transit time involved in apparent absorption process in the lung. When the formulation contains solid drug (e.g., dry powder or suspension), the MAT includes in vivo dissolution as well as absorption. If data for the same drug inhaled in solution state are available, the mean in vivo dissolution time (MDT) can be estimated as

MDTsolid = MATsolid -MATsol =MRTsolid -MRTsol C.

(13)

Linear System Analysis

Linear system analysis is a model-independent method useful in the evaluation of drug absorption processes, especially for situations where the absorption kinetics does not follow functional forms. Based on the superposition principle in a linear time-invariant system, a response C(t), to an arbitrary input f(t), of the system can be obtained using the following convolution integral (47,49): ∞

C (t ) = f (t )* C δ(t ) = ∫ C δ(t − τ ) f (τ )dτ

(14)

0

where C␦(t) is the unit impulse response characteristic of the system. For most PK applications, C(t) and f(t) represent the plasma drug concentration and the rate at which drug enters the system, respectively. The variable C␦(t) is the plasma concentration resulting from the instantaneous input of a unit amount of drug into the system. Therefore, in vivo input rate, involving dissolution and/or apparent absorption, can be estimated by deconvolution, which is the inverse operation of the convolution. Table 3 defines some possible scenarios for the application of linear system analysis to inhaled drugs. General Solution

Eq. (14) can be solved using Laplace transform based on convolution theorem,  c (s)  f (t ) = L−1 { f (s )} = L−1    c δ( s ) 

(15)

where L⫺1 denotes inverse Laplace transform operator. Because the disposition of most drugs can be described by polyexponentials n

Cδ (t ) ≅ ∑ Aie − α i t i =1

in vivo input function f(t) can be obtained using Eq. (15). For example, in the case of single exponential disposition (n ⫽ 1), C␦(t) ⫽ A1e⫺␣1t and hence, f(t) is given by

f (t ) =

[C⬘(t ) + α1C (t )] A1

(16)

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Illustration of Some System Definitions in Linear System Analysis

Unit impulse response C␦(t)

Input (t) Inhaled solutiona Inhaled formulation Inhaled formulation

IV bolus IV bolus Inhaled solutiona

Input function f(t) Absorption from the lung Drug release/absorption in the lung Drug release in the lung

a

Assumptions: (i) Similar lung deposition pattern between inhaled solution and formulation, and (ii) no precipitation of drug occurs following deposition of the solution in the lung. Abbreviation: IV, intravenous.

The amount of drug from time 0 to t, Xa(t), is then obtained by integration:

t

Xa(t ) = ∫

0

t   C (t ) + α1 ∫ C (t )dt    0 f (t )dt =  A1

(17)

t

where ∫0 C (t)dt can be obtained by integration from time 0 to time t. In actual applications, computer softwares, e.g., WinNonlin, PCDCON, PDx-IVIVC, PKQuest, etc., are available for deconvolution calculations. Model-Dependent Solution

The Wagner–Nelson equation is derived based on a one-compartment model and following mass balance:

Xa = Xt + Xe

(18)

where Xa, Xt, and Xe are amounts of drug absorbed in the body, and eliminated at time t, respectively. By derivation, the amount of drug absorbed over time T is given by T

( Xa ) = VCT + kV ∫ C (t )dt

(19)

0

This can be expressed in terms of fraction of the dose absorbed as T

Fa (T ) = FD =

( Xa )T = ( Xa )∞

C + k ∫ C (t )dt 0

∞

(20)

k ∫ C (t )dt 0

where Fa(T) or FD is the fraction of bioavailable drug absorbed at time T. It should be noted that Eq. (19) is identical to Eq. (17). Therefore, the Wagner–Nelson method represents a special case of deconvolution. When IV data are not available, the apparent in vivo fractional absorption profile can be obtained by using the terminal-phase elimination rate constant, k, and partial

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areas under the plasma concentration curve in Eq. (20). However, it should be noted that (1) k value should be derived from the true elimination phase, which may be difficult with drugs with prolonged absorption and long half-life and (2) only apparent absorption is estimated using this method. In 1983, Wagner published an exact Loo–Riegelman method for absorption analysis of multicompartment drugs (43). This is a general equation for absorption analysis of one- to three-compartment compounds. It requires IV data for the calculation of absorption profiles. For biexponential disposition, mass balance leads to: (Xa)T ⫽ Xc + Xp + Xe, where Xc, and XP are amounts of drug in the central and peripheral compartments at time t, respectively. By derivation (50). T

T

( Xa )r = CT + k12 e − k 21T ∫ Cte − k 21t dt + k10 ∫ Ct dt Vc 0 0

(21)

On the basis of mass balance, (X␣)T ⫽ Xc + Xp1 + Xp2 + Xe, similar equation can be derived for triexponential disposition: T

T

T

( Xa )r = CT + k12 e − k 21T ∫ Cte − k 21t dt + k13e − k 31T ∫ Cte − k 31t dt + k10 ∫ Ct dt Vc 0 0 0

(22)

The Loo–Riegelman method is also a special case of deconvolution, where in vivo disposition is described by two or three exponentials. Key factors influencing absorption of inhaled drugs have already been discussed. These factors should be taken into consideration when one uses a particular method for evaluating absorption. For example, lung deposition and clearance, which are dependent upon the quality of the formulation, particle size, and methods of delivery, could have potential effects on estimated bioavailability of drugs. Absorption rates from peripheral and central lung are often significantly different for certain drugs. In this situation, a compartmental model with two parallel first-order input sites and concurrent mucociliary elimination may provide a better fit to the data (43) than the one shown in Scheme 1. Difference in lung clearance kinetics between of the ultrafine and fine particles can also complicate selection of an appropriate model depending on the size distribution in the formulation or disease state of the patients (51). In linear system analysis, an inhaled solution, “may not” be acceptable as a unit impulse reference if solution and an inhaled aerosol suspension with significantly different deposition patterns are compared. Similar complications arise in comparing the kinetics of intratracheally instilled solution versus an MDI because of different deposition patterns and issues, such as drug precipitation due to low aqueous solubility and/or small volumes of liquid in the lung. In addition, application of mass-balance model-dependent techniques (e.g., Wagner–Nelson and Loo–Riegelman method) and model-independent mathematical deconvolution techniques are greatly limited by dissolution and absorption characteristics of drugs from different formulations. The in vivo input profiles are more readily defined where dissolution and/or absorption are slow. For drugs of relatively high solubility/ permeability with rapid absorption, frequent sampling in the short absorption phase is practically not

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feasible. Therefore, the use of observational, compartmental, or mean time parameters may be a better approach for evaluation of absorption. Finally, the widely used Wagner–Nelson method only provides rate profiles of the bioavailable portion of the dose. Problems can arise when comparing formulations that have different systemic bioavailability. The method also becomes invalid for evaluation of the kinetics of racemic compounds exhibiting stereoselective absorption from the lung, when the data are collected using a nonchiral assay. D.

Case Studies

Lipophilic Drugs

Inhibition of 5-lipoxygenase has many potential therapeutic benefits for conditions in which leukotriene synthesis is elevated, such as asthma. Although 5-lipoxygenase inhibitors are orally active, inhalation delivery to the lungs may offer significant reduction in the clinically effective dose, while maintaining sufficiently high local concentration of these compounds. Abbott-79175 (see structure below) is a potent 5-lipoxygenase inhibitor investigated for the treatment of AADs. It is chemically known as R(+) N-[3-[5-(4-fluorophenoxy)2-furanyl]-l-methyl-2-propynyl]-N-hydroxyurea (structure). It is a lipophilic compound with high membrane permeability but very low aqueous solubility (approximately 17 ␮g/mL). Aerosol formulations of Abbott-79175 were developed and evaluated for pulmonary absorption and bioavailability (52).

Chemical structure of Abbott-79175 (molecular formula: C15H13FN2O4; molecular weight: 304.28)

A crossover study was carried out in nine tracheostomized beagle dogs. Treatment A involved IV administration of 0.5 mg/kg drug. Treatment B involved aerosol delivery of drug, 0.5 mg/kg, formulated as an MDI using tetrafluoroethane (HFC-134a) as propellant. Five sprays were administered to the trachea of each dog. Drug levels in plasma were measured by a specific high pressure liquid chromatography (HPLC) method. Mean plasma profiles for IV delivery and the inhalation treatment are displayed in Figure 3. Key model-independent bioavailability and PK parameters are summarized in Table 4. Blood concentration profiles following IV administration were used as unit impulse responses C␦(t) in linear system analysis. Drug blood data from the aerosol formulation, C(t), were fitted to a smoothing cubic spline function and then deconvoluted with C␦(t), using PCDCON to obtain apparent in vivo drug absorption profiles (see Fig. 4). The fractional absorption, as estimated from the plateau values of the profile, closely matched the absolute bioavailability based on AUC of the formulation—i.e., 68% versus 63% (52).

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Mean blood concentration of Abbott-79175 following IV () and inhalation delivery (o) to the dogs (n ⫽ 9).

Figure 3

Hydrophilic Drugs

Sandoz 64-412 is a water-soluble compound (>5 mg/mL) that offers inhibitory effects on allergen-induced airway hyperreactivity such as asthma. In one study, administration of low (5.5 mg/kg) and high (110 mg/kg) oral doses of Sandoz 64-412 to rats demonstrated slow drug absorption such that bioavailability approximated 10% after six and 11 hours, respectively. Inhalation delivery of drug, at a dose of 5 mg/kg, yielded rapid absorption, resulting in bioavailability of 10%, 25%, and 35% in about 1.2, 2.4, and six hours, respectively (Fig. 5) (52). The relatively low oral bioavailability of Sandoz 64-412 suggests that intrinsic physicochemical properties might play a key role in absorption and disposition of this drug. Table 4 Absolute Bioavailability and Pharmacokinetic Parameters of Abbott-79175 Following Inhalation Delivery to Dogs (n ⫽ 9) Parameter AUC⬁(hr ⫻ ␮g/mL) Cmax (␮g/mL) Tmax (hr) F MRT (hr) MAT (hr) K (hr–1) t1/2 (hr) Cl (L/hr ⫻ kg) V (L/kg)

Intravenous 13.04 (⫾1.82) – – 1.00 (⫾0.00) 9.35 (⫾1.80) – 0.120 (⫾0.017) 5.89 (⫾0.85) 0.04 (⫾0.004) 0.33 (⫾0.03)

Inhalation 8.81 (⫾3.01) 0.55 (⫾0.12) 6.78 (⫾2.91) 0.68 (⫾0.23) 13.13 (⫾1.19) 3.60 (⫾1.69) – – – –

Abbreviations: AUC⬁, area under the plasma concentration–time curve from time zero to time infinity; Cmax, peak drug concentration; Tmax, time for peak drug concentration; F, fraction; MRT, mean residence time; MAT, mean absorption time; K, terminal-phase elimination rate constant; Cl, total systemic clearance; V, apparent volume of distribution.

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Mean in vivo absorption profile of Abbott-79175 following inhalation delivery to dogs (n ⫽ 9).

Figure 4

V.

Clinical and Regulatory Issues

From the discussions thus far on bioavailability of inhaled drugs, it has become apparent that development and qualification of pharmaceutical aerosols is a multidisciplinary effort. Therefore, for a new or reformulated aerosol product to successfully enter the market, interactions among many different functional areas within the basic, health, and allied sciences must occur. A typical aerosol development team would comprise chemists, biochemists, pharmacologists, biophysicists, pharmacokineticists, clinicians, physiologists, toxicologists, pathologists, engineers, regulatory specialists, and formulation scientists. Because there are many phases of the drug development effort, not all of these experts will be involved in the development activity at any one time. However,

Figure 5 Estimates of bioavailability following low-dose (5.5 mg/kg) and high-dose (110 mg/kg) oral drug administration, and as a nebulized solution (5 mg/kg) to the lungs of rats. Source: From Ref. 53.

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at a bare minimum, the principal formulation scientist and his or her clinical or medical counterpart must be involved in every phase of the process to ascertain the impact of any issue or observation that might have certain safety or clinical significance. Of these, the impact of PD response, tolerance buildup to drug dosimetry, and safety are the most crucial. These issues are briefly addressed in the following sections. A.

Developmental Pharmaceutics

Standards for pharmaceutical aerosols have been in a state of flux since about the middle of the 1990s. This is partly due to the emergence of DPIs and piezoelectric devices in clinical trials and even in commerce in Western countries. Nonspecific bioequivalence criteria, largely Cmax and AUC, as a basis for Abbreviated New Drug Application (ANDA) approvals of nondesigner products, and therefore the appearance of a large number of generics into the marketplace, may be another reason for continuing changes in inhalation aerosol drug standards. Beyond issues relating to equivalent criteria for bulk drug identity, purity, solubility, and stability, a comprehensive data package for new and reformulated aerosol products must be used as a basis for product approval. As discussed earlier, a number of physical, chemical, biophysical, and physiologic parameters can have a significant impact on the bioavailability of inhaled drugs. These data should be provided for both new and reformulated products as well as appropriate reference formulation. The data should include but not be limited to the following. Specifications for permissible levels of product contaminants such as solidand liquid-stage degradation products: These data must be included in the developmental pharmaceutics package for new and reformulated products. The specifications should include established analytical methods for extractables, reaction products, and unknown foreign substances. Specifications for extractables must include data on appropriate placebo formulations as a baseline for active product. In most cases, the placebo is water for nebulized products, and blank propellant for MDI products. Specifications for reference standards based on purity, potency, and stability: The data should include both in-process acceptance criteria and action limits for the drug. A process description package detailing critical steps and process ranges for the product: The information should include rework procedures, where excursions beyond in-process limits are allowable. In such cases, clinical lot validation data should include recommended overall batch or lot rejection standards (54). It is imperative that medical officers, regulatory staff, and project managers understand the relevance of these action limits. First, the specifications often assist the project team to distinguish between formulations aside from bioavailability results. Second, product lot approval decisions often involve a collaborative effort by all members of the development team because characterization of inhalation aerosol is often a multidisciplinary effort. Functional performance of the finished product: Functional performance data on inhalation aerosol products is directly linked to lung deposition of the

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drug. Therefore, an attempt must be made to correlate RF, mass median particle diameter, through life dose uniformity, and content uniformity results with at least in vivo performance in preclinical studies. In this sense, it is crucial that functional performance of formulations be dovetailed with both animal and clinical studies. B.

PK and PD Considerations

PK parameters are essential to the assessment of in vivo performance of drug formulations. For example, rapid absorption and high bioavailability, as indicated by high AUC, high Cmax, and short tmax, may be important for drugs intended for systemic effects only. On the other hand, if the receptor is known to be located in the airway, prolonged exposure of the drug may lead to a reduction of the clinical dose as well as a potential decrease in side effects. This can be achieved by formulation approaches, e.g., use of insoluble salts, prodrugs, or slow-release products. In this case, use of mean time parameters would be more helpful for evaluating formulation performance in the lung. In general, few inhaled drugs demonstrate a linear relationship between dosimetry and bioavailability parameters (e.g., C max and AUC). There are limited literature reports on drugs that show any proportionality of pharmacokinetic response with efficacy. Furthermore, as discussed earlier, clinical performance of pharmaceutical aerosols is dependent on a number of model-dependent and model-independent factors. For this reason, it is crucial that characterization and selection of a new or reformulated inhalation aerosol product be made with PD results, i.e., the measured clinical end point. But in the rare event, where a well-defined PK–PD relationship should exist for a drug via the pulmonary route in preclinical studies, a good argument could be made to use PK results in clinical development studies. In this case, the timed course of drug, blood, or plasma profile of metabolites, and PK parameters for the parent compound may be used to validate in vivo performance of a drug. Although such data may not be adequate for approval of a new chemical entity, the information should be sufficient for approval of reformulated products—for example, an HFC formulation of a CFC-based MDI. It is important to point out that for a drug entity under evaluation, a dose–response relationship is recommended prior to assessment of the PK–PD relationships. This is chiefly due to the fact that adverse responses to drug dosimetry in themselves could be confounding to the PD response of interest. If this is the case, then equilibration characteristics between plasma concentration and the pharmacologic end point should be evaluated. There are three basic types of equilibration relationships crucial in this type of clinical investigation: 1.

Direct equilibration without hysteresis: This kind of dose-dependent effect suggests that the measured effect is linearly reversible on the ascending compared to the descending dose curves. A drug exhibiting this kind of relationship would not display tolerance or concentration buildup during range-finding studies.

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3.

Counterclockwise equilibration with hysteresis: This type of dosedependent effect suggests involvement of a delayed equilibrium or active metabolites or receptor sensitization or a combination of some or all of these phenomena. Clockwise equilibration with hysteresis: This type of dose-dependent effect may involve sequestering of drug in tissues such as is frequently the case with tissue-extracted drugs.

In view of the possible significant effect of these confounding issues on clinical effectiveness of drugs, it is important that clinical studies be carried out with concurrent PK and PD measurements. For inhaled products, the clinical development effort must include a dose-dependent PD study, especially if any one or more of the following situations could occur: 1. 2.

3. 4.

C.

Therapeutic or toxic effects of the drug, presented as an inhalation aerosol, are indirectly related to plasma concentration. Irreversible toxicity occurs at some point in the dose–response curve. Usually, it is local toxicity of the inhaled product that is of primary concern because dilution of the drug is not possible as it is with parenteral administration. This recommendation does not suggest to waive indices for systemic toxicity, although this should not have a big impact if safety data on IV administration are available. There is evidence of PD tolerance of the drug by other routes of administration, such as IV. There are uncertainties concerning the PK–PD relationship for drug as a pharmaceutical aerosol. Bioequivalence Issues

Bioequivalence is a term used to describe two pharmaceutical equivalents, whose rate and extent of absorption are not statistically different, when administered to patients or subjects at the same molar dose under similar experimental conditions. Bioequivalence may be based on a pharmacological effect, efficacy, or PD end point (55). Because drug concentrations in blood and urine can be readily monitored, PK methods are commonly used for bioequivalence assessments. However, this assumes that equivalent PKs is indicative of equivalent clinical response (56). Unfortunately, two inhalation aerosol formulations are seldom biological equivalents of each other. This is largely because performance of inhaled drugs is directly related to physical, chemical, formulation device, and biophysical parameters associated with the product. Physiologic and lung function parameters of patients also vary widely, rendering deposition estimates that can be significantly different. Further, in some cases the PD results in themselves may be so imprecise that statistical criteria can be met only with an unreasonably large number of subjects. For this reason, bioequivalence of inhaled drugs should include more than PK estimates such as Cmax, tmax, MAT, MRT, and AUC. In fact, wherever possible, realistic estimates of PK as well as pharmacologic end points based on dose should be used as a basis for establishing bioequivalence of two products.

Bioavailability and Pharmacokinetics of Inhaled Drugs VI.

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Epilogue

Synthetic chemists in conjunction with pharmacologists are persistently exploring various ways to stabilize peptides and proteins in an effort to render them amenable to oral delivery. These fundamental studies have brought little improvement to efforts aimed at resolving the oral bioavailability problem. Furthermore, advances in peptide and protein drug research overall are presenting a number of newer potent and more complex molecules for treating still newer and difficult diseases, e.g., infections and malignancies, cardiovascular, endocrinic, neurologic, and a variety of immunologically compromised diseases such as AIDS. The absorption limitation saga applies to these newer drugs as well. Therefore, to be clinically useful by the oral route, these drugs must be administered in very high cost-prohibitive doses. Drug-delivery technologies— i.e., topical, rectal, vaginal, and the pulmonary routes of administration—may provide a solution to the problem of oral absorption with these compounds. Unfortunately, evolution of these newer routes as well as their acceptance from social and clinical standpoint will take time. Until then, historical and clinical experience with lung-delivered drugs such as antiasthmatics should make this route a rational drug-delivery paradigm for many poorly absorbed drugs. VII.

Conclusions

Pharmaceutical aerosols have been used as important drug carriers to the body for decades. Pulmonary delivery of bronchodilators, mucolytics, anticholinergics, and steroids is widely accepted in clinical practice. As drug carriers, aerosols remain unquestionably the most effective mode for treating patients with asthma and AADs. In general, however, the lung remains underutilized for systemic administration of drugs compared to the oral and injectable delivery routes. This is unfortunate and is largely a result of the human lung being a complex organ with significant perceived safety risks compared to nonconventional routes of administration. Yet the human lung presents itself as a feasible port for systemic delivery of natural products, e.g., nicotine from tobacco smoke, and a number of pharmaceuticals. Because it possesses a relatively lower degree of enzymatic activity, the lung should be effective for delivering labile and refractory compounds. Examples of this class of compounds are peptides and proteins, usually replicas or analogs of naturally occurring neurotransmitters or immunomodulators. When administered by the oral route peptides are enzymatically deactivated by metabolizing enzymes, thus reaching systemic concentrations too low for a therapeutic effect. And, being frequently unstable and sensitive to a variety of conditions in vivo, their clinical usefulness depends on concentration of intact molecules at their sites of action in the body. We have discussed a number of theoretical and practical approaches to optimize performance characterization of inhalation aerosols. Issues that can have a considerable impact on the bioavailability of biologic equivalents of inhaled drugs have also been presented. This area of drug delivery is very complex, as judged by the disciplines and formulation processes presented.

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Therefore techniques, concepts, and methods for evaluating bioavailability of inhaled drugs must have a fair appreciation for in vitro, in vivo, and biophysical parameters for these products. Acknowledgments Thanks to Dr. W. Gillespie for providing the PCDCON program. This chapter has been dedicated to millions of Katrina victims in recent hurricanes and the International Red Cross for their services. References 1. 2. 3. 4.

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8 Therapeutic Uses of Lung Aerosol

JONATHAN A. BERNSTEIN and HETAL AMIN University of Cincinnati College of Medicine, Cincinnati, Ohio, U.S.A.

I.

STEVEN J. SMITH American Medical Association, Chicago, and Chicago Medical School, North Chicago, Illinois, U.S.A.

Introduction

In previous chapters, physicochemical and physiological aspects of lung aerosol–delivery systems have been discussed. This chapter addresses the route of delivery and the various therapeutic aerosol options available for the treatment of asthma. The chief uses of aerosols designed for delivery of medication to the lungs are the immediate and prophylactic treatment of asthma and chronic irreversible obstructive lung diseases (e.g., chronic bronchitis and emphysema). This chapter compares the advantages and disadvantages of different drug–delivery systems employed for the optimal management of asthma. This chapter also compares the major aerosol systems used to deliver medication directly to the lung including metered dose inhalers (MDIs), dry powder inhalers (DPIs), and nebulizers. In addition to asthma, chronic bronchitis, and emphysema, the value of lung aerosol medications for cystic fibrosis (CF), pulmonary infections, and other disorders are addressed. The possible role of lung aerosol–delivery systems in gene therapy is also evaluated. This chapter provides an overview of delivery devices, including nebulizers, pressurized MDIs (pMDIs), and DPIs and will address factors affecting delivery of inhaled corticosteroid (ICS) therapy to the lungs. It will also address new delivery options for 219

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young children with asthma and how devices can be matched to the needs of patients to provide optimal therapeutic benefit of ICS therapy. II.

Systemic Routes of Medication Delivery for Asthma and Other Lower Respiratory Disorders

Patients often experience technical difficulty when asked to demonstrate a specific inhalational device they have been prescribed (1,2). For example, Kofman et al. interviewed 150 children and parents about aerosolized drug treatment after outpatient encounters at a hospital in Buenos Aires (4–5/2001) (3). The interview questionnaire covered parents’ knowledge about disease, utility of treatment, and cleaning techniques for relevant delivery device and children were interviewed when appropriate (3). A hands–on evaluation of inhalation technique revealed that 81% of the patients using a nebulizer did so properly but only 64% and 42% of the patients demonstrated proper use of pMDIs with spacer/holding chamber and DPIs, respectively (3). Common problems associated with using pMDIs include the inability to coordinate inhalation with actuation, inhaling too rapidly, premature cessation of inhalation because the aerosol hits the pharynx, and failing to hold a breath. Problems demonstrated for DPIs include failure to inhale forcefully and exhaling through the device before inhaling while problems that arise with nebulizers include a poor face mask seal and altered breathing patterns caused by infants in respiratory distress (1,2,4). A blowby technique often used in infants is inadequate because it can reduce active drug deposition in the lung by up to 43% (5). Furthermore, patients vary with respect to their anatomic and physical characteristics, such as airway size, inspiratory volume and flow, and to the extent of how much they breathe nasally versus through the mouth (2). Other factors that must be considered when prescribing an inhalational device include cognitive and emotional issues, which will directly effect proper compliance with treatment (2,6,7). A.

Oral Route

The safest, most convenient, and most economical method to deliver drugs to the lung is the oral route. However, the oral route is also the most unpredictable and usually the slowest of the commonly used systemic routes; oral ingestion depends on a number of factors that often cannot be controlled or predicted. Factors that can adversely affect absorption of a drug include ingestion of food or other drugs, rate of gastric emptying, destruction of the drug by digestive enzymes or low stomach pH, rapid first–pass hepatic metabolism, insolubility of drug, emesis, and need for patient cooperation and compliance. A drug, whose intended site of action is the lung, requires a much larger dose when given orally than the same drug does when administered by aerosol inhalation; the effectiveness of a corticosteroid or another drug may be greater with the oral route, but the frequency and severity of adverse reactions are also higher. In most patients with acute asthma, an orally administered bronchodilator (e.g., ␤2–adrenergic receptor agonist, theophylline) has an onset of action too slow to be of any benefit.

Therapeutic Uses of Lung Aerosol B.

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Parenteral Routes

Parenteral injection offers several advantages over the oral route. Absorption by subcutaneous or intramuscular sites is quicker, more predictable, and less erratic than oral ingestion; the rate of absorption usually varies only with the rate of blood flow through subcutaneous or intramuscular tissue. Primary disadvantages of parenteral routes are requirements for special equipment and skilled personnel trained in the administration of sterile solutions specific for the route employed. Intravenous administration of a drug provides direct access to the systemic circulation, thus eliminating absorption variables. The greatest advantage of intravenous injection over other routes occurs in emergency situations (e.g., acute asthma), where rapid onset of action and absolute control over drug dosing is essential. For these reasons, intravenous administration of a drug (e.g., a corticosteroid) is often more effective than if given intramuscularly, subcutaneously, orally, or as an aerosol. Suitable preparation of the injected solution is essential; intravenous injection of particulate material or air can cause a fatal embolism. This route requires more skill than the other common drug–delivery routes. Once injected, no method of slowing or stopping absorption is possible; for this reason, injection is deliberately slow to allow observation of the patient for signs of adverse effects and discontinuation of the infusion if such events occur. For a more detailed discussion of systemic routes of administration, see Benet et al. (8) and Levine (9). C.

Inhalant Route

Administration of drugs by inhalation directly into the lungs by use of various aerosol–delivery systems is generally preferred because higher concentrations can be delivered more effectively resulting in rapid absorption across bronchopulmonary mucosal membranes with quicker onset of action. Furthermore, systemic side effects are significantly reduced or completely avoided. As discussed in earlier chapters, drugs can be inhaled as aerosols, which are airborne suspensions of fine particles. The particles can be comprised of either liquid droplets, or solids that remain suspended long enough to permit absorption deep into the lungs. The effects produced by inhaled particles depend on their solubility and particle size. The size of the aerosol droplets or particle–containing solution must be between 1 and 5 ␮m in diameter to permit the medication to reach the bronchopulmonary mucosal surface. Particles above 2 ␮m rarely reach the alveoli, where the conditions for absorption are greatest; particles below 0.5 ␮m are exhaled without deposition in the lungs. Particle size and particle velocity both affect deposition of ICS in the lungs (10). Inhalation devices that emit aerosolized particles at a high velocity (e.g., pMDIs) generally lead to a high degree of drug deposition in the oropharynx (11). The high velocity of aerosols makes it difficult to coordinate inhalation with device actuation. Thus, the inability to coordinate inhalation with actuation results in the deposition of drug in the oropharynx (12). Reducing the speed of the aerosol particles improves

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delivery of the drug into the airways. In addition, decreasing size of the aerosol particles will improve drug delivery. The National Asthma Education and Prevention Program recommend the use of inhaled therapies for patients with any severity of asthma in any age range (13,14). Inhaled Drugs

The concept of inhalation drug therapy dates to antiquity in Asia and the Near East. The first use of inhaled medication in Western medicine occurred in United Kingdom in 1802, where Datura ferox, a congener of atropine, was employed to treat asthma. The use of the inhalant route allows easy accessibility to the respiratory tract because antiasthma drugs and other medications can be directly administered to their sites of action in the lungs. This method of drug delivery is a primary route of administration in the therapy of a number of respiratory disorders. Advantages of inhalation include: (i) Delivers medication directly to the lungs; (ii) small amounts of drug suffice to prevent or treat symptoms (8); (iii) adverse reactions are usually much less than those produced by systemic administration (9); and (iv) there is a rapid and predictable onset of action. The primary disadvantage of the inhalant route when compared to the systemic route is that some drugs have decreased therapeutic effect by this route; for example, in severe asthma and some other respiratory disorders, oral or intravenous corticosteroids are more beneficial than inhalant corticosteroids. The ability to deliver these drugs either orally or parenterally, which results in greater systemic availability, is largely responsible for this discrepancy. Moreover, patients with acute asthma tolerate oral or intravenous corticosteroids better because inhaled corticosteroids may be aggravating to their underlying bronchial hyperresponsiveness (15,16). Most of an inhaled drug that is not deposited within the lungs is swallowed and absorbed orally; the disposition of the orally absorbed fraction is the same as any orally ingested drug. Actions of the absorbed drug can contribute to both efficacy and adverse reactions. Most inhalant drugs have low bioavailability because they are poorly absorbed (cromolyn sodium), inactivated within the lung walls rapidly by lung enzymes (␤2–adrenergic agonists), and/or inactivated in the liver by rapid first–pass metabolism (corticosteroids). The disposition of aerosolized drugs absorbed from the lungs is the same as that of intravenous drugs (15,16). Available methods of aerosol administration currently employed include the use of MDIs, nebulizers, or DPIs; the amount of drug deposited in the lungs is probably 10% to 20% for all aerosol–delivery systems. For patients who are compliant with their medication and well instructed on the use of these devices, most studies show improved drug delivery or efficacy when administered by an MDI with a spacer device compared to other inhalation methods. The incidence and frequency of adverse reactions may differ depending on the specific drug utilized and patient and environmental variables. The prescription for an MDI or a DPI should always be accompanied by careful patient instruction of inhaler technique to ensure that the medication reaches its intended site of action in the lungs.

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For most drugs delivered by any method of aerosol inhalation, mouth rinsing immediately after aerosol administration reduces the systemic absorption of drug deposited in the oropharynx. Some drugs administered by oral or parenteral routes cannot be administered as an inhalant; for example, theophylline, has relatively low potency as an inhalant and an unpleasant taste (16). No standards exist for the bioequivalence of inhaled drugs, in part, because (i) numerous confounding variables are involved, which complicate the design and evaluation of such studies, and (ii) no proof exists that patients receive inadequate therapy or increased toxicity from use of a new or different aerosol–delivery system when used as directed (17). In general, nebulizers generate particles at a slow velocity and the optimal breathing technique involves slow tidal breathing with occasional deep breaths through a mouthpiece or tightly fitting face mask (18). Nebulizers are less dependent on patient coordination or cooperation than MDIs or DPIs (3). DPI particle delivery requires a high inspiratory flow rate in order to generate sufficient force to aerosolize the powder in the device, which is often problematic in young children (4 years of age) (3,19,20). Particle delivery from a pMDI is less dependent on the patient’s inspiratory flow (20). The optimal breathing technique for use of a pMDI is actuation during a slow (30 L/min) deep inhalation, followed by a 10–second breath hold (4). However, the high velocity of aerosol emitted from a pMDI often makes it difficult to coordinate inhalation with device actuation and therefore, a spacer device is often necessary to slow down drug delivery (12). Studies investigating the effect of slow versus rapid inspiratory flow rate on the distribution of radiolabeled aerosol in the lungs of patients with asthma found that the amount of radiolabeled aerosol deposited in the trachea and lungs was reduced in six of nine patients during rapid inhalation. Although the mean values were similar (5.1 ⫾ 3.0 ␮L vs. 5.2 ⫾ 2.3 ␮L) in both groups, rapid inspiration resulted in greater heterogeneity in lung deposition resulting in regions of higher–density labeling compared with slow inspiration (21). Slow inspiration resulted in better distribution of the aerosol into the lung periphery (21). D.

Metered Dose Inhalers

MDIs are the most popular device for aerosolized administration of drugs. With this method, a medication is mixed in a canister with a Freon propellant, and the preformed mixture is expelled in precise measured amounts upon actuation of the device. The Freon propellant is an inert chlorofluorocarbon. In the past few years, there has been a concern regarding the use of Freon propellants, which contain chlorofluorocarbons (CFCs) due to the risk of increased damage to the ozone layer. There are now CFC–free MDIs [known as hydrofluoroalkane (HFA)] that are commercially available, with a better safety profile for the environment. HFA formulations are available for beclomethasone diproprionate (QVAR HFA 40 and 80 ␮g), fluticasone HFA, and albuterol HFA. DPIs are becoming more readily accepted for the treatment of asthma than MDIs. Proper use of both MDIs and DPIs require that patients learn how to coordinate exhalation and inhalation with actuation of

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the device. This can be very difficult for some patients, particularly the very young and the elderly. Moreover, because several studies have demonstrated that many physicians, pharmacists, and nurses do not know how to use such inhalers correctly, it is not surprising that at least 50% to 60% of the patients are estimated to have improper inhaler technique, resulting in little or no benefit from drugs administered as MDIs or DPIs (4). Table 1 shows the correct procedure for MDI administration. To improve inhalant technique and drug delivery, spacer “devices” were introduced in 1956 (22). The concept of how spacers work is simple; after actuation, the drug is suspended within a tube or spacer device for a few seconds prior to inhalation. This results in decreased particle size of the aerosolized drug, which permits improved deposition into the lungs and reduced drug deposition in the oropharynx. The use of a spacer reduces the need for patient coordination of actuation and inhalation. The most useful spacer devices are high–volume spacers, which are often bulky; this can be a deterrent for patients who require use of MDIs outside their homes. MDI spacers vary in size and shape (4). As stated above, the amount of drug deposited in the oropharynx and swallowed is decreased significantly when spacers are used; the amount deposited within the spacer chamber varies with different spacers. The Nebuhaler and Volumatic have large chambers (about 750 mL), the Inspir–Ease is a 400 mL collapsible bag, and the Aerochamber has a volume of approximately 130 mL. Some spacers have flow indicators (e.g., Inspir–Ease). Short–tube spacers are less efficient in delivering drug to the lungs than are the large–volume spacers; however, in most patients, they probably deliver therapeutically adequate amounts. It is recommended that once patients are initially taught to use correct MDI–delivery technique (with or without spacers), they should be periodically asked to demonstrate the correct MDI technique during subsequent physician office visits. Patients may develop poor MDI technique over time, and continued reinforcement is therefore essential. Most authorities on asthma advocate that all asthma patients use spacer devices with their MDIs. Table 1 describes the procedure for the administration of pharmaceutical aerosols from MDIs. The openmouth procedure results in inhalation of smaller particles and deposition of a higher percentage of the administered drug within the lungs than the closedmouth technique (23).

Procedure for the Administration of Pharmaceutical Aerosols from Metered Dose Inhalers Table 1

Hold the mouthpiece a few inches from the mouth (openmouth procedure) or in the mouth with closed lips (closedmouth procedure) Exhale deeply Start slow deliberate inhalation and actuate the device Hold the breath for 10 sec (or at least 4 sec) Breathe normally

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Many patients actuate the device at the wrong time or inhale too rapidly. Improper administration often results in deposition of most of the drug dose in the mouth and pharynx; most of the adverse reactions produced by inhalant administration of either ␤–adrenergic agonists or corticosteroids result from this misdirected deposition (15). Breath-Activated Metered Dose Inhalers

A breath–activated MDI (Maxair Autohaler, 3M Pharmaceuticals is in St. Paul, Minnesota, U.S.A.) is currently available for delivery of measured doses of the ␤2–adrenergic agonist, pirbuterol. With this device, patient coordination with actuation and inhalation is not necessary because the actuator device releases the medication particles automatically during inspiration. This device is ineffective if patients stop inhaling at the time of actuation; therefore, patient instruction on proper use is required. Moreover, this device is being reformulated with an ozone–friendly propellant (as discussed below). Albuterol MDI is available in Europe using this same delivery system (24). Chlorofluorocarbons in Metered Dose Inhalers

MDIs have been reformulated as a result of the ban being implemented throughout the world by the United Nations on the use of CFCs. CFCs are the propellant used in all MDIs for aerosol generation. The use of these compounds has been implicated in the depletion of the Earth’s ozone layer, which protects our skin from the sun’s ultraviolet radiation. Although the medicinal use of CFCs accounts for less than 1% of the commercially produced product, an absolute ban of CFC production will soon occur (25). The United Nations Environment Program allowed for the development of suitable alternatives such as the HFA and newer and improved DPI systems, which will allow inhalant administration of all drugs presently delivered with MDIs (26). The use of DPIs will be discussed in the following section. Another alternative has been developed allowing the continued use of MDIs using the Airomir CFC–free system; in this new system, HFA 134a (HFA–134) is substituted for CFCs (27). The first drug tested with the HFA–134 propellant system was albuterol; phase I, II, and III studies have been carried out in humans. Currently, there are several HFA MDIs available for use in the United States including Proventil HFA (Schering), Albuterol HFA (IVAX), and Xopenex HFA (Sepracor) for the treatment of acute bronchospasm associated with obstructive lung diseases such as asthma. Thus far, all three of these inhalers appear to be safe and efficient in delivering medication to the lungs. A corticosteroid containing CFC–free system is also being developed. The unique physicochemical properties of HFA–134 have made it necessary to design a completely different delivery device from that currently employed with CFC–containing MDIs. The optimal doses, delivery parameters, and adverse reactions of drugs delivered with the Airomir system have yet to be assessed. The delivery of a medication with HFA–134 is considered a new drug by the FDA and will require an new drug application (NDA) (27).

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Many physicians prefer to use nebulizers for therapy of acute asthma in an emergency care unit or in patients with severe asthma at home (15,28). In jet nebulizers, the aerosol is formed by a high–velocity airstream from a pressurized source directed against a thin layer of liquid solution. Ultrasonic nebulizers are used sometimes; with this device, the vibration of a piezoelectric crystal aerosolizes the solution. Ultrasonic nebulization systems are much less popular in hospitals than jet nebulizers because they are more expensive and are not disposable; moreover, continuous atomization does not usually occur if the volume in the reservoir chamber falls below 10 mL. Furthermore, the output for ultrasonic nebulizers is decreased if suspensions and highly viscous solutions are used (11). These devices only nebulize the water and not the drug in a suspension and therefore, they may not be able to generate a spray from a viscous drug solution. Using ultrasonic nebulizers may damage the therapeutic effect of certain drugs (29,30). As with MDIs and DPIs, less than 20% of the administered dose of a jet nebulizer reaches the lungs; most of the nebulizer drug dose is deposited in the tubing and apparatus. Up to 10 times the dose delivered with an MDI is administered using a jet nebulizer; therefore, jet nebulizers can deliver more drug to sites of action in the lungs than MDIs with spacers or DPIs. Patient coordination or cooperation is unnecessary; drug delivery to the lungs is adequate in severely dyspneic patients regardless of age. Unlike MDIs or DPIs, combination of drugs can be mixed in the nebulizer solution and administered at the same time. With nebulizers, flow rate should reflect flow through the nebulizer chamber. A high flow rate reduces the nebulization time; 5 to 10 minutes is suggested as a clinically acceptable range (29,30). In comparison to MDIs and DPIs, the chief drawbacks of nebulizers have been their lack of portability, greater costs of drug delivery as a result of the need for extensive assistance from well–trained healthcare personnel, and the requirement for higher drug doses to achieve a therapeutic effect. They also require more time (about 5–10 minutes) to complete a treatment. Use of nebulizers has traditionally been limited to hospital, outpatient (physician office or emergency room), or home locations for patients who find it difficult or are unable to use MDIs with spacers (e.g., young children or patients with severe asthma). However, newer portable, battery–powered nebulizers are now available that make it easier for actively mobile patients to use. To reduce costs, many hospitals have limited the use of nebulizers because studies have demonstrated that four puffs of a short–acting beta agonist MDI results in an equivalent amount of drug delivered to the lung as one nebulizer treatment. However, there may be a psychological benefit in some patients treated for acute asthma with medication delivered through a nebulizer (29,30). Dry Powder Inhalers

Delivery of medication with a DPI requires minimal patient cooperation and coordination of breathing with actuation of this device. As discussed in earlier

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chapters, the DPIs are designed to rapidly deliver the drug powder once it is released, using the patient’s inspiratory efforts to inhale drug–containing particles. Most patients can generate the minimal inspiratory flow rate of at least 30 L/min for effective inhalation with a DPI. However, infants, some children under six years of age, mentally handicapped patients, and severely ill, semialert individuals may have difficulty achieving sufficient inspiratory flow required for proper use of DPIs (26). DPIs are suitable substitutes for CFC–containing MDIs. Delivery parameters for DPIs are discussed in detail in another chapter. The drug is mixed as a powder with a flow aid or filler such as lactose or glucose and placed in a capsule; to release the powder, the capsule is positioned within the DPI–delivery device and then pierced with a needle (Spinhaler®) or sheared in half (Rotohaler®). The energy required to disperse the powder is produced by the patient’s inspiratory effort (15). Because of the need to load each capsule individually, single–dose DPI–delivery system such as the Spinhaler and Rotohaler are less convenient to use than multidose systems (e.g., Diskhaler® and Turbohaler®). Two potential problems can occur with DPIs: (i) Most of the drug dose may be deposited within the mouth because the capsule may fail to release the drug or humid/moist conditions may cause aggregation of the powder, and (ii) the released powder may irritate the oropharynx and produce cough. DPIs are small, portable devices that are carried easily in a pocket or purse. No need exists for use of spacers. Like MDIs and nebulizers, less than 20% of the delivered drug reaches the lungs; most of the dose is deposited within the drug capsule or container. Most data indicate that if the drug–delivery technique is appropriate, no difference in efficacy exist using the same drug administered with a DPI or an MDI (15,26). For most drugs available as MDIs, specific DPI–delivery devices have been developed, although no standard design useful for delivery of all inhalant drugs is currently available. As many as 25% of the patients have been estimated to have improper DPI–delivery technique (26). Therefore, even with reduced requirements for coordination, patients should receive clear and repeated instructions on the proper use of a DPI. Patients should be told: (i) to exhale deeply (but not into the DPI); (ii) to activate the DPI and put the mouthpiece in the mouth; and (iii) to inhale as quickly and completely as possible (26). A variety of DPI devices with different designs are available in the United States, and a few others are under development. There are no significant clinical advantages of one DPI system over another. A DPI system for administration of inhalant corticosteroids is available in Europe and in the United States (26). The medications currently available in the United States for delivery with DPIs include salmeterol xinofoate (Serevent®) diskus, fluticasone propionate/salmeterol xinofoate diskus (Advair® comes in three concentrations; 100/50, 250/50, and 500/50), tiotropium bromide (Spiriva®) handihaler, budesonide (Pulmicort®) turbuhaler, fomoterol fumarate (Foradil®) aerosolizer, and mometasone furoate (Asmanex®) twisthaler.

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Aerosolized Use of Drugs in Asthma Therapy Types of Asthma and Management

Guidelines for the treatment of asthma have been presented by the National Heart, Lung, and Blood Institute and, more recently, by a joint task force convened by the Joint Council of the American Academy of Allergy and Immunology and the College of Allergy and Immunology (31–33). Regimens recommended for treatment reflect the consensus definition of asthma as a disease characterized by: (i) airway obstruction that is partially or totally reversible either spontaneously or with treatment; (ii) airway inflammation; and (iii) increased airway responsiveness to a variety of nonspecific stimuli (32). A “stepcare” pharmacologic approach to treatment of asthma is recommended. The number of medications and the frequency of their administration usually increase as the severity of asthma increases. The emphasis of treatment should be on controlling underlying airway inflammation which, if left untreated, can result in permanent structural changes in the airways. In the previous classification schemes, asthma has been characterized as mild, moderate, severe, and acute or “status asthmaticus.” Mild asthma has been defined as intermittent, brief (⬍1 hour) episodes consisting of wheezing, coughing, and dyspnea up to two times per week. Individuals with mild asthma are asymptomatic between exacerbations, but have episodic symptoms with activity such as exercise, or infrequent nocturnal symptoms of cough and wheezing (⬍2 times per month). Their forced expiratory lung volumes in one second (FEV1) during a symptomatic episode is greater than 80% of baseline but may improve by 20% or more after therapy with bronchodilators (32). In moderate asthma, symptoms usually occur one or two times per week. Exacerbations often affect sleep and/or activity levels and may last several days, necessitating occasional emergency care to control symptoms. These individuals have baseline FEV1, or peak expiratory flow rates (PEFRs), of 60% to 80% of baseline, but this could vary by 20% to 30% during symptomatic periods (32). Patients with severe asthma have continuous symptoms, which significantly limit their activity levels and cause frequent nocturnal symptoms. Emergency room treatment and/or hospitalization are occasionally required. In patients with severe asthma, the FEV1 or PEFRs are less than 60% at baseline and are subject to high variability (20–30%) after bronchodilator therapy (32). The severity of a specific exacerbation cannot be predicted by the patient’s history. Patients with mild intermittent asthma may have minimal or no symptoms; however, when some patients in this group do present with symptoms; the symptoms occur severely and decompensate rapidly. Such patients present very similarly to those with acute severe asthma, or status asthmaticus, which refers to an exacerbation of asthma unresponsive to bronchodilator medication. Patients have progressive worsening of symptoms—shortness of breath, coughing, wheezing, and chest tightness resulting in significant respiratory distress. Decreased respiratory airflow is documented by either PEFRs or spirometry. Hospitalization of these patients is usually required. Recently, many advocate that the treatment of asthma should be based on asthma

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control of lung inflammation and hyperresponsiveness rather than the severity of symptoms as in the prior classification scheme (31–33). The ultimate goals of asthma drug therapy are prevention and control of—symptoms, bronchial hyperresponsiveness, and inflammation and preservation of lung function. Treatment of acute symptoms involves the use of selective ␤2–adrenergic agonists (32,34). These medications should only be used as rescue medication and not as maintenance therapy. Another class of drugs is the long–acting bronchodilators Salmeterol xinofoate (Serevent) and fomoterol fumerate (Foradil) (35–37). These agents are also used in combination with inhaled corticosteroids (38). Long–acting bronchodilators should not be used alone for the treatment of asthma. There have been many case reports of fatality related to their use alone (39). Theophylline and other xanthine oxidase inhibitors also act as bronchodilators and can be useful for the treatment of asthma (40). However, theophylline is less commonly used than previously because of its narrow therapeutic window, increased incidence of drug interactions resulting in toxic side effects, and generally weaker bronchodilator action as compared to the selective ␤2–agonists (40). In the appropriately selected patients, theophylline (i.e., Uniphyl®) has been safely used for the treatment of nocturnal wheezing. Anticholinergic agents [e.g., ipratropium bromide (Atrovent®)] and tiatropium bromide (Spiriva) are not recommended as first–line agents in the treatment of asthma because they have a relatively weaker bronchodilator effect than ␤2–agonists. However, a recent meta–analysis has demonstrated that there is enhanced bronchodilation and improvement in spirometric parameters within one to a few hours after treatment with aerosolized anticholinergic agents with or without ␤2–agonists [Combivent® or Duo Nebulized albuterol in combination with ipratropium bromide (Nebs®)] (41). Use of DuoNebs® for the treatment of children in status asthmaticus has been adopted by several pediatric emergency rooms because they are associated with fewer side effects such as tachycardia (42). More recently, levalbuterol tartrate (Xopenex®), an isomeric form of racemic albuterol, has been approved for the treatment or prevention of bronchospasm in patients with reversible obstructive airway disease such as asthma (43,44). The primary advantage of this compound compared to its racemic counterparts is fewer side effects such as tachycardia (44,45). Long–term control of asthma in most patients involves early use of inhalant anti–inflammatory medications, i.e., the antiallergic drugs cromolyn sodium (Intal ®) or nedocromil sodium (Tilade ®), and/or aerosol preparations of corticosteroids (46,47). These drugs reduce airway inflammation and hyperresponsiveness leading to optimal asthma prevention and control of symptoms. Cromolyn sodium is an effective drug when used as a pretreatment for exercise– and cat–induced asthma. Neither cromolyn/nedocromil nor inhalant corticosteroids are effective for treatment of patients with acute asthma exacerbations. These agents should therefore not be used for immediate effects required in such patients (46). However, if patients have been treated with them prior to the acute episode, they should be maintained on them. Recommendations regarding the treatment of mild (episodic and persistent), moderate, and severe asthma have been outlined by the National Heart,

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Lung, and Blood Institute, the Joint Task Force of the Allergy societies, and Global Initiative for Asthma (GINA) guidelines; aggressive use of anti–inflammatory agents soon after diagnosis of the disease is advocated. However, by general consensus, all patients with asthma should be evaluated for possible allergic components in order to implement effective avoidance measures in home and work environments (32,33). A thorough physician–administered allergic history followed by confirmatory skin testing to common seasonal and perennial allergens is essential (32,33). For mild asthma, the GINA guidelines recommend inhaled ␤2–agonists or antiallergic drugs such as cromolyn or nedocromil sodium prophylactically to prevent or diminish a predictable asthmatic response when a patient cannot avoid asthmatic stimuli or risk factors (33). Examples of this are exercise–induced bronchospasm or cat–induced asthma. During symptomatic periods, short–acting inhalant ␤2–agonists are recommended to control acute symptoms. However, if a ␤2–agonist is used more than three or four times a day or every night, this indicates the need for additional therapy for treatment of mild, persistent, and more severe forms of asthma (33). Both short– and long–acting ␤2–agonists should never be used as a monotherapy for the maintenance of asthma control. Individuals with moderate asthma have symptomatic exacerbations of asthma more than twice a week associated with nocturnal symptoms. Requirements for frequent use of ␤2–agonists indicate the presence of significant airway hyperresponsiveness, which correlates with airway inflammation. Therefore, the use of anti–inflammatory medication such as inhalant cromolyn or nedocromil sodium and/or corticosteroids is required. Cromolyn and nedocromil sodium are virtually devoid of side effects, and either drug is very effective for prophylaxis in selected patients; neither should be used for treatment of an acute exacerbation (46). Inhalant corticosteroids have been shown to suppress airway hyperresponsiveness at doses ranging from 400 to 1600 ␮g/day and are more effective for controlling symptoms in a broader spectrum of patients (47–50). Anti–inflammatory medications must be used on a regular basis for optimal effectiveness. This recommendation requires constant reinforcement to the patient. Oral leukotriene–modifying agents as first–line treatment for mild persistent asthma are discussed elsewhere. For severe asthma, treatment is similar to moderate asthma, but higher dosages of inhaled corticosteroids, ranging between 800 and 1600 ␮g/day or nebulized budesonide are required with or without long–acting ␤–agonists (33,47–50). These individuals may also benefit from concomitant use of cromolyn or nedocromil sodium with occasional or frequent bursts of systemic corticosteroid therapy. In any patient with asthma, it must be emphasized that treatment should utilize the fewest number of medications necessary in order to optimize compliance with medication. This requires increasing each drug to its maximal dose well before adding another class of drug to the treatment regimen (33). Management of mild, moderate, and severe asthma in children is very similar to that in adults. Because the adverse reactions to the inhalational therapies discussed are very low, the benefits of treatment with these medications far outweigh the risks of therapy (46–50). Concerns about the possible adverse

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effects of inhalant corticosteroids on bone growth in children remain unresolved because some studies have demonstrated effects on growth and puberty with high doses of inhalant beclomethasone. Beclomethasone diproprionate is metabolized to beclomethasone monoproprionate, which has been shown to significantly prolong its half–life resulting in more potential systemic side effects (47,48). Lower doses of beclomethasone have been able to be used to formulate beclomethasone HFA (QVAR) because the very small, aerosolized particles (1 um in size) have been demonstrated to penetrate the large and small airways in the lung much more effectively without the need for a spacer device (47,48). However, because systemic absorption and side effects appear to be dose dependent, significant adverse effects are unlikely with the doses required to treat the vast majority of children with asthma (34). Patients with potentially fatal asthma who are at high risk from an exacerbation include those patients with a history of: (i) prior intubation for asthma; (ii) two or more hospitalizations for asthma in the past year, or three or more emergency care visits for asthma in the past year; (iii) hospitalization or emergency care visits within the past month; (iv) chronic use of systemic corticosteroids or recent withdrawal from systemic corticosteroids; and (v) serious psychiatric disease or psychosocial problems (33,51). The principles of care for an exacerbation should include reversal of airway obstruction using inhalant ␤2–agonists and early addition of systemic corticosteroids to treat the underlying airway inflammation. Supplemental oxygen may be necessary in patients who are hypoxic (33,51). The management of acute asthma or status asthmaticus requires early recognition of diminishing lung function followed by prompt communication between the healthcare provider and the patient about symptoms of clinical deterioration. Only by this interactive approach is it possible to institute early therapeutic intervention. This often requires intensification of the current antiasthma regimen, removal of an offending allergen or irritant from the environment, and the addition of systemically active corticosteroids (33,34). In summary, the ultimate goals for the treatment of asthma in adults and children are to control symptoms, improve pulmonary function, reduce PEFR variability, restore normal activity levels, attenuate or eliminate nocturnal symptoms, and reduce frequency of ␤2–agonist use (33,34). These outcome measures will reduce the number of asthma exacerbations. The resulting decrease in the frequency of emergency room visits and hospitalizations should significantly improve the quality of life for such patients and at the same time reduce healthcare costs. B.

Specific Drugs Used for Aerosol Delivery in Asthma Therapy

␤ Adrenergic Agents

The biochemical mechanisms of action of ␤2–agonists have not been completely defined, but they are known to enhance the production of 3,5–cyclic adenosine monophosphate (cyclic AMP) from adenosine triphosphate (ATP) in the presence of adenyl cyclase. Increased cyclic AMP in turn triggers other intermediate reactions, which result in bronchodilation and inhibition of mediator release in immediate hypersensitivity reactions (37).

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Pharmacologic and biochemical research has resulted in a better understanding of the adrenergic receptor system and has led to the development of drugs with more specific stimulating or blocking action. The basic molecular structure of adrenergic drugs is the catechol nucleus. This consists of a benzene ring with hydroxyl groups at positions 3 and 4 and a two–carbon (␣ and ␤) amide side chain at position 1 (34,37). Catecholamines include epinephrine, isoproterenol (Isuprel®), and isoetharine (Bronkosol®), and these agents are inactivated quickly in vivo by cellular catechol–o–methyltransferase (COMT) and sulfatases that are found in the gut and liver (34,37). The resulting metabolites are devoid of bronchodilating properties. Alteration of the molecular structure of the catecholamine nucleus has resulted in a number of antiasthma drugs. Alteration of hydroxyl positions on the benzene ring from the 3 and 4 positions makes these agents refractory to the effects of COMT and therefore longer acting. This group of compounds is referred to “resorcinol” ␤2–agonists. Resorcinols include ephedrine (Primatene®), pirbuterol (Maxair®), metoproterenol (Alupent®), terbutaline (Brethaire®), and fenoterol (Berotec®) (34,37). “Saligenins” refer to the group of compounds in which a methanol group is substituted for the hydroxyl group at the 3 position of the benzene ring. These compounds are also refractory to the actions of COMT, and therefore have a longer duration of action than isoproterenol (34,37). Included in this group are albuterol (Proventil® and Ventolin®), bitolterol (Tornalate®), and procaterol (Proair®). Levalbuterol tartrate (Xopenex®), an S–isomer–free form of racemic albuterol, is a new variation of short–acting ␤2–agonist that has fewer cardiac side effects than other similar agents. ␤2–Agonists quickly reverse bronchoconstriction and can increase the rate of mucociliary clearance, which is often decreased in asthma patients (52). The newer selective ␤2–agonists are designed to withstand immediate degradation and are thus longer acting. Most ␤2–adrenergic drugs are available as MDIs (as discussed in Introduction of this chapter), which can deliver medication directly into the bronchi in much smaller doses, thereby reducing the incidence of severe adverse reactions (52). Some ␤2–agonists (e.g., albuterol and levalbuterol) are available as solutions for nebulization (34,37). Most of the selective ␤2–adrenergic agonists have similar efficacy and potency as well as ␤2–receptor selectivity with peak effects ranging from 30 to 60 minutes (37). Duration of therapeutic effects range from two to six hours. Salmeterol (Serevent®) is a long acting ␤2–adrenoceptor agonist, which is the chemical analog of albuterol (37). Salmeterol is structurally similar to albuterol except that it has a long lipophilic N–substituent side chain, which is believed to bind the molecule firmly to the region of the ␤2–adrenoreceptor protein, allowing repeated contact of the molecule with the receptor (34,37). This higher receptor affinity results in extended duration of action (up to 12 hours) in patients with asthma (34,37). Salmeterol has also been shown to prevent bronchoconstriction induced by methacholine, histamine, allergens, hyperventilation of cold air, and exercise for up to 12 hours. Its bronchodilator effect is approximately 10 times more potent than albuterol with a much higher ␤2/␤1–selectivity ratio (50,000:1 vs. 650:1) (35). Salmeterol’s long half–life makes it valuable in the treatment of nocturnal asthma and a useful

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adjunctive medication for treatment of mild, moderate, and severe asthma (34,35,53,54). Similar effects have been reported with formoterol, which also has a long half–life, but differs from salmeterol by its quick onset of action within two to five minutes (36,55). A number of studies evaluating the safety of ␤2–agonists in recent years have reported a possible association between ␤–agonist use and increased death from asthma. Initial evidence relating the increased morbidity and mortality of asthma to ␤2–agonist use stems from two “epidemics” of asthma deaths. Increases in asthma deaths were associated with the use of high–dose inhaled isoproterenol in the United Kingdom in 1960 and, more recently, with the use of fenoterol, a potent, short–acting, nonspecific ␤–agonist in New Zealand in the 1970s (56). Increased morbidity/mortality with ␤2–agonists has been speculated to occur as a result of increased delivery of antigen into the airways, enhancing the progression of airway inflammation. It has been further hypothesized that acute bronchospastic responses may have a protective effect in shielding the lower airways from additional antigen exposure. However, this is speculation that has not been proven and remains theoretical (57). Moreover, the contention that regular daily use of these drugs is unsafe has been challenged by some authorities (57). Both of the published guidelines recommend that ␤2–agonists be reserved for symptomatic relief of bronchospasm and prevention of exercise–induced asthma (33,34). The use of ␤2–agonists on an as–needed basis may prevent tolerance, which occurs as the result of downregulation of ␤2–receptors (56). Salmeterol is being advocated for use in patients with nocturnal asthma or in patients who require frequent use of short–acting ␤2–agonists in conjunction with anti–inflammatory agents such as inhalant corticosteroids. This drug should not replace the use of anti–inflammatory medication (33–36), nor should it be used for acute asthma. Anticholinergics

Anticholinergic agents may diminish cyclic guanosine monophosphate levels and inhibit vagal efferent pathways (41,58). Atropine and analogs of this compound such as ipratropium bromide (Atrovent) have been shown to possess bronchodilating properties. In general, most studies have not demonstrated an enhanced benefit of combined use of anticholinergic agents with bronchodilators (e.g., theophylline or a ␤2–agonist) for treatment of asthma (41,58). Cromolyn and Nedocromil Sodium

Cromolyn sodium and nedocromil sodium are useful antiallergic, anti–inflammatory agents in the treatment of asthma. Cromolyn sodium is believed to exert clinical effects through several mechanisms including: (i) mast cell membrane stabilization; (ii) protein phosphorylation; (iii) inhibition of activation of inflammatory cells and the physiologic effects of their mediators; (iv) attenuation of the physiologic effects of mediators on smooth muscle; and (v) inhibition of neuroreflex bronchoconstriction by blocking the effects of neuropeptides (46).

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The majority of clinical studies have demonstrated significant decreases in bronchial hyperresponsiveness when therapeutic drug dosages are administered on a regular daily basis. Cromolyn sodium inhibits both early– and late–phase reactions induced by allergen challenge (46). Several types of asthma respond well to cromolyn sodium including seasonal or perennial allergic asthma, exercise– and cold–induced asthma, cough variant asthma, animal–induced asthma, and occupational asthma. Cromolyn sodium may reduce the need for adjunctive antiasthma medication such as theophylline, ␤2–agonists, and corticosteroids. Cromolyn sodium is a poor bronchodilator compared to ␤2–agonists and has no value in acute asthma. Because its anti–inflammatory effects are not as potent as those of inhalant corticosteroids, cromolyn is usually not useful as a single drug in patients with moderate–to–severe asthma (46). Cromolyn sodium is supplied as ampules (20 mg/2 mL) for use in nebulizers, as an MDI, and as spinhaler–powdered capsules for the treatment of asthma. Tachyphylaxis or tolerance to cromolyn sodium after chronic use has not been demonstrated, largely because of its short half–life and rapid excretion rate (46). Cromolyn sodium is extremely safe; the most commonly reported side effects include throat irritation, hoarseness, and dry mouth. This drug is also safe during pregnancy and breast–feeding; only negligible amounts cross the placenta or enter breast milk (46). Nedocromil sodium was developed in the search for a new drug with greater potency and a more diverse clinical spectrum of action than cromolyn sodium. Although nedocromil differs structurally from cromolyn sodium, it has similar pharmacologic activities and some additional attributes (46). Suggested mechanisms of action for nedocromil sodium include: (i) mast cell membrane stabilization with inhibition of mediator release; (ii) inhibition of neuropeptide release from sensory nerves; (iii) inhibition of mediator release from other inflammatory cells such as platelets, neutrophils, eosinophils, monocytes, macrophages, and epithelial cells; (iv) inhibition of protein kinase C; and (v) inhibition of physiologic effects of bioactive mediators such as chemotaxis induced by platelet–activating factor (46). Most studies evaluating the clinical efficacy of nedocromil sodium have reported that it is superior to placebo and as effective as cromolyn sodium in improving symptom scores and reducing the use of bronchodilators (46). Like cromolyn, nedocromil sodium also blocks the early– and late–phase asthmatic airway responses. Nedocromil sodium also reduces bronchial hyperresponsiveness induced by a variety of specific stimuli such as ragweed, and nonspecific stimuli such as adenosine, sulfur dioxide, metabisulfites, and exercise (46,59). It inhibits citric acid–induced cough and is more effective than cromolyn sodium in preventing sulfur dioxide–, adenosine–, and cold–induced asthma; nedocromil appears to be especially effective in treating chronic coughs associated with bronchial hyperresponsiveness (60). Administered by inhalation using an MDI, it provides 4 mg per inhalation. The recommended dose is two puffs four times a day. Like cromolyn sodium, it is safe during pregnancy and breast–feeding (46). Recent information on the mechanism of action of nedocromil sodium

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suggests that it acts by blocking chloride channels of effector inflammatory cells, thereby preventing cell activation and inflammation (61,62). Inhalant Corticosteroids

Corticosteroids are the most effective drugs for the treatment of asthma. The exact mode of action of corticosteroids in asthma is unknown but probably results from their broad anti–inflammatory effects. Oral corticosteroids reduce bronchial mucosa edema and mucous accumulation, suppress the inflammatory response, and stabilize membrane permeability (63,64). Corticosteroids also stimulate lipomodulin, which inhibits the action of phospholipase A2. This reduces the availability or arachidonic acid for synthesis of bioactive mediators such as leukotrienes, prostaglandins, and thromboxanes (63,64). Corticosteroids restore ␤2–adrenergic responses in the bronchial airways and have been shown to decrease the extraneuronal uptake of catecholamines. Corticosteroids suppress IgE–dependent late–phase allergen reactions. One week of corticosteroids can also inhibit the early response to allergen (63,64). Corticosteroids potentiate ␤2–adrenergic–stimulated increases in activity of adenylate cyclase and subsequent increase in levels of cyclic AMP (63,64). Aerosolized corticosteroids were first introduced in the United States in 1976 (65). Currently, available agents include beclomethasone dipropionate (Vanceril® and Beclovent®), triamcinolone acetonide (Azmacort®), and flunisolide (Aerobid® and Aerobid M®) (47–50,65). Beclomethasone is a nonfluorinated preparation, whereas both flunisolide and triamcinolone are fluorinated (which adds to their topical potency). Topical agents provide local efficacy while minimizing systemic side effects (47–50). Studies have shown that inhalant corticosteroids are effective in suppressing nonspecific and specific allergen–induced airway hyperreactivity in adults and children (47–50,65). Their mechanisms of action are not completely understood but are believed to reduce inflammation by: (i) decreasing capillary permeability; (ii) inhibiting mast cell mediator release; and (iii) inhibiting phospholipase A2 activity. These actions reduce bronchoconstriction, edema, and mucous secretion secondary to decreased synthesis of leukotrienes and prostaglandins (50,65). Dose equivalency studies show that beclomethasone dipropionate administered as 10 inhalations (400 ␮g/day) is approximately equivalent to 7.5 mg of oral prednisone daily (65,66). Flunisolide 1000/␮g daily (four inhalations) is equivalent to approximately 9 mg of prednisone a day (65,66). Budesonide 800 ␮g/day (available as 200 ␮g spray, DPI) has a therapeutic equivalence of 35 to–58 mg of prednisone (65,66). Mometasone fumarate (Asmanex) twisthaler is a dry powder inhaled corticosteroid with potent anti–inflammatory activity. In vitro, mometasone has a binding affinity for human glucocorticoid receptors, which is 12 times that of dexamethasone, seven times that of triamicinolone, five times that of budesonide and 1.5 times that of fluticasone. Clinical trials have demonstrated that it reduces asthma symptoms, improves lung function, and was safe and well tolerated. Asmanex can be administered as 110 ␮g one puff twice, 220 ␮g twice a day, or 440 ␮g once a day.

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Patients initiated on inhalant corticosteroids should be warned not to expect immediate symptom relief because most studies show maximum improvement by the second to third week of use. Furthermore, if prescribing an MDI, patients should be educated about the proper use of spacer devices (discussed in Introduction), which improve deposition of these medications in the airways while reducing local irritating side effects such as oral candidiasis (occurring in approximately 5.5% of the patients), hoarseness, and cough (65,67). Systemic side effects are unusual with recommended doses of inhalant topical corticosteroids. Hypothalamic pituitary adrenal (HPA) axis suppression by inhalant corticosteroids has been studied extensively (67,68). These agents generally do not cause HPA axis suppression at doses equivalent to 800 ␮g/day of beclomethasone diproprionate. There is general agreement that daily doses of beclomethasone greater than 1600 ␮g/day (40 inhalations) can be associated with measurable HPA suppression (67,68). Prolonged use of inhalant corticosteroids at doses less than or equal to 800 ␮g/day have been demonstrated to be safe in adults in several recent clinical studies (67,68). Furthermore, both triamcinolone and beclomethasone dipropionate at 400 ␮g/day (10 inhalations per day) are therapeutically effective, and they do not seem to impair bone formation, growth, or adrenal function in children at these doses (65,67,68). Inhalant corticosteroids are also safe to use during pregnancy and while breast–feeding (65,67,68). Combination of Anti inflammatory Agents

Recent studies have demonstrated an enhanced effect when antiallergic agents (i.e., cromolyn or nedocromil sodium) and inhalant corticosteroids are used together (when either agent alone does not control symptoms) (59,69). The additive effects have been attributed to their different mechanisms of action inhibiting inflammation (59,69). Nedocromil sodium is more effective than cromolyn sodium as a corticosteroid–sparing agent. C.

Aerosol Therapy for Pulmonary Infections

A number of antibiotics have been administered by aerosol to treat bronchopulmonary infections in order to decrease the systemic dosage necessary to achieve effective antibacterial control and limit systemic adverse effects (70). This is particularly important in patients with CF, who are susceptible to developing recurrent pulmonary infections secondary to bronchiectasis (see the section on CF). The results of clinical studies show that intrabronchial concentrations are lower than intratracheal instillation of these antibiotics but remain 10 to 40 times greater than with parenteral administration (41). Caution using aerosolized antibiotics should be emphasized in patients with bronchial hyperresponsiveness because bronchospasm has been reported as a potential complication of this treatment (70). The two most common medical conditions where aerosolized antibiotic therapy has been used are mucoviscidosis (CF) and bronchiectasis (70). In mucoviscidosis, these agents have been helpful in eradicating acute infections and preventing chronic colonization of difficult–to–treat organisms such as

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Pseudomonas (70). Thus far, aerosolized antibiotics have been demonstrated to be as therapeutically effective as parenteral antibiotics while causing fewer systemic side effects (70). One major advantage of aerosolized antibiotics is that it allows patients who are susceptible to recurrent chronic pulmonary infections to receive treatment at home (70). The concentration of aerosolized aminoglycosides (e.g., tobramycin) obtained by bronchoalveolar lavage has been determined. In most studies, the concentration achieved was in excess of the minimal inhibitory concentration necessary to kill most of the organisms cultured (71). Treatment of nosocomial Gram–negative infections in critically ill patients with aerosolized antibiotics such as tobramycin and third–generation cephalosporins kill over 96% of the pathogens colonizing the respiratory tract (71). Up to two months after treatment was discontinued, recolonization with resistant organisms or superinfections was not reported (71). Other antibiotics used as aerosols to treat a variety of pulmonary infections (e.g., Legionnaire’s disease and pulmonary aspergillosis) include amikacin sulfate, ciprofloxacin, erythromycin, and amphotericin B (72–75). D.

Other Medications Used as Aerosols

A number of other agents are under clinical investigation to determine efficacy for asthma when delivered as an aerosol (76). For many years, heparin has been known to have anti–inflammatory properties, which could be effective for treatment of asthma; this agent competitively inhibits inositol 1,4,5–triphosphate (InsP3) receptors in various cell types (76). Because InsP3 is a second messenger involved in activation of mast cells, heparin may inhibit mast cell activation and mediator release, which are important events occurring during exacerbations of asthma. Preliminary studies evaluating the efficacy of heparin in attenuating bronchial hyperresponsiveness have been reported in an allergic sheep model (77). Aerosolized heparin attenuated the effects on airway resistance after inhalation challenge with ascaris serum antigen; this occurred apparently in a time–dependent fashion because its effect was enhanced when administered 60 minutes before antigen challenge (77). Heparin has also been effective in preventing exercise–induced bronchospasm (78). The therapeutic effect of heparin on exercise–induced bronchospasm occurred independently of histamine–induced bronchoconstriction, indicating that different mechanisms are involved (78). In the dose used (1000 units/kg body weight), no effect was observed on partial thromboplastin time. It has been suggested, but it is unclear if aerosolized heparin is an effective steroid–sparing agent for corticosteroid–dependent patients (78,79). Lidocaine has also been evaluated for the treatment of asthma. The use of aerosolized 4% lidocaine in chronic stable asthma had no effect on pulmonary function after a 30–minute inhalation, but a slight bronchodilatory effect was observed in the large airways (79). The results of other studies have shown that an aerosolized preparation of lidocaine reduced bronchoconstriction in a significant proportion of asthma patients; the effect was unrelated to preservatives in the lidocaine preparation (80). Aerosolized lidocaine has been used primarily to

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study nonadrenergic, noncholinergic bronchodilator mechanisms in asthma. Use of lidocaine was effective in reducing airflow limitation in studies using capsaicin and substance P inhalation challenge (81,82). The medication was well tolerated and had no adverse reactions (81,82). Both aerosolized heparin and lidocaine require further double–blind, placebo–controlled trials in patients with asthma before any definitive conclusion can be made regarding overall therapeutic efficacy. The importance of maintenance of normal airway osmolarity in patients with asthma has been demonstrated; hypo–osmolar (fog) or hyperosmolar conditions can induce bronchoconstriction (83). This finding has led to investigations of inhaled natriuretic agents such as furosemide to determine their protective effect in asthma patients (83,84). Inhalation of aerosolized furosemide (40 mg) has been demonstrated to almost completely inhibit exercise–induced and ultrasonically nebulized distilled water–induced bronchoconstriction (83). It was also effective at blocking the early and late airway response to antigen, although it had no effect on methacholine responsiveness after the late antigen challenge (83,84). The effect of aerosolized furosemide has been postulated to occur through its ability to inhibit sodium and chloride exchange, thereby reducing sodium and chloride secretion into the airway lumen. This has been postulated to have an inhibitory effect on luminal mast cell release of mediators or to reduce sensory nerve output from the lumen. Interestingly, a similar effect is not observed with oral furosemide. Further studies are under way to determine if aerosolized diuretics have a role in the treatment of airway inflammation in asthma (83,84). E.

Therapy of Chronic Obstructive Pulmonary Disease

Types of Chronic Obstructive Pulmonary Disease and Management Chronic Bronchitis

Chronic bronchitis and emphysema are two distinct disease processes, but they often coexist in patients with chronic obstructive pulmonary disease (COPD). Chronic bronchitis is defined as “excessive tracheal bronchial mucous production that causes cough with expectoration for at least three months of the year for more than two consecutive years” (85). Subclassifications of this disease have been proposed to include “simple chronic bronchitis” characterized by mucoid sputum production alone and “chronic mucopurulent bronchitis” characterized by recurrent production of purulent sputum in the absence of an underlying process such as bronchiectasis (85). “Chronic asthmatic bronchitis” refers to a condition in patients who experience dyspnea and wheezing after exposure to nonspecific irritants or during an acute viral upper respiratory infection, but who demonstrate only partial reversibility of airflow obstruction after bronchodilator treatment. These patients typically have a long–standing history of cough with sputum production, which progresses to wheezing and reversible airway obstruction over time (85). It is estimated that approximately 20% of adult males have chronic bronchitis. The prevalence of bronchitis in females is rapidly approaching that in

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males, primarily because of increases in the number of female cigarette smokers. Cigarette smoking is the most common cause of chronic bronchitis, although other factors such as occupational and environmental exposures may contribute to this process (85). The pathologic characteristics of chronic bronchitis include hyperplasia and hypertrophy of the mucus–producing glands in the submucosa of the large cartilaginous airways. However, changes also occur in the small, noncartilaginous airways; these include goblet cell hyperplasia, increased numbers of mucosal and submucosal inflammatory cells, increased edema, parabronchial fibrosis, and intraluminal mucous plugging with hypertrophy of smooth muscle (85). In addition to the major contributory factors cited above, familial and genetic factors may contribute to the onset of chronic bronchitis in later life. For instance, children who are exposed to passive smoke inhalation from smoking parents are known to experience more frequent upper respiratory infections, which lead to a higher prevalence of chronic respiratory symptoms such as bronchitis or asthma. Some monozygotic–twin studies have suggested that there may be a genetic predisposition to develop chronic bronchitis, which exists independently of smoke exposure; however, this remains controversial (85). Many patients may have a significant history of cough and sputum production without a significant smoking history. Their cough seems to be present primarily in the winter months and is usually progressive over time. These patients are often overweight and cyanotic, and have no respiratory distress at rest. They tend to have an increased incidence of right ventricular heart failure with peripheral edema and have been referred to as “blue bloaters” because of the cyanosis and edema that develop secondary to right–sided heart failure (85,86). To arrive at the proper management for patients with chronic bronchitis, it is important to obtain a complete medical history and physical exam followed by objective tests to assess the degree of airway obstruction and lung damage. This is accomplished by ordering complete spirometry and a chest X ray. Arterial blood gases at room air are important for determining whether poor prognostic indicators such as significant hypoxia or CO2 retention are present in these patients (87). The most important recommendation in the management of these patients is smoking cessation, which is the only definitive action known to directly influence the progression of chronic–obstructive pulmonary disease. In addition, exposures to secondary smoke and occupational or other environmental irritants should be identified and avoidance recommendations should be made. Short courses of systemic glucocorticoids have been shown to improve symptoms be made if appropriate. Infections occur frequently in these patients. The most common bacterial causes of lung infections are “Haemophilus influenzae” and “Streptococcus pneumoniae.” Antibiotics are usually required in the treatment of acute exacerbations; a 7– to 10–day course of either ampicillin or other suitable penicillin derivative and tetracycline agents is commonly used (85). Pulmonary rehabilitation, which includes breathing exercises and recommendations on proper nutrition, improves lung

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function and quality of life in some patients (85). Bronchodilator drugs such as methylxanthines, ␤2–adrenergic agents, and anticholinergics such as ipratropium bromide (Atrovent) are the most commonly used medications to treat patients with chronic bronchitis (88,89). Theophylline is usually administered orally; in addition to its bronchodilating effects, it has other minor actions such as stimulation of central respiratory drive, increase in systemic venous return and cardiac output, increase in diaphragmatic contractility, and increase in mucociliary clearance (88). ␤2–Adrenergic agents are usually administered as MDIs and are well tolerated (85, 89). In addition to its bronchodilating effects, ipratropium bromide (Atrovent) is helpful in drying out secretions (85,89). For this reason, this drug may be useful as a first–line agent in the treatment of chronic bronchitis. Unlike asthma, combined use of ␤2–agonists and anticholinergics via MDI and spacers may be more effective than either agent alone in patients with COPD (41,89). The use of inhalant and systemic glucocorticoids in patients with chronic bronchitis remains an area of controversy; few studies have demonstrated significant therapeutic benefits of these agents in patients with COPD (90,91). However, if some evidence of reversible obstruction exists, addition of an inhalant corticosteroid to bronchodilator therapy improves lung function, alleviates symptoms, and decreases the number of exacerbations in COPD patients (90,91). Some concern has been raised that long–term therapy with inhalant and systemic corticosteroids weakens respiratory muscles. Conflicting data from clinical studies make it unclear whether a significant risk of respiratory muscle impairment exists in these patients. Because no general consensus is apparent among practitioners on the value and/or proper utilization of corticosteroids in COPD patients, it is prudent to limit the use of these agents in these patients by using inhaled corticosteroids (which are less likely to cause systemic side effects) (92,93). Bronchopulmonary drainage may be helpful in patients with copious mucous secretions, especially when the cough mechanism is ineffective. Clearance of secretions using mucolytic agents or expectorants may be effective in patients with COPD; mucolytic agents such as iodides and guaifenesin may occasionally be effective and are most often used empirically to identify those specific patients who will respond (85,94). Supplemental oxygen is the only therapy that significantly lowers mortality rates (85,87). Emphysema

Emphysema is defined as “hyperinflation of the air space distal to the terminal bronchi with destruction of the alveolar septa” (85). Prevalence data of emphysema (derived primarily from postmortem surveys) have demonstrated that the incidence of emphysema increases in the fifth decade through the seventh decade and that at death two–third of adult males and one–fourth of females have well–defined emphysema. Pathologically, emphysema is classified by the pattern of involvement of the acini in the distal lung up to the terminal bronchi (85). The two most common types are “centriacinar emphysema,” which involves the respiratory bronchi and alveolar ducts in the center of the acinus, and “panacinar

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emphysema,” that involves the entire acinus; usually both types of acinar involvement are present. Because the lung has a large functional reserve, many acinar units must be affected before lung dysfunction can be detected by spirometry (85). Patients with centriacinar emphysema have high ventilation perfusion mismatches because of absent capillaries but persistent ventilation of the acini. In panacinar emphysema, extensive involvement of the central and peripheral portions of the acinus results in reduction of alveolar capillary gas exchange and loss of elastic recoil of the lung (85). Causal factors for chronic bronchitis such as smoking and occupational and environmental exposures to irritants or fumes also increase the risk for emphysema. About 1% to 2% of patients who develop emphysema have a genetic disorder characterized by deficient or completely absent serum levels of a protease inhibitor, ␣1–antitrypsin; in healthy persons, this protein is synthesized by liver parenchymal cells and transported by mononuclear phagocytes to the lung interstitium and parenchymal epithelial lining of the lower respiratory tract (85). ␣1–Antitrypsin is an important and effective inhibitor of trypsin, neutrophil elastase (NE), and other proteolytic enzymes. Experimental evidence has shown that the structural integrity of lung elastin depends on ␣1–antitrypsin, which protects it from proteases (e.g., NE) released from leukocytes. It is now commonly believed that these proteolytic enzymes, which are derived from neutrophils and alveolar macrophages, may cause emphysema even in subjects with normal circulating levels of antiproteases (85). Many patients with genetic emphysema have a homozygous form of ␣1–antitrypsin deficiency with plasma and lung levels less than 15% of normal, which increases the risk for development of emphysema (85). Alterations in several genes are associated with ␣1–antitrypsin deficiency; the most common genes involved are the Z and S genes. Serum levels of ␣1–antitrypsin approach zero in patients who are ZZ or SS homozygous; such patients are most likely to develop severe panacinar emphysema in their third or fourth decade of life (85). Patients with COPD in whom emphysema is predominant have a long history of dyspnea upon exertion with a slight nonproductive cough. Infections occur infrequently compared to chronic bronchitis. These patients often have an asthenic bodybuild with significant weight loss (94). They use their respiratory accessory muscles during expiration; patients often have a prolonged expiration while pursing their lips, which helps to remove trapped air from their acini. These physical features exhibited by patients with emphysema have resulted in the descriptive term “pink puffer” (85,94). The usual cause of death in these patients is right–sided heart failure and hypercapnic respiratory failure. Management of patients with emphysema is similar to that for chronic bronchitis although therapeutic intervention is often less effective. Prior to developing a therapeutic regimen, it is important to assess the severity of airway obstruction with spirometry, and the degree of arterial hypoxemia while breathing room air using arterial blood gases (87). It is important to counsel patients early to stop smoking. Pulmonary rehabilitation programs that may be helpful include the recommendation of specific regimens for nutrition as well as exercises to maintain respiratory muscle strength (because respiratory

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fatigue usually leads to CO2 retention). Medications employed include methylxanthine agents such as theophylline, inhalant bronchodilators, and anticholinergic agents (88,89). These latter two agents are administered as MDIs and as solutions in simple nebulizers. Inhalant corticosteroids are helpful in improving symptoms in patients who show partial reversibility of airway obstruction. The use of chronic systemic corticosteroids is not indicated in patients with predominant emphysema; however, short courses of corticosteroids may be useful in reducing relapses in patients who experience frequent exacerbations of COPD (90,91). Continuous oxygen therapy in patients with PO2s less than 55 mmHg is recommended to prevent or delay progression to pulmonary hypertension and right–sided heart failure (85,87). Drugs and Agents Used as Aerosols for Therapy of Chronic Obstructive Pulmonary Disease Anticholinergic Agents

Cholinergic fibers in the airways are most important for controlling airway smooth muscle and bronchomotor tone. Anticholinergic agents developed in the United States for management of obstructive lung disorders include atropine sulfate and ipratropium bromide (Atrovent). More recently, tiotropium bromide (Spiriva), a long–acting, once–a–day inhaled anticholinergic has been approved in the United States for the treatment of moderate–to–severe COPD (58). Other agents are available in Europe, and some of these are being investigated in the United States (88,89). Inhalant ipratropium bromide produces peak blood levels after one to two hours; levels decline after four hours. A number of studies have compared bronchodilator actions of anticholinergic agents with those of ␤2–adrenergic agents in patients with either asthma or COPD (41,58). In patients with asthma, anticholinergic agents are slower acting than ␤2–adrenergic agents and produce less bronchodilation (41,58). However, in patients with COPD, most studies have demonstrated that anticholinergic agents such as ipratropium bromide are more effective bronchodilators than selective ␤2–agonists; therefore, practitioners usually treat patients with COPD initially with ipratropium bromide and tiotropium bromide (Spiriva) (58). Tiotropium bromide is unique in that it has prolonged M(3) muscarinic receptor blockade (58). It has been demonstrated to improve symptoms such as dyspnea, lung volumes, quality of life, and frequency of exacerbations. Clinical trials have demonstrated that tiotropium was statistically a superior or equivalent bronchodilator compared to ipratropium administered four times a day and salmeterol administered twice a day. Overall tiotropium bromide is well tolerated with the most common side effect being dry mouth (58). Beta-Adrenergic Agents

As discussed above, ␤2–adrenergic agents are less effective bronchodilators than anticholinergic agents in patients with COPD. Several studies have reported that concomitant use of different classes of bronchodilators such as ␤2–agonists and anticholinergics results in additive bronchodilation.

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Therefore, it is not unreasonable to use these agents in combination with patients with COPD. No unfavorable interactions have been demonstrated because both drugs are generally well tolerated (85). Corticosteroids

Chronic use of systemic corticosteroids in patients with chronic bronchitis and/or emphysema is not recommended (90,91). As mentioned earlier, short courses of an inhaler MDI or DPI corticosteroid, nebulized budesonide, or systemic corticosteroids may be beneficial for patients with exacerbations of COPD and prevent frequent relapses. Furthermore, some patients present with apparent COPD and severe airway obstruction by spirometry without improvement after bronchodilator therapy. A reversible component of airway obstruction (asthma) can be excluded by administering a short course of systemic corticosteroids to determine if there is subsequent improvement in FEV1. Few studies have demonstrated that inhalant corticosteroids delay progression of airway disease in subjects with COPD (90,91). Oxygen therapy is critical for treatment of patients with severe chronic bronchitis and/or emphysema (85,87). Intermittent administration of oxygen is beneficial. However, continuous use of O2 (more than 19 hours a day) is much more effective than intermittent administration in reducing progression to pulmonary hypertension and right–sided heart failure (85,87). Antiproteinases

The ␣1–proteinase inhibitor Prolastin, an ␣1–antitrypsin derived from human plasma, is now available for treatment of patients with ␣1–antitrypsin deficiency (85). Intravenous administration at a dose of 250 mg/kg four times a day has been demonstrated to be effective in selected individuals. An aerosol preparation to be inhaled once or twice a day is currently being investigated. Because this drug is expensive, it has been reserved for patients with the homologous form of ␣1–antitrypsin deficiency who are more likely to develop emphysematous lung disease (85). Many believe that evidence for emphysema should exist before therapy is administered. Prolastin has not been shown to be effective in patients with emphysema induced by smoking or other causes aside from ␣1–antitrypsin deficiency. It also has not been demonstrated to be effective in patients with heterozygous ␣1–antitrypsin deficiency or in individuals with normal or moderately reduced ␣1–antitrypsin levels (85). Lung transplantation has been established as a form of therapy for treatment of interstitial lung disease secondary to ␣1–antitrypsin deficiency–induced emphysema. One–year survival rates approaching 70% have been reported. Gene therapy is also being investigated as a technique to replace the ␣1–antitrypsin gene in patients with ␣1–antitrypsin deficiency (85) (see the section on Gene Therapy). Mucolytics

Other treatment modalities for patients with COPD include mucolytic agents, which have been used more extensively in patients with chronic bronchitis (86).

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Ambroxol and sobrerol are mucolytics in current use in Europe; the possible value of these agents and other mucolytic drugs for the treatment of COPD is unclear (86). One mucolytic agent that may be beneficial is acetylcysteine (Mucomyst). It can be administered using a simple nebulizer. Although this aerosol can be irritating to some patients, it is generally well tolerated. Long–term administration may reduce the frequency of exacerbations of chronic lung disease in patients with mucous hypersecretion that is intermittent or persistent. A suggestion has been made that acetylcysteine acts as a free–radical scavenger and prevents lung damage by oxidants in patients with chronic or recurrent infections or in those exposed to air pollution or cigarette smoke (86). Iodine has been shown to be a very effective mucolytic agent and was formerly available as an aerosol mucogenic preparation referred to as “Tergemist” in the United States. Studies have shown that this aerosolized iodide solution, like current iodide solutions available for oral use, may be as effective a mucolytic agent as acetylcysteine (86). Cysteine derivatives are mucolytic agents that are available only in Europe; these agents rupture disulfide bonds of glycoproteins in mucous secretions, dissociate mucous side chains by ionization, and solubilize their remaining fragments (86). Cysteine derivatives can be administered by inhalation in a simple nebulizer. Currently, the only mucolytic agent available in the United States for aerosolized use is acetylcysteine (Mucomyst) (86).

IV.

Cystic Fibrosis

This inherited disorder, which affects about 40,000 Americans, is the most common genetic disease in Caucasians. It is a complex disease affecting many organ systems including the lungs, pancreas, male genital tract, sweat glands, liver, and intestine. The major cause of morbidity is lung disease, which is responsible for more than 95% of mortality (95). In this disorder, a thick mucus is produced that damages the lungs and causes respiratory failure and death, usually before the age of 30 years. Until recently, the cause of CF was unclear and therapy did little more than help patients remove mucus from their lungs by coughing. In the last decade, the genetic deficit responsible for CF has been discovered and its pathophysiology defined (95–97). CF, an autosomal–recessive disease, is only expressed when the individual is homozygous for a defective gene located on the long arm of chromosome 7. The membrane protein made by this gene is the CF transmembrane conductance receptor (CFTR). This membrane protein serves as the chloride channel in cell membranes within the lungs and other organ systems and regulates transport of sodium and chloride ions. Congenital absence or a defect of CFTR in the lungs leads to decreased chloride secretion and increased reabsorption of sodium; the thick mucus produced in the dehydrated epithelium predisposes the patient to develop impaired mucociliary clearance and bacterial infections. It is not clear what specific changes in the lung cell membranes are responsible for attachment, colonization, and infection with “Pseudomonas aeruginosa”

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and other pathogens (Staphylococcus aureus, H. influenzae, Pseudomonas cepacia). Allergic bronchopulmonary aspergillosis has been reported to develop in up to 15% of CF patients. The most important pathogen in CF is P. aeruginosa; much of the pulmonary damage resulting in exacerbations of CF as well as in morbidity and mortality is due to chronic infection with this organism (98). The bacterial infection induces a host inflammatory response that, together with the bacterial products elaborated by this organism, results in destruction of the lung extracellular matrix similar to that, which occurs in emphysema. The protease NE is released by neutrophils recruited to the lung during the immune response and is believed to be responsible for most of the damage to lung structure. This protease also damages neutrophils and cleaves complement, resulting in an impaired ability to destroy lung pathogens. Secretory leucoprotease inhibitor (SLP) and ␣1–antitrypsin deficiency (AAT) are NE inhibitors present in the serum and lungs of healthy individuals, and serve as the primary protection against NE activity by irreversibly complexing to NE and blocking its proteolytic action. In CF patients, levels of natural airway NE inhibitors in the lungs are within normal concentrations, but are insufficient to neutralize the large amounts of NE released by neutrophils (99). Large amounts of DNA derived from damaged neutrophils and other anti–inflammatory cells in the lungs contribute significantly to the high viscosity of the thick mucus. The thick mucous secretions and recurrent infections experienced by CF patients lead to severe bronchiectasis of the lungs, which further predisposes these patients to develop infections and inspissated mucus plugging. A.

Therapy

Aggressive use of antipseudomonal antibiotics and physical therapy has been primarily responsible for the increase in the median age of death from under 10 to 28 years in the last three decades. Antibiotics reduce the influx of neutrophils to the lungs, which slows the rate of progression of proteolytic lung destruction by NE (100). This therapy and the introduction of newer therapies discussed below have resulted in increased survival estimates showing that the median age of death is increasing; the number of CF patients reaching adulthood is increasing by 10% a year (98). Intravenous antibiotics active against P. aeruginosa are used early on, when symptoms of an impending exacerbation first appear, and are usually given for two weeks or longer. Suggested antibiotic regimens include an aminoglycoside [e.g., tobramycin sulfate (Nebcin®) or amikacin sulfate (Amikin®)] administered with an extended–spectrum penicillin [e.g., ticarcillin disodium (Ticar®) or piperacillin sodium (Pipracil®)]. Established Pseudomonas infections are rarely eradicated by antibiotics, but chronic infections can usually be well controlled (98). Problems arise when the pseudomonal strains become resistant to the antibiotics. The popularity of aerosol administration of antibiotics is increasing; this route offers several advantages over the intravenous route: (i) The direct application of antibiotics to sites of action in the lung is very efficient as 50% less drug is used compared to intravenous injection, and (ii) systemic adverse reactions of aminoglycosides are avoided.

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In 1981, the first controlled trial of inhaled antibiotics showing effectiveness and safety was published. In this study gentamycin, an aminoglycoside, and carbenicillin, a penicillin, were administered by nebulization (101). Subsequently, a few studies have included inhalant penicillins. Two aminoglycosides, tobramycin and amikacin, are recommended in patients with CF for inhalation using a jet nebulizer and are also drugs of choice for intravenous administration (98,102). Inhaled colistin sulfate (Coly–Mycin S) is used less frequently. Other antibiotics investigated for the inhalation route in CF patients include neomycin and ceftazidime (102). The specific indications for aerosol antibiotics have not been established; it is unclear whether they should be used for deteriorating patients (exacerbations) and/or patients who require frequent intravenous antibiotics, or for patients with mild–to–moderate disease. Guidelines for proper utilization of nebulizers (discussed earlier in this chapter) must be followed in order to ensure that jet nebulizers are efficiently delivering the aerosolized antibiotics (102). Treatment of allergic bronchopulmonary aspergillosis requires systemic corticosteroids (98,103); this is the only clear indication for corticosteroids in patients with CF (75). Nebulized delivery of the antifungal amphotericin B may provide effective prophylaxis against invasive aspergillosis (102). Inhaled bronchodilators (␤2–adrenergic agonists; anticholinergics) may be useful in patients who wheeze, demonstrate an improved FEV1 or PEFR in response to a trial of bronchodilators, or have a positive methacholine challenge indicative of airway hyperresponsiveness or asthma (96). Systemic corticosteroids should be used in CF patients with documented asthma, who fail to respond to bronchodilators. Several promising approaches in the treatment of CF are being pursued. Preliminary data suggest that the daily administration of two NE inhibitors, ␣2–proteinase inhibitor (AAT) and recombinant SLP, in aerosol form reduces lung inflammation and improves lung function; this anti–inflammatory therapy may slow the progression of lung damage (99,100). Placebo–controlled studies evaluating the effects of systemic administration of other anti–inflammatory agents, corticosteroids, and nonsteroidal anti–inflammatory agents (i.e., ibuprofen) have not been completed (96). As with other prophylactic or acute therapies for CF, anti–inflammatory therapy is more likely to be beneficial in younger patients without established lung disease (104). Therefore, inhalant corticosteroids may have a significant role in prophylaxis and therapy of CF patients early in the course of their disease, before lung damage has occurred. Another potentially useful treatment is administration in aerosol form of recombinant deoxyribonuclease; this enzyme digests the DNA in the thick mucus, enhances mucociliary clearance, and improves lung function (96,105–107). A small but significant improvement in FEV1 and forced vital capacity has been demonstrated. Patients also experienced an enhanced quality of life; decreased cough frequency and severity; and improved expectoration, dyspnea, energy levels, and appetite. The effects of DNase treatment are greater in patients with more severe lung disease than in those with milder deficits in pulmonary function. Aerosol NE inhibitors and DNase appear to be safe and are well tolerated.

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The cost to administer one year of DNase therapy is about $10,000 per patient. Gene therapy and other novel treatments under investigation also promise to be expensive. It is argued that these costs can be justified because effective therapy will decrease the rate and length of hospitalization as well as eliminate the need or decrease the requirements for antibiotics and other therapy (98). Agents that act on ion channels and either block sodium reabsorption (e.g., amiloride) or increase chloride secretion (e.g., ATP and uridine triphosphate) (108–110) are in the early stages of investigation. Genetic replacement therapy is also being explored. In animal experiments, copies of the normal gene for CFTR are inserted into genetically altered, noninfectious cold viruses (adenoviruses). The altered virus carrying the gene copies is administered as an aerosol or instilled directly within the lungs. The inserted gene can be detected for up to six weeks in lung epithelia (111). In this manner, it may be possible to produce significant amounts of the CFTR and thus repair the damage that resulted from the gene mutation common to these patients. The value of a noninfectious adenovirus as a delivery vector for gene therapy is being investigated in humans; the efficacy and safety of adenovirus–CFTR DNA replacement therapy following administration of an aerosol containing adenovirus–CFTR DNA to the nasal epithelia of CF patients was demonstrated in three patients (95); these experiments are a prelude to studies investigating the effects of pulmonary inhalation of this aerosol in patients with CF. Many questions remain unanswered, including the functional half–life of the inserted DNA.

V.

Gene Therapy

The concept of gene therapy is to replace an absent or defective gene that has been identified as causing a serious genetic disorder. Specific vectors are used to carry the genetic information into the affected cells. Two genetic lung diseases may be successfully treated using this approach: (i) CF in which the protein CFTR (the chloride channel) is absent or defective from lung membranes (see the section on CF), and (ii) the genetic form of emphysema in which ␣1-antitrypsin is deflective or absent. Aerosol inhalation offers a number of advantages for CFTR gene therapy. This route of administration: (i) provides direct access to the airway epithelial cells that normally express CFTR; (ii) limits the movement of the vector containing replacement DNA; and (iii) prevents entry into the systemic circulation. Both in vivo and in vitro studies show that correction of the genetic defect in only 6% to 20% of airway epithelial cells is sufficient to restore CFTR function. A variety of vectors have been investigated for aerosol administration. At present, a noninfective adenovirus has been used most frequently as the vector carrying inserted DNA coding for CFTR as well as ␣1-antitrypsin. Other vectors under investigation for aerosol administration include adeno–associated viruses, retroviruses, and liposomes (a nonviral vector). The adenovirus–associated

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vector appears to be able to infect nondividing cells and offers the potential for stable integration into the host cell genome, which may result in long–term gene transfer and possibly a “cure” (95). The value of retrovirus vectors is less promising because they require dividing cells for successful gene transfer. Nonviral vectors such as liposomes (cationic lipids) are relatively nontoxic and do not elicit an immune response, which may offer potential advantages over viral vectors. In addition, liposomes would be easy to purify and prepare for large–scale production. ␣1–Antitrypsin replacement therapy inhibits the inflammatory lung damage produced by NE and thus acts as an anti–inflammatory agent. Aerosol administration of ␣1–antitrypsin is safe, convenient, and effective in restoring adequate concentrations of ␣1–antitrypsin within the lung (111). The same adenovirus vector used for CFTR gene transfer experiments has been investigated in transgenic animals to insert the DNA coding for ␣1–antitrypsin into liver cells that synthesize this protease inhibitor. Aerosol replacement therapy for ␣1–antitrypsin deficiency is also discussed in the section on chronic obstructive lung disease.

References 1. 2. 3. 4. 5.

6. 7. 8.

9. 10. 11. 12. 13.

14.

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PART III:

PHARAMACEUTICS AND PHARMACEUTICAL TECHNOLOGY

9 Atomization and Nebulizers

RALPH W. NIVEN

ANTHONY J. HICKEY

Innoven, Half Moon Bay, California, U.S.A.

School of Pharmacy and Medicine, University of North Carolina, Chapel Hill, North Carolina, U.S.A.

I.

Introduction

The principles of atomization are central to two of the major categories of pharmaceutical aerosol products, pressurized metered dose inhalers and nebulizers. There have been major developments in each of these technologies (Chapter 15) over the last decade, but the basic principles have not changed and remain fundamental to understanding mechanisms of droplet formation from these devices. Sprays and aerosols can be found in many forms and are often used routinely in our daily lives. Examples include rain drops, ocean spray, ink jet printing sprays, fuel injection aerosols, and of course inhalation aerosols. Most industrial uses attempt to control the nature of the sprays to fit a particular application. Insecticides sprayed from aircraft must be of a composition and size that will ensure efficient and homogeneous deposition of droplets onto crops, ink jets should provide a coherent spray with minimal extraneous droplets, fuel injectors should efficiently atomize the fuel to optimize combustion, and inhalation aerosols should be of dimensions that ensure deep lung penetration. These diverse criteria place significant demands on those responsible for the design and engineering of atomizers to fit an intended purpose. This objective is further complicated by the poor understanding of the physical phenomena that govern droplet production and spread. The process of producing sprays is termed atomization—the conversion of bulk liquid into small droplets. This phase dispersion requires energy to 253

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produce both surface area and transportation of the atomized fluid. The available internal energy associated with a fluid ejected from a nozzle is usually insufficient to convert the jet into droplets of a size range that is suitable for most applications. Therefore, assistance is often provided by a high-velocity gas stream that can pneumatically disintegrate a liquid jet emerging from a nozzle. Thus sprays can be produced and controlled for most industrial applications (1,2). Nebulizers can be considered an extension of these two-fluid or pneumatic atomizers, and will be the focus of discussion in this chapter. The objective is to give an overview of the mechanisms that control droplet, production, and breakup in atomizers and relate this information to nebulizer operation. Finally, some unique aspects of nebulizers that influence performance, when used to deliver aerosols to the lungs, will be discussed.

II.

Atomization

The basis of air-jet atomization is the interaction of a high-velocity air stream with that of a relatively slow-moving flow of liquid. The bulk liquid phase is dispersed into the gas phase and converted into small drops (3). The physical forces governing the process are surface tension and viscosity versus aerodynamic forces. Surface tension can be considered a consolidating influence that attempts to minimize the production of surface. Additionally, liquid viscosity exerts a stabilizing influence by opposing any change in the shape of droplets as they are produced. Aerodynamic forces acting on the liquid surface promote disruption by exerting an external force on the bulk liquid. The final spray depends not only on the primary droplets produced, but also on to what extent these droplets are further disintegrated (2,4,5). A general theory that describes the formation and breakup of droplets in atomizers has not been developed, although an understanding of the process of atomization in certain types of applications, such as fuel injectors, is gradually being developed. However, this necessitates a thorough knowledge of fluid and aerodynamics and the availability of mainframe computer time to evaluate complex mathematical models. To put this in perspective, it has been estimated by Reitz (6) that it would take a current Macintosh (or IBM!) computer some 28 years to quantitatively describe a spray emerging from a plain-jet orifice over a period as short as a few milliseconds. Despite this complexity, a discussion of the atomization process divided into the stages of (A) bulk fluid dispersion, (B) droplet breakup, (C) coalescence, and (D) the parameters governing droplet size will be instructive. The information will be related to nebulizer function and aqueous solutions wherever possible. A.

Bulk Fluid Dispersion

The general description for the process of droplet formation from liquid columns or jets is based on theory derived by Rayleigh (7) and applied by Castleman (8). Castleman’s work was built upon the premise that ligament

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formation is a discrete step in the reduction of a bulk mass of liquid to individual droplets, and some excellent photographs of the disintegration of laminar jets are shown to illustrate his points. The characteristic “hook” region, for example, is well defined. A sketch is shown in Figure 1 with the regions defined. This general description is still valid, although the theory is inapplicable for pneumatic and hydraulic nozzle systems because only laminar flow conditions are considered (9). The kinetics of the process have been summarized in more detail by Lapple et al. (2), as shown in Table 1. The relationship between laminar and turbulent jets is not well defined. One convenient means of distinguishing the nature of liquid jets is to plot data of the jet breakup length, or some stability parameter (10), against the jet exit velocity (Fig. 2). This shows that several transitional states exist. After first increasing, the jet breakup length decreases and then increases again as jet velocity is increased. The critical velocity (peak) will vary depending on the initial stability of the jet. The stability in turn is dependent upon the influence of primary and secondary atomization effects (11). Primary atomization refers to the breakup process affected only by internal forces, whereas secondary atomization includes the action of external aerodynamic forces (see Fig. 2). Thus the “state” of the ambient atmosphere combined with the level of internal disturbance within the liquid itself will dictate the type and extent of disintegration that occurs. Unfortunately, in two-fluid atomizers turbulent airflow is almost always combined with turbulent liquid flow, and as mentioned above, such a system defies any yet devised theoretical framework. However, to begin to appreciate the mechanisms involved, it is useful to consider a visual analogy. One that is applicable is the observation that spray is produced from wave crests on the ocean.

Depiction of the breakup of a jet emerging from a nozzle in quiescent air. Regions of instability are first generated, which is followed by extension of liquid elements and finally disruption. Source: Adapted from Ref. 8.

Figure 1

256 Table 1

Niven and Hickey Kinetics of Drop Formation

The extension of a bulk liquid into sheets, jets, films, or streams by accelerating the liquid in some prescribed manner (e.g., via nozzle or from a rotating disk). The initiation of small disturbances at the liquid surface in the form of local ripples, protuberances, or waves. The formation of short ligaments on the liquid surface as the result of fluid pressure or shear forces. The collapse of the ligaments into drops as the result of surface tension. The further breakup of the droplets as they move through the gaseous medium by the action of fluid pressure or shear forces. Source: From Ref. 2.

This mechanism may be analogous to what occurs within atomizer jets, and this is partially borne out by a study conducted by Hanratty and Engen (12), in which they examined the interaction of a turbulent airstream moving parallel to a flowing water surface. They arbitrarily divide the observed effects into several categories. At low relative velocities between the gas and liquid, such as those found on the ocean, ripples and roll waves are observed. As the velocity of the gas across the fluid is increased, the authors describe the droplets being torn from the liquid surface. This extreme is catalogued as dispersed flow. Here the Reynolds number (Re) of the gas flow exceeded 40,000, and therefore, the flow regime was turbulent and would readily generate spray. It also was noted that the Re of the liquid flow was important, and it proved more difficult to induce dispersion because the Re of the liquid was reduced. This is logical because it would be expected that liquid turbulence, manifested by surface disturbances protruding above the mean surface height, would be more readily entrained in the airstream than from a smooth surface where energy, through drag, would be necessary to generate the initial disturbances in the fluid.

The stability of a laminar jet versus the exit jet velocity. The laminar, or Rayleigh, region is described between A and B. A transitional region then runs from B to C, and then turbulent jets exist as the jet velocity increases until the spray is fully developed. Figure 2

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Within air-blast atomizers, the relative velocities between gas and liquid can be substantial. Velocities in nebulizers may reach or exceed sonic velocity (⬇350 m/sec) and hence the simple wave analogy is an oversimplification. The results nevertheless confirm that if the water is turbulent and possesses significant kinetic energy, then the atomization process will be enhanced. The disruptive process also has been nicely illustrated by Rizk and Lefebvre (13), who photographically examined sheet breakup by air-blast atomization (Fig. 3). At an air velocity of 55 m/sec, their photographs illustrate wave formation, ligament formation, and stripping, further fragmentation, and disintegration. Rizk and Lefebvre also demonstrate that as air velocity is increased, the liquid sheet disintegrates earlier, the ligaments formed are thinner and shorter, and the droplets become smaller. Many studies only consider parallel or coaxial air-liquid flow, and the manner in which fluid is presented to the airstream will impact the atomization profile (14). Most internal mixing atomizers supply liquid through a central nozzle, which will be surrounded by one or more gas inlets (Fig. 4A). This situation is reversed in many nebulizer nozzles. The fluid is aspirated in an annular ring around the air-jet emerging from a venturi and interacts while moving normal to the gas flow (Fig. 4B). The resultant interaction will be heavily dependent upon the rate of movement and the angle of entrainment of the liquid into the gas. The specifics of this type of interaction of gas and liquid in nebulizers have not been examined in detail (11,15).

Figure 3 Breakup of a liquid sheet in turbulent airflow. Wave formation, ligament stretching, stripping, and disruption or bursting are well defined. Source: Adapted from Ref. 13.

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Figure 4 (A) Types of nebulizer nozzle: Internal mixing: The gas flow interacts with the liquid prior to leaving the exit port. External mixing: Gas and liquid interact after both leave nozzle assembly. Note the use of a baffle or impinger. Liquid flow is induced by the expanding gas flow only. (B) Typical atomizer nozzles. Internal and external mix nozzles are similar to the nebulizer counterparts only in name. Usually the gas flow surrounds the liquid, no baffle is used, and the liquid feed is forced through the feed jet(s).

B.

Droplet Breakup

Once ligaments and large droplets are produced during primary atomization, unstable droplets will undergo further disintegration in an air-jet. Secondary disintegrations will occur depending upon the balance of forces within and surrounding the droplets. If we consider a single large droplet produced by the exertion of aerodynamic pressures on a liquid stream, these same aerodynamic forces, together with surface tension and viscosity, will be responsible for determining whether the droplet will undergo further breakup or remain intact. The balance of these forces has been described by Klusener (16) such that the internal pressure, PI, at any point on the surface of the drop is given by

Atomization and Nebulizers PI ⫽PA + P␴ ⫽ constant

259 (1)

where PA is the aerodynamic pressure and P␴ is the surface pressure. Thus, a droplet will remain stable in air as long as the competing forces are in equilibrium. If the aerodynamic force at one locus on the droplet surface is increased, this can be offset by a corresponding decrease in the local surface tension to maintain the internal pressure at the surface. If the aerodynamic forces become large relative to P␴, then disruption of the drop will result. This is a useful and straightforward way to envision the basic forces dictating the stability of a formed drop (Fig. 5). For water, droplet deformation is determined by the combination of aerodynamic and the surface tension forces (viscous forces being relatively small). This is given a quantitative basis by calculating values for the ratio of these forces, which is the basis of the dimensionless Weber number (We) We=

ρAU R2 D σ

(2)

where ␳A is the density of air, UR is the relative velocity between the droplet and the airstream, D is the diameter of the droplet, and A is the surface tension

Forces acting on a droplet. Disintegration is typically the combination of both aerodynamic (external), and surface tension and viscosity (internal) forces. The nature of disruption will differ depending on how the forces are applied to the drop.

Figure 5

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of the liquid. The larger the Weber number, the greater the probability of droplet breakup. The condition where the aerodynamic forces are equal to the surface tension forces is the point of breakup. This gives the critical Weber number, Wecrit as We crit =

8 CD

(3)

where CD is the coefficient of drag (⬇0.45 for turbulent conditions) (17). Droplets will disintegrate when the calculated value of We exceeds the Wecrit. The value of 0.45 for the drag coefficient suggests that the critical value will be ⬇18. However, this value is partly dependent upon the manner in which the aerodynamic forces are applied to the drop. A constant force applied to a drop can give a value of 22. A suddenly applied force can lower the Wecrit to 13 for liquids such as water (18). The size of droplets susceptible to secondary atomization can be appreciated from a single calculation using Eq. (2) with water as the example fluid. Assuming We ⫽ 20; Ur ⫽ 25,000 cm/sec; ␳ ⫽ 1.2 ⫻10⫺3 g/cm3, and ␴ ⫽ 73 dynes/cm gives D ⫽ 19.5 ␮m. It is, therefore, easy to see that shatter of preformed droplets within nebulizers can readily occur down to droplet sizes of less than 20 ␮m. A variety of descriptions exist for the types of deformation that can occur. “Bagging,” or “ballooning,” results when the drops are subjected to a steady airstream (a concave “front” to the airflow) (19). Transient bursts of airflow result in droplets forming a convex front to the airflow (20). Excellent photographs of these effects can be found in the papers of Lane (20) and Hanson et al. (21), but there are no explanations for the reversal in breakup pattern. C.

Coalescence

The general phenomenon of coagulation is an area of active study in fields as diverse as emulsion technology and cloud formation, but the subject has been largely neglected in the literature surrounding nebulizer operation. Coalescence is defined here as the coagulation of aerosolized droplets—that is, the combination of one or more droplets to form larger particles. The phenomenon occurs via a number of mechanisms, but those of interest can be considered as forced rather than spontaneous coalescence (22). Two of the mechanisms that are likely to have a dominant role during the formation of an aerosol within a nebulizer are kinematic and turbulent coalescence. The former arises through the relative difference in velocities between large droplets and small droplets upon exposure to external force fields. Turbulent coalescence will occur in turbulent airflow and arises through the sudden changes in the external dynamic forces that droplets will encounter, thereby changing their direction and causing collisions. Thus, the velocity gradient, direction, and relaxation of the airstream emerging from the exit nozzle will dictate the degree of coalescence. The probability of coalescence will decrease as the distance from the atomization center is increased and as the aerosol concentration in the continuous phase decreases. Additionally, because of evaporation, the droplet will be reduced in size, thus decreasing the probability that droplets will exceed the critical Weber number, and therefore, be subject to further disintegration. The redirection of air/fluid flow at the baffle

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surface will also contribute to the turbulence of the region. For a summary of mathematical treatments of the phenomena, readers are directed to the text of Fuchs (23). It should be recognized, however, that quantitative information on the influence of coagulation during nebulization is minimal (24). D.

Droplet Size and Prediction

Droplet size is often the most important criterion upon which atomizer performance is based, and there are volumes of work devoted to the measurement and size analysis of droplets produced by atomizers (1,2,22,25–31). Thus, the discussion here will be cursory, particularly because the primary spray emerging from an atomizer may have little bearing on the spray that ultimately emerges from a nebulizer. In addition, different measuring techniques will give different sizes for the same hypothetical droplets; the size of droplets can vary significantly within short time frames, and often there are shortcomings in the methodologies and analyses employed to measure droplet size (2). The pertinent information with respect to any future nebulizer design is contained within the empirical and semiempirical equations that have been developed to predict droplet size. These help illustrate what physical properties influence droplet size and what criteria are important for nozzle design. The often-cited work of Nukiyama and Tanasawa (32) is a good starting point because the atomizers they studied resemble those used in nebulizers today (without a baffle or impinger). Their work was one of the first attempts to identify factors controlling droplet size during air-blast atomization and they developed a model

  µ  σ  d = 585 ×   + 597 ×    UR × ρ   σ × ρ

0.45

 Q  ×  1000 × 1  Qa  

1.5

(4)

where d is the diameter of the drops (␮m ⫽ Sauter mean diameter), ␴ is surface tension (dynes/cm), Ur is the relative velocity between gas and liquid (m/sec), ␳ is the liquid density (g/cm3), ␮ is the liquid viscosity (dyne-s/cm2 ⫽ poise), and Qa and Qr are the flow rates (cm3/sec) for the gas and liquid, respectively. For water at 20° C this equation reduces to d=

Q  4998 + 904411 ×  1  Ur  Qa 

1.5

(5)

If we consider a “typical” nebulizer with Ur ⫽ 250 m/sec, Ql ⫽ 1 cm3/sec, and Qa ⫽ 200 cm3/sec, we obtain an expected diameter of ⯝600 ␮m for the primary spray. This is high but not entirely unrealistic depending upon how and where the size measurement is made. That Qa/Q1 and Ur are important can be seen in Figure 6; however, it is well to recognize that Eq. (5) is not dimensionally correct and does not include a parameter governing nozzle dimensions. Many other studies have revised and extended the use of Eq. (4) to different applications (33–35). Other observations include those of Licht (36), who determined that the maximum drop size that can be produced for a wide

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Relationship between droplet diameter and the ratio of volumetric gas to liquid flow (Qa/Q1) for different gas velocities. As both the velocity and gas to liquid flow ratio are increased, the droplet diameter is decreased. Source: Adapted from Ref. 34. Figure 6

range of atomizer designs is approximately three times that of the mass median size. One atomizer design that resembles those found in the internal mixing nozzles of nebulizers was designed by Gretzinger and Marshall (15). This pneumatic-impinger nozzle consisted of a central gas jet surrounded by annular fluid jet and an impinger situated at the exit. However, to position the impinger, a metal rod was passed through the gas orifice, and in this respect the device differs from nebulizers. Nevertheless, they were able to show that droplets with mass median diameters down to 5 ␮m could be generated by this type of nozzle without requiring excessive power input. This suggests that the gross or primary aerosol produced by nebulizers could be dispersed more efficiently. Most of the major findings from different investigations have been summarized by Kim and Marshall (35). They assume that the droplet size is a function of various factors and combine them into a series of dimensionless groupings  D D U 2ρ Q µ1 µ ρ  d = f a; 1 r a; a; ; 1 ; 1 σ T Q1 D1ρa σ µ a ρa   D1

(6)

with T denoting a turbulence term (with dimensions of length), and D1 and Da representing the dimensions of the liquid and gas nozzles, respectively. The second and fourth terms on the right-hand side represent the We number (dynamic and surface tension effects) and Z number (viscous forces), respectively. Perhaps, if applied correctly, the parameters in Eq. (6) could help in the design of improved nozzles for use in nebulizers.

Atomization and Nebulizers III.

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Nebulizers

What sets these devices apart from atomizers from a functional standpoint are, (1) the natural recycling of fluid, (2) the use of one or more baffles or impaction sites that block the primary spray emerging from the atomizer jets, and (3) the size distribution of the aerosol emerging from the device, which must be of a size range that can successfully penetrate the lung. The reason for the cycling of fluid and the baffling is to help control the size, velocity, and volume flow of the aerosol emerging from the device. Without any baffling, the spray would be delivered to the oral cavity at a forward velocity of greater than 100 m/sec at a volume rate of somewhere between 1 mL/sec and 3 mL/sec, which is somewhat impractical! Finally, nebulizer use is generally limited to aqueous-air systems, whereas, industrial atomizers are used for numerous gasliquid combinations. A variety of basic nozzle types also exist, but primarily two types are used in nebulizers: Internal mixing and external mixing, that is, the air interacts with the fluid either within the nebulizer nozzle system prior to gas expansion or outside of the nozzle system after gas expansion. This section discusses the characteristics and performance of nebulizers used for inhalation therapy. It should be understood that the division into subsections is for convenience only as the topics are clearly interrelated. A.

Nebulizer Output and Efficiency

The simplest approach to estimate output from a nebulizer is to weigh the device before and after the nebulization period. Often the density of the solute is not accounted for and output is simply expressed in terms of volume per unit time (mL/min) or volume per unit airflow (mL/L air). This is shown for a number of nebulizers in Table 2. As shown, such measurements have been performed for numerous devices by a variety of authors (37–42). The mass output is intrinsically determined by the nozzle system, baffling, and reservoir design but controlled by the applied air pressure and hence the airflow through the nebulizer. The relationship is approximately linear over a typical operating range of 10 psig to 50 psig. However, nebulizers are often “run to dryness” before therapy is halted, and a significant reduction in the output rate occurs when there is insufficient fluid in the reservoir to maintain a continuous supply of liquid to the feed jets (Fig. 7) (37). Air entrapment in the jets occurs, and intermittent nebulization will ensue (43). This can be observed and heard as “sputtering.” Output can also be affected by a number of other factors. Temperature can impact output through an increase or reduction in the evaporative content of the total mass leaving the nebulizer. If solutions are concentrated, then changes in viscosity and vapor pressure may influence evaporation as well as droplet production (44,45). However, for the majority of aqueous solutions these factors probably play a minor role. Viscosity is an important issue when trying to atomize difficult materials such as paints, molten metals, or coal slurries, and the property may have an impact on current efforts to nebulize suspensions depending on their concentration (46,47).

264

Niven and Hickey Output Air Flow Rates and Aerosol Mass Outputs of the Nebulizers

Table 2

Air pressure (psig) 10 Air flow Nebulizer “Raindrop” “Misty” “Aerotech II” “Cirrus” Devilbiss #45 closed vent Devilbiss #45 closed venta “Updraft” “Whisperjet”b “Acorn I” “Acorn II” Collison 1 jet Collison 3 jet aOutput bResults

20

Mass output

Air flow

L/min

mL/min

uL/L

L/min

4.9 4.5 4.9 4.8 6.9

0.11 0.11 0.09 0.10 0.15

21.5 24.1 19.0 21.0 21.5

7.2 6.6 7.0 6.9 10.2

–

0.39

–

2.7 4.2 5.2 3.6 1.7 4.9

0.05 0.10 0.13 0.06 0.03 0.08

19.1 22.5 25.7 15.9 15.5 16.5

40

Mass output mL/min

Air flow

Mass output

uL/L

L/min

mL/min

uL/L

0.19 0.18 0.14 0.20 0.31

25.9 27.5 19.7 43.0 30.0

11.7 10.8 11.4 11.2 16.8

0.35 0.35 0.20 0.37 0.46

29.6 32.8 17.4 33.1 27.5

–

0.87

–

–

1.20

–

4.1 5.9 7.4 5.2 2.5 7.3

0.17 0.20 0.27 0.15 0.06 0.15

41.7 33.1 36.4 23.7 22.3 20.7

6.8 9.4 11.9 8.4 4.2 12.0

0.32 0.39 0.47 0.42 0.09 0.38

46.3 40.8 39.9 50.4 21.9 31.7

airflow rates of the DeVilbiss #45 nebulizer were not measured when the vent was open. are the mean values for six “Whisperjet” nebulizers.

The “efficiency” of nebulizers has various meanings in the literature. This can mean efficiency in terms of energy use (35), function, and, more often than not, simply output. Smye et al. (48) describe a statistical efficiency model for nebulizers based on the fraction of the initial volume of solution dispensed when operated to dryness. This might seem obvious as discussed above (37–41), but they mathematically define some aspects of nebulizer performance that are not often accounted for. This includes the period of sputtering, losses to walls, and the delay period before coalescing droplets return to the main reservoir of fluid. The efficiency term, E, is given by  V + Vw (T )  E = 1 −  min  V0 

(7)

where Vmin is the minimum volume of fluid in the reservoir that ensures continuous spray production through the nebulizer jets, Vw is the volume of solution on the walls at the end of the total nebulization time, T. This last variable, Vw(T), incorporates all parameters that describe the behavior of the nebulizer during the operation. From experiments using an Acorn nebulizer they show that some 99% of the aerosol is recycled, of which 99.2% is retained on the walls of the nebulizer for 1.6 seconds with the remainder for up to 200 seconds. They also note that the output rate during the sputtering stage is ⬇25% of that during the constant stage.

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Figure 7 Mass output rate from a nebulizer versus the reservoir volume. The output rate drops slightly initially, but is relatively constant until a limiting volume is reached. Once exceeded, the output rate drops dramatically. This results because there is insufficient fluid in the reservoir to continuously aspirate liquid to the atomizer nozzle. “Sputtering” or “fizzing” ensues, which becomes progressively worse as more and more air enters the feed jets. Tapping the nebulizer can improve the situation by returning fluid, trapped on the container walls, back to the reservoir. Source: Adapted from Ref. 37.

The term “mechanical efficiency,” Em, has been coined to describe the ability of devices to produce aerosol output relative to the aspiration rate of liquid (49). Thus M  Em =  t  ×100  Qt 

(8)

where Mt is the aerosol mass output (mL/min) and Qt is the liquid throughput or aspiration rate of liquid through the nebulizer jets. To obtain such values requires an irreversible modification of the nebulizer in order to make the assessment. Unfortunately, Eq. (8) [as does Eq. (7)] looks at efficiency with respect to total output and not the more relevant solute output. The results are, therefore, of limited practical value, but Eq. (8) does serve to illustrate the poor intrinsic ability of nebulizers at producing aerosol of a size suitable for inhalation (Table 3). The low efficiency values also indicate the inherent difficulty of transforming fluid into a size range some four 1og cycles smaller than the initial dimensions of the fluid in the reservoir (5 mL has an equivalent diameter of 2.12 cm) (50). One simple approach to improve efficiency is the use of an auxiliary vent (51). A baffled vent, or vents, strategically placed in the reservoir wall can increase the mechanical efficiency of a DeVilbiss #45 nebulizer from approximately 1% to 5% (Table 3). This improvement arises because the energy used to generate the aerosol is being better utilized. In this instance, the pressure

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Table 3

Mechanical Efficiency of Nebulizers % Efficienta Air pressure (psig)

Nebulizer Puritan “Raindrop” Airlife “Misty” Cadema “Aerotech II” DHD “Cirrus” Devilbiss #45 closed vent Devilbiss #45 closed vent Hudson “Updraft” Marquest “Whisperjet”b Marquest “Acorn I” Marquest “Acorn II” Collison 1 jet Collison 3 jet

5

10

20

30

40

0.20 0.13 0.37 0.20 1.09 2.61 0.10 0.19 0.29 0.16 0.05 0.08

0.37 0.21 0.48 0.37 1.19 3.18 0.18 0.33 0.46 0.32 0.06 0.10

0.60 0.27 0.71 1.16 1.74 4.94 0.53 0.62 0.68 0.70 0.07 0.14

0.92 0.26 0.80 1.19 1.69 5.18 0.80 0.93 0.83 1.00 0.12 0.15

1.34 0.33 1.14 1.70 1.64 4.24 1.01 1.03 1.02 1.74 0.11 0.15

a The percent mechanical efficiency as defined by the percent rato of the mean aerosol output to the mean apiration rate. b The mean percent mechanical efficiency obtained for six "Whisperjet" nebulizers.

differential between the atmosphere and within the nebulizer reservoir induces additional airflow through the nebulizer reservoir, and this additional airflow is used as a carrier medium for aerosol droplets that would otherwise return to the nebulizer reservoir. This increase in output is offset by a decrease in the aerosol mass concentration, but does not appear to change the median droplet size significantly (51). To address the problem of determining drug output, many researchers simply collect the drug output on filters (52,53), or, indirectly, the output can be measured with knowledge of the concentration that occurs within the nebulizer over time. One other notable effect that is often not accounted for in experimental studies is the influence of humidity. The influence of humidity on particle size has been studied in some detail (44,53–56), but it also influences output by condensing in the pressurized airlines when the ambient humidity levels are high. Stapleton and Finlay (55) give consideration to the amount of humidity in the air entering the nebulizer but do not comment on condensation. Yet, condensation can accumulate along the fluid lines and droplets ultimately entrained in the airflow will be forced through the Venturi gas orifice. Depending upon the external humidity and temperature, this effect can be significant and the net result will be to intermittently disrupt nebulization, dilute the nebulizing solution, and reduce evaporative concentration (57). B.

Concentration Effects

It is well known that concentration of fluid within air-jet nebulizers occurs during nebulization (52,58–63) due to evaporative losses of water. Determinations of solute output from a nebulizer from weight loss become increasingly inaccurate as time of operation increases. The process of concentration was originally given a mathematical description by Mercer et al. (58). The total output from a nebulizer

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can theoretically be separated into two components: The solution output, S, and the solvent output, W (mL/L air/min). This is illustrated schematically in Figure 8. The solution output refers to droplets containing solute that will escape the nebulizer, whereas the solvent output refers to all solvent loss from the reservoir. The solute concentration C(t) within the nebulizer fluid as a function of the operating conditions changes according to  W 

   W + S  V0 C ( t ) = C0 ×   V0 − (W + S ) × F × t 

(9)

where V0 is the initial volume (mL) and F is the flow rate (L air/min). One implication of the result is that because the reservoir concentration will increase geometrically over time, the solute output should also increase nonlinearly. It would, therefore, be useful to modify Eq. (9) in terms of the solute output, Mout, because this might have practical use in predicting drug output from a nebulizer for a range of operating conditions. An expression in terms of, Mout can be derived giving Mout

 S       V   S +W   = M0 ×  1 −    V    0 

(10)

Schematic describing the output of solvent (W) and solution (S) from a nebulizer. Droplets small enough to be entrained in the outgoing airstream will be accompanied by water vapor largely derived from the evaporation of large droplets that are recycled to the nebulizer reservoir.

Figure 8

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where M0 is the initial quantity of solute in the nebulizer solution, and V is the volume at any time, t (V ⫽ V0 ⫺ [W + S] ⫻ F ⫻ t). A similar expression has also been obtained by Smye et al. (61). The influence of concentration under different evaporative conditions is shown in Figure 9. Although concentration can change dramatically with time if the evaporation rate is high, the output rate will remain approximately constant for much of the run time. C.

Temperature Effects

The evaporation of water also results in cooling of the nebulizer solutions (44,58,63). Eventually, a steady-state situation is reached where the constant rate of cooling, k0, is in equilibrium with the heat input, k, from the surrounding environment. This effect is described by ∆T (t ) = ∆TSS × (1 − e − k ×t )

(11)

where ⌬T(t) is the temperature differential between T(0) and T(t), ⌬TSS ⫽ ko/k) is the temperature differential at steady state, and k is the apparent first-order rate constant describing heat transfer from the surroundings. Thus a plot of ln(1 ⫺ ⌬T(t)/⌬TSS) versus ⫺ k ⫻ t should be linear with a slope equal to k and zero intercept (Fig. 10). The extent of cooling increases as the applied air pressure increases, and the differential can reach 12° C to 13° C in a Collison nebulizer at 40 psig. This effect has the potential to cause precipitation of concentrated medicaments at lowered operating temperature, which can result in jet blockage, changes in drug size, and reduced drop output (63). The exponential rate of

Figure 9 Simulated curves generated using Eqs. (9) and (10) for the concentration increase in a nebulizer solution (A) and the solute output from the nebulizer (B). The volume output was fixed at 0.25 mL/min for a starting solution of 10 mL and airflow of 10 L/min. The percent of solvent output was then varied between 0% and 100%. It can be seen that as the evaporative component is increased, the concentration of solution increases markedly with time. Conversely, as evaporation increases, the relative output decreases.

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Figure 10 Linearized plot of cooling curves generated during nebulization. A temperature probe would typically be placed in the reservoir fluid. The slope represents the apparent firstorder transfer of heat from the surroundings to the nebulizer solution. See Eq. (11). The slopes are similar, indicating similar heat transfer rate from the surroundings, although the extent of cooling will differ in each case.

cooling applies to the air as well as the solution temperature (64). Preheating of nebulizer solutions has, therefore, been proposed to minimize any reduction of drug output (65). Concentration changes as a function of operating temperature are shown in Figure 11. The observed data are shown together with the best fits obtained using Eq. (9). The fitted results allow the relationship between solutions (S) and solvent (W) output to be examined as a function of temperature, and it can be seen that solvent output expressed in terms of mL/min (W ⫻ F) increases linearly with the operating temperature (Fig. 12), whereas the solution output (S ⫻ F) is relatively independent of the operating temperature. This effect has also been observed for a number of commonly used clinical nebulizers (52) and suggests that temperature is influencing the evaporation of droplets produced within the nebulizer, but is not affecting their production. If, for example, the median size of the gross aerosol droplets produced within the nebulizer were substantially reduced, then more droplets containing solute might be expected to escape the nebulizer, thereby increasing solution output. A nonlinear increase in vapor pressure will occur as temperature increases (66), and if the nebulizer solution did not contain solute, this relationship might be reflected in the observed solvent output. However, a vapor pressure reduction will also accompany any increase in solute concentration and will counteract the effect of temperature (67). It is also interesting to note that the reduction in vapor pressure associated with solute concentration should also ultimately reduce W, indicating that this parameter is not constant even where temperature is fixed.

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Figure 11 The concentration increase in nebulizer solutions with time for a variety of fixed operating temperatures. [4 (), 10 (), 21 (), 29 (), 37 (), and 52 ()°C (n ⱖ 3)]. The lines represent the best fits through the observed data by nonlinear regression using Eq. (9). The observed data are the mean values of standard deviation.

The clinical use of nebulizers does not typically enable them to be used in an environment where temperature or operation will be constantly monitored. However, it is of practical use to recognize that temperature can influence drug output. For example, this might be expected if solutions are nebulized directly after refrigeration or after having been removed from the shelf at ambient temperature. If heating is necessary to prevent precipitation of drug within nebulizer solutions, then output will also be affected, as will be the time of operation before sputtering occurs. D.

Design Considerations

Current nebulizers have not been radically changed over the last century. From a mechanical and clinical viewpoint, they have some advantages and disadvantages (Table 4) compared to other forms of inhalation therapy. Several of the physical phenomena that cause problems have been discussed above. Others relate to their design and implementation, including dead volume, containment, reliability, dimensions, and noise (68). These and other issues will now be profiled. Dead volume will be approximately constant and independent of the starting volume in the nebulizer. Therefore, the smaller the starting volume, the greater the fractional waste of drug. This waste is significant enough to make physicians recommend “tapping” of the nebulizer container walls to reduce the collection of droplets (69). This problem has been partially alleviated by

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271

The relationship between solution , solvent , and total output as a function of nebulizer operating temperature. Solution output, expressed in mL/min, is independent of operating temperature over a wide range, while solvent output appears to increase in a linear manner over the same temperature range.

Figure 12

altering the shape of the reservoir. For example, the Pari IS-2 nebulizer (Fig. 13) has a reservoir shaped like an inverted cone to maximize the depth of fluid in the vicinity of the liquid feed jet (70). A greater internal wall surface area will also influence the final dead volume. Large droplets retained on wall surfaces will accumulate and therefore, become unavailable for nebulization. This leads into a consideration of container or reservoir materials. Smye et al. (43) have pointed out that the loss of drops to walls will eventually represent a finite fraction of the drug and the lag time incurred before they return to the reservoir fluid. They also demonstrate an increase in efficiency from 49% to 67% by reducing the internal surface area -by adding perspex inserts (43). A study by Furmidge (71) discusses spray retention on walls of wax or cellulose acetate, and a theory is developed to support experimental results. He surmises that the spray, the formulation, the composition of the surface(s), and the contact angle will all influence droplet retention. Any improvement in the wettability of plastic surfaces should, therefore, improve total output by minimizing drop retention. The addition of wetting agents or surfactants to nebulizer formulations should do likewise, but their use could lead to foaming as well as the issue of benefit versus risk. Although improvement in cumulative drug output might be achieved, there will be issues to resolve concerning the toxicity of inhaled adjuvants, because virtually none have been approved for use by government regulatory bodies. The space available within the nebulizer container will also influence performance. The greater the distance that the spray has to cross before

272 Table 4

Niven and Hickey Advantages and Disadvantages of Current Nebulizer Design and Use

Advantages 1. 2. 3. 4. 5.

Most nebulizers deliver droplet sizes suitable for inhalation. Units are relatively inexpensive. Wide range of drugs can be nebulized; this extends to suspensions and emulsions. Design of nebulizers is uncomplicated. Accepted form of therapy and widely used in hospitals. Only form of inhalation therapy for infants. 6. Therapy can be halted at any time. Patient can control nebulization if “thumb” vent is included in the gas line. 7. Little training required relative to propellant or dry powder MDIs. Disadvantages 1. Dead or “waste” volume can exceed 50% of starting volume and varies according to volume used. 2. Unit-to-unit output performance can vary significantly, as does performance of different nebulizers. 3. Design primarily restricted to aqueous solutions. Suspensions are increasingly being employed but may lead to variable output and wear of nozzle exits. 4. Evaporative cooling and concentration through evaporation can lead to precipitation of concentrated medication. 5. “Sensitive” drugs and formulations can be adversely affected by air-jet nebulization. 6. Most nebulizer designs cannot be tilted without spilling contents. 7. Often a poor seal exists between the body and lid of the devices, and fluid can leak to the exterior during operation. 8. Condensation in pressurized gas line (with electric compressors) can intermittently disrupt atomization and dilute nebulizing solutions with potentially contaminated water. 9. Compliance can be an issue if nebulization time is lengthy. 10. Aerosol released to atmosphere will be inhaled by others unless a relatively bulky filter and valve system is employed. 11. Output rate is not constant for the duration of nebulization. Once fluid volume is insufficient to feed the liquid intakes, intermittent nebulization will ensue and “sputtering” will be held. 12. Container components are hydrophobic, and droplets are often entrained on remote areas of the nebulizer, thus contributing to the “dead” volume.

impacting on a surface, the greater the opportunity to disperse and evaporate. And any reduction in spray velocity (velocity profile relaxation) will reduce the impaction force on the walls, so minimizing additional droplet breakup. The vertical distance to the nebulizer outlet will also influence the emerging aerosol characteristics: Droplets will evaporate even if the air is saturated with water vapor (Kelvin effect) (72). A variety of nozzle designs are used in nebulizers, but there are few supporting data supplied by manufacturers that characterize their performance beyond those of the apparent droplet size of the emerging aerosol and the total mass output over time. Several scanning electron micrographs of nebulizer gas-liquid nozzles are shown in Figure 14. These illustrate that nozzle dimensions can differ significantly and that the internal bores often contain imperfections and promentories that could influence performance (12). This may help to explain the dramatic inter- and intravariability in nebulizer performance that has been documented in several studies (45,68,73,74). The impinger, or baffle, is also part of the nozzle system, and it is surprising, given

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A cross-sectional schematic illustrating the design of the Pari IS-2® nebulizer. This modern nebulizer incorporates some useful features that overcome some of the general disadvantages of nebulizers outlined in Table 4. These include an inverted cone-shaped reservoir to maximize fluid depth in the vicinity of the liquid feed jets, an auxiliary vent to improve output, and a collector ring to entrain droplets that otherwise might trap on remote areas of the nebulizer container. The external mixing nozzle design (Fig. 14) should also help to minimize jet clogging. Figure 13

its importance, that the influence of baffle position and design has not been studied in detail with nebulizers. This is despite a study by Dalby et al. (75) showing the importance of baffle size and position at the outlet of propellant metered dose inhaler systems. It would also be helpful if nebulizers could be tilted horizontally or even inverted without affecting performance. Respiratory therapists often dose many neonates and infants during a working day, and will carry the children while doing so. It is ironic that it is sometimes necessary to tilt the infant rather than the nebulizer to effect delivery of the drug. Aerosol containment is another problem. Most of the current betaagonists and steroids are exceptionally potent drugs. Exposure to health care staff is a problem despite the availability of adapters with valves or filters to contain the aerosol. These add to the expense and may be ignored because of bulkiness and the fact that aerosol released to the atmosphere quickly evaporates, giving the impression that the medication has gone, and therefore, no longer poses a threat. This has been seen with pentamidine, where significant levels of the drug were found in the urine of hospital staff assisting with aerosol therapy (76). The use of nebulizers for a widening range of drugs including anticancer agents and recombinant proteins is, therefore, a cause for concern. One critical aspect of nebulizer performance is the time of use. The more drug that can be delivered in a short time frame, the better patient compliance will be. Means to increase the aerosol mass concentration through

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Scanning electron micrographs of the nozzle exits (gas + liquid) of several nebulizers. (A, 5 –Pari IS-2®; B, 5 Cirrus®; C, 5 Acorn®; D, 5 Whisperjet®). All are quite different, and the manufacturing techniques used to generate the nozzles appear to differ. Edges, promentories, and other imperfections will influence the output and droplet size distribution that emerges from the nebulizer. The poor surface characteristics of the Pari nebulizer were caused by sample preparation. This device contrasts the internal versus external mixing design and illustrates the smaller dimensions of the gas Venturi. Figure 14

more efficient atomization should ultimately improve dose delivery, reduce wastage, and increase reproducibility. One potential drawback, however, is that by increasing output there will be a greater probability of inducing bronchospasm in some patients. This scenario is unlikely unless aerosol concentrations can be achieved that exceed those generated by ultrasonic nebulizers, which can cause bronchospasm with distilled water (77). Any effects are also likely to be formulation dependent (drug, pH, and osmolality) (78). The need for some control over the basic performance of nebulizers is now being recognized by regulatory bodies. Revisions to the USP guidelines for aerosol testing are to be implemented (79). These will likely be duplicated in other pharmacopeias throughout the world. In addition, the British Anesthetic and Respiratory Equipment Manufacturers Association has documented a set of British standards that manufactures should voluntarily adhere to for the manufacture and testing of nebulizers (80). Notable inclusions in the draft document are guidelines for measuring the solute output in addition to the mass output: Residual volume and particle-size determinations.

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Despite all the improvements to nebulizers that could be made, they do work. Almost all produce droplet sizes that, upon leaving the device, are of an acceptable size range (1–5 ␮m) for inhalation therapy. This is probably one reason why extensive research has not been implemented on their basic design. Clearly a great deal of performance “leeway” results from the containment of most of the primary spray and the fact that only the smallest size fraction leaves the device as “useful” aerosol. This scenario is idealized in Figure 15. The baffle or impinger and container walls play the role of “equalizers” and will heavily influence the performance of nebulizers as measured by output rate and droplet size. A further reason for the lack of progress in nebulizer design may simply relate to the cost it will take to research and implement new design features. Plastic molds are expensive, and tighter manufacturing specifications may result in a company having to price itself out of a market, where the competition is already fierce. Finally, some responsibility for improving nebulizer performance lies with the formulator to develop a preparation that can be readily aerosolized without detriment to the device or to the formulation. Some consideration will now be given to these issues. IV.

Formulation

Some obvious effects have been noted in the literature with respect to temperature and concentration, as discussed. In addition, a number of other effects

Idealized drawing comparing the size distribution of the gross or primary spray generated within the nebulizer with the size fraction of aerosol likely to emerge from the nebulizer. Presumably the polydispersity can differ in many respects (hypothetical examples shown here as A, B, and C) as long as the impingers and baffle systems of the device can limit the size of the output to dimensions suitable for inhalation.

Figure 15

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must be accounted for when considering more “sensitive” drugs or formulations such as liposomes (81–83) and proteins during nebulization. It is known, for example, that air-blast atomization in spray drying can damage enzymes (84) and therapeutic proteins (85). Formulation considerations for aerosolization are in addition to those needed to maintain shelf stability. This is because the formulation literally must go through a change of state before it is introduced to the body. Development of a product without consideration of the influence of aerosolization on the stability of the drug or components may compromise in vivo efficacy. There are a variety of effects to consider in addition to those of concentration and temperature, including shear, surface area, adhesion (to surfaces), and even volume. These topics are now briefly discussed. A.

Shear

The level of shearing that occurs during nebulization has not been explored in any detail. However, shear stresses (force/area) are known to influence the stability of enzymes during filtration (86), and mild shaking or stirring can cause aggregation of proteins. Further evidence that shear stress may have a role during atomization is demonstrated by the effects of nebulization on liposomes. Gilbert et al. (87) contend that nebulization causes a size reduction of liposomes. Damage to liposomes during nebulization has also been seen indirectly through the release of water-soluble encapsulated compounds (83). The release of carboxyfluorescein from liposomes has been determined to be a function of lipid content (82), size (88), and the conditions of nebulization (89). Whether the effects on liposomes are related directly to shear stresses is unclear and cannot be divorced from surface effects (force/length). B.

Surface Area

Nebulization is obviously a highly effective means of generating surface area. The aspiration rate of fluid in the three-jet Collision nebulizer at 40 psig is over 300 mL/min (49) (⬎100 mL/min/jet). The diameter of the orifice of the exit jets is ⬇1 mm, and the distance between the exit jet and the container walls is 1.4 cm. If we assume no radial expansion of a 1-mm-diameter liquid jet emerging from the nebulizer and, for convenience, no creation of droplets, a highly conservative estimate of the rate of air-liquid interface production can be made based on the generation of a cylindrical jet of fluid. The surface area of a cylinder is 2␲rh, where r is the radius and h is the length of the cylinder. The volume of the cylinder is ␲r2h. The surface to volume ratio is there 2/r. Hence, the surface generated through each jet per minute will be 2 ⫻ volume flow rate/r ⫽ (2 ⫻ 100/0.05) ⫽ 4000 cm2/min/jet, or 200 cm2/sec for all nebulizer jets. The true surface area that is generated will obviously be much larger due to (1) droplet production and (2) bulk transfer of solution from the interior to the surface of droplets during their lifetime, but the example serves to illustrate that even this “low” rate of surface expansion will create 120,000 cm2 of surface in 10 minutes. Unfortunately, proteins are well known to be unstable at air-water interfaces (88–90), and hence the notion that nebulization can be used for administering

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proteins to the lung would appear nonsensical. However, the approach has worked well, as evidenced by the clinical use of DNase (91) and the fact that other proteins such as insulin (92,93) and ␣-l-antitrypsin (94) also appear unaffected by nebulization. But there are cases, probably the rule rather than the exception, where nebulization has not worked well. rhG-CSF, for example, is aggregated and degraded by nebulization (95), and this shows that there is the potential for both physical and chemical transformation of sensitive compounds. To help minimize any destabilization, various “protectants” have been used (96). These include several nonionic surfactants and also polyhydric alcohols such as polyethylene glycol (PEG) (96,97). The stabilizing action of the two classes may differ in that the protection of the nonionics may relate to their micellar properties, whereas the PEGs may be acting through their ability to alter the structure of water and lower interfacial tension within the ⬇10 msec time interval in which the spray (surface) exists before being destroyed against the container walls (96). If protectants are necessary, care must be taken to choose a class of compounds that will minimize the possibility of in vivo toxicity. C.

Adhesion/Binding

Proteins and perhaps peptides are also recognized to be denatured at solidliquid interfaces (98) and can readily adsorb to solid surfaces (99,100). This has implications for blood transfusion and dialysis research, where the adsorption of plasma proteins to dialyzers and container surfaces is a significant problem (101). These problems may be exacerbated by mixing, shaking, and heating, and also by nebulization. Any transient (or irreversible) unfolding induced by an input of energy to protein solutions can expose previously enclosed hydrophobic regions to the aqueous medium. Thus, increased adsorption to hydrophobic surfaces may result. As an example, consensus interferon is known to bind to glass as a consequence of nebulization (102). Samples were taken from repetitive water rinses of an emptied container after nebulization and then finally from a wash with 0.1% w/v sodium dodecyl sulfate (SDS). SDS-PAGE on these samples demonstrates a heavy band from the SDS sample, indicating that the protein is adhering to the glass surface. This was reduced, but not prevented, by the addition of low concentrations of Triton X-100 or serum albumin. Clearly, some consideration should be given to the use of nonadhesive biocompatible materials in future nebulizer design. D.

Volume

The starting volume not only affects the fractional output and temperature of the nebulizer solution (37,103), but it also can determine the extent of destabilization that will occur during a fixed time of nebulization. This arises because the aspiration rate of liquid via the feed jets is constant irrespective of the starting volume (49). Consequently, a low starting volume of fluid will be propelled through the nebulizer jets a greater number of times in a given time interval than will a large volume. This effect has been seen for lactate dehydrogenase (49) and for liposomes. In the latter case, calcein encapsulated in liposomes was used as a marker of “damage” as signified by release of the fluorochrome from the vesicles. A strong relationship between liposome stability and the starting

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volume is observed (Fig. 16). The results illustrate that the influence of volume can be significant, and, because the volume of fluid changes over time, a slow increase in the rate of any adverse effect would also be expected. E.

Suspensions

This emerging topic is of significant clinical interest in relation to steroid administration to infants suffering from bronchopulmonary dysplasia or respiratory distress syndrome (47,104,105). A delivery problem exists because there is no topical steroid salt available that is soluble in an aqueous medium. Although dexamethasone sodium phosphate can be administered in solution, it is systemically absorbed in an active form and results in side effects. Suspensions formulated for nasal delivery—e.g., Vancenase—have been nebulized with minimal success. This arises due to variable output, the potential of clogging atomizer jets, and poor nebulization efficiency (Fig. 17) (106). Cosolvent systems have been considered, but this can lead to variable output through preferential evaporation of one or other of the solvent components and eventually precipitation of the drug. One logical approach to overcome such problems is being explored by Waldrep et al. (107), where steroids and other water-insoluble drugs (108) complexed in liposomes are being used for inhalation therapy. This surmounts the problems of insolubility as well as liposome stability. Although not approved for use, the majority of commonly used lipids appear to be nontoxic via inhalation (109). V.

Summary and Conclusions

The basic mechanisms of atomization and nebulizer operation have been outlined in this chapter. Unfortunately, the phenomena governing droplet

A plot of the percent release of calcein from liposomes damaged during nebulization as a function of starting volume and time. The lower the initial volume, the greater the extent of damage for a given set of operating conditions (Collison 3 jet nebulizer; 40 psig).

Figure 16

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Preferential concentration of beclomethasone dipropionate suspension that occurs during nebulization (4 g, 10 minutes; 20 psig) of Vancenase AQ (VAQ) nasal formulation using a Marquest “Whisperjet®” nebulizer. Aerosol output is variable and low. Significant differences in steroid concentration occur compared to VAQ placebo and a solution of dexamethasone sodium phosphate (DXP). (*p ⬍ 0.05 vs. VAQ placebo.)

Figure 17

production from atomizers are not well understood: a situation that is compounded with nebulizers where there is apparently little basic science built into their design and use. This lack of information also extends to the formulations that are used with the devices. Areas that might lead to augmented performance include studies to (1) increase output via auxiliary orifices, (2) minimize temperature reduction, (3) control the liquid:gas ratio, (4) optimize the position of baffles and other impaction sites, and (5) improve the shape and size of the nebulizer. Given the time and effort being devoted to development of dry powder inhalers and powder formulations, it does not seem unreasonable to suggest that a similar effort could be devoted to nebulizers. There is certainly room for improvement.

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10 Ultrasonic and Electrohydrodynamic Methods for Aerosol Generation

BERNARD J. GREENSPAN Verus Pharmaceuticals, Inc., San Diego, California, U.S.A.

I.

Introduction

There are additional aerosol generation methodologies that have been known for many years but have gained relatively little popularity compared to jet nebulizers and metered dose inhalers (MDIs). Traditional ultrasonic nebulizers have been in commercial production since the early 1960s, and it is only recently that advanced uses in the form of vibrating plates, meshes, and membranes have emerged. The application of electrostatic forces for the production of aerosols has also been known for a long time, but has not yet gained acceptance as a commercially viable method. To some extent, this is due to the relative complexity and associated cost of the devices for use in routine patient care (1). It may also be due to the complexity of the theory and design of the devices, compared to the simple jet nebulizer. Further, application of these methods involves imparting considerable energy to the drug solutions, which, as will be seen, may cause degradation. However, as targeted delivery of medications to the respiratory tract increases, interest in these seemingly esoteric methods will continue. Ultimately, clinicians and patients will benefit from a wider variety of delivery systems. II.

Ultrasonic Nebulizers

A.

Theory of Operation

The heart of an ultrasonic nebulizer is a transducer made from a piezoelectric crystal. The properties of the piezoelectric material are such that when 285

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oscillating electrical energy is applied, the crystal will undergo small physical displacements at the same frequency as the applied energy. When this oscillatory motion is transmitted into a liquid, the energy causes waves that propagate to the surface. Investigation of the mechanisms by which ultrasound causes the formation of an aerosol began as early as the 1930s. Sollner (2,3) described the existence of surface disturbances (or “fountains,” as he called them) at the surface of a liquid when ultrasonic energy was applied. The aerosol (fog) seemed to originate from them. He proposed that cavitation effects were responsible. [The origins of the theory appear to be even earlier, as Sollner cites work by Wood and Loomis (4) from 1927.] McCubbin (5) reported the production of micrometer-sized droplets of water above the surface when 45 watts of ultrasonic energy was applied at a frequency of 2.4 MHz. He further noted that the presence of soap in the water prevented the formation of an aerosol, but the “fountain” remained. Later, Lang (6) developed a model for the formation of liquid aerosols by ultrasonic energy in the frequency range of 10 KHz to 0.8 MHz. In Lang’s model, energy contained in the surface waves caused the formation of vertical capillaries (the “fountains” previously described). When the amplitude of the applied energy is sufficient, the capillaries break up and droplets are produced. The size of the droplets (count median diameter) was found to be proportional to the capillary wavelength (A) and was a function of the excitation frequency (F), the surface tension (␴), and density of the liquid (␳). CMD ⫽ 0.34␭ ⫽ 0.34 (8␲␴/␳F2)1/3

(1)

Peskin and Raco (7) found a similar relation between the droplet diameter and capillary wavelength at the low frequencies used by Lang, but predicted a constant droplet size at higher frequencies. Lobdell (8) derived a theoretical value of 0.36 for the constant in Lang’s equation, Eq. (1). Medical applications of ultrasonic nebulizers began to appear in the mid1960s. Mercer et al. (9,10) demonstrated good agreement between the theoretical values derived from Lang’s model and experimental measurements for three ultrasonic nebulizers operating at up to 1.4 MHz. They also examined the effects of ultrasonic nebulization on the activity of medications. These results will be discussed later. Boucher and Kreuter (11) reviewed studies of the physical mechanisms underlying ultrasonic nebulization. They presented data from Gershenzon and Eknadiosyants (12) characterizing the output of ultrasonic nebulizers with solutions of differing densities and viscosities. Gershenzon and Eknadiosyants found aerosol production, in mg/sec, to be a function of the viscosity (␩), surface tension (␴), and vapor pressure (p). (Atomization rate)2 ⬀ ␲pⲐ␩␴ B.

(2)

Design of Ultrasonic Nebulizers

There are several basic designs for traditional ultrasonic nebulizers, the fundamental differences residing in the configuration of the transducer and nebulizer cup (13) (Fig. 1) and in the baffles and auxiliary airflow systems. Table 1 gives

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Figure 1 Schematic depiction of energy transfer from ultrasonic transducer to fluid medium in nebulizers. A constant liquid level is crucial in the operation of a focused transducer.

some of the more important design factors influencing the mass and particle size from an ultrasonic system (14–17). Newer nebulizers using ultrasonic vibration of a mesh or plate will be discussed later in this chapter. In the simplest design, the liquid to be nebulized comes into contact with a flat transducer, oscillating at the desired frequency. In this arrangement the energy is termed unfocused. The arrangement allows all of the liquid to eventually be aerosolized from the surface without much change in the aerosol characteristics. A second design curves the transducer to produce a focused point of energy in much the same fashion as a concave mirror focuses light at a single point. This arrangement is capable of producing a finer aerosol; however, as the liquid level drops in the nebulizer cup, the surface of the fluid moves below the focal point and the efficiency of the device decreases. Ultrasonic nebulizers with focused transducers require a separate continuous-feed mechanism to maintain the liquid Table 1 Factors Influencing the Mass Output and Particle Size Distribution from Ultrasonic Nebulizers Fluid characteristics Dynamic viscosity Density Surface tension Vapor pressure (to a lesser extent) Piezoelectric transducer Frequency of vibration Amplitude Configuration-focused or unfocused Coupling of nebulizing cup to transducer Nebulizing chamber Size of chamber Presence or absence of baffles Auxiliary air flow Source: From Refs. 14–17.

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level at the appropriate height above the transducer. The sonic energy decreases with increasing distance from the focal point (18). Devices employing flat transducers are preferred for administration of small volumes of drug (13). In some devices, the solution to be nebulized comes into direct contact with the transducer or a bonded surface above the transducer. In others, a liquid interface acts as a couplant between the transducer and the base of the nebulizer cup. This couplant, usually water (for safety reasons), allows the base of the nebulizer cup to be shaped for more efficient transfer and focusing of the energy. In common with other aerosol-generating systems, ultrasonic nebulizers produce a number of large droplets, so the physical arrangement of the nebulizer cup and the outlet allows for the trapping and recycling of the liquid. The baffles that are inserted affect the overall total output of usable aerosol and its particle size distribution. Ultrasonic nebulizers do not require a source of compressed air to operate. As such, they may be operated as breath-actuated devices, releasing aerosol only when a patient inhales through the circuit. They may also be placed in line with ventilator circuits without the complexity of recalculating the tidal volume of the patient (19). For use in humidification, many units are equipped with fans to provide auxiliary airflow, delivering the aerosol from the device to the patient or room. These distinctions have resulted in many comparative studies between ultrasonic nebulizers and jet nebulizers. These studies have examined overall performance, total output, and particle size distribution. The results of some of these investigations are summarized and discussed in the next section. In most cases, the differences seem to depend on the methods of measurement, the configuration of the devices, and the intended applications. The most significant difference between ultrasonic and jet nebulizers is the lesser extent to which the drug solution will become concentrated in the reservoir, due to the absence of compressed driving air (20) for the former device. C.

Applications

Humidification

The earliest applications for aerosols generated by ultrasonic nebulizers were in the area of humidification therapy (21). The use of steam presented the problem of added heat in order to attain high relative humidities. Ultrasonic nebulizers could be arranged to produce large amounts of water vapor and aerosol for humidification tents in respiratory care units without raising the air temperature. Herzog et al. (22) described the use of an ultrasonic nebulizer in line with a ventilator for humidification of inspired air. Stevens and Albregt (23) used a DeVilbiss model 880 to effectively deliver moisture to the lower respiratory tract of dogs. Both of these studies claimed that ultrasonic nebulizers were more effective than jet nebulizers. Modell et al. (24) and Modell (25) initiated studies of the potential adverse effects from exposure to water and saline aerosols from ultrasonic nebulizers. They found that continuous exposure of puppies for 72 hours and of mice for up to 14 days caused pathologic lesions in the lungs. Further, enhanced mortality was observed in mice exposed to normal saline aerosol for 2 weeks. This prompted a warning against continuous use of high aerosol concentrations in humidification tents.

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In addition to the potential problems of chronic exposure, several investigators have looked at the potential for bacterial colonization in the nebulizer units. Unlike jet nebulizers, which are made from injection-molded plastics that are autoclavable or disposable, ultrasonic nebulizer parts are usually not disposable, especially the piezoelectric element. Rhoades et al. (26) reported on the spread of Gram-negative bacteria (Serratia marcescens) from contaminated nebulizers to patients and from infected patients to sterile nebulizers. They reported the difficulties associated with cleaning the nebulizer units, especially the connecting tubing. Witek et al. (27) reviewed the incidence of nosocomial infection in patients receiving respiratory therapy and conducted a prospective study in 50 postoperative individuals. Minimal colonization of nebulizer circuits during the first 72 hours suggested that the CDC recommendations (28) for daily changing may be too stringent for patients with uncomplicated clinical courses. Drug Delivery and Other Clinical Applications

Textbooks on respiratory care tend to place minimal emphasis on the use of ultrasonic nebulizers for drug delivery (29,30). This may be due to the differing results of various investigators. For example, Lin et al. (31) claimed to successfully aerosolize radiolabeled sulfur and tin colloids for lung-imaging studies. Isitman et al. (32) attempted to aerosolize radiolabeled colloids, but they achieved relatively poor overall deposition. Their best results were achieved with a noncolloidal material. Wasnich (33) also reported successful results with a noncolloidal-imaging agent. He utilized an ultrasonic nebulizer operating at 4 MHz, as opposed to the earlier models, which operated at 1.35 MHz, claiming that it allowed him to attain a smaller droplet size. In general, comparison studies have found that while ultrasonic nebulizers have high outputs compared to jet nebulizers, this is often achieved at the expense of a larger median droplet diameter. Ryan et al. (34) compared the performance of seven nebulizers, including an ultrasonic nebulizer, for use in bronchial provocation challenges with histamine and methacholine. Their study emphasized the importance of inspiratory time and the constancy of nebulizer output. Because nebulizers have differing performance characteristics, it is necessary to standardize the various factors across studies. The need to monitor and control, when possible, the operating characteristics of a chosen nebulizer is a recurring theme throughout the literature. Its importance should not be overlooked. Anderson et al. (35) described changes in airway reactivity in asthmatics inhaling ultrasonically generated aerosols of hypertonic and hypotonic solutions. This was proposed as a measure of nonimmunologically mediated bronchial reactivity. However, jet nebulizers are also capable of eliciting a similar response (36,37). In a comparative study of jet versus ultrasonic nebulizers with distilled water, Baba et al. (38) found greater bronchoconstriction in asthmatic children with an ultrasonic nebulizer than with a jet nebulizer. They attributed this to the smaller particle sizes from the ultrasonic nebulizer. In contrast to Anderson, these authors concluded that chemical mediators from mast cells are involved

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in response to distilled water aerosols. More recently, Spezia et al. (39) utilized ultrasonically nebulized distilled water to induce bronchoconstriction for a comparison of the bronchodilation effects of nedocromil sodium and sodium cromoglycate. But the bronchoconstrictive effect of aerosols may not be unique to ultrasonically generated solutions. Bronchodilator administration has been studied with ultrasonic nebulizers. In a comparison between albuterol delivered from a metered dose inhaler and a solution of albuterol in an ultrasonic nebulizer (40), no change in FEV1 was seen with the ultrasonic nebulizer system. Although the authors did not present supporting data to show that the drug actually left the nebulizer, they concluded that there might have been a competing bronchoconstrictive effect from the ultrasonically nebulized material. In contrast, Ballard et al. (41) found a greater increase in FEV1 from ultrasonically nebulized albuterol than from either a jet nebulizer or a metered dose inhaler formulation. However, the authors were hesitant to conclude that the ultrasonic nebulizer system is superior, because other factors such as cost, convenience, and patient preference must be considered in selecting a delivery system. A preference for ultrasonic nebulizers for administration of salbutamol to infants was reported by Fok et al. (42). Terzano and Allegra (43) preferred jet nebulizers to ultrasonic for the administration of beclomethasone dipropionate (BDP) to mild allergic asthmatics. Their results were based on improvements in FEV1 and peak expiratory flow and were likely due to the smaller particle size obtained from the jet nebulizers in their study. Hesitancy to use ultrasonic nebulizers for drug therapy is also based on evidence that the devices can degrade the drugs in the solution reservoir and upon aerosolization. Clinical settings are usually not equipped to measure the dose, potency, or other characteristics of therapeutic aerosols. Because of this, such measurements need to be made prior to treatment of patients. Goddard et al. (10) examined the effects of ultrasonic nebulization on the activity of penicillin-G, n-acetyl cysteine (Mucomyst), and cortisone. They found no adverse effects but discussed the possibility of drug degradation in the smaller droplets. Wigley et al. (44) investigated the delivery of insulin to rabbits by nebulization. They ultimately selected a jet nebulizer, citing their own finding that insulin activity was lost by ultrasonic nebulization. They did not give any details of the ultrasonic system, however. Marks et al. (45) demonstrated the successful nebulization of surfactant lipids extracted from bovine lung lavage by both jet and ultrasonic nebulization. They found that the ultrasonic nebulizer had an advantage in that the relative humidity remained high during nebulization. Although no changes in surfactant activity were found, Marks et al. (45) acknowledged that some phospholipids exhibit a decrease in their surface activity upon vigorous sonication. Despite this, Kohler and Matthys (46), in a letter to the journal, cited loss of activity of the surfactant as a major drawback of ultrasonic nebulizers. Notter et al. (47) responded in support of ultrasonic nebulization. More recently, Schumerly et al. (48) observed efficient surfactant aerosol delivery in an isolated lung preparation, accompanied by significant improvements in ventilation–perfusion matching. Wagner et al. (49) reported a decrease in the phospholipid content of surfactant aerosol with increasing excitation frequency. In response to the letters exchanged by Kohler and Notter, Gale (50) further supported the degradation of many

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compounds by ultrasonic nebulizers. Kosugi et al. (51) observed a decrease in fibrinolytic activity of Miraclid generated from an ultrasonic nebulizer. None of these studies determined the mechanism by which the compounds lost their activity. Niven et al. (52) have demonstrated the loss of lactate dehydrogenase (LDH) activity in an ultrasonic nebulizer. This phenomenon may have been due to a combination of heating within the device and evaporation in the aerosol droplets. By cooling the nebulizer, some of the activity was retained, but the output was greatly reduced. Addition of Tween-80 to the solution helped to reduce the loss of activity in the aerosol droplets. Khatri et al. (53) also used LDH as a model hydrophilic protein to compare the suitability of jet and ultrasonic nebulizers as effective delivery devices. They compared the performance of two jet and two ultrasonic nebulizers and found dentauration of the LDH in all but one ultrasonic nebulizer. Despite the stability, the clinical relevance to protein delivery may be questionable due to the relatively large droplet size (6.1 ␮m VMD). The authors stressed that their results might not be generalized to other proteins and that the choice of nebulizer needed to be tested carefully for any particular formulation. LeBrun et al. (54), working to optimize delivery of tobramycin to cystic fibrosis patients, compared the output of a jet nebulizer and an ultrasonic nebulizer. They reported an upper formulation limit of 20% tobramycin for the ultrasonic nebulizer due to the effects of increasing viscosity of the solution. The jet nebulizer was far less sensitive to the increases in solution concentration. Asmus et al. (55), on the other hand, found little reason to prefer a jet nebulizer to an ultrasonic nebulizer for delivering tobramycin when utilizing tobramycin injection solution (40 mg/ml). Their finding was likely due to the relatively low solution concentration tested (4% vs. 10% to 20%) compared to LeBrun (54). In another comparison of jet nebulizers and ultrasonic nebulizers with sodium cromoglycate in cystic fibrosis patients, Kohler et al. (56) concluded the ultrasonic nebulizer to be at least equivalent to the jet nebulizer. Gessler et al. (57) preferred ultrasonic nebulization to jet nebulization for the treatment of pulmonary hypertension with inhaled iloprost, a prostacyclin analog. They were able to achieve efficacy by both routes but preferred the ultrasonic nebulizer due to higher output of aerosol and less waste. This was attributed to the presence of an inspiration valve on the jet nebulizer circuit. Steckel et al. (58) found degradation of aviscumine, an antitumor agent, by both ultrasonic and jet nebulization. Addition of buffering salts and surfactants to the formulation achieved some degree of stabilization. Kohara et al. (59) note the effective use of an ultrasonic nebulizer to deliver the diuretic furosemide to cancer patients with uncontrollable dyspnea. The antibacterial ceftazidime was successfully administered to healthy piglets resulting in significant local concentrations in the lung tissue (60). However, the same potentially beneficial results were not observed in piglets with P. aeruginosa-induced bronchopneumonia. This study by Tonnelier and coworkers points out that even if the formulation and nebulizer are suitable for delivering active, stable therapeutics, efficacy of the treatment will still depend on the ability to deliver the aerosol to the target tissue.

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The feasibility of aerosolizing liposomal formulations has been demonstrated in ultrasonic nebulizers using phosphatidylcholine (PC) and PC/ cholesterol mixtures (61) as well as plasmids in liposomes (62). Further work on nonviral gene delivery systems showed no differences in transfection of isolated primary alveolar or primary bronchial cells using jet or ultrasonically nebulized formulations of various polyethylenimines and plasmid DNA (63). Ultrasonic dispersion was found to be less disruptive to the carrier–plasmid complexes than jet nebulization (64). Pentamidine

While studies of ultrasonic nebulizers for the administration of bronchodilators and other therapeutic agents seem to be sporadic and not always well designed, the administration of pentamidine for the treatment of Pneumocystis carinii pneumonia in AIDS patients initiated a new wave of comparisons of ultrasonic and jet nebulizers with the same agent. In an update on pentamidine aerosol therapy, Corkery et al. (65) presented results of jet and ultrasonic nebulizer performance with pentamidine solutions. The authors identified the major factors in selecting a nebulizer system and recommended further studies, including pharmacokinetics, in order to make an informed choice. Abd et al. (66) correlated improved deposition of radiolabeled aerosol in an AIDS patient to the smaller particle size from a jet nebulizer compared to an ultrasonic nebulizer. These studies also indicate that jet nebulizers are capable of producing aerosols with smaller MMAD values than are ultrasonic nebulizers. Additionally, a study by Dolovich et al. (67) noted a concentrating effect as high as 29% for pentamidine solution in an ultrasonic nebulizer when operated with a high auxiliary airflow. In the same study, however, it was found that the median particle size of the aerosol did not increase with time during the operation of the nebulizer. Matthys and Herceg (68) demonstrated differences in pulmonary deposition of pentamidine when the solution concentration was increased from 300 mg to 600 mg in 6 mL of water. The particle size of the aerosol was not reported; however, this study again emphasizes how the physical characteristics of the solution can alter the performance of a nebulizer. Performance data for one solution may not apply when another drug is used. Performance

Over the years, nebulizer manufacturers have continued to develop newer models of ultrasonic nebulizers, changing the internal arrangements of the baffles and altering the intensities and frequencies of the crystals. The literature abounds with comparisons of jet and ultrasonic nebulizers. Table 2 is a summary of the published performance characteristics for several ultrasonic nebulizers. Nebulizer comparison studies are usually difficult to compare between laboratories because of the variety of operating conditions and test solutions that are possible. In choosing studies to include in the table, those that utilized clinical performance and patient responses were omitted because of the variability and lack of standard comparisons. There is no standard nebulizer test solution that is agreed upon for intercomparisons, although a great many investigators choose

Ultrasonic and Electrohydrodynamic Methods for Aerosol Generation Table 2

293

Output Characteristics for Numerous Ultrasonic Nebulizers a

Nebulizer

Frequency (MHz)

MMAD (␮m)

GSD

Output (mg/min)

Reference

2262 6160

107b 10

DeVilbiss® 65 DeVilbiss® 880

1.35 1.35

6.3 6.9

1.77 1.6

EuroPlus FisoNeb® FisoNeb® FisoNeb®

1.1 5.0 5.0 2.75

1.6

1.3 1.3 1.3

FisoNeb®

1.3

6.13

65

FisoNeb®

1.3

5.99

65

FisoNeb® FisoNeb®

1.3 1.3

6.13

110 111

FisoNeb®

1.3

4.7

HaloMed (Bosch®) HaloMed (Bosch®) Heyer Mono Heyer USE 77 Hico 760 E Medix

0.1

Microinhalator (Siemens®) Mist-02-Gen EN134A Mist-02-Gen EN140 Mist-02-Gen Monaghan 650 Monaghan 670 Monaghan 670 Monaghan 670 Monaghan 670b Multisonic CompactTM

108 109 63 65

2.0

112

563

0.1

17.7

1.55 2.56 1.7

6.1 4.3 6.1 6.0

1532 2085 3899

17.7

107b 107b 107b 114

113

5.7

1.4

1.4

6.5

1.4

1.35

6.5

1.4

5.2

1.6–2.0 385–420

1.4

4.3

2.1

1.65

5.7

1.87

680

107b

4.3

2.1

1590

32

3.9

1.6

3.9

1.6

1.7

1640

Minimum auxiliary air 50% max auxiliary air Maximum auxiliary air 99m

Tc—human serum albumin 60 mg/mL pentamidine in water

Normal saline

Saline (MMAD; GSD): 1.7 ⫾0.5; 1.7 ⫾ 0.1 Normal saline

15

Saline

10

1% cesium chloride 1% cesium chloride 1.9% saline solution 1% cesium chloride

9 116 9

0.815

Saline 1% cesium chloride

107b 113

1.92 2.0 2.17 1.9

Notes

Saline

117

Salbutamol

55

5 ␮g/mL iloprost (Continued)

294 Table 2

Greenspan Output Characteristics for Numerous Ultrasonic Nebulizers (Continued)

Nebulizer

Frequency (MHz)

Medisonic Spiral Mark V

MMAD (␮m)

Output (mg/min)

Reference

Notes

2.81

45.6

53

0.25

51

40 mg/mL tobramycin injection 1 mg/mL LDH (estimated output)

2382

107

GSD

Omron® U1

0.66

6.10

Omron® NE-U10 Pentasonic

1.7

6.2

Portasonic Portasonic Portasonic

2.3 2.3 2.3

1.6 1.39 2.14

Portasonic

2.3

1.93

65

Portasonic Portasonic

2.3 2.3

1.93 4.7

110 111

Pulmosonic

1.3

2.9

3.3

64

Pulmosonic Pulmosonic Pulmosonic Pulmosonic Samsonic

1.3 1.3

4.2 2.8 5.4 6.8 5.2

2.3

63 113 120 107b 114

3.4

1.25

Samsonic

118

2.2

2.06 1.9

63 119 65

930

2.7

111

⬎12

Siemens “Green Machine” Siemens Hand Held Solcovent (Solco) Sonix 2000

2.12

1.8

4.37

1 lpm auxiliary air 6 lpm auxiliary air 99m

Tc—human serum albumin 300 mg Pentamidine in 15 mL sterile water Normal saline 12 lpm airflow Saline (MMAD; GSD): 1.6 ⫾ 0.4; 1.6 ⫾ 0.1 99m Tc—human serum albumin with mechanical ventilation

63

5.0 5.0

Sodium cromoglycate 5 mg/mL in normal saline

109 1.6

108 0.46

51

1 mg/mL LDH (estimated output) (Continued)

Ultrasonic and Electrohydrodynamic Methods for Aerosol Generation Table 2

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Output Characteristics for Numerous Ultrasonic Nebulizers (Continued)

Nebulizer

Frequency (MHz)

MMAD (␮m)

UltraNebTM 99 UltraNebTM 99 UltraNebTM 99

1.6 4.6 4.5 ⫾ 0.8

Venticis® (CEA) Wisto

1.1 1.7

4.3

GSD

Output (mg/min)

Reference 119 121 114

1.9 1.9 ⫾ 0.1

Notes

Saline (MMAD; GSD): 1.5 ⫾ 0.2; 1.7 ⫾ 0.1

119 1.84

1191

107b

a

Unless otherwise noted, the solution used in the studies was pentamidine. Data from reference 107 are for maximum output conditions for each nebulizer.

b

normal saline (0.9%). And, because of the application of ultrasonic nebulizers in the treatment of Pneumocystis carinii pneumonia, nebulizer performance with pentamidine solutions is included as well. The densities and viscosities of the various solutions are different, and this must be considered when interpreting the results in the table. These data may serve as a guideline for choosing an ultrasonic nebulizer for a particular application, but one cannot overemphasize the importance of performing real-time characterization, if only to account for seemingly subtle differences in airflow or tubing configurations. McCallion et al. (69) studied the ability of an ultrasonic nebulizer to aerosolize monodisperse latex spheres of varying sizes. They found a lower limit of 1.16 ␮m and an upper limit of 11.9 ␮m. Steckel and Eskander (70) provide a detailed comparison of performance characteristics for a jet and an ultrasonic nebulizer. For the ultrasonic nebulizer, they describe an increase in the solution temperature, an increase in the concentration of the nebulizer solution (possibly because the temperature was not fixed), as well as a decrease in the solution surface tension with time. Moreover, there were differences in the droplet size distributions with time. In conclusion, the preference of nebulizer is not always obvious. When assessing performance or choosing a nebulizer for a new therapeutic agent, one must always test the output characteristics (quantity and particle size distribution) and the stability of the drug product with the chosen nebulizer under the intended conditions of use. This is because the influence of the physicochemical characteristics of individual drugs on the choice of nebulizer is not easily predicted (71). Novel Application of Ultrasonic Techniques

Nearly all ultrasonic nebulizers have required a source of line voltage to operate. This has limited their application as portable patient devices. In the past few years, a new generation of devices has emerged, combining the generation efficiency of ultrasonic energy with the portability desired by patients.

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Schematic depiction of aerosol generation from a vibrating mesh or plate. Ultrasonic vibrations of the plate cause liquid to be pumped through the tapered holes and become aerosolized on the other side.

Figure 2

One portable device was developed, tested, and marketed for a brief time (72). It used a metering system similar to an MDI valve and a battery-driven oscillator circuit (2.5 MHz). Although the MMAD of the aerosol was close to 10 ␮m, the authors claimed equivalency to an MDI formulation of fenoterol. Nevertheless, owing to a variety of considerations, commercial production of the unit was stopped not long after its release. A group of devices using vibrating meshes or membranes are now available for clinical development and use (73). A drug solution or suspension is in contact with a mesh or plate with many tapered holes. When the plate or mesh is excited at a high frequency, the solution is pumped through the holes and aerosolized (Fig. 2). An early version, described by Baker and Stimpson (74), coupled a piezoelectric element to a fine electroformed nickel mesh. The vibration of the piezoelectric transducer produced droplets of a drug solution. By controlling the size and density of the holes in the mesh, as well as the frequency of vibration, the authors were able to dispense metered amounts of several drug solutions with droplet volume mean diameters of 4.1 ␮m to 4.8 ␮m. Aerogen (now part of Nektar, San Carlos, CA, U.S.A.) developed systems with a vibrating dome-shaped plate containing tapered holes (75). Oscillating motion induced in the plate by a piezoelectric element pumps liquid through the holes, thus dispersing a fine mist. Another mesh-plate ultrasonic nebulizer (Omron, Vernon Hills, IL, U.S.A.) was demonstrated to successfully disperse an aerosol of nanocrystalline BDP (76) with greater efficiency than a standard BDP metered dose inhaler. ODEM (Royston, Hertfordshire, U.K.) has developed an atomization head which is being incorporated into their own products (TouchSpray®) (77) as well as the eFlow® (Pari, Pari-Werk, Germany) (78). These new nebulizers offer the prospects of portability, high efficiency, and greater choice for clinicians and patients for respiratory drug delivery. III.

Electrohydrodynamic Atomization

The process of electrohydrodynamic atomization (EHDA) has been known for over 100 years (79), yet its application is not as widespread as other technologies

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described in this text. There are many reasons for this, among them the complexity of the process and associated theories, the limitations on the materials and associated conditions that allow the spraying process to take place, and the small volumes of material that can be atomized. The fact that few things could be simpler than a jet nebulizer for creating a liquid aerosol has minimized interest in alternative techniques. Yet, the appeal of EHDA lies in its ability to produce near-complete atomization with virtually monodisperse, micrometer-size droplets. This section is intended to provide a brief overview of some of the theories behind the process, describe a simple apparatus and equipment that may be used by the reader to demonstrate the phenomena and perform experiments, and present some of the existing device concepts that utilize this exciting “new” technology. To date, however, there are no commercial devices for medical aerosol delivery that utilize EHDA. At least one company (Ventaira, Columbus, OH, U.S.A), however, has initiated experimental and clinical trials with a number of compounds. The process of EHDA has been known by several other names such as electrospray, electroaerosol, and electrostatic atomization. The term EHDA (80,81) has been adopted because it distinguishes the process from those that simply apply a charge to an aerosol created by other, more conventional means. In Europe, there exist aerosol therapies involving charged droplets produced by conventional nebulizers, which are considered electroaerosols. These are not produced by the processes described in this chapter and will not be discussed. A.

Theories

The theory behind the formation of aerosols from liquids subjected to highstrength electric fields is quite complex. Several textbooks have been written on the subject (82,83). There are also numerous review articles (81,84,85). While the phenomenon has been observed for over 100 years, a concise theory has yet to emerge. The most noteworthy data come from work performed over the past 50 years. Vonnegut and Neubauer (86) reported the production of uniform droplets of liquid from the tip of a capillary to which 5000 volts to 10,000 volts were applied. They theorized that the process could be understood in terms of the total energy of the system, by summing the surface energy of the droplets and the electrical energy of the applied field. In this case, the resulting size of the droplets was a function of the applied field strength and the surface tension of the liquid. Drozin (87) examined the spraying characteristics of 13 different liquids and related them to the physical properties of the liquids (specific conductivity, dipole moment, surface tension, and dielectric constant). He concluded that the critical range for specific conductivity was between 10–13 ⍀–1cm–1 and 10–5 ⍀–1cm–1 and that surface tension was not a major factor, at least for nonpolar liquids. However, Smith (88) showed that there may be no upper limit on the conductivity. Yurkstas and Meisenzehl (89) appear to be the first to have applied EHDA technology to produce solid aerosol particles. They dissolved polystyrene in acetone (0.041%) and sodium chloride in ethanol (0.015%) and successfully atomized both solutions. The aerosols produced had count median diameters between 0.11 ␮m and 0.27 ␮m with geometric standard deviations ⬍1.12. Although these sizes are not

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optimal for pharmaceutical applications, they clearly demonstrated the feasibility of the technique for therapeutic applications. In subsequent studies, others (90–92) have investigated the phenomenon of EHDA and have offered various theories on the formation of the jets and droplets. These have all been based on the properties of the liquid and the applied conditions of voltage and electric field. In their excellent review of EHDA, Grace and Marijnissen (85) examined eight differing analytic models presented since 1990. Their conclusion remains that no model exists that can predict the mode of spraying for a particular set of parameters. It is generally believed that the formation of droplets occurs when the electrical forces exceed the surface tension forces in the liquid. However, despite numerous subsequent studies (93–96), the relative contribution and means to independently control each of the influencing parameters of the liquid, electric field, and the test system (Table 3) are not completely understood. Further, the appearance of the spray (often termed the “spray mode”) has many forms. These depend on the configuration of the test system, which may vary considerably among investigators. B.

Experimental Apparatus and Devices

Despite the lack of an adequate theory to explain the behavior of EHDA, there are numerous devices described in the patent literature that avail themselves of this phenomenon. It is also relatively simple to set up an apparatus to study EHDA. The basic design of such an apparatus is shown in Figure 3. By analogy to vacuum tube circuitry, this may be termed a spray diode, owing to the two elements in the circuit. Using such a system, the reader may reproduce many of the observations and results described in this chapter. A calibrated seriological pipette serves as a reservoir for the solution and allows recording of the fluid

Table 3

Parameters Influencing Electrohydrodynamic Atomization of Liquids

Liquid parameters Conductivity Surface tension Viscosity Dielectric constant Capillary parameters Material Size Tip configuration Electric parameters Magnitude of potential Magnitude of field at liquid meniscus as it forms at the tip of the capillary and in the region where the droplets are forming Polarity Space charge Miscellaneous parameters Liquid flow rate/hydrodynamic pressure (typically the height of the liquid above the capillary tip) Atmosphere around the capillary

Ultrasonic and Electrohydrodynamic Methods for Aerosol Generation

299

Schematic diagram of a basic apparatus for demonstrating EHDA. The electrometer is optional but useful for measuring the charge conveyed by the aerosolized droplets. Typical operating voltages range to 15 kV, and a typical spacing of the capillary and groundplate is 25 mm.

Figure 3

flow rate. Alternatively, a syringe pump may be used, but great care must be taken to electrically isolate the pump from the high-voltage circuitry. (Remember that the solution will conduct electricity at the potentials involved!) The distance between the capillary tip and the collection plate determines the magnitude of the electric field at a given applied potential. For most experiments, 25 mm is a suitable distance. A direct-current variable high-voltage power supply capable of 20 kV is required to produce the electric field. Addition of an electrometer to the circuit allows measurement of the charge conveyed by the droplets. Figure 4 is a schematic representation of observations at the capillary tip. As the field strength is increased (by increasing the applied voltage), the liquid begins to drip at an increasing rate. Soon the droplets appear to form a stream. At this point, closer examination reveals the presence of a Taylor cone (97) at the tip of the capillary. As the potential is further increased, a fan of very fine droplets can be observed, first forming around the droplet stream, and then replacing it completely. It is also possible to observe spraying from multiple points on the capillary. In his text, Michelson (83) depicts this as an inversion of the meniscus of the liquid into the capillary. If the flow rate of the liquid is increased by increasing the height above the capillary, or by pumping the liquid, a higher field strength is needed to achieve the same spraying characteristics. At some point, however, depending on the geometry of the system, electrical breakdown of the air will occur, and the current may arc between the capillary and the grounded collection cup. Extreme caution must be observed, especially when working with flammable solvents.

300

Greenspan

Progressive changes in liquid droplets emerging from a capillary tip during EHDA. With no potential applied (A), the meniscus of a liquid droplet is observed protruding from the capillary tip. If the height of the liquid column is sufficient, droplets may fall occasionally. When the potential is first applied (B), droplets are pulled off the end of the capillary at a rate which increases with increasing potential (field strength). As the field is increased (C), a Taylor cone (69) is formed and the spray emerges from it. With still greater field strength, a fan spray is observed mixed with the droplet spray (D), which yields to a single fan spray (E), and sometimes multiple fan sprays (F). This progression is easily observed using the apparatus described in Figure 2. Figure 4

Variations on the simple apparatus in Figure 3 may be made to control the charge on the droplets. Addition of a third element, or grid, between the capillary and the grounded collection plate serves to accelerate the particles (98) or to shield the capillary from a secondary source of neutralizing ions (99), as shown in Figure 5. The aerosol tetrode (Fig. 6) was described by Cannon et al. (100). This system is capable of producing large quantities of monodisperse particles with a moderate degree of control of their net charge. Such a system might be configured to replace air jet nebulizers.

Figure 5 Apparatus described by Noakes et al. The grounded plate shields the capillary from the negative ions produced for neutralizing the aerosol. Source: From Ref. 99.

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Figure 6 Schematic diagram of the aerosol tetrode from the Pacific Northwest Laboratory. Source: From Ref. 100.

Booth and Rowe (101) have described an apparatus for ocular delivery of small volumes of medication. Their patent application presents data on the administration of ephidrine and pilocarpine to the eyes of rabbits. Typical volumes were 5 ␮L to 20 ␮L solution. Greenspan and Moss (102) describe a handheld apparatus for delivery of metered amounts of medication for inhalation. Their device employs a piezoelectric crystal to produce the necessary potentials to achieve atomization. Finally, Zimlich et al. at Ventaira describe a handheld device capable of delivering a number of relevant pulmonary drugs (103,104) in an ethanol-based solvent system. C.

Formulation Considerations

The best solutions for demonstrating EHDA are solvents (e.g., ethanol, methylisobutyl ketone) with low surface tensions. However, these are flammable materials, and care must be taken to avoid arcing the high voltages in the presence of such solutions. Solvents also pose potential toxicity problems. Aqueous pharmaceutical formulations are difficult if not impossible to aerosolize with the above configurations because of their extremely high surface tension. The effect of the surface tension is to increase the electric field required. In air, breakdown will occur before spraying can take place. To address this shortcoming, Tang and Gomez (105) applied a technique readily used in electrospray mass spectrometry to produce 5 ␮m to 12 ␮m droplets of water at flow rates between 6 ␮L/min and 42 ␮L/min. Their system sheathed the capillary with CO2 gas to increase the threshold at which dielectric breakdown of the air would occur. This technique might prove limiting for therapeutic applications.

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Reports of pharmaceutical preparations aerosolized by EHD are limited (103,104). Davies et al. (106) recently reported on the successful spraying of plasmid DNA in an 80% ethanol–20% water formulation. Whether the limitations of ethanol-based systems can be overcome for drug approval will be seen in the coming years. As has been seen in the case of ultrasonic nebulization, the importance of formulation cannot be overlooked. IV.

Conclusions

This brief discussion of EHDA is intended to aid the reader in understanding the basic phenomena underlying this technology. Because of the complexity of the theory and many technical and formulation difficulties, EHDA is not widely accepted in clinical settings. As further research and development is carried out, it is possible that this technology will find a permanent place in pharmaceutical aerosol applications. References 1. 2. 3. 4. 5. 6. 7. 8. 9. 10. 11. 12. 13. 14. 15. 16. 17. 18.

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Wigley FM, Londono JH, Scott HW, Shipp JC, Waldmann RH. Insulin across respiratory mucosae by aerosol delivery. Diabetes 1971; 20:552–556. Marks LB, Oberdorster G, Notter RH. Generation and characterization of aerosols of dispersed phospholipids by ultrasonic and jet nebulization. J Aerosol Sci 1983; 14:683–694. Kohler D, Matthys H. Correspondence. J Aerosol Sci 1984; 15:519. Notter R, Oberdorster G, Marks L. Correspondence. J Aerosol Sci 1984; 15:519–520. Schermuly RT, Gunther A, Weissman N, et al. Differential impact of ultrasonically nebulized versus tracheal-instilled surfactant on ventilation-perfusion mismatch model of acute lung injury. Amer J Respir Crit Care Med 2000; 161:152–159. Wagner MH, Weithoff S, Friedrich W, Mollenhauer I, Obladen M, Boenick U. Ultrasonic surfactant nebulization with different exciting frequencies. Biophys Chem 2000; 84:35–43. Gale AE. Drug degeneration during ultrasonic nebulization. J Aerosol Sci 1985; 16:265. Kosugi T, Nakamura M, Noguchi S, Huang GW. Effect of ultrasonic nebulization of Miraclid on the activity in tracheobronchial sections in rats. Laryngoscope 1989; 99:1281–1285. Niven RW, Ip AY, Mittleman S, Prestrelski SJ, Arakawa T. Some factors associated with the ultrasonic nebulization of proteins. Pharmaceut Res 1995; 12:53–59. Khatri L, Taylor KMG, Craig DQM, Palin K. An assessment of jet and ultrasonic nebulizers for the delivery of lactate dehydrogenase solutions. Int J Pharmaceut 2001; 227:121–131. LeBrun PPH, deBoer AH, Gjalterma D, Hagedoorn P, Heijerman HGM, Frijlink HW. Inhalation of tobramycin in cystic fibrosis Part 2; Optimization of the tobramycin solution for a jet and an ultrasonic nebulizer. Int J Pharmaceut 1999; 189:215–225. Asmus MJ, Milavetz G, Tice AL, Teresi ME. In vitro characteristics of tobramycin aerosols from ultrasonic and jet nebulizers. Pharmacotherapy 2002; 21:534–539. Kohler E, Sollich V, Schuster-Wonka R, Huhnerbein J. Lung deposition in cystic fibrosis patients using an ultrasonic or a jet nebulizer. J Aerosol Med 2003; 16(1):37–46. Gessler T, Schmehl T, Hoeper MM, et al. Ultrasonic versus jet nebulization of iloprost in severe pulmonary hypertension. Eur Respir J 2001; 17:14–19. Steckel H, Eskander F, Witthohn K. Effects of excipients on the stability and aerosol performance of nebulized aviscumine. J Aerosol Med 2003; 16(4):417–432. Kohara H, Ueoka H, Aoe K, et al. Effects of nebulized furosemide in terminally ill cancer patients with dyspnea. J Pain Sympt Mgmt 230; 2(4):962–967. Tonnellier M, Ferrari F, Goldstein I, Sartorius A, Marquette C-H, Rouby J-J. Intravenous versus nebulized ceftazidime in ventilated piglets with and without experimental bronchopneumonia. Anesthesiology 2005; 102:995–1000. Leung KKM, Bridges P, Taylor KMG. The stability of liposomes to ultrasonic nebulization. Int J Pharmaceut 1996; 145:95–102. Pillai R, Petrak K, Blezinger P, et al. Ultrasonic nebulization of cationic lipid-based gene delivery systems for airway administration. Pharmaceut Res 1998; 15(11):1742–1747. Dailey LA, Kleemann E, Merdan T, et al. Modified polyethylenimines as non viral gene delivery systems for aerosol therapy: effects of nebulization on cellular uptake and transfection efficiency. J Controlled Release 2004; 100:425–436. Kleemann E, Dailey LA, Abdelhady HG, et al. Modified polyethylenimines as non viral gene delivery systems for aerosol therapy: investigations of the complex structure and stability during air-jet and ultrasonic nebulization. J Controlled Release 2004; 100:437–450. Corkery KJ, Luce JM, Montgomery AB. Aerosolized pentamidine for treatment and prophylaxis of Pneumocystis carinii pneumonia: an update. Respirat Care 1988; 33:676–685. Abd AG, Nierman DN, Ilowite JS, Pierson RN, Loomis Bell AL. Bilateral upper lobe Pneumocystis carinii pneumonia in a patient receiving inhaled pentamidine prophylaxis. Chest 1988; 94:329–331. Dolovich M, Chambers C, Newhouse M. Characterization of three systems to deliver pentamidine. Am Rev Respir Dis 1990; 141:A153. Matthys H, Herceg R. Dosing strategies for aerosol delivery to the lung parenchyma, with specific recommendations for pentamidine. Respir Care 1991; 36: 989–993. McCallion ONM, Taylor KMG, Thomas M, Taylor AJ. Nebulization of monodisperse latex sphere suspensions in air-jet and ultrasonic nebulizers. Int J Pharmaceut 1996; 133:203–214. Steckel H, Eskander F. Factors affecting aerosol performance during nebulization with jet and ultrasonic nebulizers. Eur J Pharmaceut Sci 2003; 19:443–455.

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11 Spray-Drying and Supercritical Fluid Particle Generation Techniques

MICHIEL M. VAN OORT and MARK SACCHETTI GlaxoSmithKline, Research Triangle Park, North Carolina, U.S.A.

I.

Background

The ultimate aim of the various particle generation techniques used in pharmaceutical inhalation products is to create respirable particles for either the treatment of respiratory diseases and/or for systemic delivery via the lung. The object is to have a stable formulation that, in conjunction with the device, gives efficient, reproducible, and reliable delivery to the lungs over the shelf life of the product. In addition, the formulation must be able to be reproducibly filled into the device at a rate that can meet commercial demand. In order to meet all of the critical quality attributes, one must optimize and control at least two important parameters, namely the particle size and the particle surface, since these parameters determine regional deposition in the lung as well as the particle–particle interactions (both drug and excipient) and particle–component interactions (such as the primary pack, actuator, etc). The pharmaceutical industry typically uses micronization as the method of generating respirable particles, but this results in rather limited flexibility or “tunabilty” since one has relatively little control over the surface properties (energy and roughness), morphology, and density. Therefore, one must rely on the physical properties of the active pharmaceutical ingredient (API) where the only possible adjustable parameters are the polymorphic form, habit, and salt. In recent years, there has been an increased emphasis on the use of alternative particle generation techniques for which the term particle engineering has been used. Some of the potential advantages of engineered 307

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particles are improved chemical and/or physical stability, improved bioavailability (such as controlled release or better lung targeting), improved processibility on downstream unit operations, the possibility of producing respirable peptides and proteins without deactivation, and increased delivery efficiency (improved fine particle doses). This chapter provides an overview of two particle generation techniques that have been in existence for many years and are beginning to be exploited in formulating respiratory medicinal products: spray-drying and supercritical fluid methods. The intent of the chapter is not to be an extensive review of the literature in these particle generation methods because such details are published in review papers and textbooks. Instead, the purpose is to introduce the basic concepts and technologies associated with spray-drying and supercritical fluid methods, with an emphasis on formulation issues for respiratory products, such as particle size, distribution, control, and physical stability. The chapter organization will be (1) spray drying, (2) supercritical fluid technology, and (3) solid-state and powder characterization.

II.

Spray Drying

A.

Introduction

Spray drying is a one-step process that converts a liquid feed to a dried particulate form. The feed can be a solution, coarse or ultrafine suspension/slurry, colloidal dispersion (e.g., emulsions, liposomes, etc.), or paste. In a typical application, the fluid is first atomized to a spray form which is in thermal contact with a hot gaseous medium. The large surface area of contact results in rapid evaporation of the droplets to form dried solid particles/granules, which are then separated from the gas by means of a cyclone, electrostatic precipitator, or bag filter. The principal atomization methods are centrifugal (e.g., spinning disk), kinetic (e.g., pneumatic), high pressure, ultrasonic, electrostatic, and effervescent. The variety of atomization, drying, and separation techniques enables spray drying to be adapted to many applications, including low-temperature spray drying in which the drying air is maintained at low relative humidity for aqueous feeds; pulse combustion spray drying, which uses heat and acoustics for extremely rapid evaporation at low temperature; spray freeze drying (1–7); spray chilling; spray polycondensation (8); spray reaction; spray-dryer absorption; and spray concentration (9). Spray dryers can also be designed to safely accommodate flammable organic solvents and explosive powders. The flexibility of spray-drying technology has led to its use in many industries, such as pharmaceutical/biochemical, chemical (e.g., polymers, resins, ceramics, detergents/surfactants, pesticides, dyestuffs, fertilizers, etc.), food (e.g., milk, eggs, plant extracts, fruits, vegetables, carbohydrates), and environmental control. The pharmaceutical/biochemical applications include synthetic drugs (antibiotics, such as penicillin, sulfathiazole, streptomycin, tetracycline, and cefuroxime axetil), excipients (lactose, mannitol, acacia, tragacanth, sodium alginate, and gum arabic), proteins (enzymes, such as pepsin, trypsin, amylase, lipase, protease, pectinase, rennin, lactase, cellulase,

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glucose oxidase, and glucose isomerase), vitamins (A, D, B, and B12), blood serum, plasma, vaccines, microorganisms (bacteria and spores), yeast, and mycelium (9). Applications of spray drying to powder formulations for inhalation have appeared in several review articles and references cited therein (10–22). Pulmonary delivery of proteins and peptides by spray drying and other means has been noted in several review articles (13–19). Other pharmaceutical applications of this technology that are relevant to respiratory drug delivery have been reviewed and recent reviews have been published focusing on controlled release for respiratory products (10,21,22). Approaches include liposomes, polymeric systems [poly (lactic acid) (PLA)] and [poly (lactic-co-glycolic acid) (PLGA)] and nonpolymeric systems (phospholipids), and cyclodextrins (21,22). The principal advantages of spray drying with respect to pulmonary drug delivery are the ability to manipulate and control particle size, size distribution, particle shape, and density in addition to macroscopic powder properties such as bulk density, flowability, and dispersibility. The addition of excipients can further extend the flexibility of the technology to further affect the solubilitity and dissolution rate. The technology is particularly suitable to heat-sensitive compounds due to the cooling effect of rapid droplet evaporation and the short residence time of droplets in the drying chamber. Moreover, the technology is readily adaptable to compounds of low aqueous solubility, by spray drying from organic solvents. This last factor is of particular relevance when formulating new chemical entities for which aqueous solubility is poor because particles can be generated without adding surfactants or other solubilizing agents. Pharmaceutical proteins are mainly produced by lyophilization, and the powders may not be amenable to further processing, e.g., milling. Spray drying of proteins provides an important alternative to lyophilization in that the powder can be generated with the desired characteristics so that postprocessing is unnecessary. This feature is particularly significant for inhaled products, as micronization may inactivate the protein. Thus, spray drying offers an alternative method to crystallization/micronization for particle generation with tremendous flexibility. This section will be devoted to an overall description of the spray-drying process, which includes three fundamental operations: atomization, drying, and separation. The material will emphasize the crucial features of powders for inhalation with regard to the unit steps of the spray-drying process. The focus will be on the application of spray drying to control and manipulate the particle size distribution, as well as the physical stability of the generated powders, as these factors are critical to formulating inhaled products. B.

Open- and Closed-Cycle Operation

The typical spray-dryer design and components are shown in Figure 1. This particular layout is referred to as “open cycle,”because the drying gas (usually air in this mode of operation) is not recirculated but instead is vented to the atmosphere. In small-scale research, the air is exhausted to a fume hood or

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Schematic layout of open-cycle spray-dryer design.

directly into the duct work of the fume hood system. The drying gas is heated to a desired temperature which is monitored and maintained by a thermocouple and feedback system. The drying gas passes through a distributor plate and passes into the main chamber, where it is mixed with the spray generated from a nozzle or, more generally, an atomizer. The liquid feed is aqueous in the open-cycle design. The enormous surface area of the spray and its intimate contact with the heated gas result in rapid drying and solidification of the feed substance. Coarse particles may be collected at the bottom of the main chamber, but in normal operation, most of the mass passes into the cyclone where it is separated from the gas flow and collected in a vessel. A wet scrubber removes the ultrafine particles before exhausting the moist gas to the atmosphere. Although the open-cycle layout is a powerful design for aqueous feeds, it is inadequate and should not be used for organic solvents due to their flammability and risk of explosion when heated in the presence of oxygen. When spray drying from organic media, the “closed cycle” layout shown in Figure 2 is required. The principal features of the closed-cycle design are three-fold. First, the heated gas is recirculated (hence the term closed). Second, the drying gas is usually nitrogen with an oxygen content maintained less than 5% (below the level required for ignition) by an oxygen sensor; a feedback system opens a solenoid valve to increase the nitrogen concentration when the oxygen content upper limit is reached, which will occur due to leaky seals, connections, etc. Third, a condenser is employed to convert organic vapor to the

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Figure 2

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Schematic layout of closed-cycle spray-dryer design.

liquid state. Thus, as nitrogen is recirculated past the condenser, its organic vapor content is minimal, i.e., reduced to a low value depending on the trap temperature and associated vapor pressure of the solvent. As can be seen by comparing Figure 1 with Figure 2, there are many components common to both open- and closed-cycle layouts. Moreover, there are many variations of these designs, the main one being the position of the atomizer. When the spray enters the drying chamber from the top in the same direction as the gas flow, as is shown in Figures 1 and 2, the method is termed “cocurrent.” In this approach, the droplets, in their most moist state, are subjected to the highest temperatures, which results in rapid evaporation and concomitant cooling. As solidification occurs, the particles occupy positions lower down in the drying chamber, where the gas temperature is below that at the inlet. Thus, cocurrent spray drying is particularly suitable for heat-labile materials and, in the case of inhalation products, drugs and excipients with low melting points or glass transition temperatures. It should be noted that this effect is much more important in pilot- and manufacturing-scale equipment, where temperature decreases from the top to bottom of the chamber can be as

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large as 200°C (9). However, in laboratory bench-scale equipment, a typical difference between inlet and outlet temperatures is approximately 200°C, depending on such factors as the value of the inlet temperature, drying gas flow rate, liquid feed rate, and materials. In “countercurrent” spray drying, the nozzle is positioned at the bottom of the chamber, and hence droplets initially travel upwards in a direction opposite to the gas flow. This approach leads to longer residence times of droplets in the drying chamber due to the fact that they first travel upwards before descending and passing into the separator unit. However, in contrast to cocurrent spray drying, the droplets experience the lowest temperatures in their wettest state and the highest temperatures in their driest state. Thus, this method is not as suitable for heat-labile materials as the cocurrent design. The countercurrent method can lead to a product powder of higher bulk density than the cocurrent design due to the lower tendency of particles to expand and fracture. Another option is “mixed-flow,” in which gas can enter from the bottom or simultaneously from the top and bottom of the drying chamber while the atomizer is situated usually at the top. The spray can also enter from the side of the chamber at an upward angle to produce a parabolic profile. The merit of this approach is that the residence time of droplets in the drying chamber can be increased while maintaining the advantages for heat-labile materials in cocurrent spray drying. Again, it is important to mention that these distinctions are more pronounced in large-scale equipment due to the great variation in temperature in the main chamber. The other commonalities between open- and closed cycle or aqueous and organic spray drying are in the methods of atomizing the liquid feed, the drying variables, and powder separation. These fundamental operations in spray drying will be presented in the subsequent sections. C.

Atomization

Atomization is a process whereby a liquid is broken up into a collection of droplets. Since the surface area and surface free energy are increased upon spray formation, atomizing a fluid requires work. The amount of work and manner in which the work is done (e.g., nozzle design) on the liquid can dramatically affect the droplet sizes, can thus provide one with flexibility and control in generating a desired particle size distribution.a Considering the importance of size distribution in lung deposition of a therapeutic agent, it is this fundamental ability to manipulate droplet size as part of the particleformation process that makes spray drying an important technique in formulating inhaled respiratory products. There are many methods of atomizing a liquid, the chief ones of which can be classified according to the way energy is transferred to the liquid: centrifugal (rotary), pneumatic (air-assist), high pressure, effervescent, ultrasonic, and aA

distinction is made between droplets and particles. Atomizers determine the droplet sizes, while the final solid particle size distribution is controlled in part by the initial droplet sizes but, additionally, by the dynamics of the drying process. This section is concerned only with droplet formation.

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Figure 3

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Schematic diagram of spinning disk (A) and inverted cup (B).

electrostatic atomization. The first four atomizers will be considered in some detail; the last two will be discussed only briefly since the details are covered in another chapter. Rotary Atomization

Work is done on the liquid by increasing its rotational kinetic energy and providing a high (tangential) relative velocity between the liquid and the surrounding gas in the main drying chamber. The basic atomizer designs include feeding the liquid onto the top surface of a spinning disk, cup, or vaned wheel, or the bottom surface of an inverted cup (Fig. 3). There are many variations on these basic designs that can be found in books (9,23,24) and articles (25–28). However, there are common features that illustrate the flexibility of this mode of atomization with regard to manipulating droplet size. Generally, the most important variable is the disk, cup, or wheel tangential velocity at the edge, which can be broken down into disk diameter and angular velocity or rotational speed. Larger-disk diameters reduce the tangential velocity and hence relative velocity between the liquid and the surrounding gas, which results in a larger droplet size. Increasing the rotational speed of the disk, by the same rationale, increases the relative velocity and thus decreases the droplet size. At the high end of disk speeds, droplets as small as 20 ␮m are formed, and this diameter can be readily increased to over 200 ␮m at low rotational velocities (9). There are essentially three regimens of disk speed and feed rate that exert a profound effect on the distribution of the droplets. At the lower end of disk speeds and feed rates, drops detach directly from the edge, generating sprays

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of primary monodisperse and smaller satellite drops. Thus, the sprays are not monodisperse, but by suitable design of the atomizer, the parent or satellite drops can be removed by inertial impaction (25). At higher rotational speeds and feed rates, ligaments form and extend out from the disk’s edge, ultimately breaking up into primary and satellite drops. If the feed rate is increased further, the ligaments can no longer sustain the mass flow, and a sheet is formed at the disk’s periphery, which disintegrates into a spray of broad size distribution. In general, low feed rate and high disk speed are required for more monodisperse sprays, and increasing feed rate results in a large mean drop size. Other important factors that influence the mean droplet size relate to the liquid properties: such as viscosity, surface tension, and density. Increases in each of these three properties produces a larger mean drop size. Thus, when spray drying a drug at different solution concentrations, or in comparing aqueous and organic spray drying, the physical properties of the liquid must be taken into account if similar particle size distributions are required, as will presumably be the case for an inhaled product. Any changes in drop size due to feed properties can be compensated by altering the disk, cup, or wheel speed. Not only is rotary atomization flexible in terms of generating particles of desired mean size and monodispersity, but it is also versatile in its applicability to laboratory, pilot plant, and manufacturing-scale spray drying. In bench-scale research with small drying chambers and short residence times of drops, extremely fine sprays at low feed rates are required and can be achieved with disk speeds as high as 60,000 rpm (28). Spinning wheels are used in larger-scale spray-drying equipment that can accommodate production rates as large as 2 tons per hour at high wheel speeds, generating fine sprays (9). Pneumatic Atomization

This method of atomizing a liquid, also termed twin-fluid atomization, is normally done with a nozzle shown schematically in Figure 4A and B, although some applications require a rotating cup configuration (Fig. 4C). There are many variations on these basic designs, and details can be found in other reference sources (9,23). The external mixing nozzle design illustrates the fundamental mechanism of this atomization process. The liquid feed is usually pumped at relatively low flow rates through a tube and, at the discharge orifice, is impinged by a gas under pressure. The velocity of the gas is much larger than that of the liquid at the orifice, and it is this relative velocity that provides the force and work needed to create the surface area of a spray. In practice, these two variables are controlled by the flow rate of the liquid and the gas pressure for a given nozzle. On the fine end, sprays with a mean droplet size of 5 to 20 ␮m can be produced at high relative velocity, and on the coarse end a mean droplet size of ⫺200 ␮m at low relative velocity is achievable (9). Thus, the pneumatic nozzle is perfectly suited for generating particles in the size range of inhalation products for both drug and excipient. Moreover, since the cone angle of the spray can be as small as 10°, this nozzle is easily incorporated into a small chamber and is a valuable tool for laboratory-scale research.

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Figure 4 Schematic diagram of pneumatic atomizers: (A) external mixing; (B) internal mixing; and (C) inverted rotating cup designs.

Although external mixing pneumatic nozzles are limited to low liquid feed rates for the most effective atomization, they can be and are used in production-scale work by employing multiple nozzle heads in a single drying chamber. Their major disadvantage in manufacturing-scale spray dryers is their poor efficiency; it is estimated that in generating sprays of mean droplet size 10 to 20 ␮m, the efficiency is less than a tenth of 1% (9). This disadvantage is often offset by the versatility that this nozzle offers in atomizing low- to highviscosity solutions, suspensions, and pastes with the ability to manipulate and control the droplet size distribution. The internal mixing design (Fig. 4B) improves the atomization efficiency but does not offer the direct control over the relative gas/liquid velocity that is achieved by external mixing. The pneumatic internal mixing nozzle is mostly used in combustion equipment where efficiency is of paramount importance. The inverted rotating cup pneumatic atomizer, as the name implies, is a cross between rotary and pneumatic atomization. It has the advantage of reducing the outward radial velocity of the spray, so it can be used in smaller drying chambers. Its chief advantage, however, is in atomizing highly viscous, non-Newtonian fluids that cannot be easily pumped through a nozzle orifice. Just as in rotary atomization, the relevant liquid properties for pneumatic nozzles are the viscosity, surface tension, and density. In general, a small mean droplet size will result from a decrease in each of these three properties. High-Pressure Atomization

In its most basic form, the liquid feed is pressurized and forced to flow through a nozzle orifice (Fig. 5A). The smaller the orifice and the higher the jet velocity, the finer the atomization. The mechanism of droplet formation is similar to the pneumatic nozzle, except the jet velocity is much larger than that of

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Figure 5 Schematic diagram of high-pressure atomizers: (A) simple orifice and (B) swirl chamber designs.

the surrounding gas. But it is the relative velocity difference that provides the force and work needed to increase the jet’s surface area, just as for the pneumatic nozzle. The only way to produce a finer spray for a given nozzle is to increase the pressure of the liquid. For low-viscosity Newtonian fluids (e.g., water), pressures in excess of 100 atm are needed to achieve a mean droplet size of 15 to 30 ␮m, while 20 atm will generate a spray of mean size 150 to 350 ␮m (9). There are many variations on the high-pressure nozzle. The most basic improvement on this simple design is to impart angular velocity to the liquid before it exits the discharge orifice, which is accomplished by directing the liquid through tangential entry ports into a swirl chamber upstream of the orifice (Fig. 5B). The rotational motion of the liquid results in the formation of an annular, cylindrical sheet extending beyond the orifice. The thin liquid sheet provides greater energy transfer and more efficient atomization than the thick cylindrical jet of the plain orifice nozzle. However, the disintegration of a sheet into droplets produces a polydisperse spray, which may or may not be desirable in spray-drying inhaled therapeutics and excipients. The major application of high-pressure nozzles is in combustion engines, where polydisperse sprays are advantageous (23). Just as for rotary and pneumatic atomization, increases in liquid viscosity, surface tension, or density result in coarser atomization. For industrial-scale spray drying, multiple nozzles will be needed to achieve high throughput. Effervescent Atomization

In effervescent atomization, a liquid is forced to flow through a plain-orifice nozzle, but a gas is injected just upstream of the discharge tip (Fig. 6). As the

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Figure 6

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Schematic diagram of effervescent atomizer.

gas bubbles expand in the jet, ligaments are formed which disintegrate into droplets. The chief advantage of this design is that the atomization is comparable to the performance of pneumatic nozzles and the liquid pressure can be reduced from the values used in high-pressure nozzles. The sprays, however, are polydisperse in droplet size distribution (23,29). Ultrasonic Atomization

A liquid feed is passed over piezoelectric transducers vibrating at ultrasonic frequencies (e.g., 50 kHz). Droplets as small as 20 ␮m can be generated. One of the main advantages of these nozzles is their very low spray velocity, so that the droplets can be easily entrained in the gas flow of the drying chamber. This feature enables spray drying to be conveniently accomplished in small-diameter chambers in laboratory-scale research. The details of this mode of atomization are provided in another chapter. Electrostatic Atomization

Electrostatic atomization involves pumping a liquid at low flow rates through a nozzle that is at high voltage relative to ground. The electrical conductivity of the liquid causes it to absorb charge, which, in turn, drives droplet formation in an attempt to increase surface area and minimize the repulsive energy between like charges at the interface. This increase in surface area is counteracted by the liquid’s surface tension, so that for liquids of poor electrical conductivity, effective atomization requires high voltages. The smallest

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droplets are generated at the highest voltages that can be achieved before electrical breakdown of air (or gas). In general, the mean droplet size depends on the same liquid properties encountered with other atomizers, such as viscosity, surface tension, and density, but, in addition, the liquid’s dielectric constant and electrical conductivity are of importance. In the past, electrostatic atomization was limited to low feed rate applications such as spray painting, but more recent advances have enabled higher flow rates to be achieved (30). This mode of atomization is presented in detail in another chapter of this book. Summary of Atomization Methods

Each of the aforementioned methods offers its advantages and disadvantages for a particular material to be spray dried. In terms of droplet size distribution, the conventional rotary and pneumatic atomizers can be readily adapted to inhalation drugs and excipients in both bench- and production-scale work, as has been demonstrated in the scientific literature (9–22,31,32). However, the recent work in ultrasonic, electrostatic, and effervescent atomization technologies should not be overlooked in formulating inhaled products, because these methods provide alternative mechanisms of spray formation, which, for a given substance, may prove beneficial. For example, the high frictional shear encountered in rotary and pneumatic atomizers may be detrimental to some sensitive therapeutic agents, and exploring other atomization technologies may overcome this problem. D.

Drying

Once the liquid is atomized, it is in intimate contact with the heated gas, and evaporation of the drops transpires. In order to successfully collect any powder, e.g., an inhaled medicament or excipient, the droplets must dry to solid particles during their residence time in the main chamber. This drying step in the spray-drying process is critical for several reasons. First, although the atomization variables determine the initial droplet diameters of the spray, it is during the drying process that solid particles form, and the final size, shape, density (i.e., aerodynamic diameter), crystallinity, and solvent content of the particles are all affected by the conditions in the drying chamber. The important drying variables are inlet temperature (Tin), outlet temperature (Tout), drying gas medium, gas humidity,b gas flow rate, and residence time. The inlet and outlet temperatures, in normal operation, are not independently controlled. Typically, the inlet temperature is established at a fixed value and the outlet temperature is determined by such factors as the gas flow rate, chamber dimensions, and feed flow rate. However, Tout, can be treated as the independent parameter with Tin, varied until the desired outlet set point is achieved. The inlet and outlet temperatures determine the overall efficiency of the drying process. In small-scale formulation research, the efficiency is usually not an important factor. However, in scale-up of the operation, the drying b The

term “humidity” is used in its most general sense to denote water or organic vapor content per unit mass of gas.

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efficiency needs to be considered. The thermal efficiency (␩) is defined as the ratio of heat used in evaporation to the total heat input, which can be approximately calculated from T − Tout (1) η = in Tin − Tamb in which is the ambient temperature. Thus, it is clear that the thermal efficiency can be improved by spray drying at higher temperatures. In practice, particularly with pharmaceuticals, there will be a limit to the inlet temperature, since product degradation will occur, so ␩ must be maximized with this constraint. In addition to the thermal efficiency, the humidity (mass of water or organic vapor/mass of dry gas) of the gas medium needs to be considered. In aqueous spray drying in air, in normal operation the air is taken from the ambient atmosphere and hence has a humidity determined by the environment surrounding the spray dryer. Moreover, the evaporation of water from the spray further increases the humidity at the outlet of the drying chamber. Considering that, at any given temperature, matter in the solid state will have a certain equilibrium moisture content depending on the relative humidity of its gaseous surroundings, it is evident that the humidity at the outlet or in the cyclone and collection vessel must be sufficiently low to achieve the desired level of dryness of the final product. Although outlet humidities in spray dryers are adequately low to attain acceptable dryness for many substances, there are problematic materials that have high equilibrium moisture contents even at low relative humidity. In such cases, the intake air may need to be dehumidified and the outlet temperature increased to reach lower moisture contents of the powder. Failure to remove excessive moisture in hygroscopic materials can lead to agglomerated powders that, depending on the aerodynamic size distribution, may not be acceptable for inhaled therapeutics. Thus, for problematic substances, it is important to study their moisture sorption isotherms to establish the required outlet conditions before proceeding to scale-up the process. In the case of nonaqueous spray drying, which may be required for compounds of low aqueous solubility, the principles outlined above all apply except that it is organic vapor, not moisture, that must be removed from the solid. Due to the high volatility of organic liquids, it is usually not difficult to achieve completely dry products in the spray dryer, although with a particular substance it is presumably possible to trap solvent in a glassy matrix or to generate a solvate. A further drying step may be required to reduce the solvent content to pharmaceutically acceptable levels. Indeed, the pharmaceutical acceptability of the solvents themselves needs to be seriously evaluated before proceeding with organic spray drying. The condenser is crucial to successful spray drying of nonaqueous feeds because, if it operates efficiently, the humidity in the drying chamber can be maintained at a level determined solely by the amount of evaporated liquid. Usually, pure nitrogen gas is used, although there are some advantages to other gases, such as carbon dioxide (CO2), assuming they do not react with the product (9). The drying gas flow rate is important in that along with the inlet temperature, it determines the heat input rate to the drying chamber. Clearly, the heat

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input rate must be sufficiently high to evaporate the spray rapidly in the drying chamber. Equally important, the gas flow rate determines the minimum residence time of the droplets, which can be calculated from V (2) tmin = ch Q where tmin is the minimum residence time, is the drying chamber volume, and Q is the gas volumetric flow rate. The chamber volume can be estimated from its dimensions and the gas flow rate from direct measurements. Most commercial spray dryers incorporate an orifice with pressure gauge for measuring the overall flow rate. Eq. (2) illustrates that although increasing the gas flow rate will enhance the heat input rate, it will also shorten the droplets’ residence time to a value that may not permit complete evaporation. Thus, in such cases where evaporation is incomplete, increasing the heat input rate may not serve to dry the product; instead, a larger chamber may be needed, the feed flow rate can be reduced, and/or if possible, a higher inlet temperature can be used. Alternative Drying Options

Although this chapter is concerned with technology that uses a hot gaseous medium for drying the spray, it is important to mention drying alternatives. Some drug substances may have material properties that preclude conventional spray-drying operation. For example, a thermally labile drug that is hygroscopic may require a high outlet temperature and long residence time, both of which could cause chemical degradation. Other drugs and excipients may have low melting and glass transition temperatures. Considering that many substances form amorphous solid phases due to the rapidity of the evaporation kinetics (e.g., approximately 1 second in small-scale work), the value of the glass transition (Tg) is paramount in determining if spray drying is a feasible approach. If Tg is below the outlet temperature, then, in principle, the particles formed will be in a supercooled, viscous liquid state, and most likely will not be recovered by the cyclone separator as individual particles, because a film will form that may subsequently recrystallize. If only crystalline material is available with which to work, Tg can be roughly estimated from the melting point (Tm) using the semiempirical equation 2 Tg ≈ Tm 3

(3)

in which temperatures must be expressed in Kelvin units (33–39). Table 1 illustrates how Tg varies with Tm on the Celsius scale and reveals that compounds with melting temperatures below approximately 200°C may be problematic, assuming a lower estimate of approximately 40°C in low-temperature organic spray drying. In actual practice, compounds with melting temperatures less than 150°C can be successfully spray dried, because Tg > (2/3)Tm or crystalline phases may be formed. However, Table 1 can provide a good guide for action if films are observed in the cyclone or other separator.

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Values of Tg and Tm in °C

Tm (°C)

Tg (°C)

100 125 150 175 200 250 300

⫺24 ⫺0.8 9 26 43 76 109

Note: Calculated from Eq. (3).

One way to circumvent the above problem is to freeze rather than dry the spray. If the freezing medium is gaseous, the process is termed spray freeze drying and is illustrated in Figure 7. The frozen droplets are collected in a vacuum chamber for lyophilization. A variation on this technique was described in the pharmaceutical literature (7) in which the liquid feed was sprayed directly onto a bed of pulverized solid CO2. Although approximately 18 ␮m aggregates were formed, the primary particles were microscopically visible and were less

Figure 7

Schematic diagram of spray freeze dryer.

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than 3 ␮m in size. It is evident that with further study this technique could be of value for respiratory drugs. Another related alternative is to spray the liquid feed into a low-temperature fluid. This method has been used with liquids such as freon-12 (1) and liquid nitrogen (2–7). If the particles are less dense than the liquid medium, their collection is facilitated by skimming them from the top surface. E.

Powder Separation Techniques

After the spray-dried powder is formed, it must be separated from the circulating gas medium. The three relevant methods for inhalation substances are cyclone separators, bag filters, and electrostatic precipitators. Cyclones

The conventional and most widely used cyclone design is illustrated in Figure 8. The tangential entry produces a downward spiraling vortex with a central core of upward flowing gas. The circulation of the velocity field produces a centrifugal force on the entrained spray-dried particles. Thus, the particles migrate radially outward, will impact on the cyclone wall, and will be removed from the gas flow with a certain collection efficiency. The downward velocity component of the spiraling motion results in powder collection in the bottom vessel, although it is not unusual to experience significant build up of

Figure 8 Conventional cyclone design schematic: (A) cross-sectional view and (B) top view showing tangential inlet.

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powder on the cyclone wall. This problem can be particularly prevalent with the fine particles of respiratory drugs, which exhibit adhesion forces to the wall that are much greater than the shear stresses imposed by the velocity field of the gas. Techniques, such as tapping, vibration, and providing a gas sheath at the wall, can be helpful in preventing severe build up of powder, although there will always be a thin layer that cannot be dislodged by these methods. The most important factor to consider with the cyclone separator is its collection efficiency, summarized by a “grade efficiency curve” shown schematically in Figure 9. It is extremely important to note that this curve applies to a high-efficiency cyclone and that, in general, cyclone grade efficiency curves are shifted to larger particle diameters with d50 of order 5 ␮m. In spray-drying drugs for inhalation, lower-efficiency cyclones can lead to poor product recovery and severely bias the particle size distribution below 5 ␮m. With proper use of a high-efficiency cyclone, one can exploit the classifying function of this unit to effectively remove ultrafine particles by adjusting the d50 value. In general, for a given cyclone, the collection efficiency can be enhanced by increasing gas inlet velocity and increasing mass loading; further optimization of product recovery will require altering cyclone design factors such as decreasing inlet cross-sectional area, decreasing cyclone diameter, and increasing cyclone length. More details concerning cyclone design and fundamental principles can be found elsewhere (40). Advances in high-efficiency cyclone technology can be found in the scientific literature (41) and in product literature from companies manufacturing such equipment. Bag Filters

As the name implies, a bag filter is a woven fabric used to remove spray-dried particles by filtration from the gas flow. The fabric is attached to the output of the drying chamber as a stand-alone unit or is used after a cyclone in a two-stage

Figure 9

Grade efficiency curve for a “high-efficiency” cyclone.

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recovery to collect the ultrafine particles below the d50 of the latter separator. The bag filter has excellent collection efficiency, with a value of greater than 99% recovery for 1 ␮m particles. The fabrics are available in materials ranging from natural wool, to synthetic polymers (e.g., polypropylene, polyamides, polyester, fluorocarbon), to porous metals (9). Due to their high collection efficiency, bag filters should be considered when spray drying expensive drug substances (e.g., proteins) in large-scale work. Electrostatic Precipitators

The electrostatic precipitators used in spray drying are normally of parallel plate design with a discharge electrode to ionize the gas stream. The spraydried particles acquire charge and migrate to the grounded plates for removal from the gas flow. Like bag filters, the grade efficiency curves are excellent for the fine particles used in respiratory products, with greater than 99% recovery for 1 ␮m particles. Electrostatic precipitators are suitable for large-scale work in that high gas volume flow rates can be handled (42). F.

Spray-Drying Variables Affecting Particle Size, Shape, and Density

The first and most important step in successfully spray drying a drug for respiratory therapy is to control the particle size distribution. However, since it is the aerodynamic properties of the particles that are of greatest concern, one must also consider their shape and density. One can understand these factors by considering the fundamental aspects of the spray-drying process: droplet formation, solvent type and solution concentration, and droplet evaporation dynamics. As the various atomizer technologies were described in a previous section, the focus here will be on solvent/solution concentration and the evaporation process. Solvent Types and Solution Concentration

The ability to spray dry from aqueous and organic solvents is a powerful option for inhaled products, since for drugs of low water solubility, organic spray drying may be the only route available. Although a third component, such as a surfactant, can be used to improve aqueous solubility, its presence in a respiratory product will be a regulatory issue. The most important properties of the solvent with regard to the size distribution of the spray are viscosity, surface tension, and density. Considering the two most important atomizers for generating fine sprays—namely, pneumatic and rotary—increases in each of these solvent properties will result in larger drops, as has been mentioned in the section on atomization. Another property is the boiling point of the solvent, which plays a role in the dynamics of the drying process to be described shortly. Thus, the solvent characteristics affect not only the initial droplet-formation step (e.g., droplet size) but also the drying process (e.g., solid particle size). The solution concentration of the drug or excipient is another variable that can be easily manipulated to control the final solid particle size. Not only does a lower concentration lead to finer particle sizes, but it can also decrease

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Table 2 Mean Particle Size of Lactose as a Function of Solution Concentration Concentration (%)a

Mean size (␮m)b

0.1 6 15

1.0 1.8 2.3

aAqueous

solution. by scanning electron microscopy.

bDetermined

the particle density. The lower particle density will produce a smaller aerodynamic diameter for particles of equivalent geometric size, as is illustrated by Eq. (4) for perfect spheres (4) Dae ⫽ Dgeo ␳1/2 Dae and Dgeo are the aerodynamic and geometric diameters, respectively, and ␳ is the particle density. For example, Table 2 shows the particle size results of spray drying an organic substance (lactose) at several concentrations. The table demonstrates that varying the solution concentration is useful for shifting the particle size distribution in the respirable range; it is a particularly powerful variable for “fine-tuning” the mean size after the atomizer parameters have been adjusted to produce droplets of the approximate magnitude. Because it is the aerodynamic particle size that determines the regional deposition in the lung, this approach of generating geometrically large particles that are aerodynamically small has been an interesting development since the mid-1990s. Part of the interest is due to the observation that these low-density particles readily disperse and, therefore, allow the potential of developing highefficiency pulmonary delivery systems. The ease of dispersibility lies in the fact that the “effective particle diameter” is reduced, the projected area of the particle is increased, and the interstitial velocity through the powder is increased (43). Low-density particle research has been conducted by numerous individuals, but there are two companies that have been major players in the area of low-density particles, Nektar and Alkermes. One of Nektar’s approaches is to generate large porous particles that resemble whiffle balls which they call Pulmospheres. These whiffle balls have interesting properties in that they form quite stable metered dose inhaler (MDI) propellent suspensions due to the intercalation of propellent in the pores, thereby more closely matching the density of the propellent. In addition, when exiting the metered dose inhaler, the particles very rapidly decelerate due to the significant drag properties of these geometrically large particles, thus resulting in a significant reduction in throat deposition. Alkermes’ lowdensity particles resemble crumbled up paper which readily disperse because of the significant reduction in particle–particle interactions. Droplet Evaporation Dynamics

Atomization conditions influence the initial droplet sizes, but dynamic evaporation and material properties determine the final solid particle size, shape, and density. The pivotal research in this area transpired in the 1950s in a series

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of publications by Marshall and coworkers (44–47). The results of this research on evaporation of single-droplet solutions are summarized diagrammatically in Figure 10, which is approximately reproduced from the original paper (47). In addition to the shapes shown in this figure, other particle morphologies that are sometimes observed are illustrated in Figure 11, which presumably result from precipitation of solid at the surface while the drop is deformed. Indeed, observation of droplet motions in air reveal oscillations between cylinder and disk shapes depending on the Reynolds number (45). Figure 12 shows scanning electron microscopy (SEM) microphotographs of organic substances exhibiting many of the features portrayed schematically in the previous figures. The diversity of particle morphologies that can be generated by the drying process can be explained by the dynamics of evaporation and material properties. The drying of droplets can be roughly divided into two periods in a manner similar to tray drying, although there are cases in which this classification is inadequate (24). In the first period, drying of the drops is very similar, irrespective of the chemical nature of the solute. As the drops form, rapid evaporation transpires at the surface, thereby introducing concentration gradients

Figure 10

process.

Diagrammatic representation of solid particle formation during the drying

Spray-Drying and Supercritical Fluid Particle Generation Techniques

Figure 11

327

Deformed solid particle shapes sometimes observed in spray-dried products.

between the liquid surface and the interior, the former becoming more concentrated in solute. Moreover, in the early time just after the drop is created by the atomizer, there is a large difference in velocity between it and the surrounding gas. Under such conditions, boundary layer theory predicts and experiments verify (47,48) that the evaporation rate is maximum at the front edge of the drop (where impingement is greatest), decreases to a minimum just

Scanning electron microscopy photomicrographs of organic substances, demonstrating the influence of drying in determining particle size, shape, and density.

Figure 12

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beyond the equator, and increases to a second, lower maximum at the drop’s rear. Thus, there is a tangential component to the concentration gradients in solvent and solute in addition to the radial distribution. The highest evaporation rate at the front edge leads to precipitation of solute at this position once the solubility limit is exceeded, followed by complete “crust” formation. From this point onward, which is to say the second drying period, the course of events depends on the material properties of the crust, the two most important of which are its porosity and rheology. Since these material properties are determined by the chemical nature of the solute, a diversity of particle morphologies are observed. Charlesworth and Marshall (47) identified the drying temperature (i.e., inlet temperature of the spray dryer) as a critical factor relative to the boiling point of the droplet solution. If the drying gas temperature is below the solution boiling point, then three cases can be delineated, depending on material properties, as shown in Figure 10. For porous crusts, further evaporation occurs, although at a less rapid rate than in the first period, by diffusion of solvent through the pores to the interface, accompanied by no change in particle size for rigid materials. For less porous solid skins, stresses develop that fracture the crust followed by evaporation, as in the first period. Materials exhibiting viscoelasticity undergo shriveling as the particle volume decreases during evaporation, producing a “raisinlike” morphology (49). A different series of events can transpire when the inlet temperature of the spray dryer is greater than the solution’s boiling point. Just as in the case when the drying temperature is below the boiling point, no change in particle size occurs for porous solid skins. For more impervious, rigid crusts, the decreased evaporation rate after crust-formation results in a rise of the encased solution temperature to the boiling point, and subsequent bubbling results in minor or major fractures, presumably depending on the drying temperature and extent of void space available for solvent diffusion through the solid surface structure. In contrast, for low-porosity viscoelastic crusts, the positive pressure can cause inflation to sizes larger than the droplet diameter, which can be followed by collapse if the skin permeability increases to a sufficiently large value upon expansion. If the solid skin permeability does not increase significantly, ruptures can develop, leading to shriveling and in some cases a “spongy” structure if multiple inflation/collapse events occur. The SEM microphotographs in Figure 12 are examples of many of the morphologies observed as a consequence of material properties and the drying dynamics of the spray-drying process. The hollow nature of spray-dried particles is demonstrated in Figure 12, and the reason should be evident from the previous description of crust formation and subsequent evaporation through the inwardly growing shell. The above description illustrates that successful spray drying of a drug or excipient for inhalation therapy requires careful consideration of drying parameters, such as inlet temperature (i.e., drying temperature), outlet temperature, and drying gas flow rate as well as solvent type, solution concentration, and boiling point. These variables, along with the material properties of the solid crust, exert a profound influence on the particle size, shape, and density or, ultimately, aerodynamic behavior. The atomization conditions and feed

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properties, such as viscosity, surface tension, and density, control the initial droplet sizes, which, of course, also impacts on the final solid particle size distribution. The large number of variables and the interaction between them makes a statistically designed experiment approach appealing for rapid formulation development, particularly for inhaled drugs where a tight control of particle aerodynamic behavior is essential.

III.

Particle Generation Using Supercritical Fluids

The supercritical fluid can be viewed as a typical recrystallization solvent in that many of the approaches in which it is used have direct analogies to classical recrystallization methods. As is the case with traditional solvents, solutions of supercritical fluids can be used in spray-drying processes. In addition, the supercritical fluid can be used as an antisolvent that can be added to a solution (or vice versa, the solution added to the antisolvent) to force the precipitation of the solid of interest. Supercritical fluids do have some unique qualities that have led to renewed interest in exploiting this technology. Among the physical properties that the supercritical fluid has, which are not found with common solvents, are liquid-like densities with very large compressibilities and viscosities intermediate between the gas and the liquid extremes. In addition, since the density is a measure of the fluid’s solvent power, small changes in temperature and pressure can result in large changes in the solvent power, thus allowing the system’s solvent attributes to be adjusted or manipulated across a continuum. Several excellent textbooks and reviews have been published for the interested reader (50–63) as well as a wealth of information in the patent literature. Figure 13 shows a pressure–temperature phase diagram of a pure component. Where the three phase boundary lines meet is the triple point, and if you follow the liquid–gas phase line, eventually it will end at the critical point. This

Figure 13

Pressure–temperature phase diagram for a pure component.

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Table 3

Critical Parameters for Several Supercritical Fluids

Compound

Boiling point (°C)

CO2 NH3 H2O N2O Ethane Ethylene Propane Pentane Benzene Methanol Ethanol

⫺78.5 ⫺33.4 100.0 ⫺88.6 ⫺88.6 ⫺102.7 ⫺42.1 36.1 80.1 64.7 78.5

31.3 132.4 374.2 36.5 32.3 9.2 96.7 196.6 288.9 240.5 240.8

72.9 112.5 218.3 71.7 48.1 49.7 41.9 33.3 48.3 78.9 63.0

82.5

235.3

47.0

Isopropanol

Critical temperature (°C)

Critical pressure (atm)

is the temperature and pressure where the gas and the liquid can co-exist as a single phase. For pharmaceutical applications, the most popular medium is CO2 since it has a low critical temperature (Tc ⫽ 31°C) and is nontoxic, nonflammable, and inexpensive (Table 3). Although many spray-drying systems can generate particles in the respirable range (<5 ␮m), in the case of proteins, they can be denatured during the process due to the use of high temperatures to atomize and evaporate the solvent. Micronization can also adversely affect the crystallinity and chemical stability of pharmaceuticals and denaturation of proteins. This is in part attributable to the high localized temperature due to particle friction that can be minimized by reducing the temperature and operating in an inert atmosphere. Lower temperatures also make the particles more brittle, thus improving the micronization process for soft materials, and the inert atmosphere minimizes the possibility of oxidation. Lyophilization is another method of generating respirable powders for inhalation devices, but this results in broad particle size distributions and it may be very difficult to attain satisfactory accuracy and precision in the size distribution control. One method that has gained popularity in generating respirable particles is the use of supercritical fluids. As mentioned previously, the most common methods or processes where supercritical fluids are used in the generation of particles are as a solvent or as an antisolvent. As a solvent, the process is called rapid expansion of supercritical solutions (RESS), or the rapid expansion of supercritical fluid solutions; as an antisolvent it is called gas antisolvent (GAS) or supercritical antisolvent (SAS). A.

Rapid Expansion of Supercritical Solutions

RESS can be considered an extremely rapid, high-pressure spray-drying process. The general principle is to dissolve the compound in a supercritical fluid and then rapidly expand that fluid, causing the solvent to flash off and the compound to precipitate. It differs from conventional spray drying in that supersaturation can be achieved by changes in the fluid pressure (and hence solvent power) rather

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than changes in the temperature. Although this method of particle generation by decompression of supercritical fluids has been known for over 100 years, one of the first detailed investigations of the capabilities of the RESS technology was first documented in a patent granted to Smith in 1986 (64). In this patent, Smith deposited solid films or fine powders by rapidly expanding supercritical fluid solutions of organic and inorganic compounds from a region of high pressure into a region of relatively low pressure. Using temperatures in the range of 400°C to 500°C and pressures of 400 to 450 atm, he was able to deposit films of polystyrene and silica onto fused silica or platinum substrates (64,65). One of the reasons that RESS appeals to many researchers is its potential to produce uniform particles where the size distribution and morphology can be controlled. The potential for generating consistent particle sizes is attributed to the nearly uniform nucleating environment generated because supersaturation is pressure mediated and the pressure perturbation is rapidly propagated through the medium. The rapid expansion through a well-defined orifice offers some degree of control over the expansion characteristics. Morphological control using RESS has been demonstrated by numerous investigators. Smith has shown that by varying the pre-expansion temperature, polystyrene can be obtained as fibers with diameters of 1 ␮m and up to 1000 ␮m long, while at higher temperatures 20 ␮m spheres are obtained (64,65). In addition to morphological control, crystallinity can be varied, so it can be used to obtain amorphous powders and films. Selection of the supercritical fluid is based on the solute’s solubility at operating conditions; it must be nonreactive with the solute, and the apparatus must be capable of tolerating the high pressures and temperatures that may be present. The temperature and pressure requirements to achieve supercritical conditions and the desired flexibility and production output determine the configuration of the RESS apparatus. The simplest or basic system consists of a pump capable of generating pressures well above the critical pressure of the selected fluid, a heating unit to generate temperatures above the critical temperature, and a nozzle through which the solution is expanded. Figure 14 shows an example of a somewhat more sophisticated benchscale RESS apparatus (51,53). The system consists of two main units—the extraction or solubilization unit and the precipitation unit. This particular system allows the possibility to control extraction and precipitation temperature and pressures as well as the solution concentration prior to precipitation. The connections between the extraction and the precipitation units must also be maintained at the same or higher temperature than that of the extraction system; otherwise, precipitation in the lines can occur, thus causing blockage. The precipitation unit consists of a heated nozzle, a collection unit within a collection chamber. The temperature-controlled nozzle is used to define the pre-expansion temperature. Heating is necessary to prevent condensation and blockage, and the temperature is typically much greater than the extraction temperature. The particles generated from the nozzle are captured by the collection substrate, which can be by either impaction or impingement (a solid or liquid collection substrate). The collection chamber can operate either at elevated or reduced pressure, so that precipitation can be investigated at

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Figure 14

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Schematic of an experimental apparatus for the rapid expansion of supercritical

solutions.

intermediate pressures, thus allowing greater flexibility in controlling the size and morphology of the particles generated. The major variables that can be controlled in such a system are extraction temperature and pressure, postextraction concentration and composition (addition of cosolvents), pre-expansion temperature, postexpansion temperature, and postexpansion pressure. Relative to spray drying, RESS history is quite short with the bulk of the literature first appearing in the early 1980s, and in which the first extensive investigations focused on polymers and inorganics. It was not until the mid-1980s that RESS was shown to be capable of generating microfine pharmaceutical powders. In 1984, Krukonis reported generating particles of ␤-estradiol using supercritical CO2 (66). Starting with particles that were in the range of several to hundreds of microns, the resultant powder was in the range of 1 ␮m and less. Since then, RESS has been used to generate powders of phenacetin (67), lovastatin (51,52,68), and ␤-carotene (69), as well as drug-loaded polymer microspheres using poly(hydroxyacid) and poly(D,Llactic acid) (51). For lovastatin, a two-level factorial designed experiment was conducted investigating the factors of concentration, postexpansion temperature, and preexpansion temperature (68). The results indicated that the particle size of the generated powder (approximately 100–300 nm) was relatively insensitive to changes in the process conditions investigated, yet crystallinity was maintained while achieving a significant reduction in particle size. Within the same report, a similar study was conducted with supercritical CO2 and naphthalene. Naphthalene was chosen as a model compound because the thermodynamics of the system are very well known. Johnson and Penniger found that the naphthalene

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behaved differently in that the particle size of the powder generated was significantly affected by the selected operating parameters (52). RESS offers the possibility to comminute and produce mixtures in a single step under less harsh conditions than would be used in conventional procedures. Processing with RESS is dependent on finding an appropriate solvent system in which the compound of interest is soluble. It also has the possibility of eliminating the use of organic solvents such that recrystallization and size reduction can be carried out in a single step. B.

Supercritical Antisolvent Techniques

There are two approaches to performing a GAS recrystallization similar to traditional recrystallizations using antisolvents. The conventional approach is to add an antisolvent to a solution of the compound of interest, thus causing it to precipitate. The other approach is to add the solution to the antisolvent in order to precipitate the compound. The latter approach is sometimes referred to as reverse recrystallization because the order of addition of the antisolvent is opposite to that typically used. Reverse addition recrystallization is not usually used to purify the material but serves to alter or control the particle size and shape. The use of supercritical fluids as antisolvents has been referred to as GAS or SAS recrystallization. Its gain in popularity over RESS stems from the fact that it is inherently more flexible and broadly applicable; SAS requires that the compound be practically insoluble in the supercritical fluid, which is usually the case, whereas RESS is limited to compounds that are soluble in supercritical fluids. Figure 15 is a schematic of an apparatus used for SAS recrystallization (51). The solution and the supercritical fluid are pumped into the crystallizer simultaneously. The liquid solution is aerosolized by pumping through a small orifice (i.d. 20–70 ␮m), and the droplets come in contact with the supercritical fluid antisolvent. This gives rise to rapid expansion, evaporation, and dilution, causing the compound to precipitate out. The mixture then leaves the crystallizer, and the solid is filtered from the fluid. It has been shown that, using this type of SAS approach, it is possible to produce micron-sized protein particles (51,63,70,71). In one set of experiments, a solution of insulin in ethanol and water was injected concurrently with supercritical CO2 into the crystallizer held at 90 bar and 35°C. The insulin particles generated were of two morphologies: microspheres of approximately 1 ␮m or less, and thick needles which were about 5 ␮m long and 1 ␮m in width. Using the same setup, a solution of catalase in ethanol and water yielded spherical particles that were either spherical or rectangular and approximately 1 ␮m in size (51). Dimethylsulfoxide and N,N-dimethylformamide solutions of insulin concurrently sprayed with supercritical CO2 yielded insulin particles in the 2 to 4 ␮m range. The biological activity of the powders generated was tested in rats, and the results indicated that there was no difference in activity between the processed and the unprocessed insulin. The results indicate that the use of SASs has the potential of generating microfine peptide and protein particles in the respirable range without the loss of activity (70).

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Schematic of an experimental apparatus for the reverse-addition supercritical antisolvent recrystallization (solution is added to antisolvent). Abbreviation: SAS, supercritical antisolvent.

Figure 15

This approach of continuous introduction either into a chamber filled with gas or into a continuous stream of gas is effective for producing small particles and lends itself very well to a continuous recrystallization process. Drawing an analogy from classical recrystallization approaches, the procedure of injecting solution into the SAS can be called reverse-addition SAS (RA-SAS) recrystallization. The alternative approach of injecting the supercritical fluid antisolvent into a solution of the compound, which is the more conventional order of mixing, to obtain a precipitate, can be referred to as normal-addition SAS recrystallization, or normal-addition supercritical antisolvent (NA-SAS). NA-SAS is more of a recrystallization technique, whereas RA-SAS can be considered a precipitation technique. It is believed that NA-SAS recrystallization offers more control over the final particle characteristics (72). Figure 16 is a schematic of a system that could be used to carry out NASAS (72). It is probably the simplest of all the apparatuses using supercritical fluids. A solution of the compound is introduced into the vessel, which is then sealed. Supercritical fluid is slowly admitted into the bottom of the chamber and bubbles up through the solution. As the gas passes through the chamber, it dissolves in the solution, expanding the liquid phase and creating a supersaturated solution. Eventually, the solution will be diluted with antisolvent to achieve sufficient supersaturation that nucleation occurs. The precipitate solid

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Schematic of an experimental apparatus for the normal-addition supercritical antisolvent recrystallization (antisolvent is added to solution). Abbreviation: SAS, supercritical antisolvent.

Figure 16

is filtered from the solvent through the bottom of the pressure vessel, and fresh gas can be introduced to remove trace solvent. Clearly, the rate of addition will affect the quality, size, and morphology of the particles generated. Using the NA-SAS method described, a series of four steroidal compounds (prednisolone, dexamethasone, flunisolide, and triamcinolone acetonide) were studied (72). All of these compounds have very low solubilities in supercritical CO2 (⬍0.002 wt%), and therefore SAS recrystallization would be more appropriate to use than RESS recrystallization. All of the results indicated that NA-SAS was capable of generating particles in the respirable range. For example, prednisolone recrystallized from a solution initially at 5000 psi and 120°C, generated particles that had a reasonably uniform particle size distribution, and the particles appeared as loose aggregates of primary particles that were 1 ␮m or less. In addition, it should be possible to use different organic solvents to obtain different particle morphologies or crystal habits (i.e., plates or needles) similar to classical recrystallization techniques. Hybrids of using supercritical fluids and spray-drying technology are in existence, which involve using a spray nozzle where multiple streams mix in the spray head. Bradford Particle Design Ltd (now Nektar) use a process known as solution-enhanced dispersion of supercritical fluids (SEDS) to produce dry particles with a controlled size distribution. The SEDS technique involves taking an aqueous solution of the drug, then decreasing the solvating power of

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the water by saturating it with CO2 under supercritical conditions; the drug solution and a stream of supercritical CO2 are mixed using a coaxial nozzle. The high-velocity, turbulent, supercritical fluid stream breaks up the aqueous solution into very small droplets. A third stream containing an organic solvent can be used to overcome immiscibility problems between the aqueous and the supercritical CO2 phases. The three-nozzle process enables proteins to be exposed to denaturing conditions for a minimal amount of time (73). Control over the size of the particles produced is achieved by variation of a number of process variables, including the flow rates of the three input streams to the nozzle and the pressure drop across the nozzle. The SEDS technology has been used to successfully process proteins with no denaturation associated with processing (73). The production of plasmid DNA-loaded particles using the SEDS process has also been investigated. The SEDS process was destructive to unprotected plasmid DNA; however, by buffering the aqueous feed solution and using mannitol as an excipient, a recovery of 80% of the original supercoiled proportion of the DNA was observed (74). Another process that uses a mixing head/nozzle in a spray-dryer set up is called CO2-assisted nebulization with a bubble-dryer (CAN-BD) (75–79). This involves a solution or suspension in acetone, alcohol, or water that is mixed intimately with CO2 at 100 bar to form an emulsion. The emulsion is rapidly expanded to atmospheric pressure through flow restrictor to generate aerosols of microbubbles and microdroplets, where the aerosol plume is dried at 1°C to 60°C, as it mixes with nitrogen or air in the drying chamber. The mixing head is placed on a standard spray dryer, and conventional collection technology is used to collect the dry fine powders. RESS and SAS recrystallization are techniques that are capable of generating respirable particles. Many factors have to be considered in the selection of the appropriate technique. Based on the need to have reasonably high solubilities in supercritical fluids to make RESS commercially viable, it appears that SAS, SEDS, and CAN-BD are more broadly applicable. SAS recrystallization can be carried out in the reverse-addition mode (solution to antisolvent) or normal-addition mode (antisolvent to solution), and each has its own advantages and disadvantages. A recent review of super critical fluid technology in the generation of pulmonary pharmaceuticals concluded that there appeared to be no data that clearly showed that using this technology has distinct advantages over conventional methods of particle generation, and any improvement would be drug dependent (79).

IV. A.

Properties of Spray-Dried and SCF Powders and Methods of Characterization Particle Size Distribution

The most significant feature of a powder for inhalation is the particle size distribution, whether it is delivered by metered dose inhaler, dry powder inhaler, or

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nebulizer. The versatility of spray-drying and SCF methods lies in their ability to manipulate and control the size distribution in a one-step process. The principal methods for measuring the particle size distribution of powders in the respirable size range (i.e., 1–10 ␮m) are inertial impaction, time of flight, laser diffraction, SEM, electrical zone sensing, and phase/Doppler anemometry. As these methods are described in detail in this text and elsewhere (80,81), it will only be noted that all of these technologies have advantages and disadvantages, and more than one should be used to characterize the size distribution, since different methods yield complementary information. Inertial impaction and time-of-flight techniques measure the aerodynamic particle size distribution, which includes a combined effect of size, shape, and density, and in this sense these methods are the most relevant with regard to pulmonary drug delivery. Inertial impaction provides a mass distribution, and time of flight affords a number distribution function. Although transformation from one function to the other is easily made for perfectly spherical particles, errors occur for irregularly shaped particles, so it is better to measure the distribution functions directly. Fraunhofer laser diffraction measures the volume (or mass) distribution function of an optical-equivalent diameter and is well suited for studying the primary particle size distribution of the powder in its dry state due to the advances of dry powder dispersers (82). Although, in principle, SEM can measure a number distribution of the primary particles using image analysis software, in practice it is exceedingly difficult to prepare a sample of a fine powder without significant aggregation of particles, which obscures the analysis. The electrical zone-sensing method requires dispersing the particles in a liquid medium, and as a result, surfactants are required, and aggregation can be a major problem for some drugs. The phase/Doppler anemometer measures both the size and the velocity distribution of an aerosol or powder, but the size measurement is not applicable to nonspherical particles or droplets containing suspended solids. This technique is most useful, however, in characterizing the spray droplet size distribution of solutions from nozzles used in the spray dryer or SCF equipment. B.

Particle Shape and Surface Roughness

One of the features of spray-dried and SCF-generated particles is their spheroidal shape in contrast to micronized powders, which display irregular fragments of the starting crystals. These fragments can have planar faces that stack together to form strong aggregates. Spheroidal shapes have smaller contact areas between particles, which can lead to weaker aggregates than would otherwise be observed (83,84). Characterization of particle shape and surface roughness can range from qualitative description based on visual observation by SEM to more quantitative analyses using Fourier synthesis (85), asymmetry analysis using forward scattered light (86), and fractal dimensions (87–89). Although research is progressing in understanding the relationship of these static shape parameters to powder dynamics such as flow and rheological properties (90,91)

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as well as aggregation kinetics (92), which are important factors in powder processing, formulators of respiratory products are also interested in the “dynamic shape factor” which controls aerodynamic behavior. Rapid measurement techniques for the dynamic shape factor have not been developed for the fine particles used in inhalation products, but time-consuming measurements can be made for larger particles that demonstrate the significance of dynamic shape analysis (93–96). C.

Dispersion

Dispersion of a powder is a broad field of study, encompassing many areas of powder science and technology, including gas fluidization (84,97,98), adhesion (99–101), and agglomeration (102). The practical importance of dispersion with regard to inhalation therapeutics lies in the ability of a given dry powder inhaler to consistently deliver a dose of drug/excipient powder over the shelf life of the product. The two major problems are (i) efficiency of dispersion and (ii) reproducibility of dispersion. Formulators of respiratory products can overcome these challenges by improving the dispersion properties of the powders and by innovative design of devices; often the former approach is a simpler, less resource-consuming endeavor than the latter. Despite its significance in pharmaceutical performance of inhalation formulations and devices, the dispersibility of a powder is an often overlooked property, perhaps because there has been no established technique in the literature, although various methods have been explored, such as supersonic jet nozzles (103), air annular jet nozzles (104), and flow-through needles (105). This deficiency notwithstanding, the success in generating a powder by spraydrying or SCF methods can be directly assessed and compared to the micronized material by analyzing its dispersibility. One method for dispersing a powder to various degrees is shown schematically in Figure 17 (82). A chamber is pressurized to a variable extent, and the air is forced to flow through an annular gap. A powder is fed into the funnel and accelerated by the Venturi effect into the gap region where it is impinged and dispersed by the high-velocity air, ultimately forming a spray at the nozzle exit.

Figure 17 Schematic drawing illustrating the operating principle of the RODOS® dry powder disperser (Sympatec Inc.).

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The spray can then be analyzed by various means, most notably laser diffraction for which this disperser was designed. For a given powder, the dispersing air pressure can be varied to determine a relative measure of its dispersibility. For example, Figure 18 gives the results for a micronized powder of an organic substance at several dispersing pressures. SEM revealed the presence of large aggregates in the powder, and laser diffraction analysis demonstrated that a certain minimum pressure was needed to completely break up these aggregates. Such measurements provide a quick analysis of whether or not a given powder would be suitable for a specific inhalation device. Dispersion measurements can also be used to compare different powder formulations. Figure 19 shows the particle size distribution of spray dried and micronized formulations of the same substance, dispersed at the same pressure. Clearly, the spray-dried powder was more easily dispersed than the micronized material. Thus, although there is no absolute method for measuring the powder’s dispersibility, the available technology provides rapid analyses for relative determination of this important property. D.

Dryness

In the section on spray drying, it was noted that, depending on the equilibrium moisture content of a material and the relative humidity at the drying chamber outlet, it is possible to generate powder with unacceptable levels of water or solvent (in the case of organic work). Thus, the moisture or solvent content must be measured to assure that proper drying is accomplished.

Figure 18

Fraunhofer laser diffraction using the RODOS® disperser at various pressures.

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Figure 19

Fraunhofer laser diffraction using the RODOS® dry powder disperser at 3 bar.

Several techniques are available for aqueous and organic solvents, among them gravimetric loss-on-drying and thermal gravimetric analysis. For aqueous spray drying, Karl Fischer titration can be used to detect moisture levels as low as approximately 10 ppm. In the case of organic spray drying, an assay can be developed for measuring the residual solvent content in the solid, e.g., by gas chromatography headspace analysis. Hydrate and solvate formation can be detected by spectroscopic techniques, such as infrared (IR). E.

Crystallinity

In generating powders by spray drying or SCF, it has been noted in previous sections that the crystalline structure of the particles may be quite different from the recrystallized form or forms due to the rapidity and dynamics of the particle-formation processes. Indeed, the short time scales involved in spray drying, for example, often lead to amorphous solid particles. Thus, it is important to characterize the crystallinity of the generated particles (106). X-Ray Diffraction

Powder X-ray diffraction is used to determine if there is any observable crystal structure present. If there is crystalline structure, it should be compared with the structure of recrystallized forms to assess if a new

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polymorph or mixture thereof was generated. If a mixture of crystalline and amorphous material was produced, various techniques in addition to powder X-ray diffraction such as pycnometry, vapor sorption, and calorimetry can be used to quantitate the fraction of crystalline or amorphous content (107,108). It is difficult to measure less than 5% to 10% amorphous content with X-ray diffraction; for low levels of disorder, moisture sorption (108) and isothermal microcalorimetry (106,109,110) are particularly sensitive. Calorimetry

Calorimetry studies should be conducted to determine the thermal properties, most notably the thermal stability, if the substance is a glass. As mentioned earlier, glasses undergo a transition (Tg) to a supercooled liquid state (or rubbery state for sufficiently long, flexible chain polymers), which can be detected by a change in heat capacity, particularly in adiabatic calorimetry with large sample mass (38,39,106,109–112). If Tg is just above room temperature, physical stability will be a concern, specifically with regard to the primary particle size distribution, since recrystallization and particle growth will be more rapid the closer the temperature is to the glass transition. Although calorimetry studies can determine quite quickly if problems can be expected for the powder on a stability program, it should be noted that recrystallization kinetics of a glass can vary considerably depending on molecular constitution and the crystal structure to which the glass irreversibly transforms. Thus, if Tg is above the highest temperature to be examined in the stability program (namely, 40°C), depending on the chemical structure, recrystallization may or may not be a significant problem. Moisture Sorption

In general, an amorphous phase will absorb more moisture at a given relative humidity than the crystalline forms. Thus, if the spray-dried or SCF-generated material has amorphous content, it will be important to establish the amount of absorbed moisture as a function of relative humidity (i.e., the moisture uptake isotherm). The main problem with regard to respirable particles is that water can plasticize an amorphous phase (i.e., lower Tg) and thus enhance recrystallization kinetics and particle growth at a given temperature due to fusion of the particles. Even relatively low levels of amorphous material (<10%) may have a detrimental impact on the stability, manufacturability, and dissolution characteristics of the formulated drug product. Measuring the moisture sorption isotherm at an early stage in development can detect if problems are likely to be encountered with the physical stability of the particle size distribution. The molecular interpretation of moisture interactions on solidstate properties has been reviewed for pharmaceutical compounds (113). Gravimetric vapor sorption or Dynamic Vapor Sorption are powerful techniques in quantifying the amount of amorphous material present (114–121), where detection limits of much less than 1% have been reported, and through the use of relative humidity (RH) pulse technique, limits of detection of less than 0.1% have been reported (114,115).

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Miscellaneous Characterization Methods

There are many other chemical, solid-state, and powder characterization techniques that are needed to fully examine the properties of a generated powder. Such methods include impurities analysis (e.g., high pressure liquid chromatography), spectroscopic studies (e.g., solid-state NMR and IR), inverse gas chromatography, atomic force microscopy, and powder flowability (e.g., shear cell testing or bulk density measurements). The flowability of a powder is rarely improved by spray-drying or SCF methods, and this property should not be confused with dispersion, because the two properties are not directly related. For example, micronized, spray dried, and SCF-generated powders in the 1 to ⫺5 ␮m primary particle size range normally have extremely poor flow properties (unless the particle density is very large), but they may be very easily dispersed by a dry powder inhaler or other such device. Spray drying can be used to improve flow properties, but only by operating under conditions in which agglomeration occurs to produce particles of sufficiently large mass for gravity to overcome interparticle van der Waals and other forces. The relevance of flowability in spray-dried and SCF powders is much more in manufacturing-related issues, such as powder conveying, feeding, and device-filling operations, than in pharmaceutical performance of an inhaler product. References 1. 2. 3.

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Yamamoto H. Relationship between adhesive force of fine particles and their dispersibility in gas. Proceedings of the 2nd World Congress on Particle Technology, Japan, September 19–22, 1990:167–173. Buckton G, Darcy P. Assessment of disorder in crystalline powders—a review of analytical techniques and their application. Int J Pharm 1999; 179:141–158 Pikal MJ, Lukes AL, Lang JE, Gaines K. Quantitative crystallinity determinations for 0-lactam antibiotics by solution calorimetry: correlations with stability. J Pharm Sci 1978; 67:767–773. Saleki-Gerhardt A, Ahlneck C, Zografi G. Assessment of disorder in crystalline solids. Int J Pharm 1994; 101:237–247. Buckton G, Darcy P, Mackellar A. The use of isothermal microcalorimetry in the study of small degrees of amorphous content of powders. Int J Pharm 1995; 117:253–256. Fukuoka E, Makita M, Yamamura S. Some physicochemical properties of glassy indomethacin. Chem Pharm Bull 1986; 34:4314–4321. Suga H, Seki S. Thermodynamic investigation on glassy states of pure simple compounds. J NonCryst Solids 1974; 16:171–194. Takeda K, Yamamuro O, Suga H. Calorimetric study on structural relaxation of 1-pentene in vapor-deposited and liquid-quenched glassy states. J Phys Chem 1995; 99:1602–1607. Ahlneck C, Zografi G. The molecular basis of moisture effects on the physical and chemical stability of drugs in the solid state. Int J Pharm 1990; 62:87–95. Levoguer CL, Booth J. Polymorphism in Spray Dried Lactose, Surface Measurement Systems Application Notes. Mackin L, Zanon R, Park JM, Foster K, Opalenik H, Demonte M. Quantification of low levels (<10%) of amorphous content in micronised active batches using dynamic vapour sorption and isothermal microcalorimetry. Int J Pharm 2002; 231(2):227–236. Buckton G, Darcy P. The use of gravimetric studies to assess the degree of crystallinity of predominantly crystalline powders. Int J Pharm 1995; 123:265. Forbes RT, Davis KG. Water vapor sorption studies on the physical stability of a series of spraydried protein/sugar powders for inhalation. J Pharm Sci 1998; 87:1316. Shamblin SL, Hancock BC, Zografi G. Water vapor sorption by peptides, proteins and their formulations. Eur J Pharm Biopharm 1998; 45:239. Burnett D, Thielmann F, Booth J. Determining the critical relative humidity for moisture induced phase transitions. Int J Pharm 2004; 287:123. Hancock B, Zografi G.Characteristics and significance of the amorphous state in pharmaceutical systems. J Pharm Sci 1997; 86:1–12. Yu L. Amorphous pharmaceutical solids: preparation, characterization and stabilization. J Adv Drug Delivery Rev 2001; 48(1):27–42.

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12 Interfacial Phenomena and Phase Behavior in Metered Dose Inhaler Formulations

KEITH A. JOHNSON Adroit Pharmaceutical Development, LLC, Durham, North Carolina, U.S.A.

I. A c cii cij e g Ho k K Kexp Kh Kemp L Mi NA Pvap Pc r

Nomenclature Hamaker constant Concentration Cohesive energy density for pure compound i Cohesive energy density for a mixture of i and j Electron charge Gravitational acceleration Separation distance between particle surfaces Boltzmann’s constant Cohesive energy density difference between two pure components and their mixture Experimentally determined K K calculated using harmonic mean mixture rule Kh with empirically adjusted coefficient Length of adsorbed surfactant molecule Concentration of ionic species i Avogadro’s number Vapor pressure Critical pressure Particle radius 347

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R T Tc Tcs u vs VA VE Vs VT Vi x zi ␦ ⌬Hvap ␧o ␧r ␥ ␩ ␳ ␺o ␨

II.

Gas constant Temperature Critical temperature Consolute temperature Electrophoretic mobility Settling velocity Attractive potential Electrostatic potential Steric potential Total interparticle potential Molar volume of species i Distance between particles’ centers (Ho + 2r) Ion valance of species i Solubility parameter Enthalpy of vaporization Permittivity in vacuum Dielectric constant Activity coefficient Viscosity Density Surface potential Zeta potential

Introduction

A metered dose inhaler (MDI), contains a liquefied propellant held under pressure in a suitable container fitted with a metering valve. The drug is present as a solution or as a suspension of fine particles. Of particular interest in the field of colloid and interface science is the stabilization of fine particle suspensions. This chapter will focus on theory and experiments for developing suspension formulations for MDIs. To be respirable, the drug particles should have mass median aerodynamic diameters in the range of 1 to 3 ␮m (1). This size range is commonly obtained by milling drug crystals into a fine powder (2). Next, this fine powder must be dispersed in the nonaqueous propellant to form a suspension. The suspension must be stable enough to ensure accurate filling during manufacture and accurate dosing when used by the patient. Suspension stability, or resistance to flocculation, can be optimized using combinations of excipients (3). The number of excipients available to the MDI formulator is limited. Historically, three chlorofluorocarbon (CFC) propellants were approved for use: trichlorofluoromethane (Propellant 11), dichlorodifluoromethane (Propellant 12), and dichlorotetrafluoroethane (Propellant 114). However, these CFCs are being phased out of use in MDIs. Oleic acid, lecithin, and sorbitan trioleate are approved surface–active agents, or surfactants, for use in MDIs. Ethanol

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can be used as a cosolvent or a slurrying aid. Two hydrofluoroalkanes (HFAs) are being developed as “ozone–friendly” propellants for MDIs: 1,1,1,2–tetrafluoroethane (Propellant 134a) and heptafluoropropane (Propellant 227). Albuterol, beclomethasone dipropionate, fluticasone propionate, ipratropium, and levalbuterol are currently marketed in Propellant 134a MDIs. In this chapter, important physicochemical principles for formulating MDI suspensions will be reviewed. Since formulators continue to evaluate new propellants and excipients in MDIs, methods for estimating propellant physical properties and phase behavior will be presented. Experimental methods and design strategies to evaluate MDI formulations will also be discussed. Even though CFC propellants are being phased out, their physicochemical properties will be covered based on their historical importance to MDI technology.

III.

Estimating Suspension Stability

During the manufacture and use of MDIs, it is important to maintain the consistency of the suspension. Otherwise, the amount of drug filled into each inhaler may vary outside specifications. Furthermore, an unstable suspension could lead to large dose variations during use by the patient (4). There are several comprehensive reviews on suspension stability of colloidal particles (5–7). In this section, theories for estimating suspension stability will be discussed without significant attention to the derivation and assumptions required to obtain the result. Instead, emphasis will be placed on the theoretically and practically important physical and chemical properties of the formulation that an experimenter can use during product development. A.

Interparticle Potentials

The interactions between suspended particles can be placed into three categories: attractive, electrostatic, and steric (5–7). According to the theory of Derjaguin, Landau, Verwey, and Overbeek, DLVO theory, the interparticle pair potentials are calculated as a function of separation and summed to obtain the total interaction potential, VT. Next, VT is compared to the product of Boltzmann’s constant, k, and temperature, T. As the particles approach, VT should have a peak or potential energy barrier greater than 15kT for the suspension to be stable (8). Attractive

Attraction between suspended particles reduces suspension stability and causes flocculation. For nonpolar or slightly polar compounds, the main sources of attraction are induced dipole–induced dipole interactions known as the London–van der Waals forces (5–7). The attractive potential between the two particles, VA, is a function of separation (8): VA ⫽ –Ar/12Ho

(1)

where A is the Hamaker constant, r is the particle radius, and Ho is the distance between the particles’ surfaces. The Hamaker constant varies with the separation

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distance and rapidly decreases at large separation distances due to the retardation effect (9). For use in approximate calculations, we will assume that A is a constant. The form of Eq. (1) will change when two particles with different sizes interact. However, for simplicity, it is assumed that the suspended drug particles are of uniform size. The formulation variables of interest in Eq. (1) are the particle radius and the Hamaker constant. For MDI formulations, r will range from approximately 0.5 to 1.5 ␮m because this is the size range for optimal deposition in the lung (1). The other variable in Eq. (1), the Hamaker constant, depends on the molecular structure of the particle and suspending medium. Table 1 contains Hamaker constants for various substances interacting across a vacuum or air. There are two methods for estimating the Hamaker constant: the microscopic approach of Derjaguin and the macroscopic approach of Lifshitz (10,11). The suspending medium affects the value for the Hamaker constant. For material 1 suspended in medium 2, the following expression can be used to estimate A (12): A ⫽ (A101/2 ⫺ A201/2)2

(2)

where A10 and A20 are the values for materials 1 and 2 in vacuum, respectively. For example, sapphire suspended in CC14 would have a Hamaker constant of 2.2 × 10⫺20 J, which is significantly lower than the value in air or vacuum. From Eq. (2), A is smallest when the Hamaker constants of the suspending medium and particle are similar. Figure 1 contains a plot of VA, dimensionless through division by kT at 25°C, for two particles where r ⫽ 0.5 ␮m and A ⫽ 5.0 × 10⫺20 J. The absolute value of VA is very small at large distances, and becomes larger as the particles approach each other. It is important to note that VA is always negative. To prevent aggregation, there must be a repulsive potential to overcome the attraction. Electrostatic

Particles suspended in nonpolar liquids can acquire a surface charge (13–15). Such charges are possible by particle interaction with surfactants, the suspending media, trace water, or trace impurities that contain an ionizable moiety. The interaction between water and surfactant on suspension stability is particularly Table 1 Values of the Hamaker Constant for Various Compounds Across Vacuum or Air Material Teflon Water Hexadecane Polystyrene Carbon tetrachloride Sapphire TiO2 Source: From Refs. 10, 11.

A (10⫺20 J) 2.7 3.5 5.2 5.4 5.5 14.7 43

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Figure 1 Dimensionless interparticle potential, in kT units, at 25°C for two particles with r ⫽ 0.5 ␮m. V A was calculated using Eq. (1) with A ⫽ 5.0 × 10 -20 J. V E was calculated using Eq. (4) with ␧r ⫽ 2.2, ␺o ⫽ 38mV, and k⫺1 ⫽ 15 ␮m. The sum of VA and VE has a peak with a potential energy barrier greater than 15kT, which indicates that this suspension is stable.

relevant to MDIs and will be discussed later in the chapter. In this section, our focus is on the magnitude of surface charge, or potential, required to prevent aggregation of particles. If particles obtain a charge, the surface will attract counterions while repelling any ions having the same charge. This leads to the formation of a diffuse electrical double layer around the particle. The reciprocal of the double–layer “thickness,” ␬, is given by the following equation (16): ␬2 ⫽ 2e2NA⌺Mizi2/␧r␧okT

(3)

where e is the electron charge, NA is Avogadro’s number, Mi is the concentration of species i, zi is the ion valence of species i, ␧r is the dielectric constant, and ␧o is the permittivity in vacuum. The overlap of electrical double layers surrounding the suspended particles will oppose attraction. For very small electrolyte concentrations, such as those in nonaqueous media, the double–layer thickness is large. For example, if ␬ is 0.1/␮m ⫺1, the double–layer thickness is 10 ␮m. When ␬r << 1, the following equation can be used to estimate the electrostatic potential, VE, between two spheres (17): VE =

4 πεr ε o r 2 ψ o 2 exp( − kH o ) x

(4)

where ␺o is the surface potential and x is the distance between the particles’ centers (x ⫽ Ho + 2r). As with the attractive potential, r will be limited to the

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range of sizes for respirable particles. The dielectric constant can be varied by blending propellants, including blends with ethanol. ␺o and ␬ are functions of several formulation variables, such as propellant blend, surfactant level, trace moisture, and the drug’s chemical structure. A formulator needs to know the surface potential required for electrostatic stabilization in a fine particle suspension. Morrison derived an analytical expression to estimate the minimum ␺o required to stabilize nonaqueous suspensions (17): ␺o2 ⬎ 1000/r␧r

(5)

where r and ␺o are measured in microns and millivolts, respectively. The average ␧r for CFC propellants is approximately 2.2 (18). Therefore, for particles with r ⫽ 0.5 ␮m in a CFC propellant, the minimum ␺o required for stabilization is approximately 30 mV. Electrophoretic mobilities for micronized inhalation compounds in trichlorotrifluoroethane (Propellant 113) have been reported by Wyatt and Vincent (15). The absolute values of these mobilities ranged from 0.1 to 8.3 ⫻10-10 m2/sec//V. Using the following equation, one can calculate the zeta potential ζ or potential at the shear plan, from electrophoretic mobility data (16): ζ=

3uη 2εr ε o

(6)

where u is the electrophoretic mobility, ␩ is the liquid viscosity, and all units are in SI. Using Eq. (6), the physical properties of Propellant 113 [␩ ⫽ 0.68 cP, ␧r ⫽ 2.4 (18)], and the data reported by Wyatt and Vincent (15), the micronized inhalation compounds had zeta potentials ranging from 0.5 to 38 mV. Because the potential decays slowly with distance in nonaqueous media, ζ is a good approximation for ␺o. In CFCs and HFAs, electrostatic stabilization is theoretically possible for some formulations where r is approximately 0.5 ␮m because values of ␺o greater than 30 mV have been measured. Figure 1 also contains a plot of the electrostatic potential (␺o ⫽ 38 mV, ␬⫺1 ⫽15 ␮m) and the total potential, VT ⫽ VA + VE, for two particles in a CFC propellant at 25°C. VT in Figure 1 has a peak with a height greater than 15kT. Therefore, theoretically, this suspension would be stable, which agrees with the result from Morrison’s analysis (17). Even though it is theoretically possible to electrostatically stabilize a suspension of particles in nonaqueous media, the experimental results can show instability (19–21). The relatively large ratio of double–layer thickness to interparticle separation in nonaqueous media reduces the electrostatic repulsion potential. As the particle volume fraction increases, VE decreases, and the suspension stability decreases. Albers and Overbeek derived an analytical expression for the reduction of VE with the following form (19): %VE ⫽ f (r, ␬, ␾)

(7)

where %VE is the fraction by which VE is lowered and ␾ is the volume fraction of particles. Table 2 shows %VE for a suspension, where r ⫽ 0.5␮m and

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Solid Volume Fractions, ␾, Concentrations in mg/mL, and the Lowering of the Electrostatic Repulsive Potential, %VE, for a Suspension Where r ⫽ 0.5 ␮m and ␬⫺1 ⫽ 15 ␮m

Table 2

␾ 0.001 0.002 0.005 0.015 0.05

mg/mL at ␳ ⫽ 1.25

%VE

1.25 2.5 6.25 18.75 62.5

93.4 89.6 83.9 71.5 53.9

Note: %VE was calculated using the analysis of Albers and Overbeek (19).

␬⫺1 ⫽ 15 ␮m, at various volume fractions. Even at fairly low volume fractions, VE can be significantly reduced. Figure 2 shows calculations for the total potential at three volume fractions: 0 (same as Fig. 1), 0.005, and 0.01. When ␾ ⫽ 0.005, the potential energy barrier for this suspension is insufficient to prevent aggregation (i.e., maximum value for VT < 15kT). Even though ␺o can be high enough to stabilize dilute suspensions, it may not be sufficient in the more concentrated suspensions of

Figure 2 Dimensionless interparticle potential (VA + VE), in kT units, at 25°C for suspensions with r ⫽ 0.5 ␮m at various volume fractions. The parameters in the calculations are the same as in Figure 1. The reduction in VE as a function of volume fraction was calculated using the analysis of Albers and Overbeek (19). Even with ␾ ⫽ 0.005, the potential energy barrier is insufficient to prevent aggregation.

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MDI products. However, there are additional mechanisms that can increase the stability of MDI suspensions, such as steric repulsion. Steric

The adsorption of surfactants and polymers on the particle’s surface results in a molecular film that can prevent aggregation. Studies on MDI formulations have shown that steric repulsion can make a significant contribution to suspension stability (22,23). This stabilization process has two components: entropic and enthalpic. The entropic contribution comes from the decrease in configuration entropy experienced by adsorbed molecular films as the particles approach each other (24). The enthalpic contribution comes from mixing the adsorbed molecular films in the presence of the suspending medium (25). If the suspending medium is a good solvent for the adsorbed molecules, this mixing process gives a positive excess osmotic pressure that prevents particle aggregation (26). From the entropy lost when two molecular films approach each other, the repulsive potential per adsorbed molecule, Vsm, can be calculated (24): Vsm ⫽ kT ln(2L/Ho)

(8)

where L is the length of the molecule or, in this case, the length of the surfactant’s hydrocarbon chain. The number of molecules involved in this interaction is a function of their concentration on the particle’s surface and the particle size. For suspension stability, contributions from steric repulsion must give a potential energy barrier greater than 15kT. Table 3 shows the number of molecules required to give a potential energy barrier greater than 15kT for several values of L. The attractive potential was calculated using Eq. (1) with r⫽0.5␮m and A ⫽ 5.0 × 10⫺20 J. With very modest numbers of adsorbed surfactant molecules, potential barriers greater than 15kT can be generated. As L increases, the number of molecules required to create a significant potential energy barrier decreases. While Eq. (8) is a simple model for the entropic interaction, it demonstrates that small numbers of adsorbed molecules can prevent aggregation in MDI suspension formulations. Theories for the enthalpic contribution to steric stabilization require an estimate of the interaction between the adsorbed molecules and the solvent (25). If the adsorbed molecules, or the surfactants’ hydrocarbon chains, favor solvation, the surfaces will repel each other as they approach. However, if the Table 3 Number of Molecules Required for Steric Stabilization of Two Particles Where r ⫽ 0.5 ␮m and A ⫽ 5 × 10⫺20 J at 25°C

Molecule length, L (Å)

Number of molecules required for potential energy barrier in VT ⬎ 15kT

10 50 100

705 155 85

VA was calculated from Eq. (1), and Vsm was calculated from Eq. (8).

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solution properties favor a phase separation, the surfaces will attract each other, which results in aggregation. For example, silicon powder dispersions in benzene or trichloroethylene can be stabilized by an adsorbed film of polymethyl methacrylate (27). When n–hexane, which is a nonsolvent for the polymer, was added to the suspension, it flocculated. The mixing of MDI surfactants with propellants will be discussed later in this chapter. B.

Effect of Surfactant and Water in Nonaqueous Media

Surfactants in suspension formulations can aid wetting, modify surface charges, and provide steric barriers against flocculation. Lecithin, oleic acid, and sorbitan trioleate are used as surfactants in MDI formulations. In nonaqueous media, surfactants interact with water to form a wide variety of microdomains (28). While the surfactant level can be carefully controlled, trace amounts of water will be variable. Water is an impurity in the propellant at the time of manufacture. Also, it diffuses through the metering valve during storage at room conditions or at extremes of temperature and humidity (2). Since water will always be present in the formulation, it is important to evaluate and minimize its effect early in product development. Trace amounts of water can significantly affect suspension stability by increasing the interparticle attraction. Capillary condensation of water vapor increases interparticle adhesion in fine powders (29). This phenomenon also occurs in nonaqueous media, sometimes at levels as low as 20% to 40% of the saturation level (30). Water levels may also cause significant drug adhesion on the canister. On the other hand, water can improve stabilization by enabling dissociation of surfactant ions adsorbed at the particles’ surfaces (31). Water can also alter the fluidity of adsorbed surfactant layers on the surface of particles (32). The effect of water on suspension stability in nonaqueous media does not follow a simple linear relationship. Malbrel and Somasundaran studied the suspension stability of colloidal alumina in hexane, with sodium dioctyl sulfosuccinate, or AOT, as a surfactant (33). As the water level was increased from 0 to 800 ppm, the suspension was unstable, then became stable, and finally unstable again. These suspensions had maximum stability with water levels ranging from 70 to 100 ppm. For a rutile/xylene/AOT system, McGown et al. have shown that the zeta potential increased as the water concentration increased from 0 to 80 ppm, followed by a decrease at higher water levels (34). Like the results from Malbrel and Somasundaran, the effect of water at constant surfactant level was nonlinear. With both of these studies, increasing the surfactant level at constant water concentration increased the suspension stability (33) or the zeta potential (34). C.

Particle Growth in Suspension

When preparing suspensions of fine particles, it is important to control the potential for particle size changes during storage. Due to interfacial tension between the suspending medium and the particle, the solubility of small particles is higher than that for large particles. Therefore, if the particle size distribution is wide, smaller particles will tend to dissolve and precipitate on larger particles. The net effect of the process, known as Ostwald ripening, is an increase in the

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average particle size with time (35). This can be detrimental to MDI product performance because effective respiratory drug delivery is only possible in a narrow size range. Theories to predict the Ostwald ripening process have been developed (36,37). Calculations using these theories give estimated particle growth rates based on formulation properties such as width of the particle size distribution, interfacial tension, drug solubility and diffusivity in the suspending medium, and particle volume fraction. However, when the results from these theories were compared with experiment, Kabalnov et al. found that the growth rate was underpredicted by a factor of 2 (38). They noted that the theories were based on immobile particles while actual suspended particles were in Brownian motion. For their system, Kabalnov et al. found that the Peclet number, which compares the mass transfer rates by diffusion and convection, was greater than 1. This means that the mass transfer processes of Ostwald ripening would be faster than predicted by diffusion alone. To minimize Ostwald ripening, the first priority is to minimize drug solubility in the suspending medium. Techniques for controlling solubility by chemical modification, surfactant micelles, and solvents have been reviewed (39). In addition, Miller has demonstrated the importance of trace water levels on particle growth (2). Solubility experiments should be performed on the fine drug particles rather than the coarse, premilled product. Techniques for measuring drug solubility have been described by Dalby et al. (40). Phillips et al. have developed a method for early identification of ripening problems using microscopy (41). IV.

Propellant Physical Properties and Phase Behavior

A.

Physical Property Estimation

Some physical properties for CFC and HFA propellants are listed in Table 4. While the physical properties of these propellants are known over a wide range of temperatures, formulators are continuing to evaluate new propellants that Table 4

Physical Properties of Chlorofluorocarbon and Hydrofluoroalkane Propellants

Propellant

Normal boiling point (°C)

Vapor pressure (psi)

Liquid density (g/cm3)

Liquid viscosity (cP)

Critical temp. (°C)

Critical pressure (bar)

23.8

12.7

1.49

0.43

198

43.5

⫺29.8

81.5

1.33

0.22

112

40.6

3.77

26.0

1.47

0.36

146

32.1

–26.2

96.6

1.21

0.20

101

40.6

–16.5

66.0

1.39

0.24

102

29.5

11

CC13F 12 CC12F2 114 CCIF2CIF2 134a CF3CFH2 227 CF3CFHCF3

The vapor pressure, density, and viscosity are at 25°C. Source: From DuPont, Wilmington, Delaware, U.S., and ICI, Runcorn, U.K.

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give favorable physical properties to the inhaler while reducing or eliminating the use of CFCs. When evaluating new or novel propellants, data such as those in Table 4 are not always available. With minimal data, physical property correlations can estimate vapor pressures and densities for CFC and HFA propellants over a wide range of temperatures (42). Methods for estimating vapor pressure and density are discussed in this section. Vapor Pressure

The propellant’s vapor pressure, Pvap, is required to design container/closure systems and to develop manufacturing equipment and processes. Furthermore, it affects aerosol generation and the distribution of drug in the lungs (43). When Pvap is unknown, it can be estimated using physical property correlations based on corresponding–states theory (44). To use the vapor pressure correlation of Gomez–Nieto and Thodos, the propellant’s normal boiling point (Tb), critical temperature (Tc), and critical pressure (Pc) are required (45). In Figure 3, vapor pressure estimations using Gomez–Nieto and Thodos’ correlation are compared with data for Propellant 134a. There is good agreement between the correlation and experiment over a wide range of temperatures. Sometimes, the critical properties or boiling points are unknown. The boiling point can be easily measured. Critical properties and boiling points can be estimated from molecular weight and structure by group contribution methods. Using Ambrose and Patel’s group contribution method (46), the calculated Tc and Pc for Propellant 134a were 105°C and 52 bar, respectively. The vapor pressures calculated using these estimated critical properties are also shown in Figure 3. Even though the critical pressure estimate is off by 25%, the values of Pvap, especially at lower temperatures, are still reasonably accurate. The difference between the two lines in Figure 3 demonstrates that errors can arise unless critical property inputs are accurate. Propellants can be blended to obtain a wide range of pressures between those of the pure components. If a pressure correction is not required for the vapor phase fugacity coefficients, the vapor pressure of a blend, P, can be calculated by (47) ⌺xi␥i(T, P, xi)Pvap(T) ⫽ ⌺yiP ⫽ P

(9)

where xi is the mole fraction of species i in the liquid phase, yi is the mole fraction of species i in the vapor phase, and ␥i is the liquid phase activity coefficient. If the liquid phase is ideal, ␥⫽l. Otherwise, ␥ can be estimated using methods such as solubility parameter theory or UNIFAC (44). Vapor pressure data can be used to estimate other physical properties. By plotting the natural log of vapor pressure versus 1/T, the heat of vaporization, ⌬Hvap, is found using the Clausius–Clapeyron equation (48): InP vap =

−∆H vap +C RT

(10)

where C is a constant. Knowing ⌬Hvap enables calculation of the propellant’s cohesive energy density using the following equation (49):

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Figure 3 Vapor pressure for Propellant 134a as a function of temperature. Line A is the result from Gomez–Nieto and Thodos’ method (45) using critical properties from Table 4. Line B is the calculation using critical properties estimated by Ambrose’s method (46). The squares are data from ICI (Runcorn, England).

cii ⫽ ⌬Uivap/Vi ⫽(⌬Hivap ⫺ RT)/Vi

(11)

where ⌬Uivap is the internal energy of vaporization and Vi is the molar volume. Solubility parameters, ␦, are calculated from cohesive energy densities using the following equation (49): δ ii = c1ii/ 2

(12)

The use of solubility parameters to estimate phase behavior in MDI systems will be discussed later in this section. Density

The propellant’s liquid density, ␳, is another important property for developing MDI formulations. Like vapor pressure, ␳ is a function of temperature. Density establishes the ratio of mass to volume, which is important in design of container/closure systems, metering valves, and manufacturing equipment. Density affects the creaming and settling velocity of particles in suspension formulations, vs, as shown in Stokes’ equation (50): vs ⫽ 2⌬␳r2g/9␩

(13)

where ⌬␳ is the density difference between the particle and the propellant, g is gravitational acceleration, and ␩ is the viscosity. Table 4 shows that the densities of MDI propellants vary over a wide range. If the density of the drug and propellant are matched at some temperature, then the settling velocity is zero. In the absence of data, density can be estimated using physical property correlations. The correlation of Bhirud, based on corresponding–states

Interfacial Phenomena and Phase Behavior

359

theory, requires Tc, Pc, and the acentric factor to estimate ␳ (51). The acentric factor, which accounts for nonsphericity of the molecule, can be estimated from vapor pressure data (52). The results of density calculations for Propellant 134a are in Figure 4 along with data from experiments. There is good agreement between the estimation and the experiments over a wide temperature range. The density calculation using the estimated critical properties of Propellant 134a is also shown in Figure 4. Using the estimated critical properties results in a large error for the density calculation. The density estimate is more sensitive to the accuracy of critical properties than the vapor pressure estimate (Fig. 3). Two or more propellants can be blended to obtain a density between the pure components. This is very advantageous when considering the potential to minimize settling or creaming velocities by density matching with the drug particle. Like mixture vapor pressures, the mixture densities can be estimated in the absence of data. Hankinson and Thomson have developed a mixture density correlation (53). B.

Phase Behavior

When new MDI propellants are evaluated, there can be incompatibilities with approved MDI excipients. For example, it has been shown that lecithin, sorbitan trioleate, and oleic acid are incompatible with new HFA propellants (54). The combination of these ingredients at room temperature leads to phase separation rather than a homogeneous formulation. This poses a major challenge to inhaler development (55).

Figure 4 Density for Propellant 134a as a function of temperature. Line A is the result from Bhirud’s method (51) using critical properties from Table 4. Line B is the calculation using critical properties estimated by Ambrose’s method (46). The squares are data from ICI (Runcorn, England).

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Phase separation is a function of composition and temperature. If a binary system is above a certain critical temperature, known as the consolute temperature or Tcs, both components are miscible in all ratios. Below this temperature, there will be a phase separation at certain compositions. Therefore, the formulator should establish Tcs during product development. It is advantageous to have Tcs less than the lowest temperature the formulation will experience. Solubility Parameter Theory

Solubility parameter theory is a method for estimating phase behavior in liquids. For binary mixtures, Tcs can be calculated by the following expression (56): Tcs ⫽ K(V1 + V2)/4R ⫽ KVave/2R

(14)

where R is the gas constant, V1 and V2 are the molar volumes for components 1 and 2, respectively, Vave is the average molar volume, and K is a measure of the cohesive energy density difference between the pure components, c11 and c22, and the mixture, c12 (56): K ⫽ c11 + c22 ⫺ 2c12

(15)

In the previous section, it was shown that cohesive energy densities for pure components can be estimated from vapor pressure data. The cohesive energy density for the mixture is usually estimated by the geometric mean of the pure components (57): c12 ⫽ (c11 c22)1/2 By substituting Eq. (16) into Eq. (15) and remembering that ␦ii ⫽ cii some rearrangement K can be expressed as (57) K ⫽ (␦11 ⫺ ␦22)2

(16) 1/2

, with (17)

which is simply the difference between the solubility parameters of the pure components squared. The solution properties of most nonpolar and nonionic compounds can be explained qualitatively and, in some cases, semiquantitatively by solubility parameter theory (57,58). As shown above, physical properties required for the calculation of solubility parameters can be easily estimated from vapor pressure, which in turn can be estimated using physical property correlations. With this information, solubility parameter theory can be used to predict the consolute temperatures in MDI systems. Most new propellant systems are mixtures of HFA propellants, which do not contain chlorine atoms, and hydrocarbon surfactants (54,55). Therefore, results on phase behavior studies with hydrocarbons and fluorocarbons are relevant. Solubility Parameter Theory and Hydrocarbon/Fluorocarbon Systems

For mixtures with perfluoroheptane, Table 5 shows calculated values of K from solubility parameter theory and experimental values, Kexp. For benzene, carbon tetrachloride, and stannic chloride, the results from solubility parameter theory

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Table 5 Comparison of K Calculated by Solubility Parameter Theory with Measured Values, Kexp, for Various Compounds with Perfluoroheptane Compound

K

C6H6 CC14 SnCl4 n-C6H14 3-(CH3)C5H11 2-2-(CH3)C4H11 2-3-(CH3)C4H11 (CH3)8(SiO)4

10.0 6.9 7.3 1.6 1.3 0.9 0.5 0.2

Kexp 9.4 7.9 8.3 6.6 6.4 5.9 6.2 5.1

Source: From Ref. 56.

are in agreement with experiments. However, for the aliphatic hydrocarbons and methyl siloxane, solubility parameter theory does not predict the phase behavior accurately. For example, the consolute temperature for perfluoroheptane and 2–3–(CH3)C4H11 was underestimated by an order of magnitude using solubility parameter theory. This means that an experimenter could have a two–phase system when the theory predicts a single phase. Solubility parameter theory cannot predict the phase behavior of all hydrocarbon and fluorocarbon systems. Improvements have been proposed, but they require additional data such as solubility parameters determined by iodine solubility experiments (58). By modifying the mixing rule to estimate c12, it has been shown that solubility parameters can be used to more accurately estimate the phase behavior of new propellant systems (59). This method does not require any additional data and will be described in the next section. Estimating Cohesive Energy Density for a Mixture

For some hydrocarbon/fluorocarbon systems, the geometric mean overestimates the interaction between the two components (56). For systems where the geometric mean overestimates this interaction, the harmonic mean can be a better predictor for mixture properties (60). If the harmonic mean of the pure component cohesive energy densities is used to calculate the cohesive energy density of the mixture, then

c12 =

2c c 2 = 11 12 1 / c11 + 1 / c22 c11 + c22

(18)

By substituting Eq. (18) into Eq. (15), K can be expressed in terms of the cohesive energy densities as K h = c11 + c22 −

4c11c22 c11 + c22

(19)

where the subscript h indicates the use of the harmonic mean for the mixing term. Since cii ⫽ ␦ii2, Eq. (19) becomes

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K h = δ112 + δ 22 2 −

4(δ11δ 22 )2 δ11 + δ 22

(20)

Table 6 compares the results for Kh and Kexp for mixtures with perfluoroheptane. Using the harmonic mean to estimate c12 improves the predictions using solubility parameters, but there are still large deviations from experimental results. Empirical factors are often used to modify simple intermolecular interaction models (61). Fitting the coefficient on the third term of Eq. (20) to experimental results, for mixtures of hydrocarbons and fluorocarbons (56,57), gave the following equation: K emp = δ112 + δ 22 2 −

3.8(δ11δ 22 )2 δ11 + δ 22

(21)

where Kemp denotes the use of an empirically derived coefficient, 3.8, based on a harmonic mean mixing rule for c12. In Table 6, one can see that Eq. (21) is in fairly good agreement with experiments. This improved agreement between theory and experiment does not require any additional data. It is important to note that 3.8 is close to the theoretical value of 4.0. Using the Harmonic Mean or Geometric Mean for c12

As shown in Table 5, Tcs for some mixtures of hydrocarbons with perfluoroheptane can be predicted using solubility parameters with a geometric mean mixing rule for c12. When should the harmonic mean rather than the geometric mean be used to calculate c12? A contribution to the geometric mean’s overprediction of the mixture’s cohesive energy density is asymmetry in the intermolecular pair potential (56). Kihara has shown that the ratio of the triple point temperature Tt to the critical temperature is a measure of this asymmetry (62). Tt/Tc for perfluoroheptane is 0.467 (44,63). For the compounds in Table 5, Tt/Tc (44,63) was calculated and compared to the deviation from solubility parameter theory in Figure 5. As Tt/Tc approached the value of perfluoroheptane, Eq. (17) was able to

Comparison of Experimental Values (Kexp), Those from Solubility Parameter Theory Using a Geometric Mean for c12 (K), Those Using Solubility Parameter Theory with a Harmonic Mean for c12 (Kh), and Those Using an Empirically Derived Expression Based on Solubility Parameter Theory with a Harmonic Mean for c12 (Kemp), for Various Compounds with Perfluoroheptane Table 6

Compound

Kexp

K

Kh

Kemp

n-C6H14 3-(CH3)C5H11 2-2-(CH3)C4H11 2-3-(CH3)C4H11 (CH3)8(SiO)4

6.6 6.4 5.9 6.2 5.1

1.6 1.3 0.9 0.5 0.2

3.3 2.6 2.0 1.1 0.3

7.7 6.7 5.2 6.2 4.2

K and Kexp were obtained from Refs. 56 and 57.

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Figure 5 Difference between measured values of K (Kexp) and values calculated with solubility parameter theory, using a geometric mean to calculate c12, as a function of Tt/Tc for the binary systems in Table 5. Tt/Tc for perfluoroheptane is 0.467 (44,63). As Tt/Tc approached the value of perfluoroheptane, the theory using a geometric mean to calculate c12 gives a more accurate estimate of K. For compounds with Tt/Tc much less than perfluoroheptane, the theory using a harmonic mean to calculate c12 gives a more accurate estimate of K.

predict the phase behavior. When Tt/Tc was much less than the value for perfluoroheptane, there were large deviations from predictions based on a geometric mean mixing rule for c12. Therefore, Tt/Tc may be a useful dimensionless parameter to estimate when the harmonic mean mixing rule should be used to estimate c12. In general, predicting phase behavior of aliphatic hydrocarbons with fluorocarbons will require a harmonic mean mixing rule for c12. Application to HFA System

While there are no published data on Tcs for mixtures of HFAs and MDI excipients, the theory presented above should provide improved estimations of such phenomena. MDI surfactants have long hydrocarbon groups and, when mixed with Propellant 134a at room temperature, there is a phase separation (54,55). Using hexadecane as a model long chain hydrocarbon, we can apply solubility parameter theory and compare it to experimental results. Hexadecane and Propellant 134a form a two–phase system when mixed at room temperatures. At 25°C, the molar volumes for hexadecane (49) and 134a (DuPont, Wilmington, Delaware, U.S.) are 294 and 84 cm3/mol, respectively. Using 134a’s ⌬Hvap, its solubility parameter at 25°C was calculated as 7.1 (cal/cm3)1/2. The solubility parameter of hexadecane at 25°C is 8.2 (cal/cm3)1/2 (49). From Eq. (14), K should be less than 6.3 cal/cm3 if Tcs is to be less than 25°C. From Eq. (17), based on a geometric mean estimate for c12, the difference between ␦11 and ␦22 must be less than 2.5 (cal/cm3)1/2 at 25°C for K to be less than 6.3 cal/cm3. Because the difference between the solubility parameters

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of Propellant 134a and hexadecane is 1.1 (cal/cm 3)1/2, we would expect a single–phase system at 25°C. However, this estimate does not agree with experiment. Eq. (17) underestimates Tcs for Propellant 134a and hexadecane, as it does for the systems in Table 6. With the modified theory of Eq. (21), it is not easy to analytically determine the maximum difference between ␦11 and ␦22 so that K is less than 6.3 cal/cm3. Therefore, ␦11 was set to 7.1 (cal/cm3)1/2, the value for Propellant 134a, and ␦22 was varied to find the range where K was less than 6.3 cal/cm3. Using this method, ␦22 must be less than 7.7 (cal/cm3)1/2 for a single–phase system at 25°C. Because the solubility parameter for hexadecane is 8.2 (cal/cm3)1/2, Eq. (21) predicts a two–phase system, which is in agreement with experiment. This example demonstrates that a modified solubility parameter theory, based on a harmonic mean mixing rule for c12, is useful for predicting phase behavior in HFA systems. V.

Experimental Design and Methods

Theories of interparticle forces and suspension stability have been reviewed with calculations of suspension stability based on model systems. Propellant physical properties, phase behavior, and methods for predicting them have also been discussed. These theories are a foundation for designing and performing experiments on MDI suspension formulations. In this section, methods for measuring suspension stability during formulation experiments will be discussed. Practical methods for designing experiments with many variables will be reviewed. A.

Measuring Suspension Stability

Sedimentation and Sedimentation Height

Since MDI formulations are easily prepared in safety–coated aerosol vials, visual observation of suspension stability is possible (2,22). Qualitative grades can be assigned to different formulations by observing the suspension’s texture, fine or coarse, at various times after shaking. The time required for all of the material to accumulate at either the bottom of the vial or the vapor–liquid interface can be recorded to give a more quantitative measure of suspension stability. The suspension stability can be further quantitated by observing the structure of the settled or creamed layer. If flocculation has occurred, the separated particle layer will be thicker than in a more stable suspension (3). Therefore, by measuring the particle layer volume relative to the liquid volume, one obtains a dimensionless parameter of suspension stability. The higher this ratio, known as the sedimentation height, the more unstable the suspension (64). This method has been used to quantitate the suspension stability of MDI formulations (65,66). Laser Back Scattering

The particle size distribution is an important factor for suspension stability and deposition in the respiratory tract. Therefore, a method that can measure

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particle size and settling, or creaming, in MDI formulations would be useful. It has been demonstrated that both of these objectives can be obtained using a laser back–scattering instrument to study MDI formulations in glass aerosol vials (67). The Labtec 100 (Lasentec Corp., Bellevue, Washington, U.S.) can measure particle sizes from 0.7 to 1000 ␮m in suspensions as concentrated as 30% w/v. Furthermore, the instrument performs particle counting, the total number of counts being proportional to the number of particles in a region. By focusing the instrument’s probe at the vapor–liquid interface or bottom of the vial, the experimenter can measure the rate of settling or creaming, respectively. If a magnetic bar is placed inside the vial, the suspension can be homogenized during the experiment by stirring. Figure 6 shows the results from settling experiments where various size fractions of ground garnet were suspended in Propellants 11 and 12 with AOT as a surfactant (67). Garnet settled in the propellant blend, so particle counts were measured near the vapor–liquid interface. The initial particle count, or particle concentration, was obtained while the suspension was stirred. When the stirrer was turned off, the particles began to settle and particle counts at the vapor–liquid interface decreased. The number of counts decreased rapidly for 22 and 33 ␮m particles but much more slowly for 5.2 and 11 ␮m particles. The higher sedimentation rates measured for the larger particles are in agreement with Eq. (13). These suspension stability results were obtained in approximately one minute and included a particle size distribution measurement. The rate of settling or creaming from back–scattering experiments can be used as a response in formulation evaluation (67).

Laser back–scattering measurements of the particle count at the vapor/liquid interface, Counts(t), relative to the initial particle counts, Counts(0) as a function of time for ground garnet particles in Propellants 11 and 12 with AOT as a surfactant. After the stirrer is turned off, the larger particles settle more rapidly, which is in agreement with Eq. (13). This is a quantitative measure of suspension stability that can be used in formulation evaluation. Figure 6

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The delivered dose variability is a function of the formulation, the valve, and the actuator (4). To isolate formulation instability from valve and actuator performance, a suspension suitability test has been developed (4,68). In this technique, the retaining cup of the metering valve is removed so that it does not impede sampling of the bulk suspension. The inhaler is shaken, secured to a vibration–free holder, then actuated at various times after shaking. The dose collected from the metering valve is assayed for drug content. Doses obtained with no waiting time are compared to doses with 15– to 45–second waiting times. A stable suspension will show little variation in collected dose as a function of waiting time (68). The effect of delay time on delivered dose is another useful response for suspension formulation evaluations. B.

Experimental Design

Table 7 shows a list of variables for MDI suspension formulations and some of the physical parameters they influence. While the list covers most of the theoretical parameters discussed in previous sections, it does not cover canister and metering valve variables. However, this list is sufficient for starting a discussion of experimental design for MDI formulations. While there are only six variables, it is important to note that these variables have potentially significant interactions, as demonstrated in the previous sections. For example, the interaction between surfactant level and water content affects surface charge generation, suspension stability, and solvent power for the suspending medium. The formulator must identify significant variables and interactions early in the development process. The advantage of statistically designed experiments over the approach of varying “one factor at a time” has been well established (69). These experimental design techniques have been used to formulate pharmaceutical suspensions for oral administraSome Variables for a Metered Dose Inhaler Suspension Formulation and the Physicochemical Properties They Influence Table 7

Formulation variable

Some important physical properties associated with variable

Propellant or cosolvent ratio

␳, ␦, ␬, ␺o, ␧r, Pvap, A, and drug solubility

Surfactant structure

L, ␦, ␺o, adsorption on particle surface, and interaction with water at particle surface and in bulk phase

Surfactant concentration

L, ␺o, drug solubility, adsorption on particle surface, and interaction with water at particle surface and in bulk phase

Particle volume fraction

%VE

Water content

A, ␺o, ␬, surfactant adsorption, particle growth rate, and interaction with surfactant at particle surface and in bulk phase

Temperature

␳, ␦, Pvap, drug solubility, and particle growth rate

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tion (70,71). Applying these principles to MDI formulations accelerates the development process and ensures that statistically significant information is obtained with minimal experimentation (42). The most common form of experimental design is a two–level factorial. With this method, each factor is varied between two levels, one high and one low, in an orthogonal design. For three factors, x, y, and z, this requires 23 (or eight) experiments. The experiment space is a cube, as shown in Figure 7. The response for this three–factor experiment, Y, is modeled using the following equation (69): (22) Y ⫽ ␤0 + ␤1x + ␤2y + ␤3z + ␤4xy + ␤5xz + ␤6yz + ␤7xyz where ␤0 is the intercept, ␤1–3 are the single–factor or main effects, ␤4–6 are the two–factor interactions, and ␤7 is the three–factor interaction. A study designed in this way has many advantages. It is orthogonal, so all the main effects and interactions can be estimated independently of one another. It can be randomized so that true effects can be distinguished from noise. Furthermore, if a center point is added, one can estimate curvature. Replicating center points can be used to estimate experimental errors. Because it is not easy to quantify surfactant structure as a high or low level, only five of the six variables in Table 7 could be studied in a factorial plan. One surfactant could be selected and the remaining variables studied using a factorial design. To complete a factorial design with five variables, it would take 25 or 32 experiments, not including center points. This illustrates a disad-

Experiment plan for three factors in a two–level factorial design. 23 (or eight) experiments are required to complete this study that allows the experimenter to estimate an intercept, three main effects, three two–factor interactions, and one three–factor interaction. A center point can be added to estimate curvature. Figure 7

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vantage of full factorial studies: with a modest number of variables, the number of experiments required to perform the study can be quite large. To evaluate the effect of surfactant structure, the 25 experiment plan could be repeated using a different surfactant, which would then double or triple the number of experiments. However, there are practical methods for reducing the number of experiments for a full factorial study. A 25 study requires 32 experiments and allows estimation of the following parameters: an intercept, five main effects, 10 two–factor interactions, 10 three–factor interactions, five four–factor interactions, and one five–factor interaction. In most cases, the higher–order interactions are not significant. During screening studies, the main effects and the two–factor interactions are considered to be the most significant. Performing the half fraction of a 25 experiment plan, or 25⫺1 factorial study, reduces the number of experiments from 32 to 16. This is known as a fractional factorial experiment (69). These 16 experiments are used to estimate the intercept, five main effects, and 10 two–factor interactions. However, with this reduced experiment plan, the two–factor interaction estimates are confounded with three–factor interactions. This is usually not a significant problem if one assumes that three–factor interactions and higher are insignificant. A 25⫺2, or quarter–fraction factorial, further reduces the number of experiments to eight. With a 25⫺2 design, main effects are confounded with two–factor interactions. Because there are many significant interactions between the variables in Table 7, such confounding patterns should be avoided when evaluating MDI formulations. Furthermore, center points should always be added to the design when screening MDI formulations because several main effects and interactions are nonlinear (e.g., water). For most MDI formulations, the components can be varied independently of one another. For example, increasing the surfactant concentration from 0.01% to 0.02% does not significantly affect the level of propellant in the system because the surfactant is present as such a small fraction of the formulation. However, if the level of one component must be changed to accommodate a change in another component, then a mixture design plan must be used (72). Additional examples, where experiment plans were modified to fit practical constraints, are outlined by Snee (73). Computer software to aid experiment design and analysis has been recently reviewed by Studt (74).

VI.

Summary and Conclusions

Calculations using DLVO theory demonstrated that some suspensions of micronized drug particles could be stabilized in pharmaceutical propellants by their surface potentials. However, the volume fraction of particles in practical MDI suspension formulations can cause a significant reduction in stabilization from surface charge. Steric repulsion from adsorbed surfactant layers provides an additional mechanism for stabilization of MDI suspensions. Trace water levels have a significant, and usually nonlinear, effect on surface potential and suspension stability in nonaqueous media.

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Propellant physical properties can be estimated with a limited amount of data using established correlations based on corresponding–states theory. This information can be used to develop formulations and manufacturing processes, design container/closure systems and manufacturing equipment, and estimate other physical properties such as solubility parameters. Solubility parameters can be used to estimate phase behavior in MDI systems. When evaluating, combinations of long–chain hydrocarbons and fluorocarbons, the harmonic mean of the pure component’s cohesive energy density should be used to estimate the mixture’s cohesive energy density [see Eq. (21)]. Sedimentation heights, laser back–scattering, and suspension suitability tests are useful quantitative responses to measure MDI suspension stability. During screening experiments, it is important to determine significant main effects and two–factor interactions. The most efficient method for obtaining results in MDI formulations is through statistically based experiment designs. Factorials allow the experimenter to estimate many main effects and interactions. The number of experiments can be significantly reduced using fractional factorial designs. In fractional factorial designs, it is important to avoid confounding between main effects and two–factor interactions because many two–factor interactions in MDI formulations are significant. Center points should be added to all designs because responses in nonaqueous media are often nonlinear. References 1. 2. 3. 4.

5. 6. 7. 8. 9. 10. 11. 12. 13. 14. 15.

Gonda I. Development of a systemic theory of suspension inhalation aerosols. Int J Pharm 1985; 27:99–116. Miller NC. In: Byron PR, ed. Respiratory Drug Delivery. Boca Raton: CRC, 1990:249–257. Hiestand EN. Theory of coarse suspension formulation. J Pharm Sci 1964; 53(1):1–18. Schultz RK, Miller NC, Christensen JD. Method for evaluation of suspension characteristics relevant to metered dose inhaler formulations. Presented at American Association of Pharmaceutical Scientists Meeting, Las Vegas, NV, November 1990. Van der Hoeven PC, Lyklema J. Electrostatic stabilization in non–aqueous media. Adv Colloid Interface Sci 1992; 42:205–277. Russel WB, Saville DA, Schowalter WR. Colloidal Dispersions. Cambridge: Cambridge University Press, 1991. Hiemenz PC. Principles of Colloid and Surface Chemistry. New York: Marcel Dekker, 1986:611–731. Rosen MJ. Surfactants and Interfacial Phenomena. New York: John Wiley & Sons, 1978:251–270. Everett DH. Basic Principles of Colloid Science. London: Royal Society of Chemistry, 1988:35. Prieve DC, Russel WB. Analysis of the dispersion force between a layered sphere and a cylinder. Report CRDEC–CR–86038, US Army Research Office, Research Triangle Park, NC, 1986. Israelachvili J. Intermolecular and Surface Forces. London: Academic Press, 1992:176–207. Vold MJ. The effect of adsorption on the van der Waals interaction of spherical colloidal particles. J Colloid Sci 1961; 16:1–12. Novotny V. Electrical conduction in surfactant–water–nonaqueous liquid systems. J Electrochem Soc: Electrochem Sci Technol 1986; 133(8):1629–1636. Pugh RJ, Matsunaga T, Fowkes FM. The dispersibility and stability of carbon black in media of low dielectric constant. Colloids Surf 1983; 7:183–207. Wyatt DA, Vincent B. Electrical effects in non–aqueous systems. J Biopharm Sci 1992; 3(1/2):27–31.

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Hiemenz PC. Principles of Colloid and Surface Chemistry. New York: Marcel Dekker, 1986:737–786. Morrison ID. Criterion for electrostatic stabilization of dispersions at low ionic strength. Langmuir 1991; 7:1920–1922. Weast RC, ed. Handbook of Chemistry and Physics. Boca Raton: CRC Press, 1976:E34–E35. Albers W, Overbeek JTG. Stability of emulsions of water in oil. J Colloid Sci 1959; 14:510–518. Green JH, Parfitt GD. Charge stabilization of concentrated dispersions in apolar medial. In: Moudgil BM, Somasundaran P, eds. Flocculation, Sedimentation & Consolidation. New York: American Institute of Chemical Engineers, 1986:205–215. Chen CS, Levine S. The double–layer interaction of two charged colloidal spherical particles of a concentrated medium of low dielectric constant. J Colloid Interface Sci 1973; 43:599–615. Ashurst IC. Physicochemical Characteristics of Chlorofluorocarbon Based Inhalation Aerosols. Doctoral dissertation, University of Aston, Birmingham, 1985. Farr SJ, McKenzie L, Clarke JG. Drug–surfactant interactions in apolar systems. Proceedings of Respiratory Drug Delivery IV. Buffalo Grove, IL: Interpharm Press, 1994:221. Mackor EL. A theoretical approach of the colloid–chemical stability of dispersions in hydrocarbons. J Colloid Sci 1951; 6:492–495. Fischer EW. Electronenmikroskopische unterscuhungen zur stabilitat von suspensionen in makromolekularen losungen. Kolloid Z 1958; 160:120–141. Bagchi P, Vold R. Differences between fact and theory in the stability of carbon suspensions in heptane. J Colloid Interface Sci 1970; 33(3):405–419. Kerkar AV, Feke DL. Induced flocculation in sterically stabilized ceramic powder dispersions. Colloids Surf 1991; 53:363–381. Osipow LI. Surface Chemistry. Huntington, NY: R.E. Krieger, 1977:163–225. Visser J. Adhesion of colloidal particles. Surf Colloid Sci 1976; 8:3–84. Israelachvili J. Intermolecular and Surface Forces. London: Academic Press, 1992:330–333. Kandori K, Kazama A, Konno K, Kitahara A. Dispersibility of magnetic and nonmagnetic iron (III) oxides in cyclohexane and effect of water on the stability. Bull Chem Soc Jpn 1984; 57:1777–1783. Somasundaran P, Krishnakumar S, Yu X. Stabilization mechanisms of oxide dispersions in non–aqueous media. Presented at First International Particle Technology Forum, American Institute of Chemical Engineers National Meeting, Denver, August 1994. Malbrel CA, Somasundaran P. Water–induced dispersion/flocculation of colloidal suspensions in nonpolar media. J Colloid Interface Sci 1989; 133(2):404–408. McGown DNL, Parfitt GD, Willis E. Stability of non–aqueous dispersions I. J Colloid Interface Sci 1965; 20:650–664. Everett DH. Basic Principles of Colloid Science. London: Royal Society of Chemistry, 1988:149. Lifshiz IM, Slesov VV. The kinetics of precipitation from supersaturated solid solutions. J Phys Chem Solids 1961; 19:35–50. Voorhees PW. Theory of Ostwald ripening. J Stat Phys 1985; 38:231–252. Kabalnov AS, Makarov KN, Pertzov AV, Shchukin ED. Ostwald ripening in emulsions. J Colloid Interface Sci 1990; 138(1):98–104. Yalkowsky SH, ed. Techniques of Solubilization of Drugs. New York: Marcel Dekker, 1981. Dalby RN, Phillips EM, Byron PR. Determination of drug solubility in aerosol propellants. Pharm Res 1991; 8:1206–1209. Phillips EM, Byron PR, Dalby RN. Axial ratio measurements for early detection of crystal growth in suspension–type metered dose inhalers. Pharm Res 1993; 10:454–456. Johnson KA. Development and optimization of inhalation products. Proc Pharm Tech Conference Eugene, OR: Aster, 1992:66–78. Newman SP, Moren F, Pavia D, Corrado O, Clarke SW. Effects of changes in metered volume and propellant vapor pressure on the deposition of pressurized inhalation aerosols. Int J Pharm 1982; 11:337–344. Reid RC, Prausnitz JM, Poling BE. The Properties of Gases and Liquids. New York: McGraw–Hill, 1987:74–238. Gomez–Nieto M, Thodos G. A new vapor pressure equation and its application to normal alkanes. Ind Eng Chem Fund 1977; 16:254–259.

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13 Molecular Scale Behavior in Alternative Propellant-Based Inhaler Formulations

LIBO WU, ROBSON P. S. PEGUIN, PARTHIBAN SELVAM, UDAYAN CHOKSHI, and SANDRO R. P. DA ROCHA Chemical Engineering and Materials Sciences, Wayne State University, Detroit, Michigan, U.S.A.

I.

Introduction

Traditional pressurized metered dose inhaler (pMDI) formulations contain a drug either in solution or in suspension in a compressed propellant (1). Excipients commonly found in pMDIs include surfactants, which are used for valve lubrication and/or as stabilizers, and cosolvents that serve as aids in the solubilization of amphiphiles (2,3). Several alternative propellant-based inhaler formulations have been recently proposed (4). To a large extent, these novel formulations came as a response to difficulties encountered in transitioning from chlorofluorocarbon (CFC) to hydrofluoroalkane (HFA) propellants, as dictated by the Montreal Protocol. These alternatives to solution and (micronized) dispersion-based pMDIs also provide unique opportunities for the local and systemic delivery of biomolecules to and through the lungs (5). Such advances may give propellant-based inhaler formulations the necessary edge to compete with recent developments in dry powder inhaler (DPI) technology (6). The focus of this chapter is on alternative propellant-based inhaler formulations. Rather than providing an exhaustive list and a lengthy discussion of all alternative pMDIs recently reported in the literature, we concentrate on a few selected systems that help illustrate how molecular-level information can be used to guide the design of such formulations. 373

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Alternative pMDIs

We classify the alternative propellant-based inhaler formulations according to their particle size (micro- vs. nanosuspensions) and the type of dispersed phase (aqueous vs. nonaqueous). Emphasis is given to aqueous-based formulations in the form of reverse microemulsions in HFAs, which have received considerable attention from both academia (7) and industry (8). A.

Microsuspensions

Traditional dispersion-based formulations in HFAs are inherently unstable due to the cohesive forces between particles and also due to the gravitational fields. Surfactants are generally required in traditional pMDIs in order to overcome particle–particle interactions (4,9). However, the amphiphiles approved by the FDA for use in pMDIs have extremely low solubility in HFAs (10–12). The low solubility arises due to a mismatch between the hydrogenated nonpolar surfactant tails and the polar nature of the HFA propellants (3,13). Cosolvents are, therefore, generally required in order to enhance surfactant solubility. However, cosolvents may decrease the overall chemical and physical stability of the formulation (14). In order to address these limitations, several studies have focused on the design of amphiphiles for conventional HFA-based pMDIs (15–17). More recently, alternatives to traditional dispersion formulations have also been proposed (4). Nonaqueous Systems Density Matching

Polymer-Coated Spheres: The stability of insoluble drugs in HFAs can be significantly improved by incorporating the drug into a carrier, with a resulting particle density closely matching that of the propellant medium. Chitosan, a biocompatible and biodegradable polymer (18), is a potential carrier for therapeutic proteins, peptides, and DNA (19). Chitosan microspheres produced by spray-drying can have their density tuned by varying the degree of cross-linking and the type of cross-linker. The density of the non–cross-linked and glutaraldehyde cross-linked chitosan microspheres was observed to be 1.48 g.ml–1 and 1.42 g.ml–1, respectively. It closely matches the density of 1,1,1,2-tetrafluoroethane (HFA134) of 1.21 g.ml–1 at 25°C, producing a good suspension in the propellant. It was also observed that the chitosan microspheres could be easily resuspended in the propellant after a long period of sedimentation (19). These observations show that not only could the effect of gravitational fields be reduced by density matching, but particle–particle interaction could also be affected (reduced) by the modification of the drug particle surface. Porous Particles. A pMDI formulation containing micron-sized porous particles with density close to that of the propellant has been recently reported (20,21). Cromolyn sodium, salbutamol sulfate, and formoterol fumarate hollow porous microspheres were formed by spray-drying (22). A stable suspension, termed homodispersion, is formed when hollow porous drug particles are

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dispersed in HFA. The propellant-filled particles exhibit little difference in density compared to the neat propellant, thereby decreasing the tendency for the particles to cream and flocculate. Interparticle forces are also expected to be affected by changes in the surface morphology of the spheres (23). Particle Surface Modification

An alternative approach to surfactant-stabilized dispersion formulations is the modification of the particle surface before they are suspended in the propellant medium. Different from the density matching approach, the particle surface modification technique directly aims at reducing particle adhesion. Salbutamol sulfate microparticles coated with alkylpolyglycoside were made by spray-drying (24). The spray-drying technique was also used to coat budesonide microcrystals with lipids (25). Particle stability can be enhanced by increasing the interaction between the propellant and the particle surface and also by reducing particle–particle interaction. The effectiveness of the approaches discussed above on reducing the effects of gravitational fields can be easily measured by sedimentation rate experiments. Colloidal Probe Microscopy (CPM) (26,27) and Chemical Force Microscopy (CFM) (28,29) are variations of the Atomic Force Microscopy (AFM). These techniques can be used to quantitatively relate particle surface chemistry (polymer coating) and morphology (hollow porous particles) to interparticle interaction and, consequently, to the colloidal stability in the propellant medium. Recent CPM and CFM studies related to pMDIs are discussed below. Aqueous Systems (Emulsions)

Reverse water-in-perfluorooctyl bromide (PFOB) (W/PFOB) emulsions stabilized with a fluorinated surfactant were shown to be easily dispersible in both HFA134a and 1,1,1,2,3,3,3-heptafluoropropane (HFA227) (30). Such formulations can be potentially used for the local or systemic delivery of biomolecules to and through the lungs. While the perfluoroalkylated amphiphile is interfacially active at the PFOB–water (PFOB|W) interface, it cannot directly stabilize water-in-HFA (W/HFA) emulsions, indicating weak solvation of the surfactant tail by HFA and/or inappropriate surfactant balance. The ability to quantify the HFA-philicity of candidate surfactant tail groups can open up opportunities for the development of such alternative pMDI formulations. Molecular level information on solvophilicity can be obtained by a combination of CFM and nonbonded interaction energies from ab initio calculations (16). B.

Nanosuspensions

There are many advantages in delivering drugs in nanosuspension form (31). Higher dissolution rates can be achieved due to the larger surface area of small particles. This might be especially relevant when solubility-limited bioavailability needs to be overcome. Small particles also tend to have greater stability

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due to reduced gravitational effects. However, techniques for preparing nanoparticles are generally more complex and/or require large inputs of mechanical energy compared to those used in the production of micron-sized particles (4,9). Nonaqueous Systems

There have been only a few reports of nanosuspensions in HFAs (4,32). Salbutamol sulfate nanoparticles have been prepared by a combination of templated reverse-phase microemulsion and freeze drying (32). When AOT was used as surfactant to form the microemulsion, the resulting nanoparticles were not dispersible in HFA, not even in the presence of up to 10% w/w cosolvent. Nanoparticles produced from lecithin-based microemulsions, however, could be dispersed in HFA, but only in the presence of a cosolvent (n-hexane). Understanding surfactant tail solvation is a topic of great relevance to such formulations. Salbutamol sulfate nanoparticles with a biodegradable copolymer shell have been recently prepared using low energy method (33). While one of the blocks of the copolymer interacts with the polar drug, the moiety that forms the outer shell of the particle is highly HFA-philic (16). These core/shell nanoparticles have very long-term stability in a mimic HFA solvent, 2H,3H-perfluoropentane (HPFP), without the use of cosolvents (34). The enhanced stability was attributed to the strong interaction between the polymer coating the particle and the HFA solvent (34). This method represents a significant advancement compared to previously reported approaches to the production of pharmaceutical nanoparticles (35,36). It is a single-step process that requires very low energy input. Using this surface modification technique, particle cohesion can be reduced to zero, as probed by CPM in HPFP solvent medium (34). Aqueous Systems (Microemulsions)

pMDIs as Carriers for Biomolecules. Currently, there are no pMDI-based formulations commercially available for either local or systemic delivery of biomolecules. This has to do with the fact that polar compounds, including water, have low solubility in the propellants utilized in pMDIs (37,38). In order to overcome this inherent limitation, solubilization of hydrophilic molecules in reverse W/HFA microemulsions has been suggested (12). It is known that small peptides and proteins of up to approximately 30 kD, and up to a few milligrams dosages, can be administered via the lungs (39). Encapsulation and delivery of such molecules in reverse microemulsions is, therefore, potentially attractive. However, limited success has been achieved in forming and stabilizing such reverse aqueous aggregates in the propellants used in pMDIs (7,40–44). This can be attributed, to a large extent, to the lack of fundamental knowledge on the interfacial properties of the HFA|W interface (10,11). Lecithin-based reverse microemulsions were designed for the delivery of small molecules using CFCs (40). Lecithins have been also shown to stabilize biomolecule-containing reverse microemulsions in dimethylether (DME) and DME/propane propellant systems (41,42). It is

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important to notice that long-term stability and large respirable fractions were reported in those cases. More importantly, the aerosols had properties within the range of commercially available formulations (41). However, such propellants are not acceptable in pMDI formulations (45). Reverse Aqueous Microemulsions in HFA-Based pMDIs. More recently, there have been several studies on the formation of W/HFA microemulsions (7,43,44,46). However, most systems investigated have limited applicability because the fluorinated surfactants used to form the microemulsions are not accepted by the FDA (47). In the case of nonionic and ionic single-tail fluorinated amphiphiles, very large amounts of surfactant (30% w/w) are required to stabilize the microemulsions. Large concentrations of alcohol (cosurfactant/cosolvent) were also necessary (7,43,44). More important, however, is the fact that small water-to-surfactant molar ratios (Wo) were observed. It is likely that most of the water is bound to the surfactant head groups, which would significantly reduce the solubilization capacity of the reverse aggregates (42). The formation of W/HFA microemulsions with a double-tailed fluorinated sulfosuccinate surfactant, without the addition of cosurfactants, has also been reported (7). In spite of the larger Wo attained with the double-tailed surfactants, at ambient temperature, the microemulsion is reported to be stable only above the propellant saturation pressure. Here again, fluorinated amphiphiles were used to stabilize the aggregates. More recently, W/HFA227 microemulsions have been reported in the presence of an ethoxylated triblock copolymer and ethanol (46). Surfactant concentrations were once again high, and cosolvents were required to form the microemulsions. Microemulsion Formation in Compressible Solvents. Microemulsion is a liquid solution of water, oil, and surfactants, with or without additives, which is optically isotropic and thermodynamically stable (48). The free energy for microemulsion formation (Gic) is given by Gic = ATS

(1)

where is the interfacial tension, A is the change in interfacial area upon microemulsion formation (very large), and TS is the dominant favorable entropic term that arises from the dispersion of one phase on the other, upon formation of a very large number of nanometer-size droplets, the microemulsions. Clearly, for microemulsion formation (spontaneous process), a significant reduction in must occur upon addition of interfacially active amphiphiles. For the compressed CO2–water (CO2|W) interface, a reduction in interfacial tension from ~25 to ~1–3 mN/m has been shown to be a necessary but not sufficient condition for the formation of reverse aqueous aggregates (49,50). Interfacial tension versus a so-called formulation variable, as for example the hydrophilic-to-lipophilic balance (HLB), can provide insight into the preferred natural curvature of the microemulsion (51,52). Knowledge of the of the bare and the surfactant-modified HFA|W interface is, therefore, of great relevance. Microemulsion Stability in Compressible Solvents. Instability in water-inoil (W/O) microemulsions, as indicated by phase separation, can be induced by (i) curvature or (ii) attractive droplet interactions (53). Curvature effects

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arise due to the difference in surfactant head group–head group, and tail–tail interactions, which are mediated by the environment in which they are solvated (54,55). From this difference in interaction, an interfacial bending stress arises at the flat oil–water film, causing it to curve spontaneously to an optimal curvature 1/R0, provided very low can be attained. Phase separation will occur when water droplets assume a radius larger than R0; i.e., water molecules will be “squeezed out.” Equilibrium between W/O microemulsions and excess water will then exist. It is clear that the chemistry of both surfactant head and tail groups has a significant impact on the natural curvature. In this scenario, computer simulations are specially suited for probing both tail- and headgroup solvation at the microscopic level (53–55). On the other hand, phase separation due to attractive droplet interactions leads to the coexistence of two microemulsion phases of the same curvature. Such attractive interactions arise predominantly due to surfactant tail–tail interactions during microemulsion contact and, to a less extent, due to the attraction of the water droplet cores (56). The same behavior has been documented for compressible solvents (57,58). Obviously, the relative strength of tail–tail, oil–oil, and oil–tail interactions will dictate the magnitude of attraction upon collision; i.e., whether the microemulsion system is stable or not; in the case of compressible solvents, it will also dictate the minimum pressure required to stabilize such dispersions. In the case of pMDIs, however, the pressure is set to the saturation pressure of the propellant/propellant mixture. In that case, solvation of the tail groups has to be addressed either by changing the chemistry of the tail and/or by modifying the solvent environment (additional excipient or propellant mixture). The presence of additives such as alcohols (cosurfactants) at the interfacial surfactant film can reduce tail–tail, and thus micelle–micelle interactions (58). A similar effect is expected with bulkier, more branched surfactant tails (53,58,59), where tail penetration between interacting droplets is reduced due to greater steric hindrance at the interfacial region, as schematically illustrated in Figure 1.

Figure 1 Schematic diagram of the effect of tail–tail interaction during microemulsion contact with (A) single-tail and (B) double-tail surfactant. The more crowded interface (B) prevents more intimate contact between tail-group segments of surfactants adsorbed at the interface of approaching microemulsion droplets.

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Microemulsion Curvature in Compressible Solvents. Interfacial tension measurements are also relevant in understanding microemulsion curvature; i.e., whether reverse or regular aggregates are favored during the microemulsion formation process. For surfactant–water–oil systems, it has been demonstrated that there is a phase change in the microemulsion at the phase inversion point, which corresponds to a minimum in interfacial tension and zero net curvature. At this point, the surfactant has equal affinity for both the organic and aqueous phases, i.e., maximum interfacial activity. This is called the balanced state, where bicontinuous microemulsions are preferred (60–62). Moving the balance away from the phase inversion point causes the surfactant to become preferentially soluble in one of the phases, resulting in a decrease in interfacial activity and, thus, an increase in interfacial tension. This can be accomplished by changing a “formulation variable,” as schematically illustrated in Figure 2. For nonionic surfactant systems, the phase inversion point is typically tuned by varying the temperature and the chain length of the surfactant head and tail groups (i.e., HLB) (62), while salinity and pH have less of an effect (63,64). Similar behavior has been also observed in microemulsions where one of the phases is a compressible fluid such as propane or CO2 (65). For this class of solvents, pressure (density) is an additional formulation variable. Changes in pressure or temperature can have a large influence on the density and, thus, on the solvent strength of a compressible fluid (66,67). One can,

Figure 2 Schematic diagram of the interrelationship among interfacial tension (), formulation variable (x-axis), and microemulsion curvature. The formulation variable is represented here by the hydrophilic-to-HFA-philic balance (HFB).

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therefore, manipulate the surfactant phase behavior by “tuning” the interactions between the surfactant tail and the solvent in the same way that temperature can be used to tune water–head group interactions (49). Systems with compressible solvents have been shown to undergo a phase inversion density, similar to the phase inversion temperature in conventional oil–water systems (65,68,69). Therefore, surfactant interfacial activity and microemulsion curvature can also be controlled with pressure. For compressible fluids, typical formulation variables include temperature, pressure, hydrophilic to HFA-philic balance (HLB), salinity, surfactant tail length, and the presence of cosurfactant/cosolvent (49,50,69–72). It seems that the formulation variable with the most significant impact on the interfacial tension, at least for nonionic amphiphiles, HLB in the molecule (49–51,72), here defined as the hydrophilic–HFA-philic balance (HFB). As the formulation variable is changed, the microemulsion curvature can be inverted form HFA-in-water (HFA/W) to W/HFA. The discussion above clearly illustrates the link between tail solvation and tail–tail interaction, with microemulsion formation and stability. A molecular level understanding of the relationship between tail chemistry and its interaction with the solvent environment can serve as a guide to the design of not only alternative pMDIs such as W/HFA microemulsions but also traditional dispersion-based formulations. As will be discussed below, ab initio calculations and AFM experiments can be used to obtain such quantitative information.

III.

Microscopic Tools for the Development of Alternative pMDIs

In this section, we make a concerted effort to illustrate how one can take advantage of molecular level information to guide the design of alternative propellant-based formulations. While we focus on the link between microscopic information from experiments and computer simulations, and the design of amphiphiles for the formation and stabilization of aqueous dispersions in HFAs, it will be apparent that a similar approach would be equally suited for the development of nonaqueous alternative and traditional pMDI formulations. A.

Ab Initio Calculations

Nonbonded interactions play a central role in defining the molecular properties of fluids. When appropriate levels of theory and basis sets are chosen, ab initio calculations can be used to obtain quantitative information on the nonbonded properties of intermolecular complexes that range from dispersion- to electrostatic-dominated systems (73–75). The evaluation of solvent–solute and solute–solute interactions through pair binding energies (Eb) is of special relevance to the development of HFA-based formulations. It provides a simple and accurate way to quantitatively relate the chemistry of a moiety to its HFA-philicity, which in turn dictates its solubility and ability to stabilize aggregates dispersed in the propellant (16). Such knowledge has

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implications to all pMDIs, from traditional solution and dispersion formulations to alternative systems where highly HFA-philic excipients are required. Treating large systems quantum mechanically is prohibitively expensive in terms of computational time. In the case of the design of surfactant tail groups, the problem can be overcome by selecting a representative fragment of the candidate moiety (76,77). The screening process starts with the calculation of, Ebst the binding energy between the solvent (propellant) and a fragment of the candidate HFA-phile. Ebst is defined as Ebst = Est Es Et

(2)

where Est is the total energy of the solvent–surfactant tail fragment complex, and Es and Et are the energies of the isolated propellant and tail fragment, respectively. Using this definition, more energetically favorable solvent–solute complexes are indicated by more negative Ebst’s. Binding energies reported as the average of the raw and counterpoise corrected Ebst’s represent an excellent approximation to the complete basis set limit, minimizing more computationally expensive calculations (76–78). Another advantage of using the average energy is that it tends to minimize the overestimation of the energy obtained by the second-order Mller-Plesset (MP2) theory (74). Geometry optimizations and single point energy calculations should be preferentially determined at large basis sets and high levels of theory in order to accurately describe the nonbonded interactions of solvent–solute complexes. This is specially true for systems with weaker interactions [dispersion or van der Waals (VDW) forces] (75,76). In the case of HFA propellants, the presence of fluorine atoms increases the importance of the electrostatic interactions, and Ebst’s can be adequately described by the MP2 theory (79,80). The use of pair interaction energies to address HFA-philicity has been demonstrated through the calculation of the interaction between HFA134a and (i) a methyl-based fragment (CH2) and (ii) a more polar, ether-containing moiety, namely a fragment of poly(propylene oxide) (PPO) (16). CH2 was selected as the baseline fragment because it represents the tails of FDAapproved surfactants, which have limited solubility in HFA134a and have been shown not capable of forming reverse microemulsions in the semifluorinated propellants (10). The PPO tail was chosen as a potential candidate moiety based on its enhanced solubility in HFAs (81), which is likely due to the interaction between the dipole of HFA134a and the ether group. Single point binding energy calculations at the MP2/aug-cc-pVDZ, with geometry optimized at the MP2/6-31g+(d,p) level of theory, revealed that the PPO fragment interacts very strongly with HFA134a (Ebst –6.36 Kcal/mol, Fig. 3A). On the other hand, the interaction energy of the propellant and the methyl-based fragment is only –2.72 Kcal/mol (82). The presence of ether oxygen (increase in polarity) appears as an important factor for the enhancement of the tail fragment–HFA interaction. The observed interaction energy of the PPO–HFA134a complex is of much larger magnitude than those observed between CO2 and small hydrocarbon and fluorocarbon fragments (83). On the

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Figure 3 Optimized structures (MP2/6-31g+(d,p)) of the (A) PPO–HFA134a and (B) PPO–PPO complexes. The interaction points between the pairs are indicated by dashed lines. Distances are in Å.

other hand, the propellant cannot solvate CH2 well, due to the presence of weak F . . . H interactions only. This result is in agreement with the experimentally observed inability of the CH2-based surfactants in stabilizing drug dispersions in HFA (14). While Ebst can be correlated to tail solvation, a better indicator of both solubility and solvation can be obtained by understanding not only solvent–solute but also solute–solute interactions. For that purpose, the so-called enhancement factor (E) has been defined as (16) E=

E bst E btt

(3)

where Ebtt = Ett 2Et tt

(4)

and Eb is the binding energy between two tail fragments. Large (positive) Es indicate strong interaction between solvent and tail and weak tail–tail interaction. E is directly related to the solubility of the moiety and its ability to impart stability to colloidal domains in HFAs (84). At the MP2/6-31g+(d,p)//MP2/aug-cc-pVDZ level of theory, the nonbonded interaction energy between the two hydrogenated fragments was observed to be slightly more energetically favorable (EbCH2 -CH2 –4.68 Kcal/mol) than that between the PPO moieties (EbPPO-PPO –4.03 Kcal/mol, Fig. 3B) (85). The larger EbCH2 -CH2 can be attributed to the CH interactions, which arise due to the temporary fluctuation of the charge distribution originated from the electron correlation between the hydrocarbon molecules (86). The larger E observed for PPO of 1.58, compared to only 0.58 for CH2, suggests that PPO is a promising tail moiety for the stabilization of dispersions in HFAs.

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Further insight into the nature of the interaction between HFA134a and the tail fragments can be obtained by analyzing the interatomic distances and the atomic charge distribution (79,80,87). The interaction points between PPO–HFA134a and PPO–PPO are indicated by dashed lines (in Å) in Figure 3A and 3B, respectively. The HFA complex shows important interactions between the fluorine and the hydrogen atoms of the tail fragments. Also, the hydrogen atoms of the propellant are directed toward the oxygen atom of the tail fragment, indicating a stronger electrostatic C–H . . . O interaction. In addition, electrostatic potential (ESP) charge distribution (88) of the tail–HFA134a pairs was obtained at the MP2/aug-cc-pVDZ level of theory (76,77). Figure 4 shows that the highest charges among the hydrogen atoms are in PPO. In this case, more acidic hydrogens are available on the moiety for F . . . H interactions. The higher acidity of the H atoms can be attributed to the presence of the ether oxygen atom (76). From the interatomic distance and charge distribution studies, it can be concluded that the interactions between PPO and HFA134a are predominantly electrostatic in nature (76,80). In conclusion, E bst,s” have been shown to be sensitive enough to quantitatively discriminate the interaction between HFA and candidate tail fragments. Such microscopic information can, therefore, be utilized to guide the design of HFA-philic molecules. The link between the molecular-based Eb results to observable properties, such as the interfacial activity of amphiphiles containing the discussed tail groups, will be established later. B.

Chemical Force Microscopy

AFM is a powerful imaging tool for the analysis of the surface topography (89,90). Two modifications of the AFM technique are specially useful in the

Figure 4 Simple electrostatic potential charge distribution of the PPO–HFA134a complex at MP2/6-31g+(d,p)//MP2/aug-cc-pVDZ level of theory.

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development of alternative propellant-based formulations, namely Chemical Force Microscopy (CFM) (28,29) and the Colloidal Probe Microscopy (CPM) (26,27). While many articles have appeared in recent years on CPM within the context of traditional dispersion-based pMDIs and DPI formulations (23), CFM has not been explored to the same extent. The CFM technique is based on the modification of an AFM tip by covalently attaching organic monolayers with predefined functional end groups. The force of interaction between the modified tip and substrate of interest is subsequently measured in air or in solution. Similarly to the binding energy studies described above, CFM can be used to understand nonbonded interactions in condensed media. This information is critical in developing a molecular level understanding of a variety of phenomena, including colloidal stability. CFM can, in principle, measure molecular interactions ranging from weak vdW forces (<10–12 N) to strong covalent bonds (10–7 N) (23). A typical force–distance CFM curve with the complete approach–retract cycle is shown in Figure 5. The steps during approach and retraction are numbered in the figure. At large separation distances (1) there is no detectable force between tip and substrate; (2) as the tip approaches the substrate, it jumps into contact with it due to vdW attractive forces; (3) after the contact, a linear cantilever deflection is observed as the tip moves in further contact with the substrate; (4) when the cantilever deflection reaches the set point, the tip starts retracting from the substrate; (5) the force needed to pull the tip off from contact with the substrate corresponds to the force of adhesion (Fad) between the tip and substrate. The jump into contact and upon retraction happens when the gradient of the force becomes larger than the cantilever spring constant. The

Figure 5

Example of a force–distance curve obtained by chemical force microscopy.

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observed cantilever deflection (z) is converted into force using the cantilever spring constant (k) with F kz. The solvent environment plays a crucial role in the measured Fad. Large Fad will be characteristic of a poor solvent environment for the (tail) groups that modify the substrate and AFM tip. On the other hand, low Fad will be detected when the tail groups can be well solvated by the solvent environment they are in (16). CFM can, therefore, provide quantitative information on the relationship between tail chemistry and solvophilicity. The Fad, as measured by CFM, is the closest microscopic experimental analog to the Eb calculations described previously. A schematic diagram of a CFM is shown in Figure 6. It shows an AFM tip and substrate chemically modified with the same (methylbased) functional group, in HPFP solvent. The Fad of (i) a baseline alkyl moiety (CH2), representative of tails of the FDA-approved surfactants, and (ii) an ether-based tail (PPO), has been measured by CFM. The study parallels the Eb work discussed above. Octyltrichlorosilane and 3-methoxypropyltrimethoxysilane were used as representatives of the CH2 and PPO groups, respectively (91). The Fad frequency histograms are shown in Figure 7. The Fad was determined in liquid environment (HPFP), in order to evaluate the solvation capacity of HFA for the tail moieties of interest. Force–distance curves between modified tip and substrate were also determined in isooctane (ISO), in order to put the HPFP results in perspective (91). Approximately 200 measurements are shown in the frequency diagrams

Figure 6 Schematic diagram of the CFM. The modified AFM tip and substrate are immersed in HPFP.

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Adhesion force (Fad) for hydrogenated and ether-based tip substrates in HPFP or isooctane (ISO). The radius of curvature of the AFM tip is 33 nm.

Figure 7

of Figure 7, representing eight random positions on the substrate. The low Fad between PPO and PPO (tip and substrate) fragments (Fad-PPO–PPO –24.8 mN/m) indicates that the ether moiety is significantly better solvated by HPFP than the hydrogenated tail (Fad-CH2-CH2 –68.2 mN/m). It is interesting to contrast these results with Fad-CH2 -CH2 0 in ISO. It suggests that while HPFP can solvate PPO much better than CH2, it is still very far from ideal solvation. The lower Fad-PPO–PPO is in direct agreement with the more negative Eb’s determined with the ab initio calculations. CFM results provide,

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however, an important additional piece of information, which is the degree of solvation of the candidate moiety relative to ideal solvation. Combined, the molecular level information provided by CFM and Eb can be used to guide the design of HFA-philic tails for surfactants for pMDIs. In order to design amphiphiles for the HFA|W interface, however, the balance of the surfactant needs to be addressed along with the HFA-philicity of the tail group. In situ tensiometric experiments of the pressurized HFA|W interface are, therefore, discussed below. C.

Complementary Information for the Design of Surfactants for the HFA|Water Interface

The Eb and CFM results discussed previously suggest PPO to be a potential surfactant tail candidate for the HFA|W interface. On the other hand, amphiphiles with the hydrogenated fragment, the baseline tail, are expected to have poor interfacial activity at the HFA|W interface. The interfacial tension of two classes of nonionic [poly(ethylene oxide) (PEO)]-based amphiphiles, with (i) the baseline CH2-based tails (C18H35EOn), and (ii) PPO-based tri-block copolymers (EOnPO45EOn), has been reported. The tension of three FDA-approved surfactants, namely oleic acid (C17H33COOH), lecithin (C40H77O8NP), and sorbitan trioleate (Span 85, C60H108O8), has also been investigated (92). Measurements were reported at 298 K, saturation pressure of the propellant, and at 1 mM surfactant concentration—either in aqueous or in HFA phase, depending on solubility. A high-pressure pendant/hanging drop tensiometer has been developed for the measurements (49,92). As expected, the FDA-approved surfactants were shown not to be very interfacially active at the HFA|W interface (92). Oleic acid was the most active, reducing the binary tension of the HFA134a|W interface from 33.5 mN/m to 12.5 mN/m. The effect of the HFB on the activity of C18H35EOn and EOnPO45EOn at the HFA134a|W interface reveals interesting trends. The results, plotted as pressure (II 0 – , where 0 is the binary tension) versus % EO in the molecule, are shown in Figure 8 (92). The PPO class is highly active at the HFA134a|W interface, as indicated by the significantly low ’s (high II’s) compared to those seen for CH2-based tails. The HFB significantly impacts the activity of the amphiphiles. The balanced state of PPO-based surfactants is found to be at around ~50% EO, with a tension value of 1.6 mN/m (93). Even though the C18H35EOn series also takes the system through a maximum in (minimum in ), it can be argued that this value does not correspond to a “truly” balanced state, given the very poor interaction of the hydrogenated tail with the HFA134a. These results clearly indicate the fact that PPO is a better tail for HFA134a, corroborating the Eb and CFM results discussed above. The tension versus HFB (% EO) results can also be used to predict the natural curvature of the microemulsion. Surfactants that fall in the LHS with respect to the maximum in versus % EO curve are expected to form reverse aggregates. Based on this rationale, the dependence of on the concentration of EO3PO45EO3 in bulk HFA134a has been determined at 298 K, and Psat of

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Figure 8 Surface pressure () versus HFB for the surfactant modified HFA134a | W interface. Conditions are 298 K, Psat of the propellant, and 1 mM surfactant concentration.

HFA134a (93). Upon the addition of small amounts of EO3PO45EO3, HFA134a|W is substantially reduced from 31.8 mN/m to 0.5 mN/m, as shown in Figure 9. As the surfactant concentration is raised above 1.1 mM, a discontinuity in the curve of versus ln surfactant concentration (C) is observed. This discontinuity can be

Figure 9 Interfacial tension () versus surfactant concentration (C), at the HFA134a|W interface. Conditions are 298 K, and Psat of the propellant. The brake point in versus ln C curve is defined as the critical microemulsion concentration (cc).

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attributed to the formation of W/HFA microemulsions. The critical microemulsion concentration (cc) for the EO3PO45EO3 surfactant and the area per molecule (A) were reported to be 1.1 mM and 245 Å2.molecule–1, respectively (93). This value of A can be contrasted with those observed for EO13PO30EO13 (94,95), and EO19PO43EO19 at the p-xylene|water interface (94–96) of 176 Å2.molecule–1, and 215 Å2.molecule–1, respectively. D.

Atomistic Computer Simulations

While results from ab initio calculations and CFM give important insight on specific solute–solvent and solute–solute interactions, molecular dynamics (MD) computer simulations can be used to investigate the structural and thermodynamic properties of interfacial systems (97), including the formation and structure of aqueous reverse micelles (98). Study of the HFA134a|W fluid–fluid interface is of special relevance, since it can help us understand the formation and stability of macro- and microemulsion-based alternative pMDI formulations. Properties that can be determined through MD simulations include atomic density profiles, interfacial thickness, H-bonding profiles, and surface () and interfacial () tensions (92,99). There is an extensive body of scientific literature on conventional oil–water (O|W) (100–102) and compressed fluid–water (103,104) interfaces. The HFA|W interface, however, has only recently been addressed (105). Comparing and contrasting the properties of the HFA134a|W interface with those of conventional and compressible systems can help us understand the nature of the semifluorinated propellant–water interaction, as well as guide the design interfacially active amphiphiles for this interface. To a large extent, the properties of the binary HFA134a|W interface, including its free energy density, aqueous hydrogen-bonding network, and orientation of the interfacial water molecules, dictate the interfacial activity of the adsorbed amphiphiles (104). Atomistic MD simulations of the HFA134a|W interface have been recently reported (105). Equilibration of the interface was performed in the isothermal–isobaric (constant number of particles, pressure, and temperature—NPT) ensemble for at least 1 ns, at 298 K and 1.5 MPa. Production runs were carried out for at least 1 ns in the microcanonical (constant number of particles, volume, and energy—NVE) ensemble. Figure 10 is a snapshot (close-up) of the equilibrated HFA134a|W interface. Protrusions of water into HFA134a and vice versa can be qualitatively observed. The roughness of the interface arises due to thermal fluctuations (capillary waves) (106–108). The equilibrium density profiles plotted as a function of the coordinate normal to the interface (z) are shown in Figure 11. The profiles demonstrate the formation of a stable interface, with two well-separated bulk phases of HFA134a and water, and a sharp but smooth transition from one liquid phase to the other. The density profile is similar to that of the O|W (106,108) and compressed CO 2|W interfaces (103). These are indications that the chosen water and HFA potential models are capable of

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Figure 10 Snapshot (close-up) of the equilibrated HFA134a(lower phase)|Water(upper phase) interface at 298 K.

properly describing the intermolecular interactions (102). The density profile of water is smooth and fast decaying. On the other hand, HFA134a exhibits some oscillations that have been attributed to either the smaller number of HFA molecules used in the simulation and/or the relatively short

Figure 11 The density profile across the HFA134a|W interface at 298 K. The circles represent the hyperbolic tangent fit to the water profile, as described in the text.

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simulation time (109). Similar oscillations were previously reported in CCl4|W interfacial studies (106,109). The interfacial width (w) can provide significant insight into the nature of the interactions between HAF134a and water, as previously shown for O|W (102,106) and CO2|W interfaces (104). The water density profile was fit to a hyperbolic tangent functional form:  Z − Z ow   1  ρw = ρow 1 + tan h   2   w   

(5)

where wo is the water bulk density, zwo is the position of the Gibb’s dividing surface, and w is an estimate of the interfacial width. The so-called “10–90” thickness (103), t, of a hyperbolic tangent is related to w by t 2.197 w (103). The thickness of the interface is determined as the distance along the interface over which the density changes from 10% to 90% of the bulk density (102). The reported “10–90” thickness of the HFA134a|W interface is 5.2 Å (105). This value is close to 5.0 Å for n-alkane|W (102) and 4.9 Å for CCl4|W systems. As the polarity increases from alkanes and CCl4 to HFA143a, an increase in interfacial width is observed. This increase correlates with a decrease in tension from ~50 mN/m for n-alkane|W interfaces (102) to 42 mN/m for the CCl4|W interface, down to 30.8 mN/m for the HFA134a|W interface. These results agree with predictions from capillary wave theory (105,110). Therefore, systems with low s, or greater affinity to the aqueous phase, are expected to protrude more into each other’s phase. The interfacial tension was calculated from the MD simulation run using the pressure tensor equation (110): Tension =

Pxx + Pyy 1 Pzz −  2 2

  Lz 

(6)

where P

( x, y or z) is the

element of the pressure tensor, and Lz is the linear dimension of the simulation box in the z direction. The factor of ½ in the equation arises from the two liquid|liquid interfaces in the system. The reported tension of HFA134a|W from the MD simulation is 30.8 10.7 mN/m. The experimental equilibrium value of 33.5 0.3 mN/m was obtained by in situ high-pressure hanging drop tensiometry at the same temperature and saturation conditions (105). The observed HFA134a|W value is lower than that of conventional alkanes (~50 mN/m) (111) and similar to that for the compressed CO2|W interface (~30 mN/m) (112). The results reflect the increased polarity of HFA134a due to the presence of both H and F in asymmetric positions in the molecule. E.

Colloidal Probe Microscopy

The adhesive/cohesive interaction between an AFM tip modified with a single particle and the other particle, or with a planar substrate, can be measured. This technique is called Colloidal Probe Microscopy (CPM) (26,27). CPM can

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Figure 12 SEM image of the salbutamol sulfate-biodegradable polymer core-shell particles. Left lower corner: the TEM showing the core-shell structure of a nanoparticle. Right upper corner: SEM of a salbutamol sulfate microsphere attached to an AFM cantilever.

be used to understand the forces between particles, or particles with surfaces, which dominate the colloidal behavior of a great variety of materials, including paints, living cells, and drug formulations (23). Approximately 50% of currently marketed pMDIs consist of micronized drugs dispersed in HFA propellants (4). Most drug particles, however, cannot be suspended directly in HFA due to strong cohesive forces. The cohesive interaction between drug particles in HFAs has been previously investigated with CPM (23,113–115). The effects of the solvent environment and additives, such as cosolvents and amphiphiles, on traditional dispersion formulations have also been investigated with CPM (16). Such measurements are also of great relevance in alternative propellant-based formulations, where the effect of surface chemistry and morphology on colloidal stability needs to be addressed. However, adhesive/cohesive interactions are generally measured using drug crystals (probes) of irregular shape. Thus, comparison of different excipients and formulation conditions is generally hard if not impossible to be made (23,113–115). In order to address this limitation, we have recently developed a low-energy, single-step method for making smooth spherical particles of hydrophilic solutes (34). The same method can be used to make core-shell particles. Both polydispersity and particle size can be controlled by varying the degree of energy input. Using this new technique, bare and polymer-coated salbutamol sulfate spheres have been made (Fig. 12). Microspheres were attached to AFM tips and particle cohesion measured by CPM (34). This technology can also be explored to address particle/carrier cohesion/adhesion in DPI formulations.

IV.

Conclusions and Outlook

The reformulation of pMDIs with HFAs is an ongoing problem. Of noteworthy difficulty is the selection of highly soluble and interfacially active

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amphiphiles for HFA propellants. These challenges have stimulated the development of alternative dispersion-based pMDI formulations. Molecular-based information from experiments and computer simulations can be used to guide the development of such novel formulations, which hold special promise in the local and systemic delivery of biomolecules to and through the lungs. Such advances can help pMDIs compete with recent DPI technologies.

References 1. 2. 3. 4. 5. 6.

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14 Aerosol Generation from Propellant-Driven Metered Dose Inhalers

HUGH D. C. SMYTH

RICHARD M. EVANS†

College of Pharmacy, University of New Mexico, Albuquerque, New Mexico, U.S.A.

Inspire Pharmaceuticals, Durham, North Carolina, U.S.A.

ANTHONY J. HICKEY School of Pharmacy and Medicine, University of North Carolina, Chapel Hill, North Carolina, U.S.A.

I.

Introduction

Propellant-driven, or pressurized, metered dose inhalers (pMDIs) have now been used for pharmaceutical inhalation aerosols for 50 years. Their introduction, in 1956, was revolutionary in that they allowed the development of formulations for repeated, reproducible dose delivery through a metering valve for airway administration (1). Significant changes to pMDIs have transpired with the phaseout of ozone depleting propellants based on chlorofluorocarbon (CFC) chemistry but pMDIs, especially those based on hydrofluoroalkane (HFA) chemistry, remain an important component of inhalation aerosol delivery systems. Metered dose inhalers consist of active ingredient either in solution or in suspension in high vapor pressure propellants. This product is emitted through an orifice from a metering valve of known volume by means of an actuator. There is no single model for droplet formation from volatile nonaqueous propellant. Consequently, the output droplet/particle size from these metered dose inhaler systems has been, to a large extent, developed empirically. †Deceased.

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The continued importance of economical and convenient aerosol delivery systems necessitates a greater understanding of propellant droplets and their behavior. This chapter describes some general concepts that may be used as a guide to the performance of propellant systems with a particular focus on alternative propellants that will eventually eliminate CFC systems. This is an attempt to present the major factors in spray formation. Spray formation cannot be addressed without considering some aspects of the propellant system and the flow and evaporation of the droplet plume. Thus, the text begins by describing factors prior to passage of the formulation through the spray nozzle, and concludes by considering the nature of the product at a designated distance beyond the orifice. Droplet formation is significantly influenced by the vapor pressure of component propellants and by the presence of solids, surfactants, or cosolvents in the system (2,3). The vapor pressure of the propellants equilibrates with atmospheric pressure as the metering chamber of the valve is opened to the environment by depressing the valve stem using the actuator, as shown schematically in Figure 1 (4,5). This flow of liquid propellant from the metering chamber is analogous to the transient release of liquid from a pressurized vessel, but with the added complication of additional vaporization of the propellant as it passes into the expansion chamber. The equilibration with atmosphere results in rapid expansion boiling of the propellant formulation and causes propulsion of the contents of the metering chamber through the actuator orifice with concurrent evaporation of the propellant. This rapid propellant evaporation is termed cavitation and occurs because the intermolecular forces are overcome by the magnitude of the pressure gradient across the liquid.

Figure 1

dispersion.

Schematic diagram of the components of a metered dose inhaler and droplet

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Fluid dynamics of propellant within the expansion chamber and nozzle are very complicated because of the presence of pressure fluctuations that are associated with turbulent eddies and lead to unpredictable cavitation behavior (6). Domnick and Durst investigated cavitation of propellant CFC 12 when flowing through a constriction (7). Recirculation regions were identified, where bubbles are formed and grow within the constriction until they are washed downstream. Dunbar used this model to suggest similar bubble growth regions within the geometry of a typical pMDI actuator (8,9). Experimental determination of plume oscillation frequencies at approximately 700 Hz was similar to that observed by Domnick and Durst. Numerical simulation of these processes is beyond the capacity of currently available models (6). In addition to breakup within the expansion chamber and nozzle, droplets may be formed from the liquid propellant formulation due to aerodynamic loading on the surface of the liquid that overcomes surface tension. As the droplets are formed, the plume begins to expand (5,10). Figure 2 shows an instantaneous image and a time-averaged image of an HFA plume being emitted from a cut-away actuator to reveal the spray the atomization orifice. Depending on the formulation, a proportion of droplets of large mass is emitted along the axis of the actuator orifice and projects rapidly and independently of the plume formation (Fig. 3) (5,10). This was a specific issue with CFC-based systems and appears to be significantly reduced in the HFA replacement propellant systems (11). When sprayed into ambient air, the plume droplets generally decelerate rapidly due to loss of mass from evaporation and passage from a region of cocurrent to countercurrent flow. During this flow process, surface-active molecules, which may be included as dispersing agents or valve lubricants, will influence the evaporation rate of the propellants. The presence of solid particles in suspension will influence the physicochemical behavior of droplets both in particle size and distribution and also from a surface and colloidal chemistry perspective (12–14).

Infrared images of an HFA 134a plume. The left panel shows a single digital image captured of the formulation being emitted from a pMDI orifice. Aerosol eddies are observed on the edges of the plume where spray velocities are significantly reduced. The right panel shows a composite image containing approximately 60 averaged images of a single spray captured using rapid digital imaging photography.

Figure 2

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Figure 3 Dark-field photographs of a fully formed CFC aerosol plume emitted from an actuator (A) showing the heterogeneity of the droplet concentration and (B) the front of the plume showing the emanation of large droplets. Source: Courtesy of Charles Thiel and 3M Pharmaceuticals, St. Paul, Minnesota, U.S.A.

The performance and efficacy of metered dose inhalers are linked to several key patient and device-related factors. From a design perspective, the size of the aerosol droplets, the distribution, and the plume pattern are key elements influencing pMDI performance in terms of achieving respirable aerosol therapy. In the majority of metered dose inhalers the exit velocity of droplets from the actuator orifice is very large (30–200 m/sec) (6,9). The significance of this initial velocity on plume deposition will be related to the droplet size distribution, the rates of evaporation, and the plumes interaction with the inhalation air stream of the patient. Finlay has presented estimates of post-nozzle breakup of HFA droplets based on several empirical models and experimentally observed characteristics of the emitted spray (6). At lower exit velocities of around 30 m/sec droplet breakup due to aerodynamic loading is not expected for HFA droplets with initial sizes of between 10 ␮m and 30 ␮m.

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However, for higher exit velocities [200 m/sec has been estimated to be exit velocity of HFA 134a droplets (9)], droplets with initial size of between 10 ␮m and 30 ␮m will be broken down. However, Finlay estimated that breakup distances were significantly longer for 30 ␮m droplets (13 cm) compared to 10 ␮m droplets (4.2 cm), and would lead to impaction on oropharyngeal surfaces. Contrary to common belief, therefore, high exit velocities may actually be beneficial for improving lung deposition efficiency of pMDIs, provided that there is sufficient distance between the actuator orifice and the oropharynx. The initial droplet size, spray angle, and velocity contribute to the plume pattern. The dimensions of the plume are thought to be of great importance in the behavior of an aerosol in transit to the lungs, although no studies have been performed to demonstrate this. A comparison of the plume dimensions with the dimensions of confining volumes/spaces through which it passes illustrates the potential limitation of wall or oropharyngeal losses. This is particularly true of spacer devices (15–17) and mouth deposition (18). It has been suggested that actuation 10 cm in front of the mouth prior to inhalation would offer a therapeutic advantage by decreasing oropharyngeal deposition and allowing further evaporation to occur (19). Despite considerably reduced obstruction to the plume under these conditions, aerosol deposition does not improve significantly because much of the aerosol remains in the surrounding air. In addition, with HFA-based pMDIs there is some doubt over whether significant post-nozzle evaporation occurs such that spacers facilitate any reduction in droplet size (20). In Chapters 1–3, the actual and predicted effects of particle size on aerosol deposition were outlined. It is clear that particles that are >10 ␮m in diameter will rarely enter the lungs. The target size range to achieve delivery to the lungs is 1–5 ␮m. Droplet diameter at the orifice have been estimated to be greater than 15 ␮m for CFC aerosols but are significantly less in HFA aerosol systems (21,22). Regardless of system used, pMDI drug delivery depends on physical processes that occur during and following actuation in order to achieve appropriate delivery of therapeutic doses. II.

Observed Droplet Formation and Dispersion

The study of droplet formation, particularly from aqueous solutions, has been of interest for many years (6,23). It is understood that the conditions of droplet generation will influence particle size and distribution, but the mechanism of initial droplet formation is very similar from one system to another. Thin strands or sheets of fluid form as inertia overcomes the surface tension maintaining fluid integrity. These thin strands or films then break into projectiles of fluid, which, due to surface tension, take on the geometry that minimizes the surface free energy (surface area) for a given volume, namely a sphere. This sphere of fluid, illustrated schematically in Figure 4A, will then equilibrate thermodynamically with the surrounding atmosphere (23). Droplet formation from a single liquid phase moving relative to a stagnant gaseous phase has been considered above. The most common situation in pharmaceutical products is that of a liquid phase being moved in, or by, a gaseous/vapor phase, relative to a stagnant gaseous phase, as illustrated in

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Schematic diagram of the droplet formation from an orifice (A) low vapor pressure fluid and (B) high vapor pressure fluid.

Figure 4

Figure 4B. There are some regions of lower density or vapor passing through the orifice. Under these conditions, not only is the gaseous/vapor phase contributing significantly to the inertia of the liquid coaxially with the orifice, but also the turbulence of this phase gives rise to moments of inertia at angles perpendicular to this axis. Consequently, shear thinning occurs and a larger number of small droplets are formed. This may be similar to convective boiling in tubes (24). Consequently, this situation involves complex multiphase turbulent flow that has dynamic events occurring on the scale of microseconds. The cross-sectional nature of the plume can be investigated using a number of methods, including traditional thin-layer chromatography (TLC) plate exposure following actuation to a fixed distance (25) and laser light sheet digital illumination of plumes using high-speed laser diagnostic methods (26,27). These methods are not yet refined for allowing detailed understanding of the development of the aerosol plume from the orifice to downstream plume capture. Further, dimensional variations in the molding of the actuator orifice, which can be controlled by physical measurement of critical dimensions, can also influence the shape and direction of the aerosol plume. Despite the ability to control plume geometry through the establishment of in-process control of critical dimensions, regulatory agencies continue to suggest that these techniques be incorporated for routine quality control and product development studies. Studies correlating plume generation properties to in vivo measures of performance are necessary to guide formulators and device engineers toward parameters that are significant for product performance. Figure 5 illustrates plume patterns produced by standard and imperfect and/or off-centered

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Illustration of the effect of actuator orifice imperfections/centering inducing (A) left and (C) right deflections of the plume pattern (B) from a metered dose inhaler as indicated by thin-layer chromatography. Figure 5

orifices. These methods may also be able to detect regional differences in droplet density within the plume that may influence airway deposition patterns. The effect of solids in solution on spray formation has been demonstrated by Dalby and Byron (28). It is clear that the median particle/droplet size is increased by the presence of solids. This is not simply an effect of the difference between the primary particle size of the suspension and the aerosol droplet size, but may also be attributed to the tendency of particles to aggregate in suspension. Gonda and Chan have considered the mathematical probabilities of such aggregation (29,30). The effects of surfactants on interfacial properties of organic droplets have been discussed at length in another chapter. Surfactants are incorporated into nonaqueous suspensions to aid in stabilizing the system and as a valve lubricant (5,10,31,32). As the surfactants are capable of interacting to form inverted micelles (33–36), they are typically used at concentrations below the critical inverted micelle concentration (32,37,38). Because there may be multiple critical aggregation concentrations in nonaqueous solutions (39,40), it is important to define the concentration capable of solubilizing the drug. With the introduction of HFA propellant systems, surfactant use in pMDIs has declined due to poor solubility. Due to the difficultly in maintaining suspension stability without surfactants, often formulators have been forced to reformulate products as solutions instead (41). Using cosolvents such as ethanol the solubility of surfactants, drug substances, and excipients in the HFA propellants can be increased (5,32). However, surfactants are often required in cosolvent-free systems (for stability of a suspension formulation) and alternative surfactants will need to be proven safe and effective (42). As yet, no alternative surfactant compounds have been included in regulatory-approved and marketed products. A recent review by Rogueda details strategies of developing stable and reliable HFA suspension systems (43). Despite the problems of HFA suspension systems (fast separation time, aggregation, particle growth, aerosolization properties, and drug losses to adhesion), a significant number of marketed HFA pMDIs are suspensions (43).

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Several products that were suspensions in CFC propellants have now been reformulated as solutions. The positive deviation in vapor pressure from Raoult’s law observed in ethanol/HFA systems may facilitate the use of higher concentrations of ethanol for improved solubility of drug without detrimental effects on droplet size or aerosolization (44,45). However, recent evidence suggests that increasing ethanol concentrations by 10% w/w will have a significant impact on mass median aerodynamic diameter (MMAD) and droplet size (46,47). Several examples of the influence of formulation on droplet sizes and plume characteristics are well detailed in the literature by comparing CFC-based systems with HFA replacements. For example, flunisolide HFA, an inhaled corticosteroid reformulated using HFA 134a, is a solution aerosol, unlike flunisolide CFC that is a suspension aerosol. Flunisolide HFA has a smaller particle size (mass median aerodynamic diameter ⫽ 1.2 ␮m) than flunisolide CFC (3.8 ␮m) and has significantly higher lung deposition and less oropharyngeal deposition (41). This product is currently awaiting FDA approval. A significant issue for the pharmaceutical industry during the transition from CFC to HFA systems has been to match the performance of replacement inhalers to those of existing products to facilitate rapid regulatory approval. This process, sometimes called “detuning,” has been facilitated by empirical knowledge of the effect of nonvolatile components in pMDIs on particle size generation. An example of a solution HFA system designed to resemble CFC suspension system is the budesonide HFA product. The design of a budesonide HFA solution formulation that matched the marketed CFC product performance (Pulmicort™) was achieved using the addition of a nonvolatile component and modifying the actuator orifice diameter (48). These variables could influence the fine particle dose and its mean particle size in different ways such that they could be covaried to yield aerosol output with size characteristics very close to the CFC products (49). There are a number of methods for studying the phenomenon of dissolution of polar compounds in nonaqueous solution. For pressurized systems, the methods are complicated by the need to maintain propellants in liquid state. In HFA systems, there is no low-volatile component apart from cosolvents such as ethanol. Thus, several methods of solubility determination have been devised. Dalby et al. described an apparatus for the determination of solubility in CFCs mixtures, and demonstrated how limited solubility can result in Oswald ripening (50). Williams et al. (51) modified the apparatus for the measurement of solubility in hydrofluorocarbon-based systems. More recently, Gupta and Myrdal described a direct HPLC injection method (52) and found their method to be in good agreement with that of Dalby and Williams.

III. A.

Fundamentals of Droplet Formation Vapor Pressure Considerations

The vapor pressure over a propellant system is governed by Raoult’s and Dalton’s laws. Raoult’s law states that the partial pressure, P⬘A, of a solvent A

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Table 1 Some Relevant Physicochemical Properties of Chlorofluorocarbon (11, 12, 114) and Hydrofluorocarbon (134a, 227) Properties Propellant (chemical structure) 11 (CCl3F)a 12 (CCl2F2)b 114 (C2Cl2F4)a 134ab (CH2FCF3) 227c

Mol° wt. (g/mode)

Vapor pressure (psia at 25°C)

Boiling point (°C at 1 atm)

Density g/cm3 at 25°C

137.37 120.91 170.91 102.0 170.0

13.4d 94.5 27.6d 96.0 72.6

23.82 -29.79 3.77 -26.5 -17.3

1.476 1.311 1.456 1.203 1.415

in a system is equal to its mole fraction (XA) multiplied by the pure vapor pressure (PA°) or P⬘A ⫽ XAP°A

(1)

where the mole fraction is the ratio of the number of moles of A(nA) to the total number of moles (nA + nB + . . . + nn). Dalton’s law states that the total vapor pressure over the system is equal to the sum of the partial pressures or (2) P ⫽ P⬘ + P⬘ + . . . + P⬘ T

A

B

n

Assuming ideal behavior, a knowledge of the mass of propellant(s) present, their molecular weight and pure vapor pressure will allow calculation of the total vapor pressure of the system. Table 1 shows some of the physical characteristics for propellants 11, 12, and 114, the CFCs used in pharmaceutical inhalation aerosols, and also for HFA propellants 134a and 227. The vapor pressure of the system can therefore be adjusted within predictable limits by mixing propellants with different pure vapor pressures. Modulation of vapor pressure in HFA propellant systems using binary or ternary blends does not always follow ideal gas laws as described above. Specifically, HFA propellant and ethanol mixtures show positive deviations from Raoult’s law (32,53–55). As mentioned previously, it was thought that this deviation from ideality may allow increased levels of cosolvent to be used without significant vapor pressure decreases (32). This may be the case, but droplet formation appears to be influenced detrimentally by the increase in nonvolatile fraction (despite vapor pressure being maintained). Propellants offer the additional advantage of a small range of density that may be adjusted to some degree in considering quantities for mixing to balance the density of the drug particles in suspension. B.

Valves and Actuators

The metering chamber volume and actuator geometry can have a significant effect on the output from a metered dose inhaler. The volume of the metering chamber, the orifice size of the valve stem, and the orifice size and profile of the actuator dictate the volume and rate of emission of the aerosol formulation (56). The dead space between the actuator orifice and the metering chamber acts as an expansion chamber in which the propellant forms a mixture of liquid and vapor phases before exiting.

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Figure 6

C.

Structures of (A) oleic acid, (B) sorbitan trioleate, and (C) phosphatidyl choline. Surfactants

Surfactants have been included in metered dose inhaler formulations as dispersing agents and valve lubricants. The commonly employed surfactants in CFC-based pMDIs are oleic acid, sorbitan trioleate, and phosphatidyl choline, as shown in Figure 6. The behavior of these surfactants in propellant systems is known to result from polar interactions between these molecules alone, with drug particles and nonpolar interactions with the surrounding medium. Hydrophobic interactions of the aliphatic tail groups of these molecules are thought to stabilize suspended particles in the predominantly propellant systems (57). Due to the limited solubility of the traditionally used CFC-based surfactants in HFA propellants, novel surfactants are under investigation (42). In addition, valve technology has had to evolve to function without the use of these compounds (58–61). D.

Flow Through an Orifice

The orifice geometry and dimensions will have an effect on the droplet formation. Functionally, the orifice region of the actuator consists of the “sump,” a small region beneath the valve stem in which expansion occurs, and the orifice through which the mixed vapor/liquid phase is emitted, as shown in Figure 7. The dimensions of the expansion chamber and orifice may be varied. In addition, the geometry of the sump or the orifice may be varied, as shown in Figure 8. Limited work on the influence of modifying these parameters on the emitted plume and particle size characteristics had been reported in the scientific literature (56). E.

Evaporation and Shear Thinning

As the aerosol is being generated, a number of phenomena are occurring. Propellant begins to evaporate, the velocity of the product overwhelms the surface tension, and local changes in density occur. The resultant shear thinning inertia and evaporation cause the formation of droplets that continue to evaporate as they are propelled into the surrounding air.

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Photographs of actuator orifices through which aerosol is emitted, showing the symmetry of the opening.

Figure 7

IV.

Theoretical Framework Describing Droplet Formation

The effectiveness of an inhalation aerosol delivery system “depends on its ability to produce particles or droplets in a range suitable for entry into the lungs.” Predicting the spray formation produced by scientific formulations, valves, and actuators would aid in product development. An analysis of this problem may be initiated by adopting a simplified theoretical approach to predict spray formation from a single component system (62). Preliminary experiments have been performed to measure the size of particles produced by a model aerosol system for comparison with the theoretical model. A.

Transorifice Conditions

The diagram of an actuator orifice shown in Figure 9 shows the characteristic dimensions needed to evaluate droplet output. In order to assess droplet production from any nozzle, it is first necessary to know the volumetric flow rate (63):   2 ∆P Q = C D a0   ρ( 1 ( / ) − a a 0 1  

0.5

(3)

where CD is the coefficient of discharge, a0 is the orifice area, a1 is the inlet area, ␳ is the density of the fluid, and ⌬P is the pressure drop across the nozzle.

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Figure 8 Diagram demonstrating the complexity of the expansion region of an actuator. (a) upper cylinder, (b) lower cylinder, (c) lower frustrum, and (d) orifice.

The linear velocity can be determined by dividing the volumetric flow rate by the orifice area: U ⫽ Q/␣0.

(4)

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Figure 9 Schematic diagram of actuator orifice indicating dimensions required to calculate performance characteristics. P1 and P2 are canister and atmospheric pressure; d1, a1 and d0, a0 are the expansion chamber and orifice diameter, and areas, respectively.

B.

Droplet Formation

The particle size, mass median diameter (D), has been derived empirically for a plain circular orifice discharging into stagnant air (23): Dm ⫽ 6 d0 (ReL)–0.15

(5)

where d0 is the orifice diameter and ReL is the Reynolds number for flow. The Reynolds number is expressed as (64) Re =

ρd0U η

(6)

where ␩ is the viscosity of the fluid; ␳, d0, and, U are defined above. The volume median diameter may be calculated as follows: DV ⫽ Dm/␳

(7)

Using the above expressions, it is possible to derive an estimate of the mass median diameter at the orifice. C.

Evaporation

The evaporation rate, by diffusion, of a droplet can be calculated from the following expression (64): dDm 4 Dc M = dt RρDm

 P∞ − Pd     T∞ − Td 

(8)

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where Dc is the diffusion coefficient, M is the molecular weight, P⬁ and Pd are the ambient and droplet surface pressures, T⬁ and Td are the ambient and droplet surface temperatures, and R is the gas constant. The diffusion coefficient can be calculated as follows (65): 3

Dc =

1

BT 2 [(1 / M1 ) − (1 / M 2 )] 2 P(r12 )2 I D

(9)

where B is (10.85 ⫺ 2.5 (1/Ml) ⫺ (1/M2)) ⫻ T is absolute temperature, Ml and M2 are molecular weights of components 1 and 2 (propellant and air), P is absolute pressure in atmospheres, r12 is the collision diameter, and ID is the collision integral. The behavior of the droplet after leaving the orifice can therefore be predicted. Finlay has shown that evaporation estimates must be corrected for Stefan flow for most propellant systems (6). This is necessary because the assumption that evaporation does not result in bulk motion of air surrounding the droplet is not always valid. 10–4,

D.

Compositional Considerations

In the outline of some of the crucial components of droplet formation during generation from a metered dose inhaler, compositional issues have been neglected. These are clearly of great significance. The presence of solid particles or surfactants at the surface of a droplet may influence droplet formation and heat and mass transfer. This may occur as a passive surface area effect with similarities to the colligative property of vapor pressure lowering (66). It may also result in a complex thermodynamic phenomenon involving modification of the Gibbs free energy at the interface, in turn influencing the number and size of droplets formed (67). E.

Primary Atomization for HFA 134a pMDIs

Dunbar described and validated the first model of the pMDI primary atomization process and resultant spray characteristics for a pure HFA propellant (8). Termed an Actuator Flow Model, the analysis was performed by calculated multiphase fluid flow in three compartments that are active during actuation (22). In contrast to the model developed by Clark (68), internal flash and cavitation mechanisms were assumed to be the primary atomization mechanism. This was based on experimental observations that indicated that preatomization was occurring prior to the emission of the propellant from the orifice (9). In correlations with droplet size distributions, it was found that only the semiempirical model developed by Clark allowed reasonable estimation of droplet sizes exiting the pMDI nozzle. V.

Semiempirical Model of Droplet Formation

With respect to CFC-based systems, Clark has thoroughly reviewed the literature on the formation of droplets by metered dose inhalers (68). Clark then developed models in a rigorous manner based on thermal and dynamic

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equilibrium for metered discharges that were corrected empirically from studies of continuous flow. The atomization process was investigated by measuring the residual droplet diameters resulting from the atomization of propellants containing nonvolatile solutes. A correlation relating the calculated initial median droplet diameter to atomizer conditions was developed: C5

di = qem

 Pe − PA     PA 

n

(10)

where C5 is a constant, relating pressure and quality to droplet size; qe is mass fraction of vapor phase in the expansion chamber; m is a constant relating quality of flow to droplet size; n is a constant relating pressure to droplet size; Pe is pressure downstream of discharge orifice; and PA is atmospheric pressure. The form of this correlation strongly suggests that the droplet formation mechanism is aerodynamic shear between the vapor and the liquid phases rather than rapid bubble growth or flashing, as had been previously suggested. In this extensive evaluation of droplet formation from suspension formulations, the droplet dispersion was inversely proportional to spray orifice diameter, surfactant concentration, and the logarithm of the solids’ concentration, and proportional to the suspended solids’ particle size and square root of pressure. Changing the metered volume and vapor pressure has been correlated with effects on lung deposition, which is consistent with the influence of these characteristics on droplet dispersion (69,70). The computer models and droplet correlation function developed during these investigations represent powerful tools for use in the design of both current and future HFA/HFC (hydrofluorocarbon)-powered metered dose inhalers delivery systems. It may ultimately be possible to predict the formation of droplets from these devices and minimize the time required to develop suitable formulations for pharmaceutical aerosol delivery. VI.

Conclusion

Metered dose inhalers are devices for delivering small reproducible doses of drugs to the lungs. The effectiveness of these systems depends on selection of an appropriate formulation in conjunction with components suitable for delivery of the product. Traditionally, metered dose inhalers consist of surfactants and active ingredients dispersed in propellants, now HFA-based systems. If the drug is intended for solution formulation, a cosolvent may also be present—usually ethanol. These materials are dispensed through a metering valve, most frequently from an aluminum canister or glass bottle via an actuator that incorporates the patient mouthpiece and may also include a spacer/expansion chamber. The formation of droplets in a size range suitable for delivery of drugs to the lungs is the crucial aspect for the success of these delivery systems. Because the droplet output from a metered dose inhaler varies during and following

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emission, due to heat and mass transfer phenomena, it is difficult to predict or evaluate the efficacy from in vitro studies. However, it is clear that there are factors that predispose this system to disperse droplets with greater efficiency. These have been identified as propellant vapor pressure, metered volume, expansion chamber dimensions, spray orifice diameter/profile, surfactant and suspended particle concentration, and size and distribution of suspended solids. The small fraction of aerosol (as low as 10–20%) entering the lung observed with CFC-based systems (71) was consistent with visual observation of the performance of these aerosols and models intended to predict droplet behavior. With the introduction of HFA propellant systems, much greater efficiency of lung delivery can be achieved due to the different physicochemical properties and the atomization performance of smaller metering volumes that have been used for these propellants. Despite their superior performance, significant issues need to be addressed so that alternative propellant systems may have a wide range of applicability to current and new chemical entities in development. Most notably is the development of surfactant and excipient systems that confer suitable stability to solution and suspension systems while exhibiting good safety profiles. References 1. 2. 3. 4.

5. 6. 7. 8. 9. 10. 11.

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Riker Laboratories, I., Self-propelling, powder-dispensing compositions. US Patent # 837,465. 1960. Masters K. Drying of droplets/sprays. In: Spray Drying Handbook. 5th ed. Longman Scientific and Technical, New York, NY 1993:309–352. Seaver M, et al. Evaporation kinetics of ventilated waterdrops coated with octadecanol monolayers. J Phys Chem 1992; 96:6389-–6394. Hickey AJ. Aerosol filling equipment for the preparation of pressurized pack pharmaceutical formulations. In: Hickey A.J, ed. Pharmaceutical Inhalation Aerosol Technology. NY: Marcel Dekker, 2004:311–343. Smyth HD. The influence of formulation variables on the performance of alternative propellantdriven metered dose inhalers. Adv Drug Deliv Rev 2003; 55(7):807–828. Finlay WH. The Mechanics of Inhaled Pharmaceutical Aerosols. London: Academic Press, 2001. Domnick J, Durst F. Measurement of bubble size, velocity and concentration in flashing flow behind a sudden constriction. Int J Multiphase Flow 1995; 21:1047–1062. Dunbar CA, Watkins A.P, Miller JF. Theoretical investigation of the spray from a pressurized metered-dose inhaler. Atomization Sprays 1997; 7(4):417–436. Dunbar CA, Watkins A.P, Miller JF. An experimental investigation of the spray issued from a pMDI using laser diagnostic techniques. J Aerosol Med 1997; 10(4):351–368. Hallworth GW. The formulation and evaluation of metered dose inhalers. In: Ganderton D, Jones TM, eds. Drug Delivery to the Respiratory Tract. NY: VCH Publishers, 1987. Gabrio BJ, SWV Stein DJ. A new method to evaluate plume characteristics of hydrofluoroalkane and chlorofluorocarbon metered dose inhalers. International Journal of Pharmaceutics, 1999; 186(1):3–12. Crank J. The Mathematics of Diffusion. 2d ed. Oxford: Oxford Science Publications, 1975. Prandtl L, Tiejens OG. Fundamentals of Hydro- and Aeromechanics. Mineola, NY: Dover Publications, 1957 (Vol. reprinted from 1934). Tovbin YK. Theory of Physical Chemistry Processes at Gas-Solid Interfaces. Boca Raton, FL: CRC Press, 1991. Brown PH, et al. Do large volume spacer devices reduce the systemic effects of high dose inhaled corticosteroids? Thorax 1990; 45(10):736–739. Kraemer R. Babyhaler—a new paediatric aerosol device. J Aerosol Med 1995; 8(suppl 2):S19–S26. Newman SP, et al. Improvement of pressurised aerosol deposition with Nebuhaler spacer device. Thorax 1984; 39(12):935–941.

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44. 45. 46. 47.

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15 Medical Devices for the Delivery of Therapeutic Aerosols to the Lungs

RICHARD N. DALBY

SUSAN L. TIANO

University of Maryland at Baltimore, Baltimore, Maryland, U.S.A.

Schering-Plough, Kenilworth, New Jersey, U.S.A.

ANTHONY J. HICKEY School of Pharmacy and Medicine, University of North Carolina, Chapel Hill, North Carolina, U.S.A.

I.

Introduction

The last decade has seen developments in all categories of inhaler device. However, the underlying principles of aerosol delivery have changed little. Despite the numerous designs and configurations that may be employed to generate aerosols in therapeutically useful size ranges and concentrations (Chapters 9–11 and 14), there are only three basic categories of aerosol delivery systems for commercially marketed drug products. These are pressurized metered-dose inhalers (pMDIs), dry powder inhalers (DPIs), and nebulizers. These three classes of devices do not represent optimal delivery systems in terms of their ability to produce monodisperse aerosols that can be precisely dosed in a single breath but rather are examples of delivery systems that achieve minimally acceptable characteristics in a simple, convenient, inexpensive, and portable format. To be acceptable for clinical use, an inhalation delivery system must meet certain criteria: 1.

2. 3. 4. 5.

It must generate an aerosol with most of the drug carrying particles less than 10 ␮m in size and ideally in the range 0.5–5 ␮m, the exact size depending on the intended application. It must produce reproducible drug dosing. It must protect the physical and chemical stability of the drug. It must be relatively portable and inconspicuous during use. It must be readily used by a patient with minimal training. 417

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These minimal requirements alone do not guarantee commercial success. Most commercial products currently under development aim to provide multiple dosing (typically 200 doses) with minimal excipient inhalation (which can lead to poor organoleptic properties in the mouth and oropharyngeal irritation). Patient convenience, competitive manufacturing costs to pMDIs, and added value features, such as dose counters or an indication of appropriate inhalation flow rates, are also considered desirable. In this chapter, the components, designs, and operating conditions of typical inhalation products are discussed, together with the possibilities that could be realized by “next generation” aerosol delivery systems.

II.

Pressurized Metered-Dose Inhalers

Since the 1950s, pMDIs have been the mainstay of inhalation therapy, ostensibly because they were perceived to meet most of the criteria outlined above. However, over the years, a number of deficiencies have been identified. Only a small fraction of the drug escaping the inhaler penetrates the patient’s lungs (1,2) due to a combination of high particle exit velocity and poor coordination between actuation and inhalation. The unstable physical nature of suspended drug particles in propellant, combined with suboptimal valve designs, has led to reports of irreproducible dose metering following a period of rest (3). Low concentrations of potentially carcinogenic compounds were extracted from valve components by the propellant system (4) and inhaled by the patient. However, the largest threat to the continued availability of pressurized pMDIs is their dependence on chlorofluorocarbon (CFC) propellants, which have been linked to the depletion of stratospheric ozone and are now scheduled to be phased out under the terms of the “Montreal Protocol on Substances that Deplete the Ozone Layer” (5,6). Despite these concerns, new device designs, improved formulations and valves, and a switch to “environmentally friendly” propellants continue to keep the pMDI in common use. The FDA has recently described its position on the phaseout of CFC propellants in pMDI products (7). This is the first step to limit the use of propellants in future products. The modern pMDI shown in Figure 1 is little changed from its predecessors and contains the same three basic ingredients, drug, one or more propellants, and in most cases a surfactant. A liquefied propellant serves both as an energy source to expel the formulation from the valve in the form of rapidly evaporating droplets and as a dispersion medium for the drug and other excipients. A surfactant is typically present to aid with the dispersion of suspended drug particles or dissolution of a partially soluble drug and to lubricate the metering value mechanism. In some formulations, a surfactant is reported to be unnecessary (8). Drug can be dissolved in the liquefied propellant/surfactant combination, with or without the aid of a less volatile cosolvent (9,10), or suspended in the form of micronized particles (11). In all currently marketed formulations, drug dissolution necessitates the use of an ethanolic cosolvent. Flavors (such as dissolved mint extracts) and suspended sweeteners (for example, micronized

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Figure 1 Diagram of a typical pressurized metered-dose inhaler showing mechanism of particle formation.

saccharine) may be present to combat the unpleasant taste associated with significant oropharyngeal deposition following inhalation. To enhance chemical stability, antioxidants (ascorbic acid) or chelating agents (EDTA) may be present in formulations in which the drug is dissolved. The popularity of traditional CFC propellants stemmed from their low pulmonary toxicity, high chemical stability and purity, and compatibility with commonly used packaging materials. In addition, they are nonflammable. Combinations of the three most widely used CFCs, trichlorofluoromethane (CFC-11), dichlorodifluoromethane (CFC-12), and 1,2-dichlorotetrafluoromethane (CFC-114), are typically combined in varying ratios to achieve a desirable combination of vapor pressure, liquid density, and solvency (12). Following a long search for alternative propellants with similar characteristics to CFCs, 1,1,1,2-tetrafluoroethane (HFC-134a) emerged as the primary replacement, and commercial formulations containing this propellant have gained or are awaiting marketing approval in several countries (13). In addition, 1,1,1,2,3,3,3-heptafluoropropane (HFC-227) is being actively investigated. In the recent past, numerous other propellants were investigated (14), and formulations containing them have been described in the patent literature (15). However, the expense of commercial development, primarily due to chronic toxicity testing, appears to have reduced the interest of many companies in these propellants. The physical and chemical properties of both CFCs and the emerging propellants are widely reported. Table 1 of Chapter 14 provides information on the important properties of the propellants mentioned in this chapter. Gaseous propellants and nonpressurized sprays are not used due to decreasing internal pressures during use and the inability of pumps to produce small enough droplets, respectively. Historically, propellant blend compositions appear to have been somewhat arbitrarily selected, as indicated by multiple drugs from the same company being delivered using the same propellant and surfactant combinations. This practice, while minimizing change over time during manufacturing, is unlikely to yield formulations that provide optimal drug delivery. The

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primary formulation characteristic dictated by the propellant is product vapor pressure. This determines the size and exit velocity (16) of the emitted spray and the rate at which propellant evaporation occurs. Careful optimization is essential. High vapor pressures produce faster propellant/droplet exit velocities, which may lead to enhanced oropharyngeal deposition. However, such blends also yield smaller, faster evaporating droplets that facilitate production of smaller inhaled particles or droplets, which can enhance lung penetration. Conversely, a less volatile blend may minimize impaction in the throat at the expense of producing larger, slowly evaporating droplets that are prone to impaction high in the respiratory tract. This is apparent when the smaller particle size of an aerosol in which the compound is dissolved in pure CFCs is compared to one containing significant amounts of less volatile ethanol (10,11). The effect of vapor pressure on pMDI leakage rates, transportation regulations, and filling process must also be considered. The propellant blend also dictates the product density because the other excipients are present at low concentrations. Large density differences between the propellant blend and the true density of suspended drug particles are known to cause erratic dosing if there is a delay between shaking and actuation of the pMDI (11), owing to the nonhomogeneous drug distribution within the canister due to drug sinking or floating in the propellant. This can be minimized by matching the drug and propellant density or by facilitating deflocculation of the bulk suspension with an appropriate surfactant. Propellant blending also allows the solubility of a drug or surfactant in the propellant system to be manipulated because some propellants (such as CFC-11) are often better solvents than others (such as CFC-12). However, it is important to remember that density and solvency cannot be manipulated independently of vapor pressure, so in practice the changes that can be made are limited. One problem with the availability of only one alternative propellant (HFC-134a) is the loss of flexibility associated with the use of blends. Several papers and patents address this issue with the inclusion of ethanol as a cosolvent (15,17). While this facilitates slurry production during manufacturing, reduces the vapor pressure, and enhances the solubility of several surfactants in HFC-134a, it may result in a reduced fine particle fraction following pMDI actuation. Oleic acid, soya-derived lecithin, and sorbitan trioleate were widely used as surfactants in CFC-based formulations. The choice of concentration is usually determined by experimentation in an attempt to maximize the fine particle fraction and dosing reproducibility. Concentrations as high as 2% w/w have been used in concentrated suspension formulations, presumably to facilitate reliable valve operation in the face of an abrasive sprayed product. Because surfactants are nonvolatile, they increase the diameter of drug particles allowed to evaporate to dryness following aerosolization and may also reduce the evaporation rate of propellant from sprayed droplets (9). High concentrations are best avoided because of their unpleasant taste and tendency to accumulate excessively around spray orifices. Much lower concentrations (0.01–0.1% w/w) are often sufficient to facilitate homogeneous dispersion of suspended drug in propellant, following shaking. Less time consuming, surrogate tests to identify appropriate surfactant concentrations do not always

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correlate well with the observed fine particle fraction following spraying (18). All the surfactants listed above have been shown to exhibit maximum solubilities less than 0.02% in HFC-134a, which has spurred a search for alternatives (19). The patent literature contains numerous references to potentially useful surfactants for use with HFC-134a. These include polyethylene glycol, propoxylated polyethylene glycols, perfluoroalkanoic acids, and numerous others, all of which exhibit enhanced solubility. Byron et al. have suggested that solubility may not be essential to the development of a successfully formulated suspension product if the drug particles are coated with an apparently insoluble surfactant that permits easy resuspension (19). Drug is either dissolved or suspended in a pMDI formulation. The equilibrium size of a sprayed droplet containing dissolved drug depends on the starting droplet size and the concentration and density of the dissolved, nonvolatile ingredients it initially contained. In such a system, it is theoretically possible to alter the drug concentration to achieve a wide range of sizes. If the droplet does not evaporate to dryness, which is likely if it contains a nonvolatile cosolvent, then the evaporation rate becomes the primary determinant of droplet size. In such formulations, it is not necessary to reduce the size of the drug particles prior to incorporation into the formulation because the drug is a molecular homogeneous dispersion in solution. Solution formulations are also easier to manufacture and do not exhibit the drug-sedimentation–related problems described above. Because of the intimate molecular interaction between the dissolved drug and excipients, reduced chemical stability compared to a suspension formulation has been observed. Thus, supersaturation and precipitation at low temperatures must be avoided. Partitioning of drug into valve elastomers has also been noted with a corresponding decrease in dose delivery per actuation (20). The aerosolized drug output from a suspension formulation cannot be smaller than the original particles used to prepare the suspension. Micronization, spray drying, or controlled crystallization is therefore essential to produce appropriately sized particles. Gonda has shown that dilute suspension concentrations in the absence of dissolved excipients should theoretically produce sprayed output of a similar size to the original drug particles (21). As the suspension concentration increases, multiple particle inclusion in the sprayed droplets leads to the production of aerosolized aggregates with a larger size than that of the starting drug. This effect becomes more significant at suspension concentrations greater than 1% v/v and will also be exacerbated by the inclusion of other nonvolatile excipients. Experimental observations of model suspensions show the same trend [(21) and references therein]. In accordance with Stokes’ law, smaller drug particles separate from the bulk suspension more slowly than larger ones (22). This offers a means of reducing the physical instability outlined previously if the suspension cannot be adequately stabilized using surfactants. Suspension formulations can deliver larger doses than solutions because the formulator is not limited by the drug’s maximum solubility in the propellant blend. However, high concentrations are likely to cause valve blockage and may abrade valve elastomers during filling (if the slurry is pumped through the valve) or patient use. Partial solubility of a

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drug in propellant has been associated with an increase in the size of the suspended particles. Dalby et al. have reported a method of conveniently measuring drug solubility in pressurized systems (23), and Phillips et al. (24) have documented its ability to predict crystal growth. As with other inhalation delivery systems, it is inappropriate to evaluate the formulation independently of the “packaging.” The essential components are the container, metering valve, and actuator. Containers for suspensions are typically one-piece aluminum canisters, with a 20-mm external crimp (cut edge or rolled top) (25). Some products (e.g., Azmacort) utilize more attractive epoxy-coated aluminum canisters. Some suppliers recommend anodized aluminum canisters for solution formulations, although most of these products are packaged in plastic-coated glass bottles. While aluminum is inherently opaque, the coating on glass bottles is usually opacified when used with light-sensitive drugs such as epinephrine. The container is typically required to withstand internal pressures up to 180 psig without distortion. The formulation type, labeling, aesthetic requirements, and need for inprocess fill weight monitoring usually dictate the choice of container. Plastic containers (e.g., polyethylene terephthalate, PET) are also available but are not utilized in any marketed inhalation products. Metering valves (Fig. 2) are designed to release a fixed volume of product during each actuation. Assuming that the valve fills with a homogeneous drug solution or suspension, the metered dose is the product of the valve volume and drug concentration. Usual valve volumes range from 25 ␮L to 100 ␮L, although larger volumes are available. Typically, valves contain a prefilled metering chamber that is isolated from the bulk reservoir, as the chamber empties through the valve stem. This is initiated by pressing the stem into the body of the valve. When pressure on the stem is released, an internal spring returns the stem to its rest position, and the metering chamber refills through one or more channels from the reservoir. Such a valve is described as “holding its prime,” indicating that the product that will be sprayed next is already located in the metering chamber prior to actuation and must be retained there if the next dose is to be complete. If drug escapes from the metering chamber during a period of quiescence (perhaps due to propellant drainage or suspension instability), the following actuation will release a smaller dose. Drug escape from the metering chamber of pMDIs during periods of nonuse has been associated with a low subsequent dose of albuterol. For this reason, alternative valve designs are available, although not without their own limitations (26). While drug concentrations present in the canister are altered by temperature-dependent changes in propellant density and propellant leakage, the past history of under- or over-dosing by the valve should also be considered. Valve volumes are approximate due to propellant-induced swelling of the metering chamber elastomers and mechanical distortion during crimping and repeated use. It is, therefore, prudent to assess valve performance with specific formulations and not infer their reliability from successful use with a previous product. All metering valves on inhalation products are designed to operate in the “valve down” or “inverted” position and do not utilize a dip tube associated with continuous valves operated by applying constant pressure to the actuator for

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Figure 2

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Schematic diagram of a metering valve.

the required period of dose delivery (typical of large canisters intended for topical application to skin). This is because it is difficult to redisperse drug particles that sediment or float in a narrow tube. In addition, the valve must contain the pressurized product and retard the ingress of moisture or oxygen. Most valve bodies are constructed from plastics and resins, although metal body valves do exist. An aluminum ferule allows the valve skirt to be crimped on the canister. Return springs are usually stainless steel. Valves contain elastomeric seals in the metering chamber and between the canister and the valve. Elastomer composition is frequently proprietary and sometimes unknown to even the valve manufacturer. Low concentrations of several potentially harmful chemicals are known to be extracted from valve elastomers by propellant systems (4). This has led to the use of elastomer extraction procedures to reduce the concentration of these materials prior to valve assembly and in the development of “cleaner” elastomers. Elastomer performance is critical to the functioning of a valve because in combination with other factors this determines the propellant leakage rate, metering reproducibility, and the speed and reliability of stem return following actuation. Limited elastomer swelling may be considered beneficial because it helps ensure a good seal. However, excessive swelling results in a nonfunctional valve. Thus, the advent of new propellant systems has necessitated the development of new elastomers. The actuator is frequently the most visible part of the pMDI. Its function is to allow actuation of the valve, direct the spray into the patient’s mouth, and provide the orifice through which the metering valve discharges its spray. A well-designed actuator with a separate or integral (to prevent loss or accidental inhalation) dust cap protects the valve stem from damage and keeps it aligned

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with the seat. This prevents the accumulation and subsequent inhalation of dust and provides a place for patient use information and product identity. The spray orifice and valve stem seat are arguably the most critical parts of the actuator. Larger spray orifices are often used in combination with large volume metering valves to spray concentrated suspensions with a reduced likelihood of blockage. Large orifices ensure fast emptying. When small volumes of dilute solutions and suspensions are sprayed, smaller spray orifice diameters may be preferred because they generate smaller droplets (16). To avoid leakage, a tight fit between valve stem and actuator seat is essential. Additionally, it is advisable to minimize sharp turns and blind ends in the path the product follows from the valve stem to the spray orifice to prevent accumulation, and subsequent blockage, by nonvolatile drugs and excipients. With the exception of Azmacort® (originally Rhone-Poulenc Rorer’s product, Collegeville, PA, U.S.A, now Kos Pharmaceuticals, Cranbury, NJ, U.S.A.) and Astra Zeneca’s Breathancer® (Lund, Sweden), all pMDIs are supplied with a molded plastic actuator that positions the patient’s lips very close to the spray orifice (if they use the closed-mouth method of inhaler use). This provides a short distance between the spray orifice and oropharynx and necessitates excellent coordination between actuation of the pMDI and inhalation by the patient if almost complete oropharyngeal deposition is to be avoided. Spacer devices (Fig. 3) were developed to increase this distance, allowing the rapidly advancing aerosol cloud to decelerate before reaching the throat (12). This makes perfect synchronization between actuation and inhalation slightly less important. In addition, spacers allow more time for propellant evaporation, resulting in the formation of smaller droplets or particles and less reflex coughing and exhalation due to local cooling of the throat by impacted, evaporating droplets. A large proportion of drug that would otherwise deposit in the oropharynx is retained in a spacer. This reduces systemic drug levels and minimizes local side effects. A distinction is made between spacers and holding chambers (also called reservoirs). Spacers are essentially hollow tubes through which a patient should have started inhaling prior to actuating the pMDI. They are designed to empty in a single inhalation. Reservoirs are typically larger in

Figure 3

Photograph of an Aerochamber® spacer with a generic pMDI attached.

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diameter, frequently conical or pear-shaped devices, and are designed to permit actuation of the pMDI prior to initiating an inhalation. Their larger size is designed to reduce drug losses on the interior wall of the reservoir due to impaction and sedimentation. Reservoirs usually contain a one-way valve to prevent an inadvertent exhalation from flushing a previously aerosolized dose from the device. Delays between actuation and inhalation, making multiple actuations in the reservoir, and emptying the reservoir over several inhalations, all reduce the efficiency of aerosols delivery to the lungs. Because larger spacers and reservoirs enhance lung delivery compared to smaller ones, collapsible designs are common (e.g., InspirEase®, Schering Corporation, Kenilworth, NJ, U.S.A.). Other designs direct the emerging aerosol spray in the opposite direction to the inhaled airstream in an attempt to increase the flight time while minimizing device size (OptiHaler®, Healthscan Products, Cedar Grove, NJ, U.S.A.). Spacers and reservoirs may also contain flow restrictors to control the patient’s inhalation rate and have mechanisms to coordinate inhalation with pMDI actuation (OptiHaler). Many audibly warn the patient when they are inhaling too fast [Aerosol Cloud Enhancer® (ACE), DHD Diemolding Healthcare Division, Canastota, NY, U.S.A; AeroChamber®, Monoghan Medical Corporation, Plattsburgh, NY, U.S.A; InspirEase]. Newer devices are typically transparent to encourage regular cleaning, and some are designed to fit into ventilator circuits. Baffles located within several small actuators have been shown to yield many of the same advantages (27). Because spacers and reservoirs are often designed to fit multiple pMDIs (which may have significantly different compositions, valves, and actuators), their ability to equally enhance the delivery of all products has been questioned. The Autohaler® (3M Pharmaceuticals, St. Paul, MN, U.S.A.) is a small device that uses a mechanical vane to detect when a patient’s inhalation rate is appropriate for automatically firing the proprietary pMDI it contains. While achieving excellent coordination between inhalation and actuation, it does not produce the other advantages associated with a spacer or reservoir. The pMDIs from other manufacturers cannot be used in the Autohaler. Electronic devices capable of more sophisticated flow monitoring, by programming actuation at different flow rates or points in the breathing cycle, and which can record the history of patient compliance have been developed (28).

III.

Dry Powder Inhalers

DPIs offer a unique opportunity for the delivery of drugs to the lung as aerosols. These devices combine powder technology with device design in order to disperse dry particles as an aerosol in the patient’s inspiratory airflow (29,30). Powders have been insufflated for medical purposes throughout history (31). It is only recently that efforts have been made to establish the dispersion properties of particles and their impact on therapeutic effect (32,33). Progress has been made in recent years as the emphasis has changed from unit-dose systems, employing only the patient’s breath to generate the aerosol, to multiple-dosing reservoir devices that actively impart energy to

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the powder bed to introduce drug particles into the inspiratory airflow. Chapter 16 covers the topic of DPIs in more detail. All DPIs have four basic features: (1) a dose-metering mechanism, (2) an aerosolization mechanism, (3) a deaggregation mechanism, and (4) an adaptor to direct the aerosol into a patient’s mouth. The major components of a dry powder inhaler are the drug powder, and other powdered excipients where necessary, a drug reservoir or premetered individual doses, the body of the device, and a cover to prevent ingress of dust or moisture. To introduce drug particles into the lung, they must be ⬍5 ␮m in aerodynamic diameter (34–36). This is generally achieved by milling the powder prior to formulation (37,38). In recent years, there has also been some interest in spray drying powders to achieve the same end (39–43). Small particles are notoriously difficult to disperse (44,45). The forces governing dispersion are well documented and consist mainly of electrostatic, Van der Waals, and capillary forces (46,47). Knowing that these forces exist has not facilitated aerosol generation to any great extent. One approach that has been taken to improve the dispersion of dry powders is the inclusion of an excipient (48,49), notably lactose (50,51). The lactose particles are intended to act as carrier particles for the drug and as such are in a much larger size range, 60–80 ␮m (52–54). Drug particles are theoretically stripped from the surface of the lactose particles to which they are loosely attached during the generation process (55,56). This process is illustrated schematically in Figure 4. Thus, the drug particles are dispersed and can traverse the upper respiratory tract while the excipient particles do not pass beyond the mouthpiece of the device or the mouth and throat of the patient. In the devices that have long been approved for use in the United States, the Spinhaler® (Fisons, Rochester, NY, U.S.A.) and Rotahaler® (GSK, Inc., RTP, NC, U.S.A.) unit doses of Intal® (disodium cromoglycate) and Ventolin®

Schematic diagram of the stripping of drug particles from carrier particles on the inhaled airstream.

Figure 4

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(albuterol sulfate), respectively, are packaged in hard gelatin capsules. The dose of powder itself is delivered from the gelatin capsule by different mechanisms. The Spinhaler has a mechanism for piercing the capsule (57). The cap of the capsule fits into an impeller that rotates as the patient breathes through the device projecting particles into the airstream, as illustrated in Figure 5. The Rotahaler, shown in Figure 6, has a mechanism for breaking the capsule in two pieces. The capsule body containing the dose falls into the device while the cap is retained in the entry port for subsequent disposal. As the patient inhales, the portion of the capsule containing the drug experiences erratic motion in the airstream, causing dislodged particles to be entrained and subsequently inhaled. A number of other devices have been approved for use in other countries. The Inhalator® (Boehringer Ingelheim, Ridgefield, CT, U.S.A.) has a mechanism for piercing the ends of a hard gelatin capsule containing the dose of fenoterol (58), as shown in Figure 7. The inspiratory flow then passes through the capsule. The Diskhaler® (GlaxoSmithKline, Inc., Research Triangle Park, NC, U.S.A.), shown in Figure 8, employs packaging consisting of individual doses of albuterol sulfate in blister packs on a disk cassette. Following piercing, inspiratory flow through the packaging depression containing the drug induces dispersion of the powder. One of the more sophisticated systems approved for use in Europe is the Turbuhaler® (Astra Zeneca Lund, Sweden) shown in Figure 9 (59). This device employs a multidose reservoir of terbutaline sulfate. The dose is metered into small conical cavities by twisting a grip at the base of the device. When the patient inhales, air ducted through the cavities dislodges a dose of drug. In addition to the devices mentioned above, many others patented for use are described in the literature (although none have so far received regulatory approval in the United States). However, the novelty of these systems is typically associated with their mechanism for aerosolizing the powder rather than the way a unit dose is packaged or separated from a bulk gelatin cap reservoir. Reproducible dose metering remains the most difficult challenge in device design. For this reason, prefilled doses in gelatin capsules or multiple

Figure 5

Diagram indicating the essential components of a Spinhaler®.

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Figure 6

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Diagram indicating the essential components of a Rotahaler®.

depression blister packages or extraction of a specified volume of free flowing powder from a multidose reservoir are likely to remain common. The volumetric metering of powder poses some unique problems because powder flow and dispersion behavior is complex (60,61). There are many factors related to the powder itself that influence the clinical effect of the product. Particle size is of primary importance in defining the location of lung deposition of aerosols and hence their effectiveness (62). Metering and dispersion characteristics are affected not only by particle size, but also by rugosity (15), shape (63,64), moisture content (65), surface chemical composition (66), and charge (57,67). Reported exceptions to the above metering mechanisms are the Easyhaler which has a rachet metering system to remove discrete doses from a powder bed (68); a rotary planar device, which has a compact powder bed from which a helical blade removes a dose of drug (69,70). A device that involves dislodging powder from a specific length of dimpled tape (71) has also been suggested. To protect the drug reservoir from aggregation or chemical degradation, a desiccant is incorporated into most devices that contain bulk drug rather than individually protected doses.

Figure 7

Diagram indicating the essential components of an Inhalator®.

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Figure 8

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Diagram indicating the essential components of a Diskhaler®.

Safeguards are necessary to ensure that the desiccant cannot be inhaled. All dry powders require a mechanism to disperse the drug as an aerosol. For example, the Bernoulli or Venturi effect has received some attention (72) as has compressed gas propulsion (73). Large, excipient particles (typically lactose), to which drug is loosely bound, are used to aid dispersion. Their improved flow properties compared to cohesive powder beds of smaller particles allow easier separation of individual particles into the airstream. Most devices have additional mechanical methods of introducing energy into the powder bed to facilitate dispersion. The Spinhaler has an impeller that rotates to aid dispersion. Clearly, the rotation speed will vary according to patient effort (airflow rate) and through the breathing cycle (74–76). The Rotahaler to some extent employs rotation as a means of dispersion because the capsule containing drug rotates, albeit erratically, in the device (77). The Inhalator is based on the principle that a large volume of air will pass through the powder bed by inducing a large pressure drop across the capsule. This pressure drop is

Figure 9

Diagram indicating the essential components of a Turbuhaler®.

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Diagram indicating the essential components of a Handihaler®.

brought about by the patient’s inspiratory flow. Consequently, the patient’s inspiratory effort and inspiratory duration may influence aerosol generation (78). The Diskhaler (79,80) and Turbuhaler (81–85) also employ a pressure drop to introduce drug particles into the patient’s inspiratory flow. Novel approaches have been suggested to put energy into the powder bed for dispersion. Increased dose delivery has been achieved by tapping a dimple in the packaging material or by using compressed gases to assist in the generation of powders (86). Two other methods of note involve the continuous input of energy. One method employs a small motor and impeller to disperse the powder (87). A second method utilizes a gas-assisted approach (88). The mouthpiece can simply be a tube through which the patient inhales, as is the case with the majority of DPIs. The Turbuhaler, however, has a series of tortuous channels through which the aerosol passes before entering the mouth of the patient. These channels offer the advantage that poorly dispersed particles will either impact in transit and become re-entrained as smaller particles, or be

Figure 11

Diagram indicating the essential components of an Aeolizer®.

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permanently removed from the airstream. Provided dose uniformity is maintained, either of these options is desirable. Impaction followed by re-entrainment helps ensure the correct particle size for lung deposition, while permanent impaction will remove particles destined to deposit in the mouth or throat of the patient. The induction of turbulent flow in narrow tubes is also associated with enhanced deaggregation. However, narrow tubes may produce a high-resistance device through which patients may find it difficult to breathe. It has been demonstrated that the materials that the drug powder comes into contact with during processing have a significant impact on its electrostatic properties (89). Indeed, these acquired properties may influence their subsequent interaction with construction materials present in the device. In extreme cases, drug powders may fail to leave the device. With a growing acceptability of refillable devices, the appearance of metal devices or those composed of conductive plastics would not be surprising. The internal geometry of the device is of great importance to the generation of the aerosol. The dimensions of channels through which the inspired airflow passes dictate the pressure drop across the device. Accumulation of drug particles is often associated with severe changes in the direction of airflow flow. In addition, device design determines whether the dose will empty as a bolus or as a continuous stream of powder over the course of an entire inhalation. This has implications both for the ultimate location of drug deposition within the lung and for defining the flow rate at which the device should be tested. Most of the devices have to some extent mimicked the external appearance of pressurized pMDIs. The size and orientation in use of pMDIs and many DPIs are very similar. This reflects the popularity of pMDIs and the recognition that any viable DPI must remain small and simple to operate. Because DPIs employ a variety of operating mechanisms, standard protocols for their use have not been developed as in the case of pMDIs (90). For this reason, patients must be taught the correct operation of each device. DPIs are typically comprised of multiple components, which are machine- or handassembled. It is vital to ensure that components are designed to facilitate correct assembly (both at the production facility and following cleaning by a patient) to avoid either failure to deliver a dose or accidental inhalation of part of the device (91). Additionally, in excipient-free devices, a patient may be unaware of whether he has in fact inhaled a dose or not. This has led to the incorporation of dose counters on newer reservoir devices. In most cases, the breathing pattern of the patient defines the operating conditions of the DPI. pMDIs (which contain high vapor pressure propellants) and nebulizers (which are electrically or pneumatically driven) utilize patientindependent energy sources to generate an aerosol. In contrast, the effectiveness of a DPI is much more susceptible to the vagaries of age, gender, disease, and breathing cycle of the device user. The breathing cycle has been the subject of several studies in the recent literature (75,76,92,93). The inspiratory flow during moderate effort can be maintained constant throughout the majority of the inhalation cycle. In contrast, at maximum effort, constant flow is not achieved throughout the inhalation cycle.

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There has been significant discussion regarding appropriate volumetric flow rates at which to assess DPIs for realistic demonstrations of their potential clinical performance. It seems clear that more than one flow rate must be adopted to evaluate performance adequately (94–96). The flow rate should be selected based on the internal resistance of the device. In this way, a device will be tested at a flow rate that a patient can reasonably expect to achieve through the device during normal use. Indeed, the Food and Drug Administration have suggested that flow test should be limited in total volume sampled (2 L) to simulate in vivo conditions (7). The recent approval of Exubera® (Pfizer, Groton, CT, U.S.A.), an inhaled insulin product, has extended approved DPI to include an active dispersion mechanism, and systemic drug delivery via the lungs. It may be anticipated that such novel products will lead to review of pharmacopeial and governmental standards.

IV.

Nebulizers

The CFC content of pMDIs and perceived technical difficulties associated with the development of DPIs have led to a resurgence of interest in nebulizers. However, nebulizers and their associated respiratory solutions have been used in inhalation therapy since the early 19th century (97). Although early nebulizers were considered cumbersome and unreliable compared to pMDIs and DPIs (98), recent innovations have changed this perception. The popularity of nebulizer use is currently increasing due to their ability to generate small droplets capable of penetrating deeply into the lung, their high dose-delivery capacity, miniaturization of hardware, and the development of high-output nebulizers, permitting shorter treatment times and increased efficiency of drug delivery to the patient (99–101). Coordination between aerosol generation and breathing, which is required for successful pMDI use, is not essential for nebulizers, making them useful for treating hospitalized, very young, and elderly patients (102,103). Nebulizers also represent ideal delivery systems for biotechnologically derived proteins and peptides because aqueous solutions suitable for atomization are frequently much easier to formulate than either pMDI or DPI systems. This is exemplified by the recent introduction of rhDNase (Pulmozyme®, Genentech, San Francisco, CA, U.S.A.) in nebulized form and the ongoing development of nebulizer delivery systems for insulin, factor IX, and a host of other macromolecules for systemic and local delivery (104–106). Additionally, nebulization can be used to deliver more complex dispersed systems such as liposomes (107) or microspheres (108). These advantages are tempered by the larger size, more obtrusive dosing, constant cleaning requirements, and unit-dose nature of nebulizers. Perhaps, the most important limitation to their expanded use is the fact that nebulization therapy usually requires a nebulizing device, respiratory solution or suspension, auxiliary tubing, and mouthpieces or face masks. These components are often assembled and used arbitrarily, making it almost impossible to specify the actual dose of drug inhaled by the patient. This leads not only to interdevice variability (109) but also to a degree of variability within the same device design (110).

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Marketed respiratory solutions are generally composed of drug dissolved in aqueous, isotonic solvent systems that may contain preservatives to reduce microbial growth. However, in some cases, the presence of preservatives such as sodium metabisulfite (111), benzalkonium chloride, and ethylene diamine tetraacetic acid (EDTA) (112) has caused coughing and bronchoconstriction. Thus, some “preservative-free” products are on the market. Isotonicity is frequently achieved using sodium chloride. Other excipients are minimized and typically limited to pH adjusters and antioxidants. Respiratory suspensions are not currently marketed in the United States, but suspensions such as budesonide are found commercially in Europe (113). The density, surface tension, and viscosity of the respiratory solution or suspension, in combination with the nebulizer design and its operating conditions, determine the atomization characteristics of the system. Respiratory solutions are supplied in multidose bottles or unit-dose plastic, or glass ampules. In response to specific patient needs, dilution with either sterile water or normal saline may be required. There are two common types of nebulizer, air-jet and ultrasonic, with numerous commercially available examples of each type. The operation of an air-jet nebulizer requires an external gas supply (usually compressed air or oxygen). This gas supply is the driving force for liquid atomization. Figure 12 shows the typical components of the air-jet nebulizing device. Respiratory solution or suspension is placed into the reservoir. Minimization of the “dead volume” (solution that will not exit the nebulizer) is of critical importance since dose volumes of 1–3 mL are typical, while dead volumes of 0.5–2 mL or more have been reported (114,115). This accounts for the conical design of many reservoirs. Compressed gas is forced through the jet, causing a region of negative pressure to develop because the gas is

Schematic of a typical air-jet nebulizer. A, “t” piece; B, dilution or make-up air (x L/min); C, baffle; D, reservoir; E, recirculating droplets; F, respiratory solution or suspension; G, compressed gas (y L/min).

Figure 12

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expelled at high velocity through the spray orifice. This low pressure, aided by capillary action, causes the solution to travel up channels parallel to the gas supply tube where it becomes entrained in the gas stream and is sheared into a liquid film. The film is unstable and collapses under surface tension to form droplets. The larger droplets (15–500 ␮m in diameter) return to the reservoir to be recirculated. This is achieved by placing a baffle or the reservoir wall in the path of the air–liquid stream to induce droplet impaction and coalescence. This accounts for the majority of nebulized solution. It is important to minimize the internal surface area of the nebulizing chamber because a large percentage of the fill volume can adhere to the wall and be unavailable for recirculation. The smaller droplets (1–10 ␮m in diameter) that escape impaction on the baffle leave the nebulizer through a connecting “t” piece carried by the nebulizing airflow out of the nebulizer (116,117). The separation between the spray nozzle and the baffle, and the droplet exit velocity from the nozzle, is of primary importance in defining the size of droplets that escape impaction. The “t” piece allows for the introduction of dilution or “make-up” air and provides a place to attach auxiliary tubing, face masks, or mouthpieces, which aid in diverting the aerosol to the patient. The use of dry compressed nebulizing gases and make-up air accelerates droplet evaporation, aiding in reducing the droplet size. In some nebulizers, a large evaporation chamber further enhances droplet drying. Evaporation also extracts heat from the recirculating droplets, causing solution remaining in the reservoir to cool (a drop of approximately 5–9°C from the initial temperature). Solvent losses due to evaporation cause solution remaining in the reservoir to concentrate (118,119). This increase in concentration can potentially lead to drug or excipient crystallization within the reservoir, which could theoretically clog the nebulizing air jets and reduce the total aerosolized output of drug. In most situations, crystallization does not occur because reservoir fill volumes are small and rapidly exhausted, and drugs are typically hydrophilic and exhibit high solubilities (120). However, it has been reported that increases in solution concentration may change the total output and output rate of aerosolized drugs, as well as influence the equilibrium droplet size of the aerosol (119,121). The major controllable operating variable associated with air-jet nebulizers is the airflow rate through the nebulizer, which is critically dependent on the driving pressure, except when the spray orifice acts as a critical orifice. Because most nebulizers are operated from small compressors that do not permit selection of specific driving pressures or flow rates, the nebulization conditions are usually not well defined. Additionally, many marketed compressors for domiciliary use are incapable of attaining adequate pressures or flow rates to produce an aerosol from demanding air-jet nebulizers (122). These problems are compounded by a lack of manufacturer-supplied specific operating conditions for most nebulizers and have led to the recent trend of approving respiratory solutions with specific nebulizers/compressor combinations, as in the case of rhDNase. Other variables, such as relative humidity, ambient temperature, and inspiratory flow rate, are known to influence nebulizer performance, but are difficult to standardize under patient use conditions.

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Air-jet systems are further subdivided into “single-use” and “multiuse” categories. Most marketed nebulizers fall into the single-use category, designed to be used once and then discarded, and as such are molded from inexpensive plastics at high speed. However, to minimize costs, patients typically reuse these devices even though it is unclear to what extent a nebulizer can be considered reliable if reused (123,124). One study reported that patients reuse a disposable air-jet nebulizer in excess of 100 times before obtaining a new one (125). Recently, nebulizer manufacturers have introduced multiuse nebulizers, which are designed for repeated use and are easy to clean and assemble reproducibly. This is significant because cleaning regimens, manufacturer specified or otherwise, and alterations in baffle or capillary feed tube position within the nebulizer have been reported to significantly alter nebulizer output characteristics (126–128). Other innovations include the development of dose-sparing mechanisms such as interrupters, which manually or automatically block airflow to the nebulizer, and therefore aerosol generation, during patient exhalation. This increases the total inhaled dose received by a patient. Other designs divert the inspiratory airflow of the patient through the nebulizer, to enhance the volume of air available to entrain aerosolized droplets, or use valves in the mouthpiece to prevent exhaled air from passing through the nebulizer and carrying the aerosol to waste. Yet another modification collects aerosol generated during exhalation in a rubber balloon. Droplets in the balloon augment newly generated droplets during the next inhalation. While in most nebulizer systems droplets are unlikely to evaporate to dryness prior to inhalation, systems employing dilution air and producing small droplets may have their equilibrium size limited by the concentration of solutes in the initial respiratory solution, rather than the evaporation rate of the droplets. Air-jet nebulizers are known to be capable of successfully delivering respiratory suspensions (129). Unlike the air-jet devices, an external gas supply is not required for operation of electrically or battery-driven ultrasonic nebulizers. Figure 13 shows the typical components of an ultrasonic nebulizing device. Respiratory solution or suspension is atomized by means of a piezoelectric crystal transducer. An alternating current causes the shape of the crystal to alternately shrink and expand, causing a vibration that is amplified by a stainless-steel shim. This vibration is then transferred to the solution or suspension in the nebulizer reservoir. Two mechanisms have been proposed to explain how these oscillatory waves traveling through the bulk liquid generate an aerosol cloud. The first mechanism (Fig. 14A) proposes that the crystal transducer, operating at a low frequency, vibrates the bulk liquid causing the formation of cavitation bubbles. As the air bubbles, which have a short life span, rise toward the air–liquid interface, the internal pressure of the cavitation bubbles equilibrate with the atmosphere, causing the bubbles to implode. As the bubble bursts at the air–liquid surface, portions of the bulk liquid break free from the turbulent bulk liquid and form droplets. The second mechanism (Fig. 14B) involves highfrequency crystal vibrations, causing formation of capillary waves in the bulk liquid. The capillary waves constructively interfere to form peaks and a central geyser. The geyser increases in amplitude until it becomes unstable under the

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Schematic of a typical ultrasonic nebulizer. A, Face mask or mouthpiece; B, baffles; C, geyser of respiratory solution or suspension; D, piezoelectric crystal; E, internal fan; F, battery or electric source.

Figure 13

influence of continuing excitatory vibrations, and portions of the peak break loose from the bulk liquid and are ejected into the air as droplets (130,131). Regardless of the mechanism, once liquid breaks free from the bulk liquid and forms a droplet cloud, larger droplets either impact on baffles or return to the reservoir surface under the influence of gravity, to be recirculated. Smaller droplets leave the nebulizer aided by an internal fan. Fan speed and adjustable flow restrictors can alter the air velocity over the reservoir surface, influencing both the droplet size and the aerosol output rate. The reduced volume of dilution air compared to that passing through an air-jet nebulizer results in a more concentrated aerosol cloud with less evaporative capacity. What evaporative cooling does occur is more than offset by heat conducted to the reservoir from the vibratory system. This heat results from frictional forces induced by movement of the transducing crystal. The increased reservoir temperature has not been associated with altered aerosolization characteristics over short nebulization intervals (132). However, heating may be detrimental to thermolabile formulations of proteins and cause odor changes in antibiotic solutions (133). Excessive heating can be reduced by the presence of substantive heat sinks and sensory feedback systems that reduce the oscillation frequency in response to higher temperatures. Such systems can automatically turn off nebulizers that have emptied their reservoirs. Other systems under development include multiple-dose nebulizers that atomize solution from a self-contained reservoir for inhalation in a single breath. This can be achieved by dropping drug solution onto the surface of a vibrating platform or through a vibrating grid. Ultrasonic nebulizers do not reliably incorporate suspended drug particles into the aerosolized droplets they produce and may therefore be considered unsuitable devices for aerosolizing respiratory suspensions, microspheres, and other

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Figure 14 Proposed ultrasonic nebulizer aerosolization mechanisms. (A) Cavitation bubble formation at low frequency; (B) capillary wave formation at high frequency.

colloidal systems (134). Breath-controlled delivery of nebulized aerosols is a recent development. The two prominent examples of this approach are the Halolite® (Profile Therapeutics, West Sussex, U.K.) device that monitors breathing cycles and delivers the aerosol on the first portion of the breath (135) and Akita® (InaMed, Gauting, Germany) that controls patient breathing to enhance lung deposition (136). Unlike pMDIs and DPIs, there are no established or proposed USP functionality standards to characterize the nebulized output of a respiratory solution or suspension used in combination with a nebulizer. By analogy to testing prescribed for other inhalation dosage forms, it seems appropriate to develop methods and agree on specifications for nebulizers. It is obviously vital to know the dead volume, quantity of aerosol emitted from the nebulizer, the duration of aerosol generation, and to have an indication of the aerosol size or fine droplet fraction. Standardization of the methods necessary to perform these tests would represent a substantial step in the future utilization and development of nebulizers.

V.

Selection of a Delivery System

A number of factors impact on the decision to develop a drug in combination with a specific inhalation delivery system. In some situations, a case can be made for delivering the drug using any of the previously described systems, based on marketing preference or regulatory climate. However, in other cases, one delivery system offers clear advantages over the others. When the site of action is known for a particular agent, the optimal delivery system to reach that location should be used. pMDIs and DPIs are typically incapable of depositing a large percentage of their emitted dose in

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the peripheral airways and are therefore unsuitable for delivery of drugs, such as ribavirin, which is active against viral infections located in the terminal airways. To enhance peripheral deposition, the Small Particle Aerosol Generator® (SPAG-2, ICN Pharmaceuticals, Costa Mesa, CA, U.S.A.) uses a highly baffled air-jet nebulizer, coupled with additional drying air, to produce small particles or droplets of approximately 1.2 ␮m. Nebulizers producing small droplets may also be considered well suited to delivering molecules intended for systemic activity because better absorption is likely to occur from the terminal airways in which smaller aerosols tend to deposit. However, Inhale Therapeutic Systems (Palo Alto, CA, U.S.A.) is known to be developing a dry powder inhaler intended for the delivery of insulin and other high-molecularweight proteins that act systemically, so in the future, deep lung targets might become more accessible to other types of delivery systems. Conversely, pMDIs and DPIs are ideal for delivering drugs such as bronchodilators, to the upper airways, because this is the main site of airway resistance in asthmatics and where the receptors are located. In addition, pMDIs and DPIs provide drug as a concentrated bolus in a single breath, providing the rapid bronchodilation needed to relieve a severe asthma attack. In comparison, the SPAG-2 may take up to 18 hours to deliver its drug, while a more conventional nebulizer would require 5–10 minutes (137). Biotechnologically derived drugs, which can be stabilized in an aqueous solution, but lose their activity on drying or in the presence of organic solvents, are probably most suited to delivery by a nebulizer. Because aqueous solutions are often easier to formulate, this route to commercialization may also be expedient. Conversely, drugs susceptible to hydrolysis are more suitable candidates for incorporation into pMDIs and DPIs. pMDIs also provide protection from oxidation. Drugs with low aqueous solubilities are good candidates for delivery by DPIs. However, they must be readily convertible into small, discrete particles by micronization, spray drying, or controlled crystallization, and ideally be relatively nonhygroscopic. This is also true of drugs that are insoluble in propellant blends, if one is to manufacture a suspension-type pMDI. Drugs that are soluble in propellant, with or without the aid of a cosolvent, can potentially be formulated into solution pMDIs. Chemically unstable compounds, or ones likely to react with or partition into packaging materials, would not be good choices for such a system. While these observations may serve as a useful guide, they are no substitute for a comprehensive preformulation evaluation of a potentially useful inhalation drug candidate.

VI.

Future Developments

The systems described above represent what is achievable today, or in the near future. It is interesting to speculate on what possibilities will arise in the long term. Next-generation devices will probably include features that indicate how many doses remain to be inhaled or when a dose is due. They could easily measure and record the flow rate and inspired volume under which doses were inhaled, as an aid to monitoring disease progression or patient compliance.

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Devices that will actuate a pressurized pMDI at a specific point in the breathing cycle if preset parameters are achieved already exist. Single-use DPIs are under development. Beyond these innovations, new methods of aerosol generation are likely to be developed or existing ones miniaturized to clinically useful sizes. One can easily envisage hybrids of pMDIs and ultrasonic nebulizers, in which large propellant droplets are broken into more respirable sizes by impaction on a vibrating transducer or grid, rather than simply being wasted within the actuator. pMDI valve stems could be transiently heated immediately before actuation to expel droplets at higher pressures and facilitate faster evaporation, without causing drug degradation. Manual pumps capable of generating respirable sprays from aqueous solutions are closer to reality every year. Dry powders could be deaggregated by vibrational elements, rather than relying on turbulent airstreams, or generated electrostatically. Inhalers employing environmentally acceptable supercritical solvents to dissolve and atomize drugs at enormous pressures have already been described in the patent literature (138). These opportunities, combined with the increasingly precise identification of pulmonary target regions, the prevalence of respiratory disease, and a continuing need for a viable delivery system for potent biotechnology-derived drugs are likely to drive development of inhalation systems as the first century of the new millennium begins. References 1. 2. 3. 4. 5. 6. 7. 8.

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16 Next Generation Dry Powder Inhalation Delivery Systems

ANTHONY J. HICKEY

TIMOTHY M. CROWDER

School of Pharmacy and Medicine, University of North Carolina, Chapel Hill, North Carolina, U.S.A.

Oriel Therapeutics, Inc., Research Triangle Park, North Carolina, U.S.A.

I.

Introduction

A.

Historical Perspective

Dry powder inhalers have been through a number of periods of development over the last half century. The first period of the 1960s and 1970s saw early passive inhalation devices, which lacked sophistication and performed poorly, using the patients’ inspiratory flow to disperse the drug, as alternatives to the then dominant propellant-driven metered dose inhaler (MDI) products. In the 1980s and 1990s, sophisticated passive inhalers and active systems, the latter using applied energy for dispersion of drug, were developed. These products were unique in that they were, in some cases, first market entry products for asthma compounds or were developed for the new market for systemic delivery of drugs. Early indications, from the recent introduction of combination products, are that dry powder inhalers may supersede MDIs and potentially dominate the inhaled drug market. Since the beginning of the millennium, new strategies have emerged to improve dry powder inhaler performance and it is anticipated that the next generation of devices will be suitable for the treatment of a variety of diseases in a technologically efficient and clinically effective manner. B.

Lungs as a Route of Administration

The lungs are the site of gas exchange and a sophisticated organ for the filtration of air. Their use as a route of drug administration or site of delivery 445

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requires an understanding of mechanisms of deposition of particles in the lungs and site of absorption or action. The four major mechanisms of deposition in the lungs are inertial impaction, interception, sedimentation, and diffusion. Because drug particles are often uniformly shaped with symmetry in many planes and rarely have large masses of submicron-sized particles, they are subject to the dominant effects of inertial impaction and sedimentation. Interception is important for the deposition of elongated particles, or fibers, which are rarely included in pharmaceutical products. In fact, there are pharmacopeial specifications limiting particles having dimensions greater than 10 ␮m (1). The small mass of submicron particles may be subject to diffusional effects, which are modest in the context of the behavior of a majority of the aerosol. The recent interest in nanotechnology and in particular, nanoparticles may lead to erroneous conclusions regarding aerosol delivery. Nanoparticles dispersed as dry powders form stable micron-sized aggregates that behave aerodynamically according to the aggregate size. At best, these particles will deaggregate on deposition, which has implications for the biology of their disposition. Two blood supplies serve the lungs, one to the airways and a second to the periphery (2). The barrier to entry to the airways blood supply is a small surface area, ciliated columnar epithelium with periciliary aqueous and mucus layer. The barriers to entry to the peripheral blood supply are planar–alveolar epithelial cells alone and a very large surface area. It is postulated that the periphery of the lung will give rapid access to the circulation for systemic delivery. Local action is achieved by identifying the target receptors or cells. ␤-adrenergic receptors are known to exist in the largest quantities in the periphery of the lungs. Cholinergic receptors are located in the central airways. Pseudomonas aeruginosa lives as a plaque on the sticky mucus of cystic fibrosis (CF) patients and therefore ciliated airways are a target for antibiotics. Some respiratory infectious microorganisms reside in the alveolar macrophage and the periphery of the lungs would be the target for therapy. C.

Target Disease States

There are a number of pulmonary diseases that have been targeted for aerosol therapy including asthma, chronic obstructive pulmonary disease (COPD), emphysema, CF, pulmonary hypertension, and cancer. In addition, a number of systemic diseases have been targets including: Pain management, migraine, diabetes and osteoporosis. Each of these diseases represents a unique opportunity for local targeting, unique bioavailability, etc. The differential diagnosis of asthma and COPD can be difficult. While COPD is more likely to be associated with smoking and the elderly than asthma, some of the important symptoms of disease may be similar. However, asthma is considered a reversible airways disease, while COPD is not. Emphysema is characterized by an imbalance in the enzyme elastase, which results in destruction of peripheral airways and alveoli. CF is a systemic disease. Its effects in the lungs relate to chloride-ion transport. A result of this imbalance is sticky, immobile mucus that clogs the airways and is an ideal medium for the

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growth of the microorganism, Pseudomonas aeruginosa. Consequently, CF patients have difficulty breathing, clearing mucus, and have a chronic lung infection. Pulmonary hypertension involves vasoconstriction in the lungs resulting in high blood pressure. This disease is difficult to treat as systemically acting hypertensives can induce hypotension, which would be potentially life threatening. There are a number of local cancers that occur in the airways. Because the airways can be accessed directly using aerosols these cancers are obvious targets for aerosol therapy. Examples of relevant targets include diffuse adenocarcinoma, an insidious disease that is not amenable to chemotherapy and requires surgical resection. D.

Therapeutic Agents

A range of therapeutic agents is available to treat these diseases and more are under development. Table 1 lists a range of drugs available for the treatment of the diseases noted. While the lungs have seen an increase in their potential as a route of administration, it is evident that a very small number of agents are available for delivery by this route in comparison with others. E.

Characteristics of Dry Powder Inhalers

Dry powder inhalers are characterized by three key components: A powder formulation, a metering system and an aerosol dispersion mechanism. Dry powder formulations may contain drug alone, following a unique manufacturing approach or in a blend with a carrier particle, usually lactose (25). As new products are being considered, greater attention is being paid to matrix particles with multiple components and additional substances, notably sugars, which can be used as carrier particles (26).

Table 1

Drugs Delivered as Aerosols for the Treatment of a Variety of Diseases

Disease Asthma

Chronic obstructive pulmonary disease Emphysema Cystic fibrosis

Cancer Diabetes Osteoporosis Pain management

Drug ␤-Adrenergic agonists Anticholinergics Corticosteroids Tiotropium Salmeterol Fluticasone ␣-1-Antitrypsin Amiloride UTP analog Tobramycin DNAse Doxorubicins (Adriamycin) Insulin Calcitonin Parathyroid hormone Morphine Fentanyl

References (3,4) (5,6) (7) (8–10) (9,10) (11) (12) (13) (14,15) (16) (17,18) (19,20) (21) (22) (23) (24)

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Metering systems fall into three categories: Unit dose (capsules), multiunit dose (blisters and strips), and reservoirs. Each of these offers advantages and limitations and corresponds to a level of sophistication of the devices they serve. The advantage of unit- and multiunit-dose devices is the isolation of each dose establishing uniformity and facilitating control of storage stability. However, there may be patient compliance issues especially regarding unitdose systems because multiple steps are involved administering the aerosol. The converse is true of reservoir systems. Improved patient compliance may be achieved as fewer steps are involved in dose delivery. However, dosing uniformity and storage stability are difficult to achieve when sampling from a bulk powder. Aerosol dispersion mechanisms are based on the fundamentals of fineparticle physics. Few methods exist for imparting energy to powder beds to aid in the dispersion of particles into the atmosphere. These include pneumatic, mechanical (impact), and vibrational. Section 4.2 illustrates the range of methods that employ these principles. A variety of inhalers have appeared that utilize these mechanisms and all, in one form or another, are under development. II.

Formulation

A.

Desirable Physicochemical Properties (Stokes’ Law)

In the relatively short history of inhaled dry powder drug delivery, certain key properties have been identified as critical to successful aerosol delivery. Particle size and distribution are important because of their role in lung deposition and regional targeting. Particle morphology is important for reproducible dispersion and lung deposition. Recently, the influence of density on particle dispersion and aerodynamic behavior has also been identified as significant. Density impacts on van der Waals forces and on aerodynamic size according to Stokes’ Law. Stokes’ law for terminal settling velocity takes the form (27):  g ρC (De )  2  g ρ0C (Da )  2 VT =   De =  18 η  Da  18 ηκ   

VT ⫽ terminal settling velocity ␩ ⫽ air viscosity ␳ ⫽ particle density De ⫽ equivalent volume C (De), C (Da) ⫽ slip correction factors

g ⫽ acceleration due to gravity ␬ ⫽ dynamic shape factor ␳0 ⫽ unit density Da ⫽ aerodynamic diameter

It is also important that the particles exist as a stable polymorphic form of the drug. Susceptibility to moisture uptake, hygroscopicity, is an undesirable property that frequently leads to chemical and physical instability. Figure 1 shows particle morphology, which reduces density by introducing intraparticulate space. These forms of particles exhibit large geometric size but small aerodynamic diameters.

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Different types of particles: (A) hollow particles; (B) nanoparticle aggregates; (C) low-density highly porous particles. Source: Adapted from Refs. 28–30.

Figure 1

B.

Particle Manufacture

A number of techniques for particle manufacture have been evaluated for aerosol delivery. The dominant method for producing commercially available products is air jet, attrition, or milling. This process requires large particles to impinge on each other at high velocity (pressure), thereby reducing their size as they are broken. This technique often results in surface amorphous content for crystalline materials. The large surface area per unit mass results in large electrostatic charge and potential for moisture uptake. Each of these items is a source of instability for the product and must be controlled closely. Spray drying has become a popular alternative to jet milling (Fig. 2). It is another destructive method in that jets of liquid containing the drug are atomized creating droplets with a large surface area to mass ratio, specific surface area. On drying, these droplets form small solid particles, which, depending on the nature of the drug, additives, and processing conditions, may be different sizes and structure (crystalline vs. amorphous). Spray drying has been viewed as a logical alternative to jet milling for the production of protein and peptide particles and has found a unique application in the production of low-density particles. Supercritical fluid manufacture of particles has been evaluated for the production of respirable particles. This technique will allow the preparation of particles with well-defined morphological and size characteristics. This method has been described in detail elsewhere (31).

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Figure 2

C.

Common methods of particle manufacture.

Excipient Selection and Properties

There are a number of excipients available for use in inhalers but those suitable for dry powder delivery are limited (Table 2). Excipients may serve different purposes in the product. Excipients might be used as diluents, dispersing agents, matrices, formulation aids, stabilizers, and bioavailability modifiers. Lactose, the most long-standing additive to dry-powder formulations, serves several purposes. As a diluent it allows metering of small doses of drug in manageable quantities of powder. The blend with lactose also reduces drugparticle aggregation and during the inhalation cycle helps fluidize the powder. Table 2

Excipients

Excipients Sugar Lactose Trehalose Mannitol Sucrose Glucose Maltodextrin Cyclodextrin Surfactant Phospholipid Oleic acid Magnesium stearate Protein and peptides Human serum albumin Poly(leucine) Poly(lysine)

References (34,35) (34,36) (34,37) (36,38) (39) (40,41) (42) (35,43) (44) (45) (34) (46) (47)

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It should be noted that a proportion of the drug remains attracted to the large lactose particles upon aerosol generation and contribute to residual inefficiency of most devices. As materials of biological origin proteins, peptides, and nucleic acids are considered as therapeutic agents. Sophisticated formulation approaches are required to maintain their stability and hence their activity. Unlike small molecular weight molecules, these macromolecules are often more stable in the amorphous state. Additional sugars may be considered in this context such as mannitol or trehalose (32,33). Other formulation aids such as surfactant may be used to improve the dispersion properties of the particles or to enhance the stability of the product. Finally, additives that modulate the dissolution properties of the drug may be employed to influence bioavailability, e.g., magnesium stearate (48).

III. A.

Metering Systems Unit Dose (Capsules and Blisters)

The original unit-dose metering system was a direct derivative from orally ingested products intended for GI delivery, the capsule. The majority of capsules are composed of gelatin. Gelatin is a protein derived from the bones and skin of cattle or pigs by an elaborate process involving acid and alkali extraction (49). Capsules occur in a number of sizes each intended to deliver different quantities of drug. The #3 capsule is the most frequently employed for aerosol product and contains up to tens of milligrams of material. A disadvantage of gelatin as a capsule material is its propensity for water association. Five to fifteen percent of the mass of gelatin is attributable to water at equilibrium with atmosphere. For hygroscopic drugs, moisture transfer from gelatin to the drug may occur, causing instability. As a consequence other materials such as hydroxypropylmethyl cellulose (HPMC) have been used in capsules (50). For those who cannot ingest substances of animal origin, starch is also available but this has little practical value for inhaled products. B.

Multiunit Dose (Blister Disks, Strips, Dimples, and Tubes)

Unit-dose blisters, which are usually composed of aluminum alone or in conjunction with a plastic liner, are considered barrier packaging (51). The entry of moisture can only occur at the point of the seal, which may be achieved by the application of pressure or heat and may require adhesive. Blisters can also be arranged in multiunit-dose systems. Multiunit-dose blisters may be arranged on disks or in strips. Examples of these types of metering system include the Diskhaler® (Relenza, GSK) and Diskus® (Advair/Seretide, GSK) C.

Reservoir

Reservoir systems employ a hopper of powder, which feeds a metering system that is indexed into the air stream upon device actuation. These systems require

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engineering know-how to avoid moisture ingress into the reservoir and to maintain potency throughout the use of the device. One strategy to avoid moisture uptake is to include a desiccant in the packaging. A second method not unique to this metering system is to over wrap the entire product to present a barrier to moisture. Example: Turbuhaler® (Astra Zeneca) and Clickhaler® (Innovata Biomed) (Fig. 3). Examples of each of these metering systems are currently available in the United States and globally. IV. A.

Dispersion Mechanisms Passive Systems (Fluid Dynamics)

Passive inhalers are so called because they do not contribute energy to the dispersion of drug. The energy is imparted by the inspiratory flow of the patient. All of the currently marketed devices are passive inhalers. They vary enormously in internal geometry and pressure drop generated upon inhalation. The paradox of dry powder inhaler performance is that high-pressure drop, consistent with high shear, is an efficient means of powder dispersion but the resistance experienced by the patient may be at best, uncomfortable and at worst, an obstacle to the asthmatic patient. In order to allow lower-pressure drops to be adopted tortuous paths, gratings, impellers, and a variety of baffles have been adopted (Fig. 4). B.

Active Systems (Applied Energy Systems)

Active systems introduce energy (electrical or mechanical) into the powder bed to disperse powders as aerosols or to introduce them into inspiratory airflow. Active systems can potentially reduce patient-variability effects relative to passive systems because passive systems require the patient to provide all of the

Figure 3 Illustration of metering system from Clickhaler®. Source: Courtesy of Innovata, Nottingham, U.K.

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Illustration of mechanisms of dispersion of particle: (A) Spinhaler®, (B) Rotahaler®, and (C) Turbuhaler®

Figure 4

energy necessary to disperse the powder dose, while active systems provide some of this energy through an internal, stored energy source. Active dry-powder inhalers (DPIs) powder dispersion is limited to a handful of physical or electrical mechanisms available for use in a handheld device. Mechanical mechanisms include vibration, impact force, compressed air, and impellers (52). All of the mechanical actions can implemented strictly mechanically or through electrically driven means. In addition to the electrical generation of mechanical forces, other electrical means include electric field manipulation of charged particles. Active dispersion in a DPI increases the overall device cost and complexity. Assuming a performance benefit, these costs will be more easily absorbed in more expensive therapies. It is telling that the first (and currently only) approved active device is Pfizer’s Exubera® product for dry-powder insulin delivery because diabetes therapy can support slightly greater expenses than treatment of respiratory diseases. The Exubera DPI uses compressed air generated by the patient through an on-board pump to disperse the powder dose through a “transjector” nozzle into an antistatic holding chamber. Active DPIs for respiratory therapy will have to address costs issues in order to be commercially viable. Table 3 lists active DPIs. The use of electronically driven dispersion provides an opportunity to include additional features in the inhaler. Recent FDA guidance makes an

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Table 3

Active Dry-Powder Inhalers

Company

Means of active dispersion

AC Pharma AirPharma Dura Microdose Nektar Oriel Therapeutics 3M Vectura

Mechanical or electronic

Compressed air Compressed air Impeller Vibration Compressed air Vibration Impact force Compressed air

Mechanical Mechanical Electronic Electronic Mechanical Electronic Mechanical or electronic Mechanical

References (53) (54) (55) (56) (57) (58) (59) (60)

integral dose counter a necessary part of any DPI design (61). DPIs incorporating electronics can use the electronic system to provide dose counting. Other features of electronic-based DPIs could include dose reminders, audible or visual cues of successful dose delivery, measurement, and tracking of inhalation flow-rate parameters. Although no electronic device has yet been approved, the prevalence of electronics in society and in medicine suggests that their eventual approval is inevitable. V.

Inhaler Designs (Patent Review)

These are given in Table 4 and Figure 5. VI.

Characterization of Dry-Powder Inhaler Products

The characterization of aerosols has always been difficult because of their dynamic and heterogeneous nature. The conventional methods of particle-size analysis focus on aerodynamic behavior, which is of significance for the entry of particles to the lungs. These methods include inertial impinger and impactor methods (71). These instruments are calibrated with monodisperse spheres and are good for sampling ambient aerosols, which are equilibrated with atmosphere. For pharmaceutical aerosols, which do not reach equilibrium in the sampling time frame, the data must be interpreted in the context of the experimental design (72). Table 4

References for Commercially Available Inhaler Designs

Device Diskus® Diskhaler® Clickhaler® Ultrahaler® Turbuhaler® MAGhaler® Easyhaler® Twisthaler® Cyclohaler/Aerolizer®

Company

Area of action Drug-delivery method References

GSK GSK Innovata Biomed Aventis Astra Zeneca Boehringer Ingelheim Orion Schering-Plough Novartis

Locally Locally Locally Locally Locally Locally Locally Locally Locally

Strip pack Blister package Powder Powder Dosing Wheel Dosing Wheel Powder Dosing wheel Powder

(62) (63) (64) (65) (66) (67) (68) (69) (70)

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Relationships between (A) device (simple/complex) and mechanism of dispersion (passive/active), (B) device (simple/complex) and formulation (simple/complex), and (C) formulation (simple/complex) and mechanism of dispersion (passive/active) for a number of devices shown in the key.

Figure 5

A.

Current Methods

Inertial samplers are operated at designated airflow rates that are an important part of the performance of the system and underpins the validity of the calibration. To some extent, theoretical prediction of particle sampling on impactor stages, based on curvilinear motion, may be employed as approximation to performance (73,74). However, the FDA guidance document has indicated that flow-rate studies be considered under conditions that bring the calibration of impactors into question and render stage comparisons the only valid approach to quantification. As a consequence particle sizing estimates would be inaccurate. There are some practical considerations in estimation of particle size using these techniques. For solid particles, their momentum may result in bounce and reentrainment (75). Porous or hollow particles will behave as smaller aerodynamically sized particles as will the elongated particles (76). Because impactors are not calibrated with porous, hollow, or elongated particles, other phenomena such as image charges or interception may play a role in their deposition, which is unaccounted for in testing. The data must, therefore, be viewed as an approximation and full consideration given to physical properties of particles that might influence their sampling. To avoid misinterpretation of data, it is customary to employ more than one method to evaluate particle size. Microscopy is used as a confirmatory tool to establish both size and morphology. Laser-sizing techniques such as laser

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diffraction, velocimetry, and time of flight may be used to give data on the broad distribution including large aggregates delivered from dry powder inhalers. B.

Important Questions that Will Arise with New Systems

In light of the recent approval of Exubera (Pfizer, Groton, CT, U.S.A.) to deliver insulin systemically following inhalation, other performance criteria may have importance in the future. The bioavailability of drugs depends on physicochemical properties (crystallinaty, polymorphism, hydrophobicity, solubility, and partitioning) and physiological parameters (inspiratory flow, depth and timing of breathing, age, and disease state). Particle size has been characterized for a number of years, but dissolution rate is rarely characterized and has never been part of the approval process for aerosol products. This varies change as the future yields more inhaled aerosols intended to achieve systemic pharmacological activity.

VII. A.

Conclusions Regulatory Considerations

DPIs constitute a unique problem with regard to conventional guidelines on generic products problems. In order to be considered generic, the components of a drug-delivery system should be essentially the same as the Innovator product and the performance should, within statistical limits, be similar. However, DPIs have very few, if any, components in common and may involve very different mechanism of drug dispersion. The most optimistic view of the regulatory review process might allow abbreviated submissions, where a new drug formation is used in an existing approved system. It is unlikely that a novel device, with clearly different components, would be viewed as equivalent by current definitions. However, these observations raise an important point of discussion. Do the components play as important a role in the decision-making process as performance criteria? Is it important for a generic product to appear the same to the patient population familiar with the Innovator? Do the converging issues of ozone depletion and global warning, with their impact on propellant-driven MDIs, the need for novel drug-delivery systems, and the advances made in DPI technology require more visionary thinking to keep the marketplace competitive, while (but) retaining sufficient regulatory oversight to ensure the quality of the product and safety for the patient population? In response to these questions the following speculative remarks are offered: 1.

2.

Clearly the most important factors in establishing a generic product’s specifications relate to comparable performance to the Innovator. The importance of componentry relates to impact on performance assuming failure. Providing the performance of the “generic” device parallels the Innovator even in failure mode, the components are incidental. It is important that the patient should not be confused by having a generic “equivalent” that operates in a significantly different manner to

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the Innovator. The ergonomic question of “significance” requires debate and discussion. This may pose the most significant barrier to applying the term “generic” to DPIs. There is no question that much greater flexibility of thinking will be required if the next generation of inhalers are to make an impact in disease control. We appear to be in the paradoxical situation that much improved systems are available and could be adopted, but the concept of “equivalence” ties their performance to outdated and inefficient criteria. This simple observation is not intended to trivialize the issue, but the subject is worthy of debate.

It may be the case that time and an accumulating database will ultimately be required to resolve these questions. In the meantime, the quality, efficiency, and cost of products may stagnate. The next generation of inhalers will undoubtedly require approaches to regulation based on sound science, engineering in relation to safety and efficacy (clinical sciences). VIII.

General Conclusion

The ability to design and manufacture highly efficient and reproducible drug products will shift the focus, in the delivery of aerosols to the lungs, from engineering principles and design to disease therapy. Ideally, the delivery will be something that is sufficiently reliable to become a negligible consideration in the efficiency of the drug. The biology of the lungs and the nature of the disease will then become the hurdle to effective therapy and the delivery system will require similar considerations to any other damage form (i.e., for oral or parenteral administration). Future diseases are likely to target new diseases. Obvious examples are infectious diseases, cancer, and diabetes. They will employ either highly engineered formulations or devices. In the long run as the costs of production decrease, it is likely that combinations of these engineered formulations and devices will be evaluated. Acknowledgment The author is grateful to Matt Robinson for assistance with the preparation of the manuscript. References 1. 2. 3.

Aerosols. In: United States Pharmacopeia. Webcom Limited: Toronto, 2006:2617–2636. Hickey AJ, Thompson D. Physiology of the airways, In: Hickey AJ, ed. Pharmaceutical Inhalation Aerosol Technology. New York: Marcel Dekker, Inc., 2004:1–29. Windom HH, Burgess CD, Siebers RW, Purdie G, Pearce N, Crane J, Beasley R. The pulmonary and extrapulmonary effects of inhaled beta-agonists in patients with asthma. Clin Pharmacol Ther 1990; 48(3):296–301.

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75. 76.

Index

Acidopnea, 139 Adenylate cyclase, 112 Adult respiratory distress syndrome (ARDS), 187 Aeorlizer, components of, 430 Aerobid1, 192 Aerochamber1 spacer, 424–425 Aerosol(s), 127, 447 acetic acid, 140 administration, 172 applications, 172 respiratory tract infections, 171 delivery systems components, 400 dry powder inhalers (DPIs), 417 effectiveness, 409 inhalation, 399 medical devices for, 417 nebulizers, 417 pressurized metered-dose inhalers (pMDIs), 417 deposition, 162, 403 in environmental situations, 79 model, 55, 78 in nasal airways, 62 in upper respiratory tract (URT), 64 deposition mechanisms, 448 Brownian diffusion, 60 impaction, 60 sedimentation, 60 distribution, 128 droplets, 130 drug delivery, 161, 447

[Aerosol(s)] drug hygroscopicity, in human airways, 43 dry powder, 154 electric charges, 9 environmental, 79 fibrous, airborne behavior of, 9 formulation, 174, 175 generation, methodologies of, 285 hygroscopic, 32 hygroscopicity, in airways, 32 hypertonic and hypotonic solutions, 289 inhalation therapy, 104 inhaled, simulation of, 10 interaction with pulmonary endothelium, 164 isotonic, 17 metered dose inhaler (MDI), 77 mucociliary clearance (MCC), 162 nebulized, 437 pharmaceutical, 188 bioavailability of, 188 deposition of, 2 standards for, 212 polydispersed, 62 respiratory deposition of, 56–57, 76 mathematical model for, 57 sampling conventions, 73 to respiratory deposition, 73–75 size-fractionating characteristics of, 75 therapeutic, 76, 78 therapy, 2

461

462 [Aerosol(s)] bronchopulmonary infections, 236 pulmonary infections, 236 thermodynamics, theoretical models of, 32 transport equations, in airways, 70 Aerosol cloud enhancer (ACE), 425 Aerosolization, 405 Aerosolized drug, 1, 420, 421 delivery, 3D simulations of, 21 deposition model for, 2 in mouth and oropharyngeal region, 3 See also Individual drugs Air Liquide–CyberMedicine collaboration, 20 Airway(s), 140, 161, 446 airflow velocities in, 70 alveolar macrophages, 91 bifurcation and gravity angles, 4 cholinergic receptors in, 446 conductance, 104 deposition, 15, 404 cigarette smoke, 9, 50 cloud motion, 9 interception, 9 particle, formulas for, 8 disease, 2, 15 asthma, 22 cancer, 22, 447 chronic obstructive pulmonary disease (COPD), 22 cystic fibrosis, 22 heterogeneity of, 2 obstructive, 101 epithelial lining of, 87 generation of, 86 head and throat, 2 hygroscopic particle behavior in, 43 innervation of adrenergic, 92 afferent sensory nerve, 91–92, 115 efferent motor nerve, 91–92 mast cells, 91 mechanical properties of, 100

Index [Airway(s)] mucus, 89 peripheral, 446 physical simulator of, 43 resistance, 104, 107 smooth muscle cholinergic innervation of, 113 regulation of, 113 relaxation, 91 surface epithelial cells, 88 irregularities, 8 tracheobronchial (TB), 2, 161 Akita device, 437 Allergic bronchopulmonary aspergillosis, 244 Alpha-adrenoceptor, 115 Alpha-proteinase inhibitor, 245 Alveolar–capillary barrier, 132 Alveolar epithelium, 163 cells, 133 monolayer, 168, 174 permeability, 164 Alveolar macrophages (AMs), 150, 198 Alveolar ventilation, 98, 99 Amines alkaloidal, 156 drugs, 156 pulmonary uptake of, 156 Aminoglycosides, aerosolized, 237 Angiotensin-converting enzyme (ACE), 157 Anion exchanger (AE), 133 Antagonism, competitive and functional, 110 Antibiotics aerosolized, 236 antipseudomonal, 245 Arachidonic acid metabolism, 117 Asthma, 169, 445, 446 classifications of, 228 definition of, 228 disease characterization of, 228 forced expiratory lung volumes (FEV), 228 peak expiratory flow rates (PEFR), 228

Index [Asthma] therapy aerosolized drugs for, 228 anticholinergic agents, 233 beta-adrenergic agents, 231 cellular catechol–o–methyltransferase (COMT), 171, 232 cromolyn sodium, 233 management, 228 nedocromil sodium, 234 treatment of, 170, 228 Asthma allergic airway disease (AAD), 187 Atomization, 312–318 air-blast, 257 definition of, 253 methods effervescent, 316–317 electrostatic, 317–318 high-pressure, 315–316 pneumatic, 314–315 rotary, 313–314 ultrasonic, 317 principles of, 253 Atropine, 113 1 Azmacort , 422, 424

Beclomethasone dipropionate (BDP), 153 administration of, 290 Benzalkonium chloride, 433 Bernoulli effect, 429 Beta–adrenoceptors, 115, 446 Blisters, 451 Blood vessels, pulmonary, 93 Blue dextran, 157 Bradykinin homologues, 157 metabolism, 157 1 Breathancer , 424 Bronchiole, characteristics, 88 Bronchoconstriction, 109 of aerosol, 290 Bronchodilator drugs, 239 inhaled, 246

463 Bronchomotor tone, human, central reflex control of, 114 Bronchopulmonary drainage, 240 Bronchopulmonary dysplasia, 278 Bronchus, characteristics, 89 Budesonide, 433

Capacitance heat-transfer model, lumped, 38 Capacitance mass-transfer model, lumped, 39 Carboxyfluorescein, 276 Catechol I-methyltransferase (COMT), 171, 232 Ceftazidime, antibacterial, 291 Centriacinar emphysema, 240 Chitosan microspheres, 374 as therapeutic proteins carrier, 374 Chlorofluorocarbon (CFC), 348, 352 aerosol plume, 402 flunisolide, 406 in pharmaceutical inhalation aerosols, 407 physicochemical properties of, 407 propellants, 373, 418 properties of, 419 surfactants in, 420 properties of, 349 suspension systems, 405 Cholinergic fibers, 242 Cholinoceptor antagonists, muscarinic, 113 Chronic bronchitis asthmatic, 238 cigarette smoking, 239 classifications of, 238 definition, 238 pathologic characteristics of, 239 Chronic obstructive pulmonary disease (COPD), 238, 446 anticholinergic agents, 242 antiproteinases, 243 beta-adrenergic agents, 242 corticosteroids, 243

464 [Chronic obstructive pulmonary disease (COPD)] cysteine derivatives, 244 diagnosis, 446 mucolytics, 243 Chymase, 116 Ciliary beating frequency (CBF), 161 Clara cells, 89, 150, 167 Clausius–Clapeyron equation, 357 CO2-assisted nebulization with a bubble-dryer (CAN-BD), 336 Colloidal probe microscopy (CPM), 375, 391 Computational fluid dynamics (CFD) simulation, in human proximal airways, 19 Corticosteroids aerosolized, 235 inhalant anti-inflammatory effects, 235 treatment, of asthma, 235 oral, 235 Cromolyn sodium, 233, 374 anti-allergic agents, 233 anti-inflammatory agents, 233 bronchial hyperresponsiveness, 234 therapeutic drug dosages, 234 Cyclooxygenase metabolites, 117 Cystic fibrosis (CF), 2, 446 autosomal-recessive disease, 244 Cystic fibrosis transmembrane conductance receptor (CFTR), 133, 244 Cytokines, 138

Dalton’s law, 406 Derjaguin, Landau, Verwey, and Overbeek (DLVO) theory, 349 Desmethylimipramine (DMI), 164 Dexamethasone sodium phosphate, 278 Dichlorodifluoromethane, 419, 348 Dimethylether (DME), 376 1 Diskhaler , 427 components of, 429 Droplet–air model, 40 Droplet

Index [Droplet ] dispersion, 402 components, 400 evaporation dynamics, 325–329 size, prediction of, 260 Droplet formation, 400 fundamentals evaporation and shear thinning, 408 flow through orifice, 408–409 surfactants, 408 valves and actuators, 407 vapor pressure considerations, 406 mechanism of, 253, 315 performance characteristic dimensions, 411 semiempirical model of, 412–413 theoretical framework, 411–412 Drug(s) absorption, 147 administration, 170;172, 445 aerosolized. See Aerosolized drug bioavailability, 173, 456 delivery aerosol, 149 by dry powder inhalers (DPIs), 3 hygroscopicity, 32 by metered-dose inhalers (MDIs), 3 by nebulizers, 3 pulmonary, 148 dispersion mechanisms, 452 hydrophilic, 164 hygroscopic behavior of, 42 lipophilic, 164 metabolism, 147 nonlipophilic, lung absorption of, 189 physicochemical characteristics of, 51 See also Drugs, inhaled therapeutic, 187 therapy, 171 transport, 173 Drugs, inhaled advantages of, 222 bioavailability of, factors affecting, 194 bioequivalence of, 223 biophysical factors of, 198

Index [Drugs] clinical and regulatory factors of, 211 bioequivalence, 214 developmental pharmaceutics, 212 pharmacodynamic parameters, 213 pharmacokinetic parameters, 213 hygroscopic aerodynamic diameters of, 47 deposition mechanism, 48 thermodynamics of, 42 in vitro factors, 195–196 in vivo factors, 196–198 lung airways characterization, 199–200 pharmacokinetic. See Pharmacokinetic inhaled drugs Dry powder aerosols, 154 Dry powder inhalers (DPIs), 3, 189, 417, 445 characteristics of, 447 components, 426 delivery parameters for, 227 design, 454 dispersion forces, 426 mechanisms, 452 effectiveness, 431 excipients, 450 features, 426 jet milling process, 193 metering system, 448, 451 multi unit, 451 reservoir, 451–452 unit-dose, 451 operating conditions of, 431 particles, 449 regulatory considerations, 456 spray drying process, 193 technology, 373 Duo-Medihaler1, 193

Eel calcitonin (ECT), 158 eFlow1, 296

465 Elastomer, 423 Electrohydrodynamic atomization (EHDA) definition of, 297 process for, 296 Electrostatic forces, application of, 285 Electrostatic precipitators, 324 Emphysema, 240 Emulsion technology, 260, 376 Endocytosis inhibitors, 137 Endothelial cells, 129 Enzymes, 148 degradations, 148 distribution metabolism, 149 flavin-containing mono-oxygenases (FMO), 152–153 glutathione S-transferase (GST), 149 kinetics, 155 in lung, 149 proteolytic, 168 systems mono-oxygenase systems, 151 pulmonary CYP-450, 151 Ephidrine and pilocarpine, application of, 301 Epinephrine, 422 Epithelial cells, 129 Epithelial lining fluid (ELF), 130 airways, 138 pH of, 132 redox potential of, 135 reducing activity of, 135 thickness of, 131 tubular myelin structures within, 130 Epithelial sodium channels (ENaC), 133 Esterases, lung, 153 Ethanolic cosolvent, 418 7-Ethoxycoumarin (7-EC), 151 Ethylene diamine tetraacetic acid (EDTA), 433 Excipient, selection, 450 Exhaled breath condensates (EBC) cytokine concentrations in, 138 dilution, 139 patients with acidopnea, 139 pH of, 139 respiratory droplets, 138

466 Extrathoracic region particle deposition in, 17 steady inspiratory flow, 19 Exubera1, 173, 432, 453, 456

Flavin-containing mono-oxygenases (FMO), 149, 152, 153 Flunisolide chlorofluorocarbons, 406 Flunisolide hydrofluoroalkane, 406 Fraunhofer laser diffraction, 337

Gas antisolvent (GAS), 330 Gene therapy, 247 Gentamycin, 172 Gibbs free energy, 412 Gland cells, 89 Goblet cells, 88 G-proteins, 112–113 receptor activation of, 112 subunits of, 112 G–protein-coupled receptors, 111 Gram-negative bacteria, 289 Guanine nucleotide–binding regulatory proteins. See G-proteins

Haemophilus influenzae, 239 Halolite device, 437 1 Handihaler , components of, 430 Heptafluoropropane (propellant 227), 349, 374, 419 Hexadecane, 363 Human primary microvascular endothelial cells (HPMEC), 169 Humidification, 288 Hydrofluoroalkane (HFA), 349 aerosol systems, 403 chemistry, 399 flunisolide, 406 metered dose inhalers delivery systems powered by, 413 physicochemical properties of, 407 plume, 401 propellants, 373, 407 properties of, 356

Index [Hydrofluoroalkane (HFA)] reverse microemulsions in, 374 suspension systems, 405 Hydrofluoroalkane-134 (HFA-134), 225 for chlorofluorocarbons (CFCs), 225 physicochemical properties of, 225 propellant system, 225 Hydrofluorocarbon (HFC), 406 metered dose inhalers delivery systems powered by, 413 Hydrophilic–HFA-philic balance (HFB), 380 Hydroxypropylmethyl cellulose (HPMC), 451 Hygroscopic growth, of saline droplet, 40 Hygroscopic pharmaceuticals, inhaled, 31 Hypertonic and hypotonic solutions, aerosols of, 289 Inhalation drug-delivery systems application criteria, 417 metering valves on, 422–423 selection, 437–438 types of, 189 Inhalator, 427 components, 428 Inhibitors aprotinin, 174 leupeptin, 174 protease, 174 Inositol 1,4,5–triphosphate (InsP3) receptors, 237 In silico aerosol dosimetry model, 21 1 InspirEase , 425 Insulin delivery of, 290 therapy, inhaling, 135 1 Intal (disodium cromoglycate), 426 International Commission on Radiation Protection (ICRP), 17 lung-deposition model, 67 respiratory tract model, 67 1 Ipratropium bromide (Atrovent ), 233 Isooctane (ISO), 385

Index Karl Fischer titration, 340 Labtec 100, 354 Laminar jets disintegration of, 254 versus exit jet velocity, stability of, 256 Laser back-scattering, 364–365 Laser diffraction, Fraunhofer, 337 Lecithin, soya-derived, 420 Lecithin-based reverse microemulsions, 376 Lidocaine aerosolized, 237 bronchodilatory effect, 237 treatment of asthma, 237 Lipophilic drugs, lung absorption of, 189 Liposome-encapsulated ciproflaxin, 17 5-Lipoxygenase inhibitors, 209 Lung(s), 127, 445 aerosol delivery to, 417 aerosol deposition in, 8 air movement in, 100 air velocity profiles, 6 airways, dimensions of, 4 alveolar compartment, 127 aspiration of strong acids in, 133 blood supplies to, 446 breath condensates, 137 bronchial circulation, 94 bronchioles of, 89 cancer, 149 cell, 167 of children, 72 compliance, 102 deposition models, 60–61, 446 for radiation protection, 63, 65 drug absorption in, 147 drug delivery to, 2 drug metabolism in, 147 in vitro methods, 165 in vivo methods, 165 models of, 164–165 drug uptake by, 166 dynamics, regional depositionfraction curves of, 62

467 [Lung(s)] elasticity, 103 esterases, 153 expansion, 99, 102 fiberoptic bronchoscope of, 3 function flow–volume measurements for, 106 forced expiratory flow measurements for, 105–106 peak flow measurements for, 105 for heavy-activity breathing conditions, 14 humidification, 138 isolated perfused, 165 mechanical properties of, 100 mono-oxygenase systems in, 151 morphology, symmetric and asymmetric, 4 nonrespiratory functions, 95 particle deposition patterns, three-dimensional (3D), 15, 17–18 peptides, 154 degradation, 156, 159–160 perfusion apparatus, 166 peripheral, 10, 163 alveolar epithelium, 163–164 physiology of, 84 pressures, 100–101 proteins in, 159 enzymatic degradation, 159–160 pulmonary blood flow to, 93 pulmonary deposition fraction in, 73 pulmonary function tests for, 97 redox indicators, 135 respiratory functions, 95 acid–base balance, 96 gas exchange, 95 respiratory tract, 149–150 enzymatic systems in, 149 for sedentary breathing conditions, 12, 17 simulator, operating details of, 44 slices, 166–167 preparation, 166

468 [Lung(s)] thickness, 166 structure, 63, 84 systemic blood flow to, 94 therapeutic agents, 135, 447 tracheobronchial deposition fraction in, 72 transplantation, 243 ventilation of, 96 volume changes in, 102 volumes and capacities, determination of, 97–100 water transport in, 134 Macromolecules absorption rate, 136 diffusion, 136 vesicular transport, 136 Macrophages, alveolar, 198 Mass median aerodynamic diameter (MMAD), 162, 406 Medication delivery, systemic routes of inhalant, 221 disadvantages of, 222 intravenous, 221 oral, 220 parenteral, 221 Metered dose inhalers (MDIs), 189, 285, 325, 336 aerosol administration, 222 aerosol device, 77 breath-activated, 225 chlorofluorocarbons (CFCs), 191, 225 components, 400 design, experimental, 366–368 dimethyl ether for, 191 dosimetry, 189 drug delivery to lungs by, 413 excipients for, 192 hydrocarbon/fluorocarbon systems, 360–361 hydrofluoroalkane (HFA), 191, 223 liquefied propellant gas (LPG), 191 nonsoluble compressed gasses for, 192 performance and efficacy of, 402 products, 445

Index [Metered dose inhalers (MDIs)] propellants for, 191, 348 for pulmonary drug delivery, 189 soluble compressed gasses for, 192 solvents for, 190 spacer devices with, 3 surfactants, 190, 363, 408 suspension formulations physical parameters for, 366 variables for, 366 suspension stability, estimation of, 349–356 Metering systems, categories, 448 Methacholine, 113 Methylene blue, oxidization, 135 Microemulsions in compressible solvents, 377 in HFA-based pMDIs, 377 Micronization, 307 Monoamine oxidase (MAO), 149 distribution, 155 mitochondrial, 155 physiological role of, 156 Monoamines inactivation of, 155 oxidase activity, 154 Mononuclear phagocytes, 241 Mucociliary clearance (MCC) drug effects on, 162 regulatory control, 161 Mucociliary escalator, 89–90 Multiple-path particle dosimetry (MPPD), 70–73 lung deposition fraction by, 73 regional deposition fraction by, 71–72

NADPH cytochrome P-450 reductase, pulmonary, 152 Nasal airways aerosol deposition in, 62 morphometric model of, 65 Nasal breathing, extrathoracic airways for, 50 Nasal cavity, 19 enzyme activity of, 150

Index Nasal passages, deposition models for, 64 Nasal-pharyngeal-laryngeal-tracheal zone, 57 Nasopharyngeal airway, aerosol deposition in, 62 National Council on Radiation Protection and Measurements (NCRP), 65 lung-deposition model, 65 respiratory tract model, 66 Nebulizer, 189, 336, 417 1 Acorn , 264 advantages and disadvantages of, 270 aerosol droplets, medium for, 266 air-jet categories, 435 components of, 433 applications of, 226 clinical use of, 270 cosolvents, 193 delivery systems for insulin, 432 proteins and peptides, 432 in inhalation therapy, 432 jet versus ultrasonic, 289 liquid-vapor formulations, 193 mechanical efficiency, 266 design of, 273 nozzle, 257 types of, 258 output and efficiency of, 263 particle size, 266 scanning electron micrographs of, 274 solution, concentration of, 295 statistical efficiency model for, 264 surfactants, addition of, 271 therapy, of acute asthma, 226 types, 433 ultrasonic. See Ultrasonic nebulizer Nedocromil sodium, 234 bronchial hyperresponsiveness, 234 bronchodilation effects of, 290 clinical efficacy of, 234 effector inflammatory cells, 235 mechanism of, 234 Nerves, cholinergic, 109

469 Neuropeptides, sensory, 115 Neurotransmitters, 113 N-lauryl-beta-D-maltopyranoside, 193 Norepinephrine (NE), 155 Normal-addition supercritical antisolvent (NA-SAS) recrystallization, 334 experimental apparatus for, 335 Nosocomial infection, incidence of, 289

Oleic acid, 387, 420 in CFC-based pMDIs, 408 structure, 408 Oligonucleotides, 173 O-methylisoproterenol sulfate, 171 1 OptiHaler , 425 Ostwald ripening process, 356, 406

Panacinar emphysema, 240 Pentamidine, deposition of, 292 Peptides degradation endogenous, pulmonary control of, 156–157 proteolytic and antiproteolytic balance, 157–158 metabolism, in respiratory tract, 158 therapeutic, 187 Perfluoroheptane, 360, 362 Phagocytes, mononuclear, 241 Pharmaceuticals, hygroscopic inhaled, 31 Pharmacokinetic inhaled drugs, 200–211 compartmental models, 201–202 drug absorption evaluation, 200 hydrophilic drugs, 210 linear system analysis applications of, 206 drug absorption process, 206 Laplace transform method, 206 Loo–Riegelman equation, 208 superposition principle, 206 Wagner–Nelson equation, 207, 208 lipophilic drugs, 209

470 [Pharmacokinetic inhaled drugs] non-compartmental models absorption rate of, 204 bioavailability, 203 drug disposition, 203 mean arrival time, 205 mean residence time (MRT), 205 Phosphatidyl choline in CFC-based pMDIs, 408 structure of, 408 Piezoelectric transducer, vibration of, 296 Platelet–activating factor (PAF), 116 Plume cross-sectional nature of, 404 generation properties, 404 geometry, 404 oscillation frequencies, 401 Pneumatic atomization, 314–315 Pneumocystis carinii, treatment of, 292 Polyethylene terephthalate (PET), 422 Poly(propylene oxide) (PPO), 381–382 Powder(s) dispersion of, 338–339 dynamics, 337 flowability of, 342 SCF. See SCF powders separation techniques, 322–324 spray-dried. See Spray-dried powders X-ray diffraction, 340–341 Precipitators, electrostatic, 324 Pressurized metered-dose inhalers (pMDIs), 373, 417 actuation, 424 amphiphiles in, 374 CFC-based phosphatidyl choline in, 408 sorbitan trioleate in, 387, 420 development tools ab initio calculations, 380–383 atomistic computer simulations, 389–391 chemical force microscopy, 383–387 colloidal probe microscopy, 391 surfactants design, complementary information for, 387–389

Index [Pressurized metered-dose inhalers (pMDIs)] with HFA, 374 molecular scale behavior in, 373 particle formation mechanism of, 419 surfactants in, 374 Propellant-based inhaler formulations, 373 classifications microsuspensions, 374–375 nanosuspensions, 375–380 Propellant-driven metered dose inhalers (pMDIs) aerosol generation from, 399 lung deposition efficiency of, 403 primary atomization, 412 Propellant(s), 348–349 evaporation, 400 gaseous, 419 liquefied, 418 liquid density, 358 properties, physical, 356–364 density, 358–359 vapor pressure, 357–358, 414 Proteins, 159, 160 degradation, 159–161 pulmonary absorption, 158 Eel calcitonin (ECT), 158 insulin, 158 therapeutic, 187 1 Pulmicort , 406 Pulmonary blood vessels, 93 Pulmonary cytochrome P-450 (CYP-450), 148 and hepatic enzymes, 151 substrate specificity of, 151 Pulmonary diseases, 446 therapeutic agents, 447 Pulmonary vasoconstriction, hypoxic, 94, 96 1 Pulmozyme , 432

Radioaerosols, 128 droplets, administration, 131

Index [Radioaerosols] indicators 99m Tc-DTPA, 128 99m TCO4 , 128 Raoult’s law, 406 Rapid expansion of supercritical solutions (RESS), 330–333 Rat pulmonary type II cells, 169 Redox dye, 135 Respiratory diseases, 79 Respiratory system, particle deposition in, 17 Respiratory tract anatomical scheme of, 60 deposition efficiency, 60 deposition model, 56, 79 Findeisen model, 57 Landahl model, 58–59 electric charge effects in, 9 extrathoracic region of, fractional deposition curves for, 67–68 morphological structure, 56 temperature (T) and relative humidity (H) environment of, 46 RESS. See Rapid expansion of supercritical solutions Reverse-addition supercritical antisolvent (RA-SAS) recrystallization, 334 Ribavirin, 438 1 Rotahaler , 426 components, 428

Salbutamol, administration of, 290 Salbutamol sulfate, 374 microparticles, 375–376 Salbutamol sulfate-biodegradable polymer core-shell particles, 392 1 Salmeterol (Serevent ), 232 Sandoz 64-412, 210 SAS. See Supercritical antisolvent SCF methods, versatility of, 336 SCF powders

471 [SCF powders] characterization methods, 342 properties crystallinity, 340–341 dispersion, 338–339 particle shape and surface roughness, 337–338 particle size distribution, 336–337 Secretory leucoprotease inhibitor (SLP), 245 Sensory neuropeptides, 115 Serratia marcescens, 289 Silicosis, 74 Small particle aerosol generator (SPAG-2), 438 Sodium cromoglycate, bronchodilation effects of, 290 Sodium metabisulfite, 433 Solubility parameter theory, 360 and hydrocarbon/fluorocarbon systems, 360–361 Solution-enhanced dispersion of supercritical fluids (SEDS) technique, 335–336 Sorbitan trioleate, 387, 420 in CFC-based pMDIs, 408 structure, 408 Spacer devices lung delivery by, 425 and reservoirs, 424 1 Spinhaler , 426 components of, 427 Spray-dried powders characterization methods, 342 properties crystallinity, 340–341 dispersion, 338–339 particle shape and surface roughness, 337–338 particle size distribution, 336–337 Spray dryer design, structure of, 310–311 Spray drying, 308 advantages of, 309 applications of, 309

472 [Spray drying] open- and closed-cycle operation, 309–312 technique, 336 applications, 374–375 variables, 324 Spray freeze dryer, 321 Stokes’ law, 448 Streptococcus pneumoniae, 239 Sulfotransferase, 154 Sulfuric acid esters, 154 Supercritical antisolvent (SAS), 330 techniques, 333–336 experimental apparatus for, 334 Supercritical fluids parameters for, 330 particle generation using, 329 Surface tension effect of, 301 structure of, 254 Surfactants adsorption of, 354 in suspension formulations, 355 and water, in nonaqueous media, 355 Surfactant-stabilized dispersion formulations, 375 Suspension formulations, 421 stability estimation of, 349–356 factor for, 364 measurement of, 364–366 test, 366 water, effect of, 355

Tachykinins, 92 Tachyphylaxis, 234 Technegas, 129 1,1,1,2-Tetrafluoroethane, 349, 374, 419 Theophylline, 240 Thin-layer chromatography (TLC), 404 Tiotropium bromide, 242 Tobramycin, 172

Index [Tobramycin] delivery of, 291 injection solution, 291 TouchSpray1, 296 Trachea, 58 air velocity in, 37 turbulent flow field in, 37 Tracheobronchial (TB) airways deposition mechanism inertial impaction, 7 particle deposition in, 18 steady inspiratory flow in, 20 Tracheobronchial (TB) tree fluid dynamics environment, 7 smooth muscle in, 90 Tracheobronchial tube structure, 86 Trichlorofluoromethane, 348, 419 Tryptase, 116 1 Turbuhaler , 427 components of, 429

UDP-glucuronyltransferase (UDP-GT) enzymes, 154 Ultrasonic nebulizer, 226, 285, 290, 433 aerosolization mechanisms, 435–437 auxiliary airflow systems, 286 components of, 436 design of, 286 effects of, 286 Lang’s model, 286 medical applications of, 286 piezoelectric material, properties of, 285 therapeutic effect, 226 versus jet nebulizers, 226, 288 UNIFAC theory, 357 Upper respiratory tract (URT), aerosol deposition in, 64

Vancomycin, 172 Vascular endothelial cells, pulmonary, 94

Index Vasoactive intestinal peptide, 114 Ventilator circuits, 288 Ventilatory parameters, 2 1 Ventolin (albuterol sulfate), 426 Venturi effect, 429 Vesicles, intracellular, categories, 136

473 Water-in-perfluorooctyl bromide (PFOB) (W/PFOB) emulsions, 375 Weber number, 259 Xanthines, 118

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Lung Biology in Health & Disease Volume 186 Pleural Disease

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Lung Biology in Health & Disease Volume 151 Fetal Origins of Cardiovascular and Lung Disease

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