3.3.1  Mechanisms of Particle Deposition in the Respiratory Airways

Three major mechanisms are responsible for particle deposition in the lungs: inertial impaction, sedimentation, and Brownian diffusion (16). The deposition mechanism directly correlates with the particle diameter and determines the deposition of the particles in a particular area of the respiratory airways (26).

3.3.1.3  Brownian Diffusion

This is the predominant deposition mechanism for particles with an MMAD of less than or equal to approximately 0.5 μm. These particles move arbitrarily with gas molecules and collide against the airway walls. This is the mechanism through which respiratory gas exchange happens within the functional units, acini, of the lung. Hence, it is hypothesized that aerosol particles depositing via Brownian motion deposit in the acinus part of the human respiratory apparatus. A contrasting characteristic of Brownian deposition is that as particle size decreases, deposition increases. Particle deposition via Brownian diffusion is directly proportional to the diffusion constant DB

$$\mathrm{DB}=\left(\mathrm{ckT}\right)/(3\pi \mu d_p),$$

where c is the Cunningham constant that factors the influence of air resistance on the extreme small size of aerosols, k is the Boltzmann constant and T is the absolute temperature. Around 80% of particles with an MMAD of less than or equal to 0.5 μm are removed out of the respiratory tract during the exhalation process (29). The pattern of deposition of aerosol particles in the body is shown in Figure 3.3.

Figure 3.3   Pathway of Aerosolized Drug Particles in the Body (16)

3.3.2  Factors Affecting Particle Deposition

Various physicochemical properties, physiological, and anatomical factors affect the deposition of aerosol particles in the bronchial tree. Among the physicochemical factors, aerodynamic characteristics such as particle size, particle shape, density, and aerosol velocity are perhaps the most important determinants of aerosol deposition. Lung physiology can also impact aerosol deposition and the depth of penetration into the respiratory tract. Pulmonary function tests (PFT) are used to evaluate the impact and progression of pulmonary diseases on lung physiology. PFTs are used to distinguish between asthma and COPD, and are also used to categorize/stage the severity of the condition. The severity of the pulmonary disease determines the therapies recommended to treat the condition and when it is necessary to adopt changes in treatment regimens. Four general parameters should be considered to evaluate the size and morphology of an aerosol particle:

1. Mass Median Diameter (MMD) is the diameter of the particles of which 50% w/w of particles have lower diameter and 50% w/w have a higher diameter.
2. Percentage of weight of particles with a geometrical diameter of less than 5 μm.
3. Geometric Standard Deviation (GSD) is the ratio of the diameters of particles from aerosols corresponding to 84% and 50% of the cumulative distribution curve of the weights of the particles.
4. Mass Medium Aerodynamic Diameter (MMAD) describes the size and morphology of the aerosol’s particles by considering their geometrical diameter, shape, and the density: MMAD = MMD x Density^½ (16).

3.3.3  Effect of Particle Size

The particle size of an aerosol has a crucial impact on its mechanism and site of deposition in the lung. Large-size particles with diameters of (>10 μm) that come in contact with the upper airway tract are removed by mucociliary clearance. Aerosols with a particle diameter range from 0.5 to 5 μm deposit through sedimentation mechanism (16). Aerosol particles that are intended to penetrate the lung were determined to be in the size range around 2–3 μm (30). Particles with very small diameter may be exhaled before depositing in the lung; however, holding the breath can prevent this. Extremely small diameter particles (<0.1 μm) are not easy to manufacture, though they efficiently settle by Brownian diffusion mechanism. Human studies performed using particles in the size range of between 0.005 and 15 μm demonstrated that minimum deposition was seen at a particle size of 0.5 μm. For particles greater than 0.5 μm the deposition increased due to an increase of inertial and gravitational deposition, whereas deposition increased as the particles were progressively smaller than 0.5 μm due to an increase in diffusion transport. It has also been observed that at a particular particle size, higher deposition is seen with an increase in tidal volume, a clear indication that lung physiology and depth of inhalation can have an impact on aerosol deposition. Nevertheless, researchers have not been able to confirm an exact geometrical diameter that results in deposition of inhaled particles because even large particles that have a porous internal structure can penetrate and deposit in the lungs (31).

3.4.1  Mucociliary Clearance

This is the respiratory system’s unique mechanism against materials that enter the respiratory tract from the outside environment during breathing. The mucociliary clearance is an efficient self-respiratory cleaning besides other clearance mechanisms such as cough and alveolar clearance (16). From the trachea to the terminal bronchioles, the ciliated epithelium is extended and covered by a double-layered mucus blanket: a low-viscosity periciliary sol layer covered by a high-viscosity gel layer. The mucus is secreted by airway epithelial goblet cells and submucosal glands. The upward movement of the mucus clears the trapped insoluble particles toward the pharynx from where it can go into the gastrointestinal tract (28). The efficiency of mucociliary clearance is significantly diminished in respiratory diseases like asthma and cystic fibrosis (32).

3.4.2  Alveolar Clearance

Absorptive and non-absorptive clearance mechanisms clear deposited particles from the terminal airways (33). The absorptive mechanism involves either direct penetration into the epithelium cells or uptake and clearance by alveolar, interstitial, intravascular, and airway microphages. The alveolar epithelial surface where respiratory gas exchange occurs is populated by alveolar macrophages that engulf and remove undissolved particles that deposit in the alveoli. The uptake by macrophages can severely limit the systemic bioavailability of drugs contained in such particles due to enzymatic degradation inside the macrophages. The non-absorptive mechanism carries particles from the alveoli to the ciliated area from which the mucociliary clearance process further removes the particles from the conducting airways (16).

3.5.1  Pressurized Metered Dose Inhalers

The pMDI revolutionized pulmonary drug delivery. The pMDI consists of a pressurized canister that contains the drug in the form of a solution or suspension along with a suitable propellant. Initially chlorofluorocarbons (CFCs) were used as the propellant in pMDIs. But due to the deleterious impact on the ozone layer and its subsequent banning, CFCs are no longer used in pMDIs. Currently, hydrofluoroalkanes (HFA) are used as the propellant in marketed formulations. A dose metering valve is a part of the sealed end of the canister that inserts into the actuator. As the actuator is depressed, it presses the valve stem assembly down through an attached spring mechanism. As the valve stem moves down it disconnects the metering chamber, allowing the propellant and formulation to eventually emanate through the valve orifice. This releases the pressurized propellant along with the drug solution or suspension into the actuator. The unique design features of the actuator extend out as the mouthpiece delivers the high-velocity aerosol spray. The mouthpiece has a removable cover that prevents dust and particle contamination. The basic aspects of this system have not undergone considerable change; however, the methods that are used in formulating the drug suspension have undergone some changes. MDIs have to be primed before using, and when the formulation is a suspension, the system has to be shaken well prior to use (35).

Numerous studies have been done to evaluate how much of the emitted dose eventually deposits in the lungs. It has been established that the lung deposition of the emitted dose from a pMDI is between 10 and 20%. The low lung deposition is partly due to oropharyngeal deposition of about 50–80% of the emitted dose (28). Additionally, efficient delivery of the aerosol from the pMDI into the lungs significantly depends on the use of proper technique when self-administering the dose. Some common mistakes that were found among patients are errors related to the breathing pattern and inspiratory air flow. A considerable number of patients were found to hold their breath after emitting the dose from the device leading to loss of drug in the oropharyngeal region. When inspiring the emitted dose from a pMDI, the patient is expected to breathe in slowly and deeply. If the patient breathes in rapidly this alone will cause the aerosol particles to deposit mostly in the oropharyngeal region. After proper breathing in of the aerosol, most products require the patient to hold the breath for 10 seconds, which in itself may not be always possible due to the decline in respiratory anatomy and physiology in advanced disease states (36). Numerous patients have difficulty with the hand–mouth coordination that is necessary while using the product.

Accessories that can be attached to the pMDI mouthpiece, such as spacers and valved holding chambers, have helped to alleviate the problems associated with hand–mouth coordination. These devices function as an aerosol reservoir into which the dose is actuated and which then can be inhaled, reducing the problems related to hand–mouth coordination. Another innovation that was developed in pMDIs to address hand–mouth coordination is the breath actuated pMDI (BApMDI). BApMDIs use an inspiratory airflow triggered system that releases the dose when the patient inhales after the device mouthpiece is placed appropriately (37). The inhalation causes the firing of the device and no hand-based actuation is needed. Despite all these disadvantages, pMDIs are still the most popular method for pulmonary drug delivery due to the small size, convenience of handling, and its sheer presence and optimization in the pulmonary drug delivery realm for over 65 years.

Table 3.2 shows various pMDIs currently in the US market along with certain characteristics associated with each product. As is evident from the table, the two most commonly used propellants are HFA-134a and HFA-227. Most products have minimal amounts of other excipients when present in the formulation. All the products except ciclesonide (Alvesco®) are suspensions. Ciclesonide is a solution and the product does not have to be shaken prior to use. The Bevespi Aerosphere® product uses a co-suspension of the active ingredients in porous phospholipid particles prepared from 1,2-distearoyl-sn-glycero-3-phosphocholine (DSPC) stabilized by calcium chloride. The aerosphere technology coupled with the co-suspension delivery technique is a notable advancement among the pMDI product platforms. The aerosphere (also referred to as pulmosphere) technique is used to first manufacture highly porous phospholipid microparticles (17). An emulsion of DSPC in perfluorooctyl bromide (perflubron) and water stabilized by calcium chloride is atomized and sprayed into a dryer. As the drying proceeds, the perflubron evaporates, leaving behind nano- and micrometer sized pores on the dried DSPC microparticles. The DSPC microparticles can then be blended/mixed with one or more micronized active ingredients within a single product. The drug crystals interact strongly with the DSPC particles allowing for formulating a combination product with no drug–drug interactions that may destabilize the product. This technique also allows for delivering a higher than conventionally possible dose of the active ingredients due to the unique particulate microstructure, physicochemical stability, and suspension uniformity. The aerosphere-co-suspension technology has at least partially addressed some of the challenges traditionally attributed to pMDIs, such as low lung penetration, dose inconsistency, poor suspension stability, etc. The highly porous DSPC particles enter the lungs with ease through the inspiratory airflow where they absorb moisture and collapse into a mucinous mass that readily assimilates into the pulmonary membrane. This causes the active ingredient(s) to be released locally where it produces its biological action.

3.5.2  Dry Powder Inhalers (DPIs)

DPIs were first introduced in the market in the 1970s as single dose systems. The medication was present in gelatin capsules that were placed in the device and then inhaled. One important distinction between DPIs and pMDIs related to product usage is that the patient is expected to inhale forcefully and deeply when using the DPI. DPIs do not need any propellants and use inspiratory air to deliver the aerosol into the lungs. Thus, high-velocity inspiratory air will enable efficient entry of the particles into the lungs. Additionally, forceful inspiration will produce a turbulent stream of air that is capable of breaking down solid agglomerates propelling the aerosol into the lungs, thereby minimizing oropharyngeal deposition (38). Approximately 12–40% of the emitted dose from a DPI deposits in the lungs. A significant problem associated with DPIs is the loss of about 25% of the emitted dose within the device. Issues related to poor lung deposition from DPIs are partly due to the inability of the inspired air to de-aggregate large solid particulates in the formulation (39). The design of the device helps to address this issue to a certain extent, but the aerodynamics of the inspiratory air flow, humidity in the air flow, and temperature changes within the respiratory tract can determine how well large particulate agglomerates break down when a DPI is being used. The device itself can be an impediment to generating an optimal inspiratory air flow. Research on approaches such as the use of compressed air and powered propellers to disperse the aerosol powder have not yet lead to commercialization. The DPI products currently in the market are all inhalation driven and are available in different designs. Table 3.3 lists most of the currently available DPIs in the United States. One of the first modern DPIs approved in early 2000s was the Advair Diskus®. This product was one of the most significant innovations among pulmonary drug delivery systems to obtain widespread approval and market penetration since the first MDI was introduced in 1956.

Soon after the Diskus® entered the market, the Turbuhaler® design, which was a reservoir-type multi-dose non-reusable device, became available. Additionally, one of the first inhalation driven, single-dose reusable types of DPI designs, the Handihaler® device, was introduced in the market at the same time. Based on the nature of the DPI devices and their design, the currently available products can be categorized as non-reusable or reusable types. And within the non-reusable type there are two categories: the ones that contain a premeasured dose of medications placed in blister strips and those that have a dry powder reservoir that can be metered to deliver the appropriate dose. The Diskus® device contains a blister strip of medication and the Turbuhaler® and Respiclick® designs are examples of the reservoir type. The reusable DPIs consist of a device into which a capsule or cartridge containing the dry aerosol powder is placed. The capsule is then pierced using needles connected to hand activated buttons on the device, after which the powder is inhaled (40).

An important improvement in DPI design was the introduction of the Ellipta® design. This device contains two blister strips delivering two or more different medications that cannot be combined in a single blend due to incompatibility or other technical difficulties. Another notable development among DPI products was the introduction of Diskhaler® technology (41, 42). This is a breath-actuated multi-dose reusable type product. An important and distinct characteristic is that the medication is placed in a disc shaped strip with each strip containing four blisters containing the antiviral medication zanamivir mixed with lactose. The disk is loaded into the device prior to use, punctured using the piercing mechanism in the device, and the patient inhales the powdered aerosol through the mouthpiece. Most DPIs use lactose monohydrate as the sole carrier excipient and some formulations contain magnesium stearate as a lubricant (43). The tobramycin Podhaler® uses DSPC microparticles as the carrier (44). A remarkable breakthrough among DPIs came through the introduction of fumaryl diketopiperazine (FDKP) microparticles. FDKP (bis-3,6(4-fumarylaminobutyl)-2,5-diketopiperazine) particles are manufactured using the patented Technosphere® technology that leads to the formation of highly porous and homogenous microparticles with an aerodynamic diameter between 2 and 2.5 μm. Over 90% of FDKP particles prepared in this manner are in the respirable size range (0.5–5.8 μm). FDKP, a fumaramide derivative of diketopiperazine at first self-assembles into planar nanocrystals in an acidic medium (pH < 5.2). A small amount of Tween 80 is also added in the liquid medium, and residual Tween 80 is present in the final product. The planar nanocrystals then further aggregate to form three-dimensional spherical aggregates with high porosity. The drug can be mixed into the initial excipient solution which gets adsorbed and incorporated in the final porous crystalline microparticles. The highly biocompatible preparation method led to the first successful inhaled insulin product, Afrezza® (an earlier inhaled insulin product, Exubera®, was withdrawn in 2007). Afrezza® is a tremendous technological achievement because of the non-invasive pulmonary method used to deliver a fast-acting insulin. The highly porous and homogenous FDKP particles containing insulin are capable of penetrating deep into the lungs where they readily dissolve at pH close to neutral, releasing the loaded insulin (45, 46).

DPIs are popular and very effective means of delivery inhaled therapeutics for systemic and local activity. Despite its unique advantages, there are several aspects related to the DPIs that have to be addressed. One particular issue is the loss of aerosol particles within the device. Because of the need for forceful inhalation during their use, DPIs are not recommended for children under 5 years of age. Even among older and adult patients significant errors related to proper use of DPIs have been reported. Some of the most commonly seen errors include incorrect breathing pattern during usage and wrong method of loading the inhaler. The DPIs also suffer from large inter-individual variabilities related to deposited dose, since the optimal delivery is mostly dependent on the patient’s respiratory pattern during product use.

3.5.3  Nebulizers

Modern nebulizers use external energy to produce a fine mist that is inhaled by the patient. Most of the formulations for nebulization are sterile solutions, except for Pulmicort Respules®, which is a sterile suspension. Table 3.4 lists various nebulizer formulations currently available in the US. Traditionally nebulized formulations are administered using jet nebulizers or ultrasonic mist nebulizers (47). Jet nebulizers use compressed air that is passed using tubing into the nebulizing device chamber in which the solution or suspension formulation is placed. The gas is made to pass through a narrow opening, creating a huge pressure differential, increasing the air velocity. At the point where the high-velocity air exits the opening of an inner tube in the nebulizing chamber, a thin film of the liquid formulation is provided through a capillary effect operated feed. At the narrow orifice a negative pressure is produced, and the liquid film gets drawn into the path of the air where it collapses into fine mist droplets due to liquid surface tension. A significant portion of the mist developed consists of very large and non-respirable 15–50 μm droplets. A suitably placed baffle removes such large droplets, feeding the liquid back into the nebulizing chamber. Smaller respirable droplets are allowed to ensue out of the nebulizer into an appropriate mouthpiece or a mask that can be placed on a critically ill patient. This method of delivering inhalation mist is extensively used in institutional settings where compressed air outlets are standard fixtures by patient bedsides. Jet nebulizers are popular in home settings and are used in conjunction with a portable air compressor.

Jet nebulizers are available in the reservoir, breath-enhanced, and breath-actuated types. The reservoir type employs continuous mist generation and cause over 80% of the aerosol to be lost to the environment. The breath-enhanced nebulizer uses two valves to minimize loss of the generated mist when the patient exhales. When in use, as the patient inhales, one valve opens on the vent through which air flows into the nebulizer and is carried through the mist and the mouthpiece. As the patient exhales, a valve present in the mouthpiece expels the exhaled air and mist out, but the vent valve closes, minimizing the loss of nebulized mist to the environment. Breath-actuated nebulizers have a valve that opens and aerosol generation is activated only when the patient inhales. Ultrasonic nebulizers use a piezo-electric plate that vibrates at ultrasonic frequency when electric current flows through it. The ultrasonic energy is transmitted to a mesh that vibrates when activated by the sound energy. The drug solution is in contact with the piezo-electric unit and the vibrating mesh. As the solution is vibrated, as well, cavitation causes water to break apart into small mist droplets as it extrudes out through the vibrating mesh. Some disadvantages related to nebulizers have been the need to clean the parts such as the reservoir regularly to prevent microbial contamination. Additionally, these devices also suffer from high loss of generated mist out into the environment (48).

The soft mist inhaler (SMI) Respimat® uses mechanical energy that is generated as a spring contained in the device is twisted and released to generate the mist. A cartridge containing the drug solution is loaded into the device. Following this step, the bottom of the device is twisted 180°, enabling the spring to be compressed to the desired tension. The twisting action also delivers an accurately metered volume of the drug solution into a pump cylinder. Next the patient presses the dose release button that causes the tension in the spring to be released and the energy from the spring forces the metered solution through one of the most crucial components in the system, called a uniblock (49). The uniblock is a precision engineered microchannel system comprising a silicon wafer bonded to a borosilicate glass plate. The silicon wafer has inlet, outlet, and filter channels etched on it using technology similar to that used to engineer microchips. Through rigorous and meticulous engineering, the channels have been produced to generate two fine jets of the drug solution that converge at a preset angle to generate a slow-moving cloud of mist (1/10^th the speed of aerosol emitted from an MDI). The metered dose generated from a Respimat^® device is expelled over ~1.2 s compared to the 0.1 s from an MDI. Although the breathing technique when using the Respimat^® is similar to that of pMDIs, the lung deposition of various medications were found to be, in some instances, more than twice that obtained from pMDIs (50). The Loxapine (Adasuve^®) oral inhalation product listed in Table 3.4 is unique in many respects. It is a single dose, breath-actuated, non-reusable mist system. After the device is properly activated by pulling and removing a tab, the device mouthpiece is placed in the mouth. The patient inhales through the device and the inspiratory air activates a drug-coated heat source that vaporizes the drug in less than 1 second. The drug vapor condenses into an aerosol mist with an MMAD between 1 and 3 μm (51).