Dry-powder inhalers (DPIs) are complex multifaceted systems designed to reproducibly deliver an efficacious dose of medicine in an appropriate, aerodynamic range of sizes to enable the treatment of systemic or topical diseases (see Figure 1). Variables that influence the performance and stability of DPIs include the formulation, device, and processes by which the formulation is prepared (e.g. blending, spray-drying) and filled into the device. 

 Figure 1: Composition of dry-powder inhaler products (FIGURE COURTESY OF THE AUTHOR.) 

Most DPI formulations contain a mixture of active pharmaceutical ingredient (API) and an inert carrier material such as lactose. As shown in Figure 2, various attributes of the powder formulation must be considered and managed through the preparation processes and when filling the material into the relevant DPI device or primary device pack (e.g., capsule, preformed blister, blister disc, or geometrical array). In the ideal world, inhalation powder formulations are insensitive to atmospheric and environmental conditions such as light, oxygen, and humidity; are uniform in concentration; flow well; have a low tendency to agglomerate; and do not compact during filling. Under such circumstances, the API might be expected to readily disperse under low energy conditions to produce a uniform and respirable aerosol with an acceptable fine-particle dose. For topical inhalation aerosols, aerodynamic particle-size distributions in the region of 2–5 µm are typically targeted. 

Figure 2: Characteristics of dry-powder inhaler formulations. (FIGURE COURTESY OF THE AUTHOR.) 

However, powders do not always exhibit these ideal properties. More often, powders are susceptible to moisture which can lead to changes in surface morphology and/or aggregation. In addition, non-uniformity and concentration gradients can result from segregation within the powder bed. These and other factors must be effectively managed during the preparation of formulations and their filling into DPI devices, regardless of device format. 

 The DPI device into which the formulation is filled can also affect the selection of filling method and equipment. The types of DPIs range from simplistic, low efficiency, passive devices to high(er) efficiency, active delivery devices in which the aerosolization process is driven by an external energy source. In terms of powder filling, the most relevant variables are the size, shape, and array of cavities into which the powder is to be deposited. Reservoir-based devices contain a relatively large powder hopper compared with devices designed to deliver pre-metered powder doses, typically in the 1–30 mg range, usually into either capsules or preformed blister cavities. In this context, the dispensing of powders in the fill weight range of 1–30 mg is defined as 

Cavity arrangement can affect productivity. For example, the unit operations required to fill and index a circular array of cavities in a disc might require more time than filling a blister strip that indexes linearly. To overcome such device complexities, unique and tailored filling solutions are beneficial. 

Although in simple terms, the process of filling DPI products merely requires the reproducible dispensing of powder aliquots into a cavity, the microdosing of ordered mixtures containing irregularly shaped particles of different particle-size distributions and mixed surface morphology represents a significant technical challenge. The need for device-specific equipment for high-speed operations and consistent processing adds to this challenge.  

 Figure 3 demonstrates the different factors to be considered (e.g., equipment, speed, scale) when developing either a clinical product and/or a commercial product. Figure 3 also indicates potential batch sizes or the number of filled doses required to support a product evaluation during its development. In formulation feasibility (i.e., pre-investigational new drug/investigational medicinal product dossier), for example, up to and including clinical proof-of-concept (i.e., Phase IIa) batch sizes in the region of 100–5000 doses are adequate not only for enabling clinical evaluation but also for the execution of pharmaceutical development studies. Consider also that fully automated filling solutions are expensive and, as mentioned, tend to be designed to accommodate a specific device format. 

Figure 3: Considerations of development stage in filling dry-powder inhalers. IND is investigational new drug. Ph is Phase. NDA is new drug application. MAA is marketing authorization application. (FIGURE COURTESY OF THE AUTHOR.) 

The challenge, therefore, has been in the development or identification of fit-for-purpose, manual or semi-automated filling equipment with the capability and functionality to handle inhaled powders in support of early DPI development studies. Although commonly used for early feasibility studies, manual dispensing of single-digit milligram quantities of powder is extremely labor-intensive and requires great manual dexterity. The nature of such manual operations can result in significant fill-weight variability, thereby causing an unnecessary number of rejected doses. Manual operations for inhalation delivery systems also have the potential to cause the powder to compress or to compress in an inconsistent manner. Such factors can contribute to variable delivery from the DPI device causing variability in key product-performance measures such as delivered dose uniformity and fine-particle dose. Such variability can lead to poor decision-making on the part of the inhalation development scientist regarding the suitability of a formulation or product for further development. 

To address this capability gap, bench-scale powder handling technologies such as Xcelodose (Capsugel, Cambridge, UK), Powdernium (Symyx Technologies, now Accelrys Inc, San Diego, CA), and the Quantos Perfect Dosing System (Mettler-Toledo, Columbus, OH) have each found niche application in the small-scale microdosing of inhalation powders. These technologies are being used in research and development (R&D) studies as well as in current good manufacturing practice (CGMP) production of Phase I and II clinical-trial materials (1). The data shown in Figure 4a, which are derived from studies conducted in Catalent's Research Triangle Park, North Carolina, facility using the Mettler-Toledo Quantos system demonstrate that powder blends containing micronized API and inhalation-grade lactose can be accurately and reproducibly delivered across, in this instance, a target fill-weight range of 1–25 mg without any powder segregation (2, 3). This assertion is demonstrated by the linear correlation between API assay value per powder aliquot and the total mass of powder dispensed. In this study, percent relative standard deviations (%RSD) of < 1% were achieved for fill weights in the range 5–25 mg and < 2.5% for fill weights between 1 and 2.5 mg (n=10 replicate measurements). These results were obtained without any optimization of the filling process, thereby suggesting that with further optimization, filling precision could be increased. 

Figure 4a (left): Effect of powder fill weight on blend concentration-segregation. Figure 4b (right): Effect of filling process on Ã¦rosol performance-compaction. NGI is Next Generation Pharmaceutical Impactor (Westech Instruments, Marietta, GA). (FIGURE COURTESY OF THE AUTHOR.) 

Additionally, the author received a fine-particle dose of > 50% (i.e., defined as the proportion of the dose delivered from the device with an effective cut-off diameter (< 5 µm), using inertial impaction for the formulation delivered from a simple capsule-based device (i.e., Monohaler, Plastiape, Italy), generated using a Next Generation Pharmaceutical Impactor (NGI, Westech Instruments, Marietta, GA). These data are presented in Figure 4b.  

In this latter study, the dose from a single capsule was discharged into the impactor at a flow of 100 L/min. Hence, it can be implied from the comparatively high aerosol efficiency of this simple formulation/device combination that the dry-powder formulation must have been dispensed into the device during the filling process with minimal compaction of the powder blend, thus allowing the API to be readily aerosolized from the carrier material on actuation of the device. Furthermore, it should be noted that through the judicious application of the same processing-equipment configuration, filling procedures, and the use of common personnel, it was possible to achieve effective process transfer and ensure excellent correlation in performance of the clinical product with respect to the early R&D development lots.  

 Although it is possible to conduct effective fit-for-purpose filling trials using small-scale semi-automated equipment, such equipment is not scaleable in a viable manner to allow application for the onward development of a commercial product (1). A commercial product must be produced in large quantities at high speed and with the appropriate degree of consistency required to ensure reproducible product quality. Therefore, one cannot develop a product beyond clinical proof of concept without having in place an effective scale-up path. 

As depicted in Figure 3, it is important to bridge the gap from clinical products to commercializable products filled on high-speed, semi-automated filling equipment. It is also crucial to consider the effect of changes to the filling process, equipment, formulation, or to the type of device on the product's pharmaceutical performance (e.g., dose delivery, fine-particle dose), especially when progressing to later stages of clinical development. For example, in these later development stages, higher speed equipment that is tailored to a particular device or product will need to be introduced.  

If a simple, low-cost, off-the-shelf, capsule-based inhaler is used for early clinical studies, a more sophisticated inhalation device may be selected for the registration and commercial-product stages. In such a case, it is likely that the new device might generate different aerosol performance and produce differences in dose delivery and fine-particle mass. Technical and clinical studies would be needed to bridge the early clinical and product development data to that of the product under development for commercialization.  

In a similar manner, changes made to the mode of device filling or filling equipment can also affect powder flow, segregation within the powder stream or bed, powder compaction, and so forth. Any one of these variables, or many other potential sources of variability, can confer significant changes in the initial performance of the product or in its performance over time (i.e., stability and shelf-life).  

Ideally, the inhalation development scientist should consider the selection of a mode or mechanism of powder-formulation filling for early formulation feasibility and then replicate that same unit filling operation throughout the product's development life cycle. Figure 5 provides one example of a scale-up pathway for DPI filling. In this scenario, the unit operation of filling is conducted using a fixed-volume cavity precision bored into a rotating drum. As the formulation is fed to the filling head, an aliquot of powder is drawn into the drum bore through the application of a small, negative air pressure. This negative pressure is maintained as the drum is rotated until the filled cavity is in the six-o'clock position, at which point a small positive pressure is applied to dispense the powder aliquot into the respective dose packaging format (e.g., capsule, preformed blister). 

Figure 5: Approaches for filling dry-powder inhalers. IND is investigational new drug. Ph is Phase. NDA is new drug application. MAA is marketing authorization application. Omnidose is a tradename. (FIGURE COURTESY OF THE AUTHOR.  FIGURE 5 IS REPRODUCED FROM HARRO HOFLIGER AG/COURTESY AUTHOR.)  

Appreciable development work can be required to optimize the cavity geometry and operating pressures for a particular powder (e.g., active concentration, flow, bulk density) and its target fill weight. However, in general terms, the effect of these factors on the filling process are consistent irrespective of process scale. When scaling the filling process, if the unit filling operation is maintained as constant, the major considerations for the process engineer can reduce to those factors affecting throughput product (e.g., uniformity of powder feed, form/seal/cut configuration, number and array of filling heads, in-process verification of accuracy and reproducibility) as opposed to optimization of the unit-filling process. As a result, the likelihood of consistent product quality, consistent product performance and potential to shorten development timelines is greatly enhanced. 

In the laboratory scale equipment presented in this example, the process is operator driven (i.e., the powder is dispensed into the small hopper located above the single cavity filling drum). A vacuum line is attached to the filling drum, and using a foot pedal, the operator applies vacuum to the powder to fill the cavity. The operator manually rotates the drum to the dispense position and, again using the foot pedal, applies positive pressure to dispense the powder into the capsule or blister. Operation at this scale requires blisters to be preformed and sealed manually. Intermediate scale semi-automated equipment operates in a similar fashion, however the drum will include multiple filling cavities and contain automation features.  

Typically, at this semi-automated scale of operation, the device cavities may be pre-formed, sealed, and cut manually off-line (if blisters) with statistical off-line fill-weight verification. Although significant operator intervention still occurs at this scale, filling throughput in the region of 1000–2000 doses per hour is feasible depending on the device configuration. In the high-speed scenario, all operations occur on-line under high throughput.  

Figure 6 contains some examples of alternative proven methods for the filling of powder-based pharmaceutical products along with an opinion on suitability for the processing of inhalation powders. In this context, the dosing principle refers to the specific mode within which the powder is dosed rather than providing any reference to the powder feeding mechanism.  

Figure 6: An integrated dry-powder inhaler (DPI) filling process scale-up path. (FIGURE COURTESY OF THE AUTHOR.) 

In conclusion, accurate, reproducible, and cost-effective, equipment-based solutions are available to support formulation feasibility and proof-of-concept clinical studies for molecules in DPI devices. However, application of such approaches requires the developer to bridge the technical and clinical programs when working toward product commercialization. Although  more expensive, the most robust approach to filling process development, enabling a more rapid path to market, is to adopt a scaleable process from the outset right through the product development life cycle. 

This article represents a partial synopsis of a presentation given by Dr. Davies-Cutting at Management Forum, Dry Powder Inhaler Conference, London, July 2010. Special thanks to Harro Höfliger for the use of certain images and data.  

 is director of inhalation products and technologies, at Catalent Pharma Solutions in Research Triangle Park, North Carolina, tel. 919.465.8430, fax 919.481.4908, 

 1. Micro-dosing Equipment Fills Niche in R&D, Clinical Trial Materials, 

 2. L. Mao et al., proceedings of Drug Delivery to the Lungs XIX, Edinburgh, UK, 2008, pp. 195–199. 

 3. L. Mao et al., "Fast into Man Model for Dry Powder Inhaler (DPI) Development," presented at AAPS, Los Angeles, CA, 2009. 

Articles

The author reviews key considerations for formulating powders for use in inhalers. This article is part of a special Drug Delivery issue.

Considerations and Approaches for Filling Dry-Powder Inhalers

In the course of three generations, ultrasound contrast agents have developed from simple air bubbles without shells to precise imaging devices. Their half-life in the bloodstream has grown from several seconds to longer than 15 minutes. Along with this great improvement, microbubbles have become able to reach not only the heart, but also organs more distant from the injection site such as the brain, the kidneys, and the liver. These advances have inspired the industry to consider using ultrasound contrast agents as targeted drug carriers (1–2).  

Ultrasound drug targeting entails an "aim-and-shoot" conception. That is, the target site can be visualized by diagnostic ultrasound imaging before an operator triggers drug administration by increasing the intensity of ultrasound waves, thus destroying the microbubbles and releasing the drug. The cavitation energy released by bursting the microbubbles can render biological barriers such as cell membranes, the capillary endothelium, and the blood-brain barrier temporarily permeable for large and polar drug molecules (3–4). Some of the most promising therapeutic research on microbubbles today involves targeting chemotherapeutics to solid tumors and nucleic acids to cells for gene therapy.  

One of the major challenges for the development of drug-loaded microbubbles has been to attain high drug-loading amounts. In response to this challenge, scientists have developed new structural designs of ultrasound contrast agents that offer enhanced loading potential. 

Structural design of ultrasound-triggered drug carriers

Ultrasound-triggered drug-carrier particles of several distinct structures can achieve targeted drug delivery. The particles' common feature is a great difference between the propagation velocity of ultrasound within the particles and within the surrounding medium.  

In general, such great structural diversity of ultrasound-active drug carriers (UAC) exists that the most popular terms in the literature (i.e., "microbubbles" and "ultrasound contrast agents") do not adequately describe them anymore. By way of example, many acoustically active constructs such as phase-shift nanoemulsions do not share the structure of a microbubble. Yet only a few of these constructs can scatter ultrasound, a characteristic that is required of contrast agents.  

So far, no clear nomenclature has emerged that distinguishes ultrasound active particles' application as drug carriers versus contrast agents and takes their structure into account as well. UACs may be divided into six categories, based on the the particles' formulation design (see Table I).

 Table I: Classification of ultrasound-active drug carriers according to their formulation design. 

In the recent past, UACs' structural designs have evolved to incorporate therapeutic molecules with a broad array of physicochemical properties (see Figure 1). Polar drug molecules, such as small interfering RNA, may be connected to the microbubble shell through electrostatic interactions or hydrogen bonds (see Figure 1a). Amphiphilic molecules may intercalate between shell phospholipids (see Figure 1b), and oil-soluble drugs may be included in an oil layer (see Figure 1c).

Figure 1: Structural design of ultrasound active carriers (UACs) and methods of drug incorporation (in red). The drug can be attached to the outer surface (e.g., by electrostatic interactions or hydrogen bonds) (a), an amphiphilic drug may intercalate between the monolayer phospholipids (b), a lipophilic drug may be incorporated into an oil layer (c), secondary-carrier associated microbubbles may carry the drug (d), a drug may be physically encapsulated into a polymeric shell (e), a protein may be surface loaded onto microcapsules (f), drug may be loaded onto the entire shell volume of protein microcapsules (g), and protein microcapsules may be loaded layer by layer (h). Adapted from Ref. 33. (IMAGE IS COURTESY OF THE AUTHOR) 

Drugs and nucleotides may be complexed into nanoparticles (i.e., secondary carriers) that are assembled in their turn onto the microbubble shell (see Figure 1d). The drug also may be encapsulated into a polymeric shell and covered by a layer of biocompatible protein, as is Point Biomedical's (San Carlos, CA) CardioSphere (see Figure 1e). Microcapsules with protein shells may be loaded by surface adsorption, by incorporating the drug into the entire shell, or by the layer-by-layer approach (5). 

Microbubbles are appropriate carriers for several drugs. Scientists have loaded the particles with large molecules such as oligonucleutides, plasmids, and proteins, as well as with small molecules such as doxorubicin, docetaxel, and dexamethasone (6–10). Research also has focused on delivering formulations including paclitaxel and resveratrol using acoustically active lipospheres. In addition, secondary-carrier-associated microbubbles have been introduced for the delivery of plasmid DNA and doxorubicin (11–15). 

In many experimental applications, microbubbles have proven to be suitable for drug targeting. The recent generation of microbubbles can be administered into the vein by infusion or by a bolus injection, and they possess a considerable half life of more than 15 min. (see Figure 2). During administration, diagnostic ultrasound is externally applied to the site of disease as during a common imaging procedure. Microbubbles scatter the ultrasound back and serve as contrast agents, thus allowing the clinician to find the target precisely and focus the ultrasound beam. Furthermore, the intensity of the scattered signal provides information about the amount of microbubbles that reside at the target site. 

Figure 2: The mechanism of action of ultrasound-active carriers (UACs). After being applied to the patient, UACs can be excited on the target site using low-intensity ultrasound (black solid lines) and visualized through the backscattered signal (gray dashed lines). Clinicians burst the UACs by increasing the ultrasound intensity (red solid lines). The cavitation energy perforates the capillaries and increases cell permeability for the released drug (red dots). Adapted from Ref. 33.  (IMAGE IS COURTESY OF THE AUTHOR) 

The clinician then applies a pulse sequence of ultrasound with a higher intensity to excite the microbubbles beyond their cavitation threshold. The drug carriers burst into fragments of drug-bearing shell debris and small gas bubbles, which dissolve and are exhaled through the lungs. 

The applied acoustic energy of ultrasound waves is transformed into cavitation energy and causes hydrodynamic phenomena such as microjetting, microswirling, and shock waves. These cavitation forces can cause the sonoporation of biological barriers by opening transient wall gaps and making cells and tissues permeable to the drug. After treatment is complete, the cells and vessels recover within seconds or minutes, and therapeutic effects begin. 

The size range of drug molecules or drug-loaded particles that can be administered by sonoporation is limited. Capillaries and the blood-brain barrier can be opened for particles as large as several hundred nanometers (e.g., liposomes and polyplexes) (16).  

The permeability of cell membranes, however, is restricted to particles of 2–3 MDa in mass (e.g., micelles, proteins, and plasmids) (17). Secondary-carrier-associated microbubbles provide one probable way to circumvent this restriction (see Figure 1d). These hybrid constructs combine site targeting, the ability of UACs to open capillaries, and the capacity of nanoparticles to enter cells by phagocytosis. The microbubbles carry the nanoparticles to the target site and open the capillary wall. The released nanoparticles penetrate the tissue and are taken up by the cells through phagocytosis, thus acting as secondary carriers for the drug (see Figure 3). Within the cell, the nanoparticles break the endosome defense mechanisms (e.g., by osmotic rupture or the proton-sponge effect) and release the plasmid into the nucleus.

Figure 3: Plasmid DNA uptake and release after capillary break and extravasation into the tissue. The secondary carriers (e.g., polyplexes) are attached to the microbubble wall (bottom left corner). After ultrasound-mediated microbubble destruction the polyplexes leave the vessel and penetrate the tissue. They are taken up by cells and escape enzymatic digestion. The therapeutic plasmid is then released to the nucleus and expressed. Adapted from Ref. 33. (IMAGE IS COURTESY OF THE AUTHOR) 

Sonoporation allows the efficient targeting of delicate organs such as the brain and the heart. Unlike liposomes, ultrasound-mediated targeting is independent from the availability of capillary fenestration or specific cell ligands, which are necessary for targeting by enhanced permeability and retention or for affinity targeting.  

Current applications of ultrasound drug targeting

Ultrasound-active drug carriers have been applied in numerous fields of medical research. To date, the areas of greatest interest appear to be gene delivery and tumor therapy.  

Microbubbles provide a method to focus the action of chemotherapeutic drugs onto the tumor without requiring an invasive intervention. Anthracyclines are perhaps the mostly widely endorsed candidates for targeted chemotherapy. They have been included in formulations such as poly(lactic acid) microcapsules, soft-shelled phospholipid microbubbles, secondary-carrier-associated microbubbles, and as phase-shift nanoemulsions in microbubbles (8, 9, 14–15, 18–19). Delivery of taxanes such as paclitaxel and docetaxel has also been investigated (9, 11). Microbubbles also have been used to overcome the blood-brain barrier with promising results (16, 20–21). Before studies on large animals can begin, however, appropriate medical ultrasound devices must be developed. 

Some gene therapy applications use mild diagnostic ultrasound to treat sensitive tissues such as the myocardium. Ultrasound-targeted gene therapy mostly focuses on promoting myocardium revascularization after ischemic stroke, or on antisense therapies of vascular hyperplasia (22, 23–24). Studies indicate that microbubbles are likely to become an effective tool for preventing or treating ischemic stroke. Furthermore, microbubbles have successfully targeted and destroyed blood thrombi, thus helping to restore the perfusion of the affected myocardium (25). 

Future directions for ultrasound drug targeting

The wide and established application of ultrasound contrast agents has provided a solid foundation for research into ultrasound drug targeting. Since the US Food and Drug Administration published safety recommendations for the use of diagnostic ultrasound in 1976, a considerable amount of data has been collected and evaluated. Researchers now must bridge the gap between diagnostic and therapeutic applications. 

More than 20 years of clinical experience with contrast imaging has helped scientists map the body organs most accessible to ultrasound. These organs represent suitable targets for microbubble targeting.  

Although FDA has approved ultrasound contrast agents only for a few limited cardiac-imaging applications to date, the agents are used much more broadly in Europe and Asia (e.g., to image the liver). A growing regulatory acceptance of contrast applications is necessary to prepare the ground for clinical studies on therapeutic drug targeting. 

Several safety concerns with regard to microbubbles still pose open questions to the authorities. In 2008, FDA issued an alert about Optison (GE Healthcare, Waukesha, WI) and Definity (Bristol-Myers Squibb, New York), the two microbubble contrast agents approved in the US. The authorities revised the contraindications and warnings for these products after several cases of severe cardiopulmonary reactions occurred in patients with pulmonary hypertension shortly after the contrast agents were applied. Recent postmarketing observations, however, provided reassurance about the safety of ultrasound contrast agents (26). 

Those safety incidents relate to microbubbles' ability to pass through narrow lung capillaries. Third-generation microbubbles are smaller than red blood cells and pass through the lung capillaries without impeding circulation. Severe adverse effects after microbubble administration have been observed in patients with multiple morbidities and large cardiopulmonary disturbances. Nevertheless, these concerns may potentially affect microbubble applications such as drug administration, where fairly enhanced dosages may be needed to reach therapeutic concentrations at the target site. 

One question relevant to all drug-targeting approaches in general is how to estimate a drug's pharmacokinetic therapeutic window. In a targeting approach, the active drug must be enriched at the disease site, and the therapeutic window must be reached there rather than in the systemic circulation.  

An important milestone on the way toward creating a successful product is the establishment of a lean and scalable manufacturing process that conforms to good manufacturing practice. The formulator should bear regulatory requirements in mind as early as in the prepatent stage of product development. If a production process is not designed to allow terminal sterile filtration, for example, during a drug's preformulation stage, it is problematic to redesign the process during the production of clinical batches. 

With their dimensions of between several hundred nanometers and 5–6 µm, UCA cannot be subjected to sterile filtration. Their sensitivity to pressure excursions is also a hurdle to terminal steam sterilization. Sterilization through gamma irradiation is, in most cases, not an option because of the sensitivity of glass and most active ingredients to gamma rays. A sterile filtration of the contrast agent Definity is possible before the fill–finish step because the microbubbles are formed at the patient's bed side by mechanical agitation in the final container. If the product is supposed to include preformed microbubbles, however, some aseptic processing steps seem to be inevitable. 

The development of techniques for safe and precise ultrasound focusing in sensitive areas such as behind the cranial wall remains a vital need. The ultrasound cycle should correspond to the maximal tissue replenishment achieved with microbubbles (e.g., by using electrocardiographic triggering to target the myocardium and the coronary vessels). 

Finally, microbubbles pose a serious challenge for establishing methods for quality control and particle characterization. The final product often does not contain a single particle species; it includes at least two. Phospholipid microbubbles, for example, are always accompanied by liposomes. Furthermore, depending on their diameter, microbubbles rise upwards at various speeds. This feature requires scientists to consider the analytical approach for particle sizing and determining Zeta potential.  

Undoubtedly, many aspects of UACs still must be sufficiently explored before an advanced drug-targeting strategy can be developed. Nevertheless, the industry has witnessed the emergence of the first fully established and marketed approach for active drug targeting: antibody drug conjugates. Inevitably, this approach will be followed by other, even more sophisticated means to guide the drug through the body to the desired site of action. The field of ultrasound drug targeting will remain active and exciting. 

This article was funded by the BioFuture grant of Germany's Federal Ministry of Education and Research (Raffi Bekeredjian, Heidelberg, Germany). 

 is an associate professor of pharmaceutical technology and biopharmaceutics, and 

 is a professor of pharmaceutical technology and biopharmaceutics, all at Ludwig-Maximilians University Munich, Butenandtstr. 5-13, D-81377 Munich, Germany, 

 is an associate professor of cardiology at Ruprecht-Karls University in Heidelberg, Germany.

*To whom all correspondence should be addressed.

 3. P. Li, W.F. Armstrong, and D.L. Miller, 

Online Exclusives

Microbubbles can temporarily open many biological barriers for polar molecules, macromolecules, and particles. Scientists have brought well-known contrast agents back to the laboratory and redesigned them as drug carriers. This article is part of a special Drug Delivery issue.

Drug Targeting by Diagnostic Ultrasound Contrast

Reproducible sustained delivery of a drug at a target site is one of the main themes in controlled drug-delivery systems. The most commonly used drug-delivery systems, which can release drugs longer than one week, are parenteral injections and implants. Certain implant systems can deliver drugs for more than one year, and the longest drug delivery can be achieved by biodegradable or nonbiodegradable implant systems. Some examples of US Food and Drug Administration approved long-acting products are listed in Table I.

(GEORGE DIEBOLD, PHOTODISC, GETTY IMAGES)  

Long-acting injectable formulations offer many advantages when compared with conventional formulations of the same compounds. These advantages include the following:  a predictable drug-release profile during a defined period of time following each injection; better patient compliance; ease of application; improved systemic availability by avoidance of first-pass metabolism; reduced dosing frequency (i.e., fewer injections) without compromising the effectiveness of the treatment; decreased incidence of side effects; and overall cost reduction of medical care.

 Table I:  Examples of US Food and Drug Admistration-approved long-acting formulations on the market.*  

This review focuses on the current status and explores long-acting injectables with special attention to marketed products. Injectable routes, types of long-acting injectables (i.e., oil-based injections, injectable drug suspensions, injectable microspheres, and injectable 

 systems), drugs and polymers for depot injections, commercially available depot injections, and future injectable sustained-release drug-delivery systems are also discussed. 

Types of injectable routes of administration

It is well recognized that the advantages of parenteral injections are immediate systemic drug availability and rapid onset of action. Another significant and unique advantage of parenteral injection is a long-term drug delivery by the formation of a depot or reservoir at the injection site after drug administration. As depicted in Table I, intravenous (IV) injection can be used for certain prolonged acting drugs due to the drugs' long half-lives in the body after IV administration. The sustained release of drug from these preparations is a result of the long-acting property of drug and its residence in the bloodstream or the bone. 

In general, there are two routes by which long-acting parenteral injections are most frequently administered: intramuscular (IM) and subcutaneous (SC). To determine the injectable route of administration for long-term delivery formulations, many factors should be considered such as safety profile, ease of administration, patient's limited mobility, area for target injection sites, quality of life and cost of therapy (1). In many cases, SC is the preferred route for administering a drug by injection because of greater area for target injection sites, use of shorter needles, ease of self-administration, less discomfort and inconvenience for patients, and better safety profile (1). Various insulin products are given SC, and this route of administration presumably continues to represent the primary route of delivery for protein-based drugs. However, the volumes of SC injection are usually limited to no more than 1–2 mL, and only nonirritant substances can be injected by a SC route because irritants can cause pain, necrosis, and sloughing at the site of injection. On the other hand, greater injection volumes (2–5 mL) can be given by the IM route. Mild irritants, oils, and suspensions can be injected by IM route in the large skeletal muscles (i.e., deltoid, triceps, gluteus maximus, and rectus femoris) because these muscles are less richly supplied with sensory nerves and are more vascular. Therefore, a few SC injections for long-term release can be found on the market (i.e., Depo-SubQ Provera 104, Pfizer (New York); Nutropin Depot, Genentech (South San Francisco, CA), and Eligard, sanofi-aventis (Paris), and many long-acting IM injections are available on the market (oil-based injections, injectable drug suspensions, and injectable microspheres).  

Sustained-release properties of injectables

Sustained-release parenteral injections can be divided into several types: oil-based injectable solutions, injectable-drug suspensions, polymer-based microspheres and polymer-based 

 formings. Oil-based injectable solutions and injectable drug suspensions control the release for weeks while polymer-based microspheres and 

 gels are claimed to last for months (1, 7). 

Oil-based injectable solutions and injectable drug suspensions.

 Conventional long-acting injections consist either of lipophilic drugs in aqueous solvents as suspensions or of lipophilic drugs dissolved in vegetable oils. The administration need for these long-acting formulations only takes place every few weeks or so. In the suspension formulations, the rate-limiting step of drug absorption is the dissolution of drug particles in the formulation or in the tissue fluid surrounding the drug formulation. Poorly water-soluble salt formations can be used to control the dissolution rate of drug particles to prolong the absorption, and olanzapine pamoate is an example of a poorly water-soluble salt form of olanzapine. Certain drugs for long-acting formulations are synthesized by esterification of the parent drug to a long-chain fatty acid. Based on its extremely low water solubility, a fatty acid ester of a drug dissolves slowly at the injection site after IM injection and is hydrolyzed to the parent drug. Once the ester is hydrolyzed intramuscularly, the parent drug becomes available in the systemic circulation. The release rate of paliperidone palmitate from long-acting injectable suspension is governed by this mechanism. In many formulations, a fatty acid ester of a drug is used to prepare an oil-based parenteral solution, and the drug-release rate from solution is controlled by the drug partitioning between the oil vehicle and the tissue fluid and by  the drug bioconversion rate from drug esters to the parent drug. However, several other factors such as injection site, injection volume, the extent of spreading of the depot at the injection site, and the absorption and distribution of the oil vehicle 

 might affect the overall pharmacokinetic profile of the drug. Decanoic acid esters of antipsychotic drugs are widely used for these oil-based IM injections. 

 The development of polymer-based long-acting injectables is one of the most suitable strategies for macromolecules such as peptide and protein drugs. Advantages of polymer-based formulations for macromolecules include: 

 stabilization of macromolecules,  improvement of systemic availability, extension of biological half life, enhancement of patient convenience and compliance, and reduction of dosing frequency. 

Among the various approaches to deliver macromolecules parenterally, biodegradable microsphere systems are the most commercially successful. The most crucial factor in the design of injectable microspheres is the choice of an appropriate biodegradable polymer. The release of the drug molecule from biodegradable microspheres is controlled by diffusion through the polymer matrix and polymer degradation. The nature of the polymer, such as composition of copolymer ratios, polymer crystallinities, glass-transition temperature, and hydrophilicities plays a critical role in the release process. Although the microspheres structure, intrinsic polymer properties, core solubility, polymer hydrophilicity, and polymer molecular weight influence the drug-release kinetics, the possible mechanisms of drug release from microsphere are as follows: initial release from the surface, release through the pores, diffusion through the intact polymer barrier, diffusion through a water-swollen barrier, polymer erosion, and bulk degradation. All these mechanisms together play a part in the release process (2). 

Another intensively studied polymeric injectable depot system is an 

forming implant systems are made of biodegradable products, which can be injected 

 a syringe into the body, and once injected, congeal to form a solid biodegradable implant. This article will briefly summarize the types of 

forming implants because the topic has been intensively reviewed elsewhere (3–5). Biodegradable injectable 

forming implants are classified into five categories based on the mechanism of depot formation: thermoplastic pastes, 

 polymer precipitation, thermally induced gelling systems, and 

 solidifying organogels. The mechanism of depot formation of thermoplastic pastes is to form a semisolid upon cooling to body temperature after injection into the body in the molten form. Cross-linked polymer networks can be achieved 

 in various ways, forming solid polymer systems or gels. Methods for 

 cross-linked systems include free radical reactions, usually initiated by heat or absorption of photons, or ionic interactions between small cations and polymer anions. 

formings can be produced by causing polymer precipitation from solution. A water-insoluble and biodegradable polymer are solubilized in a biocompatible organic solvent to which a drug is added which forms a solution or suspension after mixing. When this formulation is injected into the body, the water-miscible organic solvent dissipates and water penetrates into the organic phase. This leads to phase separation and precipitation of the polymer forming a depot at the site of injection. This method has been designed as Atrigel technology (QLT, Vancouver, Canada), which used as a drug-carrier system for Eligard. Thermally induced gelling systems show thermo-reversible sol/gel transitions and are characterized by a lower critical solution temperature. They are liquid at room temperature and produce a gel at and above the lower critical solution temperature.

 solidifying organogels are composed of water-insoluble amphiphilic lipids, which swell in water and form various types of lyotropic liquid crystals. 

Drugs delivered as sustained-release injectables

Various drugs are investigated for sustained-release injectable delivery systems for controlled drug delivery as recently described by these authors (6). These systems include small molecular drugs and protein/peptide drugs. Examples of drugs for sustained-release injectable delivery systems include: hormone therapy (i.e., human somatropin) (7, 8); protein therapeutics such as the analog of glucagon-like peptide-1 (9); recombinant human bone morphogenetic protein-2 (10); superoxide dismutase (11); salmon calcitonin (12, 13); insulin (14–16); gene delivery such as plasmid DNA (17–19); cancer therapeutic agents such as bleomycin (20), paclitaxel (21), cisplatin (22), a peptide-like antineoplastic agent (23); postoperative pain therapeutic agents such as ketorolac tromethamine (24); schizophrenia drugs such as aripiprazole (25), olanzapine (26); contraceptive peptide vaccine (27); drugs to treat alcohol dependence such as naltrexone (28); and immunosuppressive drugs such as rapamycin (29). 

Despite a number of parenteral depot studies using a variety of drugs, only drugs in limited therapeutic areas are available on the market. Antipsychotic drugs and hormones have been used for more than five decades in the field of schizophrenia and hormone replacement therapy. Since the first launching of microsphere formulation, Lupron Depot (Abbott, Abbott Park, IL) for the palliative treatment of advanced prostate cancer in 1989, several microsphere formulations and 

forming implants have been released on the US market. The therapeutic indications and drugs of commercialized products include: the palliative treatment of advanced prostate cancer (leuprolide acetate and triptorelin pamoate); the treatment of acromegaly (octreotide acetate and lanreotide acetate); the long-term treatment of growth failure (somatropin-rDNA origin); the treatment of schizophrenia (risperidone); and the treatment of alcohol dependence (naltrexone). 

As recently described by these authors (6), a variety of biodegradable polymers for controlled drug delivery intensively studied over the past several decades include polylactides (PLA), polyglycolides (PGA), poly(lactide-co-glycolide) (PLGA), poly(ε-caprolactone) (PCL), polyglyconate, polyanhydrides, polyorthoesters, poly(dioxanone), and polyalkylcyanoacrylates. Among the various approaches to deliver macromolecules parenterally, injectable biodegradable microspheres are the most successful systems (30). Many microsphere research reports have demonstrated the usefulness of biodegradable polymers such as PLGA microspheres (31–38), PCL microspheres (39), polyanhydride microspheres (40), polyorthoesters microspheres (41), and polyalkylcyanoacrylate microspheres (42, 43). 

The Atrigel technology that is used in Eligard containing leuprolide acetate and PLGA is a once-monthly 

forming implant for the palliative treatment of advanced prostate cancer. Many reports have been published on novel biodegradable 

forming polymers such as multiblock poly(ether ester urethane)s consisting of poly-[(

)-3-hydroxybutyrate] (PHB), poly(ethylene glycol) (PEG), and poly(propylene glycol) (PPG) polymer (44), PEG-grafted chitosan polymer (Chitosan–PEG) (45), methoxy poly(ethylene glycol)–poly(sebacic acid–D,L–lactic acid)–methoxy poly(ethylene glycol) triblock copolymer (mPEG–poly(SA–LA)–mPEG) (46), PCL–PEG–PCL triblock copolymer (47), and PLGA–PEG–PLGA triblock copolymer (48).  

Commercialized polymer-based injectable depot systems have used polymers or copolymers composed of monomers of lactic and glycolic acid. These polymers have the advantages of being semipermeable, biocompatible, and biodegradable, which makes them universally acceptable as injectable materials for drug-depot systems (49). 

Commercially available injectable sustained-release drugs

The list of commercially available injectable sustained-release drug delivery systems available on the market as pharmaceutical products is shown in Table II. Parenteral long-acting formulations (oil-based solutions and drug suspensions) have been in clinical use for many decades in the field of hormone replacement therapy. Sesame oil-based injection containing testosterone enanthate (i.e, Delatestryl, Endo Pharmaceuticals, Chadds Ford, PA) and castor oil-based injection containing estradiol valerate (Delestrogen, Monarch Pharmaceuticals, Bristol, TN)  were approved by the US Food and Drug Administration in the 1950s, and drug suspension for injection containing medroxyprogesterone acetate (Depo-Provera, Pfizer) was approved by FDA in September 1960. The administration route of these products is IM injection, and all of these products are still available on the market. In 2004, long-acting SC injection of medroxyprogesterone acetate (Depo-SubQ Provera 104, Pfizer), which is equally effective despite an almost 30% reduction in the dose, was approved by FDA. The first long-acting injectable microspheres of recombinant growth hormone (Nutropin Depot, Genentech) received approval from FDA for pediatric growth hormone deficiency (GHD) in December 1999. Nutropin Depot is designed to be administered by SC injection once or twice monthly.

Table II. Commercially available injectable sustained-release drug-delivery systems. 

In the 1960s, parenteral depot formulations of typical antipsychotic drugs were introduced for clinical use in Europe (50). Although long-acting typical antipsychotic formulations are widely used in Europe, clinicians in the United States have thus far been reluctant to use them despite their potential advantages because of several reasons such as concerns about increased adverse effects compared with oral therapy and the belief that patients do not accept or tolerate depot formulations as well as oral agents (51). Therefore, many oil-based depot formulations containing typical antipsychotic drugs (haloperidol decanoate, flupenthixol decanoate, fluphenazine decanoate, zuclopenthixol decanoate, and pipothiazine palmitate) are available on the market in Europe, Canada and Australia, but only haloperidol decanoate (Haldol Decanoate) and fluphenazine decanoate (Fluphenazine Decanoate injection) formulations are available in the US. In 2003, the long-acting formulation of risperidone (Rispedal Consta, Janssen, division of Ortho-McNeill Janssen Pharmaceutical, Titusville, NJ) became the first depot atypical antipsychotic drug to become available in the US (51). Rispedal Consta is formulated as an aqueous suspension of biodegradable microspheres. Water is added to the vial of microspheres, and the aqueous suspension is injected intramuscularly every 2 weeks (52). FDA approved paliperidone palmitate long-acting injectable suspension (Invega Sustenna, Janssen) for the acute and maintenance treatment of schizophrenia in July 2009 (53–56). Paliperidone palmitate, an atypical antipsychotic agent, is the palmitate ester of paliperidone and is the major active metabolite of risperidone (9-hydroxy-risperidone) (53). Paliperidone palmitate was formulated as an aqueous drug suspension with a specific particle-size distribution that has sustained-release properties and thus facilitates monthly dosing (53). In 2009, a long-acting depot formulation of olanzapine pamoate (Zyprexa Relprevv, Eli Lilly, Indianapolis, IN) was approved by FDA for the US market (57–59). Zyprexa Relprevv is an aqueous drug suspension containing a salt of pamoic acid and olanzapine (olanzapine pamoate monohydrate) for deep IM gluteal injection (60). 

Lupron Depot (leuprolide acetate) is the first marketed injectable PLGA microspheres in the US (approved in 1989) (61). Lupron Depot provides fairly constant release of the peptide during 1 month or 3 months in humans after IM injection and show sufficiently reliable efficacy for the treatment of patients with hormone-dependent cancers such as advanced prostate cancer (61). Encouraged by success of Lupron depot, several PLGA microsphere formulations have been investigated, and Trelstar (triptorelin pamoate, Watson Pharmaceuticals, Corona, CA) received approval from FDA for the palliative treatment of advanced prostate cancer in June 2001. Trelstar is designed to be administered by a single IM injection in either the buttocks, and the dosing schedule (one-, three- or six-month) depends on the product strength selected. In January 2002, the first parenteral 

forming formulation, Eligard was approved by FDA for the US market. Eligard uses the Atrigel technology, and Atrigel is a polymeric (nongelatin-containing) delivery system consisting of a biodegradable PLGA polymer formulation dissolved in a biocompatible solvent, 

-methyl-2-pyrrolidone (NMP). Eligard is administered subcutaneously, where it forms a solid drug-delivery depot and is designed to deliver leuprolide acetate at a controlled rate during a one-, three-, four- or six-month therapeutic period. 

In April 2006, FDA approved naltrexone extended-release injectable suspension (Vivitrol, Alkermes, Waltham, MA) for the treatment of alcohol dependence. Vivitrol is supplied commercially as a microsphere formulation of naltrexone for suspension, to be administered by intramuscular injection every four weeks. 

Sandostatin LAR Depot (Novartis, Basel, Switzerland) received approval from FDA in November 1998 for the treatment of acromegaly, a chronically disfiguring and debilitating hormonal disorder. Sandostatin LAR Depot is a sterile PLGA microspheres formulation of octreotide acetate for IM injection at every four weeks. Although it is not available in the US, a prolonged release PLGA microsphere formulation of lanreotide acetate for injectable suspension (Somatuline LA, Ispen Pharmaceuticals, Kleve, Germany) is available as a commercially available pharmaceutical product on the market in Europe. The indication of Somatuline LA is the same as that of Sandostatin LAR Depot, and Somatuline LA is designed to be administered by IM injection every two weeks. 

FDA approved Somatuline Depot for the long-term treatment of acromegaly in August 2007. Somatuline Depot formulation consists of a unique supersaturated concentration of lanreotide acetate (24.6% w/w lanreotide base), and contains only water for injection as an excipient (62). It is thought to form a precipitated drug depot at the injection site due to the interaction of the formulation with physiological fluids because Somatuline Depot can produce a stable gel when mixed with water at a specific temperature and pressure. The most likely mechanism of drug release is a passive diffusion of the precipitated drug from the depot toward the surrounding tissues, followed by the absorption to the bloodstream for a month. The administration route of Somatuline Depot is deep SC injection. 

Injectable sustained-release drug-delivery systems in clinical trials

Several clinical trials of injectable sustained-release drug delivery systems are currently conducted in the US. Some examples of injectable sustained-release drug delivery systems currently in clinical trials are listed in Table III.

Table III. Examples of injectable sustained-release drug delivery systems in clinical trials. 

Phase I pharmacokinetic-pharmacodynamic studies are ongoing for microsphere formulation of progesterone to establish the minimum effective dose of progesterone microspheres suspension, for weekly intramuscular injection. Phase III clinical studies of aripiprazole for once monthly IM depot administration are ongoing to evaluate efficacy, safety, and tolerability. Phase I trials are in progress with once-monthly IM injection of octreotide pamoate to investigate safety and tolerability of an octreotide extended long-acting formulation after a single dose in humans. Phase III studies of pasireotide long-acting release formulation are underway to evaluate the efficacy and safety of pasireotide LAR. 

forming system, and this system consists of sucrose acetate isobutyrate (SAIB), a pharmaceutically acceptable solvent, and one or more additives. One characteristic of the system is that a SAIB/solvent mixture has a low viscosity, but upon injection, the viscosity increases substantially as the solvent diffuses away from the SAIB (63). After dissolving or dispersing the drug in the SAIB/solvent solution, this solution is injected subcutaneously or intramuscularly. Upon injection, the solvent dissipates from the SAIB, and the increased viscosity controls the release of the drug from the gel. SABER-bupivacaine is designed to continuously deliver bupivacaine, a common local anesthetic, up to 72 hours to treat local post-surgical pain. This system injected at the surgical site prior to the wound closure and is currently in Phase III clinical studies in the US. 

ReGel (BTG, London) is a thermally reversible gelling system and is based on biodegradable triblock copolymer composed of PLGA–PEG–PLGA. Immediately upon injection and in response to body temperature, an insoluble gel depot is formed. OncoGel (BTG) is supplied as a frozen formulation of paclitaxel in ReGel and is entering Phase II trials. OncoGel is being injected directly into the tumor for oesophageal tumors, and the gel disappears in four to six weeks as it releases the paclitaxel. 

As evident by the growing number of sustained-release injectable pharmaceutical products on the market, injectable depot systems are becoming one of the most effective systems for long-term drug delivery. Owing to the enhanced quality of life and cost of therapy supported by the advances in drug formulation and polymer science, more sophisticated injectable depot systems will be developed and commercialized in the near future. Moreover, the introduction of more potent drugs and protein/peptide drugs are particularly good candidates for formulation as long-acting parenteral depot systems. Polymer-based injectable depot systems for protein/peptide drugs have many advantages such as protection of sensitive proteins from degradation, prolonged or modified release, pulsatile release patterns, and enhancement of patient compliance. These important and unique advantages offer potential commercial success of future sustained-release injectable pharmaceutical products that have novel active pharmaceutical ingredients, including therapeutic proteins and peptides. 

 is research associate professor at Sungkyunkwan University, School of Pharmacy, Suwon, Gyeonggi-do, Republic of Korea. 

 is a postdoctoral fellow and visiting scholar, 

 is an assistant professor of pharmaceutics and pharmaceutical technology, all at the University of Kentucky College of Pharmacy, Department of Pharmaceutical Sciences–Drug Development Division, 789 S. Limestone St., Lexington, KY 40536-0596, tel. 859.257.1571, 

. Heidi M. Mansour is also a member of Pharmaceutical Technology’s edtorial advisory board. 

A review on the current status of long-acting injectables, including commercially marketed products. This article is part of a special Drug Delivery issue.

Sustained-Release Injectable Drug Delivery

The production of minitablets for the sustained release of a drug is a new and promising area in pharmaceutical research. Minitablets (also known as microtablets) have a diameter equal to or smaller than 2–3 mm and can be filled into hard gelatine capsules (1). 

Minitablets are multiple-unit dosage forms, therefore, they present all the advantages of these systems over the single-unit dosage forms. These advantages include a low risk of dose dumping, less inter- and intra-subject variability, and a high degree of dispersion in the digestive tract, thus minimizing the risks of high local drug concentrations and reproducible bioavailability.

Minitablets are good substitutes for other multiparticulate dosage forms (e.g., granules and pellets) because they can be manufactured by direct compression. As a consequence, the use of solvents is avoided and high production yields are obtained compared to extrusion and spheronization processes. Furthermore, due to the manufacturing process, defined size and strengths can easily be produced, with small variability within and between batches (1, 2). 

Sustained-release minitablets can be produced either by matrix or by coating them in a fluid bed or coating pan (3–6). Unfortunately, little published information is available on both minitablet-coating technologies and the coating of minitablets in a pan (5, 6). However, from an industrial point of view, the coating of minitablets in a pan should be extremely convenient because it provides higher production capabilities, lower waste of coating materials, and faster equipment cleaning time with respect to the fluid bed.  

The aim of this research was to investigate the possibility of manufacturing a sustained-release, multiple-dosage form by coating minitablets, produced by direct compression, in a solid wall pan. Theophylline (TH) was used as the model drug, and an innovative ready-to-use pigment dispersion was tested to evaluate the possibility of reducing production time when compared with a traditional coating formulation. Finally, coated minitablets were filled into hard gelatine capsules. 

 TH (particles lower than 100 µm) was purchased from BASF (Ludwigshafen, Germany). Avicel PH 102 was supplied by FMC Biopolymer (Brussels). Spray-dried lactose, magnesium stearate, talc, titanium dioxide, trietilcitrate and yellow quinoline (all European Pharmacopoeia-grade) were purchased from Polichimica (Bologna, Italy) and used as received. Eudragit RL, Eudragit RS (Evonik, Darmstadt, Germany, polymer conforms to Ammonio Methacrylate Copolymer, Type A 

 31) and WAS (a product containing talc, titanium dioxide, yellow quinoline and trietilcitrate) were supplied by Rofarma Italia (Milano, Italy). Hard gelatine capsules size 0 were purchased from Capsugel (Colmar, France). 

TH, Avicel PH 102, spray-dried lactose, and magnesium stearate in four different ratios (total 4 kg per batch) were blended using an IMA Cyclops blender (IMA, Ozzano, Italy).  

Direct compression was performed using a tablet press machine for research and development and small batches (IMA Pressima equipped with 2 EU-D punches fitting 24 minipunches, 2 mm diameter each, (see Figure 1). Compression force used to obtain 9 mg minitablets was maintained at 20 kN for all batches with fill-shoe speed adjusted to 25 rpm.

Figure 1: Special tooling for minitablets (IMA, Ozzano, Italy). (FIGURE COURTESY OF THE AUTHOR) 

. The coating process of 10 kg of minitablets for each batch was carried out in an IMA GS 25-L solid wall pan (see Figure 2). Two different coating formulas were used (see Table I): a traditional one containing single components to be previously dispersed in water (A) and the ready-to-use pigment dispersion WAS (B). 

Table I: Comparison of the two coating formulas. 

Coating dispersion A was prepared by homogenizing, in a separate vessel, pigments, plasticizer, and part of the solvent for 10 min. The dispersion obtained was then added to a main vessel containing Eudragit RL/RS and the remaining water and stirred continuously. Coating dispersion B was prepared by simply adding WAS to the main vessel containing Eudragit RL/RS and solvent without the need of predispersion. The coating suspension was than stirred continuously. Both formulas were tested at different polymer amounts expressed in total solids' weight-gain percentage.

Figure 2: Schematic of the solid-wall coating pan. (FIGURE COURTESY OF THE AUTHOR) 

The coating process was started by loading the minitablets in the pan manually and without any drum rotation. A preheating phase was carried out at the minimum drum speed to ensure the proper temperature of the cores prior to the spray phase. When the minitablets reached the proper temperature, all the coating dispersion was sprayed by a dedicated compressed-air-fed gun. At the same time, the pan speed was adjusted to the ideal value to avoid the cores sticking or being damaged; mixing was ensured by six shark-fin-type mixing baffles placed in the center of the drum. 

During the process, drying was ensured by a special paddle system blowing the correct amount of inlet air at the desired temperature and humidity. Cores were dried from the surface and from the bottom, improving the uniformity of evaporation. 

Table II: Optimized process parameters for formula A. 

When all the coating dispersion was sprayed, a drying step was performed to reduce the humidity of the cores, then the minitablets were unloaded manually after a short cooling phase. Optimized process parameters for formula A and B are described in Table II and III, respectively.

Table III: Optimized process parameters for formula B. 

 The coated minitablets were then filled (n = 42) into size 0 hard gelatine capsules using a Zanasi Lab capsule filler (IMA Ozzano, Italy), which has a maximum speed of 12000 capsules/h, equipped with a minitablets dosing group (Figure 3). The dosing group is composed of a minitablets hopper, a special wheel with a predefined number of holes (the dimensions of which are dictated by the size of the minitablets) and a drum with pushers for the discharge of the minitablets in the capsule body.

Figure 3: Capsule filler for the minitablets dosing group (IMA, Ozzano, Italy). (FIGURE COURTESY OF THE AUTHOR) 

The minitablets enter into the wheel's holes by means of vacuum. The wheel rotates, the vacuum is cut off and the minitablets are dosed into the capsules by gravity fall. 

The minitablets dosing unit is equipped with a camera to check that all the holes are filled with minitablets. If one or more minitablets are missing, the corresponding capsule is rejected. 

 The minitablets were characterized as follows:  

 The weight variation was determined weighing 20 minitablets. 

 The friability of minitablets was determined by introducing 6.5 g of them in a Roche friabilator (Erweka, Düsseldorf, Germany) at a rotational speed of 25 rpm. After 100 rotations, the minitablets were sieved over a 250 µm sieve. The friability value was calculated as the percentage of the final weight after sieving to the initial weight of the minitablets. 

 The crushing strength was tested (n=20) using a TBH 200 (Erweka). 

 The coating surface and the cross-section slice of minitablets were examined using scanning electron microscopy (ESEM-FEI QUANTA 200, FEI, Eindhoven, The Netherlands); the samples were previously sputter-coated with gold. 

 The analysis of the drug content was carried out by dissolving  30 mg of minitablets in 100 mL of pH 7.4 buffer. The amount of drug was then spectrophotometrically determined after 24 h (UV2 Spectrometer, Unicam, Cambridge, UK) at 271.8 nm. Each sample was analyzed at least in triplicate. 

 dissolution tests were performed using the USP Apparatus 2 (paddle) (Pharmatest, Steinheim, Germany) rotating at 50 rpm. As a dissolution media, 900 ml of pH 7.4 phosphate buffer was used at a temperature of 37 ± 0.5 °C. Samples of uncoated and coated minitablets containing an amount of drug chosen to assure sink conditions (C < 0.2 Cs) were added to the dissolution medium. The solution was filtered and continuously pumped to a flow cell in a spectrometer (UV2 Spectrometer, Unicam, Cambridge, UK) The amount of drug dissolved was analyzed at 271.8 nm. The dissolution tests were performed at least in triplicate. 

The target of this study was to investigate from the industrial point of view the possibility of manufacturing a final solid-dosage form containing a multiparticulate system providing a controlled release of the drug. The work was divided into three steps: first the feasibility of producing minitablets by direct compression was evaluated, secondly the coating of the produced minitablets in a solid wall pan was studied, and finally the process of filling the minitablets into hard gelatine capsules was checked. 

Four batches of minitablets containing an increased amount of TH from 50% to 80% (w/w) were produced and characterized in terms of weight variations, friability, and crushing strength. These four batches were labelled mT1, which contained 50% (w/w) TH; mT2, which contained 60% (w/w) TH; mT3, which contained 70% (w/w) TH; and mT4, which contained 80% (w/w) TH. Because the second step of the production was a film coating in a pan, particular attention had to be paid to friability and crushing strength to assess that the cores had enough mechanical resistance to tolerate the stress involved in pan coating, especially in production size. 

The results showed that all the formulations considered were successfully compressed into minitablets having small weight variations, high crushing strength, and very low friability (see Table IV). Among the produced batches, formulation mT4 was selected for the subsequent steps due to its high TH content, which enables the production of a 300-mg dosage unit. 

Figure 4: Scanning electron microscopy images of the uncoated batch labeled "mT4" that was prepared by the authors. The images show formula A at different coating levels and formula B at different coating levels. (FIGURE COURTESY OF THE AUTHOR) 

The second part of the study was aimed at demonstrating the feasibility of coating minitablets in a solid-wall pan. The first challenge was to avoid coating defects. In fact all coating processes, if not properly controlled, could lead to defects such as sticking or twinning of two or more cores together, nonuniform distribution of the coating, peeling of the coating surface, weakness of the coating, especially cracking on the tablet edges, and coating roughness/orange-peel effect. For this reason, preliminary tests (not reported) were performed to find the best coating conditions. Different combinations of pan speed (range 15–24 rpm), atomization pressure (range 0.8–1.5 bar), inlet air temperature (range 55–65 °C), inlet air quantity (range 200–250 m3/h); cores' temperature (range 30–35 °C) and spray rate (range 30–45 ml/min.) were investigated to create process harmony. 

The optimal combination of parameters found for the traditional coating formulation A are reported in Table II. Subsequently, the possibility of replacing the standard coating preparation with a ready-to-use system (formulation B) was tested with the aim to reduce both total process time (starting from the coating preparation) and the need for ancillary equipment. 

Comparing Tables II and III, it is evident that for both formulation A and B the coating conditions were identical, which demonstrates the interchangeability of the traditional formula containing the single components with the ready-to-use pigment dispersion at least regarding the process parameters. Two different coating levels were tested for both coating formulas to check the influence of the amount of polymers on drug release. 

To assess the coating quality, the surface characteristics and the cross-section slides of the coated minitablets were observed using scanning electron microscopy at different coating levels for either minitablets coated with the standard formula A and the innovative formula B (7). The results in Figure 4 show that a complete and uniform layer was achieved using both formulas even at low coating levels.

Figure 5: Dissolution of uncoated and coated minitablets A and B at different total solid-weight gain. TH is theophylline, Awg is weight gain for formula A, and Bwg is weight gain for formula B. (FIGURE COURTESY OF THE AUTHOR) 

Dissolution tests were then performed on mT4 uncoated and mT4 coated with formula A and formula B at two different coating amounts. These tests aimed to confirm the absence of coating defects such as cracking and to evaluate the effect of the amount of polymer on drug release. In addition, the equivalence of both coating formulations was checked. As expected, Figure 5 shows that the uncoated mT4 released 100% of the active ingredient in less than 90 minutes, clearly demonstrating the need for film coating to achieve a sustained release. The good coating quality was demonstrated by obtaining a sustained release behavior using both formulas even at the lowest polymer amount. Furthermore, the release data showed the influence of the coating level on TH release. The dissolution profiles of both formulas at the same coating levels were not significantly different, which is evidence that  the ready-to-use system is equivalent to the traditional one.

Table IV: Characteristics of the four formulations used in minitablet compression.  

Finally, coated minitablets were filled into hard gelatine capsules size 0. The maximum speed achieved for the production was 12000 capsules/hour, representing 80% of the highest speed of the capsule filler when dosing standard volumetric pellets. Production was successfully carried out for 2 h without any rejected capsules, thereby confirming the system reliability. 

Minitablets containing a high amount of TH, low friability, and high crushing strength were obtained by direct compression and then coated in a solid-wall pan. The use of a ready-to-mix pigment dispersion represents an interesting alternative to shorten coating-suspension-preparation time when compared with traditional formulations. The coated minitablets were then successfully filled into hard gelatine capsules. In conclusion, the results demonstrated that the innovative approach used in this study could be potentially useful for the industrial-scale manufacture of a sustained-release, oral-solid-dosage form. 

The authors would like to thank Rosamund Cervo for her kind support in checking the English. 

The authors describe an alternative approach to compressing and coating minitablets for use in a sustained-release, solid oral-dosage form. This article is part of a special Drug Delivery issue.

Minitablets Coated in a Solid-Wall Pan for Theophylline Sustained-Release Capsules

Liquid-filled hard-capsule technology is well-established for formulation and manufacture (1). In addition to being a convenient product format, it is well recognized as offering a wide range of applications, including for use in compounds with low bioavailability or with low-dose uniformity issues. Liquid-filled hard capsules can provide products with various release profiles and may be used to overcome stability concerns such as moisture uptake. Capsule-in-capsule formats represent a significant further development in the product type. In this format, a formulated capsule (i.e., liquid-filled or dry-filled) is nested in an outer liquid-filled capsule. Each capsule achieves a particular formulation aim (e.g., prompt or controlled release). This article will describe formulation approaches and provide examples of 

 investigations of a capsule-in-capsule system that enables the presentation of multiphase or multidrug formulations in a single dosage unit.

Figure 1 suggests a range of possible formulation aims that capsule-in-capsule technology may achieve. This drug-delivery system has beneficial flexibility. The formulation aims for each component capsule may be achieved by selecting the carrier suited to the particular delivery goal (e.g., a prompt or slow-release profile). Available carriers include the following options:  

Nonaqueous fluid, possibly thickened with a thixotropic agent (e.g., colloidal silicone dioxide) 

Thermosoftening excipients with various solution characteristics that can be matched to 

The resulting formulations may be nonaqueous solutions, suspensions, semisolid matrices, or microemulsions, all of which are suited to liquid-filled hard capsule formats. 

Figure 1: Target areas for the capsule-in-capsule platform. CPD is compound. (FIGURE IS COURTESY OF THE AUTHOR) 

Common suitable nonaqueous fluids include pharmaceutical or nutraceutical oils (e.g., medium-chain triglycerides or soya oil), which are selected according to technical or market needs. Individual materials may require the inclusion of antioxidants or filling under nitrogen to minimize oxidation. Thermosoftening carriers are commonly fully synthetic [e.g., the high-molecular-weight polyethylene glycols (PEGs) or poloxamers], selected according to melting point and viscosity. Thermosoftening carriers also may be semisynthetics such as the macrogolglycerides (e.g., lauroyl, linoleoyl, and caprylocaproyl), selected according to melting point and hydrophilic–lipophilic balance (HLB). Varying carrier HLB values (i.e., 2–14) enable the solubility profile of the active ingredient to be matched to the required release profile. For example, a soluble compound may be formulated with a low-HLB carrier to give a slow-release profile, and a high-HLB material may improve the dissolution rate of a poorly soluble active. 

Self-emulsifying drug-delivery systems also may be used. These systems generally contain an oil and a surfactant and also may contain a cosurfactant. The systems are suitable for compounds whose log 

 value is 2–4. For compounds with higher log 

 values, oils should be used. The related microemulsion products are homogeneous, clear fluid systems with small droplet size (i.e., 100–600 nm) that comprise an aqueous phase, an oily phase, a surfactant, and a cosurfactant. They form spontaneously on gentle mixing, are thermodynamically stable, and have low viscosity. The microemulsions include nonionics as surfactants, derivatives of propylene glycol or polyglycerols as cosurfactants, and vegetable oils or fatty acid esters as oily components. The proportions required for the various components may be established using pseudoternary phase diagrams to map out the microemulsion domain where the fluid remains clear. Typical applications are in increasing bioavailability, varying the release profile, improving stability, and handling potent or toxic compounds.

Figure 2: Dissolution profiles for series 1 doses. The outer Labrafil (GattefossÃ©, St-Priest, France) capsule contains caffeine (caff.), and the inner polyethylene glycol 6000 contained nicotinamide (nic.). Results after 0, 2, 3, and 5 months' storage are shown. (FIGURE IS COURTESY OF THE AUTHOR) 

The author outlines how to choose carriers and capsule shells according to dosage requirements and intended use. This article is part of a special Drug Delivery issue.

Pharmaceutical Technology-11-01-2010

Considerations and Approaches for Filling Dry-Powder Inhalers

Drug Targeting by Diagnostic Ultrasound Contrast

Sustained-Release Injectable Drug Delivery

Minitablets Coated in a Solid-Wall Pan for Theophylline Sustained-Release Capsules

Capsule-in-Capsule Technology

Melt Extrusion: Shaping Drug Delivery in the 21st Century

Transdermal Delivery of Vaccines and Therapeutic Proteins