Table of Contents



Optimization of Protein Release from a Bone Fracture Repair System

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By Amanda Walker

Mentored by Ashley Aston and Prasad Shastri

Table of Contents

I. Abstract

II. Introduction

a. Market Analysis

b. Problem Assessment and Goals

c. Proposal

d. Industrial Overview

III. Methodology

a. Product Specifics

i. Components

ii. Manufacturing

iii. Degradation Process

b. Constraints and Alternatives

IV. Results

a. Data and graphs

b. Safety issues and Economics

V. Recommendations and Conclusion

Abstract

It has been estimated that there are about 1.5 million bone fracture diagnoses per year in the United States. The estimated cost of a bone graft or implant is between $4,000 and $26,000 dollars (depending on length of hospital stay, surgeon and assisting physician fees, and other variable charges), while the estimated cost of a spinal fusion procedure is about $50,000. Degradable polymer implants have many benefits over other types of treatment, but the best type has yet to be discovered. In order to create the optimum bone cement, Dr. V. Prasad Shastri is developing a unique polyanhydride that could be injected into the fracture site as a paste and polymerized in vivo using blue light from a dental lamp. The material is made of a newer type of crosslinked polymer that has improved mechanical strength and degradation characteristics. Furthermore, the material will be able to deliver proteins to the surrounding tissue, such as growth factors, that will enhance the rate and quality of bone growth. My individual contribution to this ongoing project has been the testing of new porogens and porosities in order to optimize the release rate of the therapeutic protein. Ideally, the protein will be released continually over the course of the degradation. During this year, I analyzed the release of protein from the product using two different porogens, NaCl and gelatin. The results of the NaCl study seemed to indicate that it would not be a beneficial component of the product, but the results of the gelatin porogen study indicated that it could potentially be used successfully.

Introduction

Market Analysis

Bone fracture is not a new health problem, but it is by no means a trivial one. It has been estimated that there are about 1.5 million bone fracture diagnoses per year in the United States. This estimate includes about 42,409 skull fractures, 255,826 facial fractures, and 700,000 spinal fractures. It is the cause of about 11,696 deaths per year.[i] One of the main causes for the high number of fractures is the United States’ high incidence of osteoporosis. About 1 in 3 U.S. women and 1 in 8 U.S. men are or will be affected by this condition, which is responsible for $13.3 billion in healthcare costs per year.[ii] The mean length of stay in British hospitals for rib, sternum, or spinal fractures is 10.8 days.1 The repair of these types of fractures, which cannot be fixed with a typical cast, may require bone grafts, synthetic implants or cements, or spinal fusion procedures. The estimated cost of a bone graft or implant is between $4,000 and $26,000 dollars (depending on length of hospital stay, surgeon and assisting physician fees, and other variable charges),[iii] while the estimated cost of a spinal fusion procedure is about $50,000.[iv]

Problem Assessment and Goals

There are many current methods for correcting bone defects. The three main categories (excluding casts) are autograft, allograft, and synthetic implant materials. Autograft transplants are taken from the body of the patient needing the surgery. The bone is removed from one area and transplanted into another. This is the preferred treatment, but the amount of tissue a doctor can harvest is limited.[v] Allografts are pieces of bone taken from a cadaver. The main problem with these is the potential for disease transmission.5 Synthetic implants, therefore, have advantages over autografts and allografts.

Synthetic bone implants can be classified as degradable and non-degradable. The latter frequently takes the form of metal or ceramic nails and screws, or of poly (methyl methacrylate) (PMMA) bone cement. Non-degradable materials can cause stress on the surrounding bone, and eventually may break due to fatigue.[vi] PMMA has certain advantages, such as introduction to the injury site as a paste that can be chemically hardened in the operating room. However, the heat generated during this reaction can cause tissue necrosis of up to 5 mm surrounding the implant. The ensuing tissue inflammation can impair the proper healing of the bone.[vii]

Degradable implants reduce some of the problems caused by non-degradable implants. An example is shown in Figure 1. However, inadequacies still exist. The most commonly used degradable bone plates and bone screws are constructed of poly (lactic acid) and poly (glycolic acid) copolymers. They have the advantage of eliminating the need for a second surgery. However, these copolymers undergo bulk erosion. This causes a loss in their mechanical integrity because the inside erodes at the same rate as the outside. Finally, the degradation products may cause an increase in the local pH, which can cause inflammation that impairs healing.7

Our goal, therefore, is to create the optimum bone cement. In order to reach this goal, several factors must be taken into account. First of all, the material must degrade at the proper rate of bone healing. Secondly, it must remain strong for the time that it takes the bone to grow in. Third, the material needs to be biocompatible, causing little to no detrimental effects on the surrounding tissue. Fourth, it needs to be easily cured in vivo. Finally, the material should be able to deliver proteins to the surrounding tissue, such as growth factors, that will enhance the rate and quality of bone growth.

Proposal

To accomplish this goal, Dr. V. Prasad Shastri is developing a unique polyanhydride that could be injected into the fracture site as a paste and polymerized in vivo using blue light from a dental lamp. The curing process takes just two and a half minutes. The material is made of a newer type of crosslinked polymer that has improved mechanical strength and degradation characteristics. The material degrades by surface, not bulk, erosion, thus maintaining its mechanical integrity throughout the healing process. The differences in surface and bulk degradation are demonstrated in Figure 2.

In addition, two additives will be mixed into the polymer which will assist the healing of the defect. The first, bone morphogenetic protein, will be capable of spreading from the cement and into the surrounding bone tissue. The second, a porogen, will dissolve shortly after the polymer hardens. The holes that it creates (called the porosity) enable the protein to float out of the polymer and into the surrounding tissue, where it will react to speed the healing process.

My individual contribution to this ongoing project has been the testing of new porogens and porosities in order to optimize the release rate of the therapeutic protein. Ideally, the protein will be released continually over the course of the degradation. If too much protein is released in the first few days after surgery, then none will be left to release toward the end of the implants degradation. This problem is called a “burst effect”.

Industrial Overview

The International Standardization Organization set the standards for implantable medical devices in a set of documents called ISO 10993. These documents set the standards for such issues as interactions with blood, local effects after implantation, and general toxicity. In addition, there are several documents that relate directly to implantable, degradable devices. Part 13 is titled the “Identification and quantification of degradation products from polymeric medical devices”, while part 16 is titled “toxicokinetic study design for degradation products and leachables”, and part 17 is “establishment of allowable limits for leachable substances.” [viii]

In 1997, the FDA established its own guidance document for “testing biodegradable polymer implant devices”. It requires:

- a quantitative analysis of the composition and molecular structure of the components

- explanation of the physical properties and thermal properties of the component materials

- a protocol for strength testing of the device in a physiological solution at 37 degrees Celsius

- shelf life of the final product

- biocompatibility, including identification of the degradation products and their metabolic pathways.

Researchers must adhere to these requirements in order for a polymeric implantable device to receive FDA approval.[ix]

Products similar to ours can be found on the market today. A line from a 1993 patent states, “it is therefore an object of the present invention to provide a method for preparation of highly pure anhydride copolymers having a controlled composition, especially for use in biomedical applications.”[x] This patent allowed Guilford Pharmaceuticals to market a chemotherapeutic device called Gliadel®, which has since revolutionized the post-operative prognoses for glioblastoma patients. The copolymer consists of poly bis(p-carboxyphenoxy)propane: sebacic acid in a 20:80 ratio. It can be used to deliver a potent chemotherapeutic agent called BCNU. The product comes in a wafer form, which is placed in the hole left by the tumor extraction (see Figure 3). After the surgery, the wafers begin releasing their drug, killing most leftover tumor cells. This ingenious product can be used as a model for other types of implants that use anhydride copolymers.

The ETEX corporation also manufactures a similar product. ETEX utilizes a bioresorbable, biocompatible, poorly crystalline apatitic (PCA) calcium phosphate that stimulates osteoclast activity. The kit comes with additives that can alter the resorbability and strength of the material. It also can be injected during surgery.[xi] However, the material is not as strong as our crosslinked copolymers. The compressive strength of the ETEX bone substitute is 12 MPa.[xii] Crosslinked MSA and MCPH are 34 and 39 MPa, respectively. [xiii] Trabecular bone’s compressive strength is between 5-10 MPa, while cortical bone’s is between 130 and 220 MPa. Therefore, the crosslinked anhydrides are more like bone than the ETEX bone substitute. Furthermore, the chemically induced curing reaction that takes place in ETEX bone substitutes takes between 15 and 20 minutes,12 as opposed to our two and a half.

Methodology

Product Specifics

Components

The materials that are used to create the polymer are:

▪ two monomers: sebacic acid (MSA) and 1,6-bis(p-carboxyphenoxy) hexane (MCPH)

▪ Porogen

▪ Protein (ultimately, bone morphogenetic protein)/sugar protector

▪ Camphorquinone photoinitiator

▪ poly (ethelene glycol) diacrylate (PEGDA)

The monomers, MSA and MCPH, have methacrylate end groups. These are very reactive and help to ensure the photopolymerization of the substance. They produce highly crosslinked polyanhydride networks, which are stronger by nature than linear polymers (see Figure 5). In fact, even after the copolymer has degraded by more than 50%, it still retains over 70% of its original tensile modulus.[xiv] The polyanhydride is a hydrophobic material that degrades by hydrolysis of its anhydride linkages, ensuring that the material undergoes surface and not bulk erosion. Furthermore, since MCPH is more hydrophobic than MSA, the degradation rate can be controlled by varying the ratio of the two monomers. In this study, we used a ratio of 70:30 MCPH to MSA. The crosslinking that occurs between the monomers provides a mechanical strength that is not present in linear polymers.

The porogen is the aspect of the project that I have investigated over the past year. Porosity is an important aspect in this project, because the holes are so necessary for its success. Primarily, it is important that the proteins are able to escape from the polymer. Secondly, the gaps will allow the bone cells to penetrate the implant as it degrades.[xv] The porosity must be modulated in order to obtain the optimal protein release rate. Ideally, the same amount of protein would be released every day, so that a graph of time versus cumulative quantity of protein released would be linear. The release would continue until the implant was fully dissolved. However, this is difficult to maintain. Often, too much protein is released initially, leaving none for later. This is called a “burst effect” and it is something that needs to be eliminated. The burst effect is demonstrated in Figure 6. The porogens that I tested were NaCl, which is commonly used for this purpose, and gel microspheres, which were created in the Shastri lab.

Because bone morphogenetic protein (BMP) is expensive, we chose to use a substitute for the testing phases. Horseradish peroxidase (HRP) is cheap and easily detectable, so we decided to use it for the degradation studies. However, all proteins are easily denatured when they leave their natural environments. For this reason, it is necessary to protect the protein by mixing it with β-lactose and gelatin. This not only protects it from light, but also makes its quantification more accurate.[xvi]

The camphorquinone initiator is the substance which causes the hardening (curing) of the polymer. It is widely used in dentistry, because it eliminates the need for detrimental UV light. Camphorquinone, when exposed to blue light, forms free radicals. These radicals attack the carbon-carbon double bonds present in the acrylate groups of the MSA, MCPH, and PEGDA. The chain reaction that ensues causes the polymer to harden in about two and a half minutes. The prevalence of double bonds in the mixture ensures the crosslinking of the monomers with each other.

Furthermore, the blue light that causes the reaction also turns the camphorquinone white and into a scattering agent. This allows for very good penetration of the light. For example, with this system, Anseth et al. photocured a 4 cm long screw by shining light only on the head. The scattering of the blue light allowed polymerization to occur even in the dark regions of the screw, like the threads. The PEGDA, with its extra acrylate groups, helps to achieve polymerization at greater depths.13 It also makes the uncured mixture less viscous and easier to manipulate.

Manufacturing

In order to make the samples, it was first necessary to formulate a stock of the HRP/ β-lactose mixture. Before mixing in the protein, the lactose was granulated to a diameter of 250 micrometers. Granulation is a process by which the mixture is forced through a sieve with holes of a certain diameter. Using the Biotek Synergy HT spectrophotometer with KCB software, protein uniformity assays were conducted on the stock to ensure that the protein was uniformly dispersed within the powder (see Appendix A). Ten different powder samples were taken, diluted in a solution, tested for the amount of protein that they contained, and then compared against each other

Next, the porogens had to be made. For the NaCl porogens, the only preliminary preparation was to grind the salt crystals. The diameter of the salt porogen was determined to be less than 106 microns. The diameters of small particles are measured using a set of sieves. They are stacked so that the sieve with the holes of the largest diameter is on the top and the sieve with the smallest is on the bottom. The powder is poured onto the top sieve and the sieve that it falls to is recorded. The diameter is reported as a range: the smallest it could possibly be is the diameter of the holes of the sieve below it, and the largest is the diameter of the holes above it.

The gel porogen manufacturing process was a little more complex. The protocol involved some organic chemistry techniques that were conducted over the course of 2 days. After they had been fabricated, the diameter of the gelatin microspheres was determined to be between 106 and 180 microns.

Once the protein stocks were ready and the porogens had been fabricated, I began making the samples for the degradation studies (see flowchart in Appendix B). We decided to try samples of 25% porosity for the initial degradation studies. First, I mixed 500 mg of the NaCl stock with 200 mg of the HRP/sugar stock. Then, I added 100 mg of the PEGDA and mixed on a watchglass. Next, 770 mg of MCPH were added and the ingredients were mixed, and finally, 330 mg of MSA were added and mixed. This puttylike substance was then placed into the cylindrical mold and illuminated for 2 minutes and 30 seconds with a blue dental lamp such as the one shown in Figure 7. The hardened sample was then placed into a labeled vial along with 10 mL of PBS (as shown in Figure 8). Four samples of each porogen and four control samples were made in this manner. The manufacturing of the gel porogen samples followed the same protocol, but using 500 mg of gel microspheres instead of 500 mg of NaCl. The average height of the samples was 8.24 mm, average width was 3.86 mm, and average mass was 231.55 mg.

Degradation process

The process of testing the protein release included making samples with varying porogens. In the beginning, we also made some with no protein, as a control. It was quickly determined that the controls were not releasing protein and so the studies with the porogen samples were initiated.

In order to track the release of protein, we had to quantify it using the BioTEK Synergy HT spectrophotometer with KCB software. Each sample was placed in a vial containing 10 mL of phosphate buffer solution (PBS). Every time the PBS was changed, a sample of it was saved to see how much protein it contained. Samples were typically taken every 1-3 days. Using an enzyme kit that turned the HRP blue, I was able to detect the amount of protein in each PBS sample with the aid of the spectrophotometer. The initial expectation was that all of the protein would be released within about 40 days, so the plan was to stop data collection at that point. The data was recorded and saved in an excel file for later analysis.

Apart from the protein release studies, there are several other engineering analyses that must be performed on these types of materials before they are ready to be commercialized. For example, the mechanical characteristics are important to the success of the implant in vivo. The tensile modulus of the material has been tested using a dynamic mechanical analyzer (DMA-7; Perkin Elmer, Norwalk, CT). Sample slabs were polymerized and degraded in PBS (pH 7.4, 37 C۫). Experiments were performed to measure the moduli as a function of temperature and degree of degradation by applying a sinusoidal stress at a frequency of 1 Hz. Compression testing was performed using the Micro MT100 (MonoResearch, Williamsville, NY). Small cylindrical samples were placed under a 5 kg compression that increased at a rate of 1mm/min until failure. 13

Furthermore, before marketing the product it will be necessary to conduct studies on the depth of curing. Fourier transform infrared (FTIR) spectroscopy will allow the researcher to look for the disappearance of carbon-carbon double bonds. This indicates polymerization, since the acrylate double bonds link together to form single carbon-carbon bonds.

Constraints and Alternatives

My most serious constraint for this project was the time allotted to complete it. Each degradation study took longer than expected, and several problems were encountered. For example, over 10 hours were lost in testing a protein that was found not to work because its background signal was too high for it to be detected. On top of that, new samples had to be made with a different protein, which cost about 2.5 hours. Learning to run the assays was also a difficult task. A few beginner mistakes were made, each costing between 1 to 2 hours. Altogether, between 18 and 22 hours of work were lost. Learning to analyze the data was also a time consuming process.

On the other hand, money was not a serious constraint, due to the fact that I was working through a lab. Dr. Shastri had already taken care of writing grants and obtaining funding. All I had to do was use the chemicals and resources available to the best of my ability.

Some different proteins and porogens were considered as alternatives. Many previous studies used different types of porogens, such as fluorescein bovine serum albumin (FITC-BSA),[xvii] or an ammonium bicarbonate gas forming porogen.[xviii] However, the FITC-BSA is not applicable to our project because its acidity could harm the physiological environment, and we did not have time to try the ammonium bicarbonate method. Other potential alternatives were different polymeric structures. Many researchers are using poly (lactic acid) and poly (glycolic acid) copolymers. However, it has been shown that these systems rapidly lose their mechanical strength due to inadequacies in their degradation styles. 13

Results

Data and graphs

After conducting the degradation studies with the NaCl and gelatin porogens, I found that the gelatin porogen displayed a more favorable protein release rate. For the NaCl porogen degradation study, only 11.2% of the protein that had initially been incorporated into the polymer had been released by day 55. I continued this study longer than the expected time of 40 days in hopes that more protein would be released.

In the gel porogen degradation study, about 56.7% of the protein that had initially been incorporated into the polymer had been released by day 35, which is a huge improvement over the NaCl porogen’s release profile. However, 45.2% of the protein had been released by day 3 of the gel degradation study. This is a prominent burst effect which will be necessary to reduce before manufacturing. I would have liked to continue the study longer, but I ran out of time. The results of both studies are shown in Figures 9 and 10.

Safety issues and economics

A thorough analysis of the safety issues involved in the design, manufacture, and use of this product can be seen in the designSafe report. Most hazards can be reduced by perfection of the design, proper implantation, and precautions taken by the user. The main issue is the degradation of the product. If it degrades too fast, it will release too much acid into the blood stream. Speeding of degradation could possibly be induced by fatigue, crushing, or uncommon reactions within the body. The product is still in the testing phases to ensure its safety when it is released to the market.

At this point, it is difficult to estimate the cost of the product. Cost analysis would include the price for the individual monomers, the protein/sugar, the porogen, the syringe that would administer the drug, the manufacturing labor costs, the packaging costs, the research and development costs, and any profit that the company hoped to make. The only way to currently estimate the cost of the product would be to compare it with similar products and situations. Of course, the area and severity of fracture must be taken into account as well.

In February of 2006, PR Newswire Europe Limited, released a report that stated the cost of treatment with Gliadel Implants to be around $7,587.42 (United States currency).[xix] This is only a rough estimate, but since the components and usage is similar, it seems feasible. A 1997 study found that the average cost of a metallic bone implant for an ankle fracture was $8,064.26, while the same fracture treated with biodegradable poly (glycolic acid) screws cost the patient $6,258.95. Based on these analyses, I would estimate that the treatment would cost somewhere between $5,000 and $10,000 dollars.

Recommendations and Conclusion

During this year, I analyzed the release of protein from the product using two different porogens, NaCl and gelatin. The results of the NaCl study seemed to indicate that it would not be a beneficial component of the product, but the gelatin porogen study indicated otherwise. Nevertheless, more studies on the gelatin are necessary before deciding to use it. For example, I used 25% porosity, and obtained a less than optimal protein release profile. Perhaps another porosity percentage, or even another porogen altogether, would provide the optimal protein release.

Furthermore, I would recommend that the degradation studies take place over a longer period of time. A collection of data about the time it takes a bone to heal is shown in Figure 11. The average is about 60 days. Therefore, a standard should be made that the protein be released linearly over a longer period of time (our initial speculation was 40 days). This will ensure that the product is successfully cooperating with its environment.

In conclusion, we propose a new biodegradable bone fracture implant. The implant will degrade slowly as the patient’s bone grows in around it. The material is stronger than competing products and it retains its strength as it degrades, which is uncommon. It can be hardened quickly and harmlessly during surgery, which is a benefit that many of our competitors cannot boast. Furthermore, it releases therapeutic proteins that assist the healing process. The release of the protein will optimally be complete and continual over the course of the bone degradation.

Appendix A

Process Diagram for Protein Quantification

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Appendix B

Flowchart for making the polymer

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Appendix C

Time Spent on Project

|date |time spent (hours) |task type |

|late october |3 |research of papers |

|late october/early november |1 |interviewing Ashley |

|3-Nov |0.5 |planning and labwork |

|7-Nov |1 |labwork |

|8-Nov |1 |labwork |

|10-Nov |1.5 |labwork |

|29-Nov |2.416666667 |labwork |

|1-Dec |1.75 |labwork |

|6-Dec |1.5 |labwork |

|8-Dec |1.667 |labwork |

|9-Jan |3.5 |labwork |

|10-Jan |2 |labwork |

|13-Jan |1 |labwork |

|19-Jan |3 |labwork |

|25-Jan |0.833333333 |labwork |

|26-Jan |3.333 |labwork |

|27-Jan |0.333 |labwork |

|28-Jan |0.333 |labwork |

|29-Jan |0.333 |labwork |

|30-Jan |0.333 |labwork |

|31-Jan |2 |labwork |

|1-Feb |0.333 |labwork |

|2-Feb |1.5 |labwork |

|3-Feb |0.333 |labwork |

|7-Feb |2.5 |labwork |

|8-Feb |0.166666667 |labwork |

|9-Feb |2.5 |labwork |

|10-Feb |0.16667 |labwork |

|13-Feb |0.16667 |labwork |

|14-Feb |0.416666667 |labwork and planning |

|15-Feb |0.1667 |labwork |

|16-Feb |1.5 |labwork |

|17-Feb |0.16667 |labwork |

|20-Feb |2.6667 |labwork |

|23-Feb |1.3333 |labwork |

|24-Feb |2.3333 |labwork |

|27-Feb |0.33333 |labwork |

|28-Feb |1.75 |labwork |

|1-Mar |0.16667 |labwork |

|2-Mar |2.75 |labwork |

|13-Mar |2 |labwork |

|14-Mar |2 |data analysis |

|15-Mar |0.5 |data analysis |

|16-Mar |3.5 |data and labwork |

|23-Mar |0.16667 |labwork |

|27-Mar |0.16667 |labwork |

|28-Mar |1.5 |labwork |

|2-Apr |3.5 |data |

|3-Apr |1.25 |lab and data |

|4-Apr |3.16667 |labwork |

|5-Apr |3.083 |data analysis |

|6-Apr |3 |data |

|8-Apr |2 |data |

|total hours ---> |79.41435333 |  |

Appendix D

Ideation Workbench

Ideation Process

 

Project Initiation

Project name: Porosity in Biodegradable Photopolymerizable Anhydrides

Project timeline:

November 29th- Christmas Break- Practice salt degradation study. My regular lab hours will be Tuesday and Thursday from 10 AM-12 PM.

January 7th- February- Trial run on real degradation study. Will chose regular lab hours for second semester.

February-April- Real degradation studies

April- Poster and finishing touches

 

Project team: Amanda Walker, working with graduate student Ashley Aston and mentored by Prasad Shastri

 

Innovation Situation Questionnaire

 

Brief description of the situation

I am studying a biodegradable polymer that has enough mechanical strength to be useful in bone repair surgeries. Protein released from the polymer will help the bone to heal faster. The polymer degrades as the bone grows in. My studies consist of making samples of the product and collecting data of the amount of protein released as I vary contributing factors like porosity. Optimally, I will get the protein to be released for 3-4 weeks.

 

Things that I need to work on are...

Improve mechanical strength

Reduce time wasted

 

 

Detailed description of the situation

Ideal vision:

Magician- I would have him reduce the amount of time it takes to do the Coomassie assay. Also I would get him to do the protein granulation and gelatin granulation for me.

Mini Problem- The protein doesn't want to last for the full 3-4 weeks or I get a burst effect.

Supersystem

 

What changes to the supersystem(s) would resolve the situation?

The person would not get a broken bone or a bone disease.

What changes to one or more of the subsystems within the system would resolve the situation?

The first porosity I try would be just right to obtain the perfect protein release profile.

 

System name

The light hardened biopolymer

 

System structure

The polymer is putty like before it is hardened with the blue light. So we mix the proteins and the porogens into the putty, then harden it. When the polymer is exposed to water, the porogens dissolve, leaving holes to increase the surface area that the protein can be released from.

 

Supersystems and environment

The system environment is bone. Other nearby systems are cartilage, blood, marrow, etc. The temperature is 98.6 degree celcius. The system will interact with it's environment in that the blood will infuse the biopolymer to allow the proteins to escape.

 

Systems with similar problems

The precedent for this study is a similar substance. It is a wafer containing BCNU, a therapeutic agent. The wafer is placed in the hole left by the removal of a glioblastoma. The BCNU is released to kill the leftover brain cancer cells. Aspects of this problem could be used for my project, but we'd have to change things like the material strength, etc.

 

Input - Process - Output

The problem is getting too much or too little protein release over the time of 3-4 weeks. It is possible to solve the problem by influencing the inputs. I will be varying the concentration and material of the porogen to solve the problem. The output leaving the system is the protein release.

 

System inputs

monomers, proteins, porogens and initiator

System outputs

the polymer

 

Cause - Problem - Effect

 

Problem to be resolved

The problem is getting too much or too little protein release over the time of 3-4 weeks.

Mechanism causing the problem

Having too much or too little porosity.

Undesirable consequences if the problem is not resolved

Run out of protein too early or not enough is delivered. url_IPS/IPS/2244_Other_problems.htm

Ashley is working on the other areas of the project, like making sure the polymer degrades in the right amount of time.

 

Past - Present - Future

 

History of the problem

THe major problem, broken bones and bone disease, has been around forever. The recent problem, getting optimized protein release from biopolymers, has only been around for the past few decades. Crosslinked polymers like this one have no previous data on protein release.

Pre-process time

You have to mix the protein with the sugar, granulate the protein, run a Coomassie assay on it, quantify the porogen or make it if you are using the gelatin microspheres.

Post-process time

Not really possible, since the putty is hardened in the bone during surgery, and then it is not accessible after the surgery. I guess you could get the patient to take something that will get into the blood and modify the degradation, but that might not be safe.

 

Resources, constraints and limitations

How will improvement be measured?

•              Polymerized samples will be degraded in 37 degree PBS.

•              PBS will be changed at allotted times.

•              The old PBS will be analyzed using the Coomassie assay.

•            This will determine how much protein was released over the last time period.

 

Available resources

I have all the material and equipment necessary to run the experiment from the lab.

Allowable changes to the system

Changing the porogen.

Constraints and limitations

•              Constraints

•           Release=3-4 weeks

•           Avoid burst effect

•              Limitations

•           Time- 4 months

•           Money- no real constraints

 

 

 

Problem Formulation and Brainstorming

 

Diagram 1

[pic]

12/1/2005 4:37:25 PM.

 

Find an alternative way to obtain releases protein into the bone environment

 

Synthesize the new system

You could have a system that has protein reservoirs that open as the protein degrades but I don't think it would be as effective.

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In Text References

[i] "Statistics of Fractures." . 15 Mar 2006. . 24 Apr 2006 .

[ii] Erickson, Kelley, Susan Baker, and Jason Smith. "Kyphoplasty—minimally invasive vertebral compression fracture repair - Home Study Program." AORN Journal Nov (2003): 2.

[iii] Christenson, Lisa. "Bone Grafting." Health A to Z. Dec. 2002. Gale Group. 24 Apr. 2006 .

[iv] "Information about Spinal Fusions: 2003." AAOS Department of Research and Scientific Affairs. 2003. American Academy of Orthopaedic Surgeons. 24 Apr. 2006 .

[v] Hedberg, Elizabeth, Andrew Tang, Roger Crowther, Darrell Carney, and Antonios Mikos. "Controlled Release of an osteogenic peptide from injectable biodegradable polymeric composites." Journal of Controlled Release 84 (2002): 137-150.

[vi] Ishaug, Susan L., Genevieve M. Crane, Michael J. Miller, Alan W. Yasko, Michael J. Yaszemski, and Antonios G. Mikos. "Bone formation by three-dimensional stromal osteoblast culture in biodegradable polymer scaffolds." Journal of Biomedical Materials Research 1997: 17-28.

[vii] Shastri, Venkatram Prasad, Robert F. Padera, Peter Tarcha, and Robert Langer. "A preliminary report on the biocompatibility of photopolymerizable semi-interpenetrating anhydride networks." Biomaterials 2004: 715-721.

[viii] "TC 194: Biological evaluation of medical devices." List of technical committees. 2003. International Organization for Standardization. 24 Apr. 2006 .

[ix] "Guidance Document for Testing Biodegradale Polymer Implant Devices." ODE Guidance Documents. 3 Sep. 1997. U.S. Food and Drug Administration. 24 Apr. 2006 .

[x] Domb, Abraham. "Fatty acid terminated polyanhydrides ." USPTO Patent Full-Text and Image Database. 12 Jan. 1993. United States Patent and Trademark Office. 24 Apr. 2006 .

[xi] Lee et al. , Dosuk D. "Bioceramic compositions." USPTO Patent Full-Text and Image Database. 6 Dec. 2005. U.S. Patent and Trademark Office. 24 Apr. 2006 .

[xii] "Alpha-BSM Bone Substitute Material." Product Overview. Sep. 2004. Etex Corporation. 24 Apr. 2006 .

[xiii] Anseth, Kristi, Venkatram P. Shastri, and Robert Langer. "Photopolymerizable degradable polyanhydrides with osteocompatibility." Nature 1999: 156-159.

[xiv] Muggli, Dina Svaldi, Amy K. Burkoth, and Kristi S. Anseth. "Crosslinked polyanhydrides for use in orthopedic applications: degradation behavior and mechanics." Journal of Biomedical Materials Research May 1999: 271-278.

[xv] Mikos, Antonios G., and Johanna S. Temenoff. "Formation of highly porous biodegradable scaffolds for tissue engineering." Electronic Journal of Biotechnology 15 Aug 2000. 24 Apr 2006 .

[xvi] Baroli, Bianca, V. Prasad Shastri, and Robert Langer. "A Method to Protect Sensitive Molecules from a Light-Induced Polymerizing Environment." Journal of Pharmaceutical Sciences June 2003: 1186-1194.

[xvii] Lu, Lichun, Georgios N. Stamatas, and Antonios G. Mikos . "Controlled release of transforming growth factor beta-1 from biodegradable polymer microparticles." Journal of Biomedical Materials Research March 2000: 440-451.

[xviii] Nam, Yoon Sung, Jun Jin Yoon, and Tae Gwan Park. "A novel fabrication method of macroporous biodegradable polymer scaffolds using gas foaming salt as a porogen additive." Journal of Biomedical Materials Research Jan 2000: 1-7.

[xix] "NICE Poised to Deny Brain Tumour Patients Access to Latest Breakthrough Treatments." 22 Feb. 2006. PR Newswire Europe Ltd. 25 Apr. 2006 .

References for Images

[1] From Discovery to Innovation. National Research Council Canada. 25 Apr. 2006 .

[2] Whittaker, Andrew. "Project C5: Studies of the Degradation Rates of Biodegradable Polymers." Honours and PhD Projects. 29 Sep. 2005. University of Queensland. 25 Apr. 2006 .

[3] Neuroanatomy and Neuropathology on the Internet. Department of Neurology University of Debrecen, Hungary. 25 Apr. 2006 .

[4] Huang, Xiao, and Christopher S. Brazel. "On the importance and mechanisms of burst release in matrix-controlled drug delivery systems." Journal of Controlled Release June 2001: 121-136.

[5] "Curing Light Unit." 2005. Bonart. 25 Apr. 2006 .

[6] "Browse Our Patient Education Library." AAOS- Your Orthopaedic Connection. American Academy of Orthopaedic Surgeons. 25 Apr. 2006 .

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Figure 2- materials that degrade by bulk erosion lose their mechanical integrity over time, as is demonstrated on the right hand side of this figure [2]

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Figure 5- cross linked polymers are mechanically stronger than their linear counterparts.

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Figure 4- chemical components of the polymer backbone

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Figure 4- the top data set in this graph demonstrates an undesirable “burst effect”. The bottom data set demonstrates a more desirable, linear release profile. This graph is unassociated with my data, and is only for demonstration purposes. [4]

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Figure 1- a photograph of a degradable implanted polymer [1]

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Figure 5- a blue dental lamp that could be used for curing the polymer [5]

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Figure 6- cylindrical polymer sample in degradation mode

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Figure 3- hole left by the removal of a glioblastoma [3]

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Figure 9- typical time to heal for fractures in different area [6]

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Figure 7- graph of protein release for NaCl porogen degradation study

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Figure 8- graph of protein release for gel porogen degradation study

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