Evaluation of Absorbable Poly(ortho Esters) for Use in Surgical Implants

A. U. Daniels,* Kirk P. Andriano,* W. Paul Smutz,* Melissa K. 0. Chang,* and Jorge Hellert

*Orthopedic Bioengineering Laboratory, Division of Orthopedic Surgery, University of Utah School of Medicine, Salt Lake City, Utah and +Controlled Release and Biomedical Polymers Department, SRI International, Menlo Park, California

Recent reports describe an unfavorable noninfective inflammatory response to acidic degradation products in clinical applications of bone fixation devices fabricated from bulk hydrolyzing polyglycolides and polylactides (PGA and PLA). The work described here suggests that poly(ortho esters) (POEs) offer an alternative. By comparison, hydrophobic POEs degrade predominately via surface hydrolysis, yielding first a com- bination of nonacidic degradation products, followed by alcoholic and acidic products gradually over time. POE specimens proved acutely nontoxic in United States Pharma- copeia tests of cellular, intracutaneous, systemic, and intramuscular implant toxicity.

Hot-molded specimens degraded slowly in saline, retaining 92 % initial stiffness (1.6 GPa flexion) and retaining 80% initial strength (66 MPa flexion) in 12 weeks. Degradation was almost unaffected by decreasing saline pH from 7.4 to 5.0. This demonstrated the relative hydrophobicity of POEs, since incorporation of small amounts of acid within the polymer markedly increases the degradation rate. Degradation rates were increased substantially by dynamic mechanical loading in saline. This may be true for other degradable polymers also, but no data could be found in the literature. Presumably, tensile loading opens microcracks, allowing water to enter.

Solvent cast POE films were strong in tension (30+ MPa tensile yield) and reasonably tough (12-15% elongation to yield). Higher molecular weight films (41-67 kDa) showed no degradation in mechanical properties after 31 days in physiological buffer at body temperature. A 27-kDa film offered similar initial strength and stiffness but began showing mechanical degradation at 31 days. The films showed a decrease in weight with exposure time but no change in either molecular weight or water absorption at 31 days, further supporting the observation that POE degrades by surface hydrolysis rather than by bulk hydrolysis. 0 1994 John Wiley & Sons, Inc.

INTRODUCTION

Background

Metallic implant devices have been used for bone frac- ture fixation for many years. These devices align bone fragments, bring their surfaces into close proximity and control the relative motion of the bone fragments so that union can take place. However, load sharing between the device and bone is in proportion to device and bone structural stiffness. Complete healing of the bone requires normal loads and is therefore prevented as long as the device is present and bears part of the load usually seen by the Also, sudden removal of the device can leave the bone temporarily weak and subject to refracture. On the other hand, some device structural stiffness is

Requests for reprints should be sent to A.U. Daniels, Ph.D., 8 Blackfoot Circle, Wayne, NJ 07470.

Journal of Applied Biomaterials, Vol. 5, 51-64 (1994) 0 1YY4 John Wiley & Sons, Inc. CCC 1045-4861/94/010051-14

necessary to limit bone motion at the fracture site since gross motion is known to result in nonunion.

Consequently, replacing a metallic fracture fixation device with a bioabsorbable polymer or composite device that has an appropriate combination of initial strength and stiffness is of considerable interest because subsequent absorption has two very important advantages. First, as absorption reduces device cross section and/or the materi- als elastic modulus, the load is gradually transferred to the healing bone. Second, because the device eventually will be completely absorbed, surgical removal is not necessary for complete bone healing.

The first polymer used as an absorbable fracture fix- ation device was poly(DL-lactic acid): and since then, many other absorbable polymers and composites have been investigated. Two detailed reviews of this field have been published

In addition, other recent review^^,^ indicate absorbable polymers have promise for use in surgical applications other than absorbable sutures. For example, copolymer films of poly(1actic acid) (PLA) and poly(capro1actone) implanted into canines were successful in preventing

--

52 DANIELS ET AL.

postoperative pleural and pericardial adhesion.' Also, absorbable sheaths of oxidized regenerated cellulose were found to diminish significantly the extent and severity of intertendinous adhesions between the plantaris and Achilles tendons in rabbits.'" And finally, artificial nerve grafts prepared from tubes of poly (glycolic acid) (PGA)/trimethylene carbonate copolymer, and filled with collagen have been used successfully for regeneration of 5-mm gaps in the rat sciatic nerve."

In general, research into candidate materials for use in temporary surgical implants has focused on synthetic ab- sorbable polymers that are hydrophilic such as poly(1actic acid) and poly(glyco1ic acid), which degrade primarily by bulk hydrolysis.I2 However, recent clinical studies of absorbable polymeric bone fixation devices fabricated from PGA and PLA have identified a potentially serious problem. Although fracture healing occurred with these devices, there were reported rates between 7 and 48% of an associated noninfectious foreign body inflammatory response requiring clinical inter~ention.'"'~ Depending on the polymer, the time for manifestation of this local toxic response was from 7 weeks to nearly 3 years after implantation. Presumably, the bulk degradation mecha- nism of these polymers generates large amounts of acidic degradation product^,'^-'^ which may exceed local tis- sue clearance capabilities when end-stage degradation is reached. 18-2n

A hydrophobic absorbable polymer that undergoes lim- ited hydration so that hydrolysis occurs mainly at the surface of the polymer, releasing degradation products gradually over time, has not been evaluated for potential use in surgical implants serving a temporary function. This article describes preliminary studies on the fabrication, sterilization, biocompatibility, mechanical properties, and degradation rates of such a polymer, poly(ortho ester) (POE), fabricated both by hot compression molding and solvent casting.

The results of preliminary investigations suggest POEs will be an important addition to the choice of degradable materials for the design and development of temporary surgical implants. Although POEs are not as strong and stiff as the highly crystalline PGA and PLA materials, their mechanical properties appear sufficient for many low-load applications. In addition, POEs proved easy to process and were acutely nontoxic in standard in vitro and in vivo protocols of cellular, intracutaneous, systemic, and muscle implant toxicity.

MATERIALS AND METHODS

Polymer Background, Synthesis, and Degradation

POEs are a family of synthetic absorbable polymers that have been under development for medical applications for a number of years2' This family of highly hydrophobic absorbable polymers is completely amorphous.22 Hy- drolytic degradation occurs predominately by surface

h y d r ~ l y s i s , ~ ~ in contrast to the bulk hydrolysis of more hydrophilic PGAs and PLAs. POEs have been successfully used as erodible matrices for drug delivery of therapeutic agents.24 Polymer degradation rates can be accelerated by adding acidic excipients to the bulk polymer or retarded by adding basic exc ip i en t~ .~~ Degradation rates also increase with decreasing molecular weight.26

The synthesis of POEs has been described p r e v i ~ u s l y , ~ ~ and is shown in Figure 1. Briefly, the rigid diol, trans-cyclohexanedimethanol and the flexible diol, 1,6- hexanediol, in the desired ratio are added to tetrahy- drofuran and the mixture is stirred until all solids dissolve. A solution of the diketene acetal 3,9- bis(ethy1ene 2,4,8,10-tetraoxaspiro [5,5] undecane) is then added, and the polymerization is initiated by the addition of a catalytic amount of p-toluenesulfonic acid dissolved in tetrahydrofuran. After the initial exotherm subsides, the mixture is stirred for about 2 h, and polymer is isolated by precipitation into a large excess of methanol containing a small amount of triethylamine stabilizer, followed by filtration and vacuum drying at 60 "C for 24 h. Polymers prepared by this procedure typically have molecular weights in the 80-100 kDa range. By giving special attention to exact stoichiometry and using very pure reagents (>99.5%), molecular weights in excess of 180 kDa have been achieved. Lower molecular weight materials (20-30 kDa) can be made where more rapid degradation is desired.

Polymers with a wide range in glass transition temperature (T,) can be synthesized by varying the ratio of rigid to flexible diols (Fig. 2). But the effect on elastic modulus is modest for POEs reported here because the Tg is well above room tempera- ture and the service temperature for implants, that is, 37 "C.

When bulk POE of the type described above is exposed to an aqueous environment, the polymer undergoes an initial hydrolysis of the ortho ester linkages to yield a mixture of the diols used in the synthesis, and the mono- and dipropionates of pentaerythritol. These mono- and diesters eventually hydrolyze to pentaerythritol and propionic acid p(Fig. 3).24

Even though ortho ester linkages are much more labile than the ester linkages, the polymer is highly

/OCH2\ ,CH20\ CH3CH=C C C-CHCH, + HO-R-OH

\ OCH; \CH20/

Figure 1. Synthesis of poly(ortho ester).

POEs FOR USE IN ,

QLASS TRANSITION TEMPERATURE (C) 120

SURGICAL IMPLANTS 53

0' J 0 20 40 60 80 100

PERCENT TRANS-CYCLOHEXANDIMETHANOL

Figure 2. Effect of different molar ratios of trans-cyclohexandime. thanol and 1,6-hexanediol on the glass transition temperature of POE.

hydrophobic, so that initial hydrolysis is largely confined to the outer surface of a polymer specimen. Therefore, degradation of this polymer may not generate acidic products rapidly at the implantation site, provided that the water soluble, low molecular weight products of the first hydrolysis at the solid/liquid interface can diffuse away. This could be a significant advantage over the bulk hydrolysis of polyesters that hydrolyze directly to yield large amounts of acidic compounds, both soluble and insoluble when end-stage degradation is reached.

Polymer Characterization

POEs having 60:40 and 90:lO molar ratios of trans- cyclohexanedimethanol and 1,6-hexanediol and various weight average molecular weights were used for this work. As an aid in developing processing conditions for this family of polymers, the rheological and thermal properties of 60:40 POE (M, = 86000) and 90:lO POE ( M , = 73 000) polymers were determined. Dynamic

CI HOCH; 'CH~OH OR

HOCH2 CH2OH \ /

+ CH3CH2COOH CH3CH2COOH + C HOCH; \ CH~OH

Figure 3. Hydrolytic degradation of poly(ortho ester).

rheometry was used to determine the viscoelastic properties of the polymer samples in their melt phases. In addition, differential scanning calorimetry (DSC) at a heating rate of 10 "C/min was used to confirm the reported Tgs of the polymer samples and their propensities for thermal degradation above their Tgs . Instrumentation included a Rheometrics Inc. Model RDS-I1 Dynamic Spectrometer and an Omnitherm DSC 700.

Specimen Fabrication

Hot-Molded Bars. Quantities of both the 60:40 and 90: 10 POE polymers were reduced to predominantly coarse powders having a maximum particle diameter of less than 250 microns. This was accomplished by milling the polymer in a Thomas-Wiley Intermediate Mill at room temperature in open air and passing the material through a 40-mesh screen.

Test specimens of 60:40 and 90:lO POE (38.1 X 12.7 X 1.6 mm) were prepared by hot compression molding 0.9 g of powdered polymer in a dual-plunger stainless steel die at either 115 "C for 60:40 POE or 120 "C for 90:lO POE and 4000 psi for 5 min using a hydraulic press with heated and water cooled platens. Polymer specimens were then cooled to ambient temperature while maintaining pressure at 4000 psi.

Solvent Cast Films. 60:40 POE polymers with mo- lecular weights of 18,27,41, and 67 kDa were dissolved in methylene chloride at concentrations of 20.0,17.5,15.0, and 9.1 wt%, respectively, and stabilized with trace amounts of triethylamine (less than 0.1 wt% of the final polymer solution). Solutions were left to stand covered at ambient temperature for 12 h, then gently stirred and covered for another 12 h for degassing. This was followed by cooling the polymer solutions to 5 "C and storing for 24 h covered.

54 DANIELS ET AL.

The solutions were then hand-drawn on a polished glass plate with a variable height casting blade set at 1.27 mm. Prior to drawing, both the glass plate and casting blade were cooled to 5 "C. Af- ter casting, films were stored under a glass cover with a desiccant (CaCI2) at ambient temperature for 48 h to allow initial evaporation of solvent. Films were removed from their glass plates using a ra- zor blade, and residual solvent was removed by vacuum drying at 0.05 torr and room temperature for 72 h.

Micro-dogbone tensile specimens were die cut in accor- dance with ASTM D638 (type V).27 Specimen thickness ranged from 90 to 210 microns.

Acute Toxicity Studies

Tissue Culture Agar Overlay. Acute cytotoxicity of POEs was evaluated by standard tissue culture agar over- lay assay,28 using L929 mouse fibroblast cells and the direct cell contact method. Appropriate amounts of ethy- lene oxide sterilized (ETO) hot compression molded 60:40 and 90:lO POE polymer samples were aseptically placed onto the solidified agar, along with positive and neg- ative controls (natural black rubber and polyethylene, respectively). Samples and controls were tested in trip- licate and incubated for 24 h. Neutral Red was added to 2 to 3 h.

The cells were then evaluated using an inverted micro- scope. The results were scored by determining the area of decoloration (zone index) and the area of lysis (lysis index). The final result recorded was the response index, which is the ratio of zone index to lysis index.

USP Class VI-Systemic and lntracutaneous In- jection. Acute toxicity of tissue responses to POEs was evaluated by the United States Pharmacopeia (USP) Class VI systemic and intracutaneous tests.29 Hot compression molded 60:40 and 90:lO POE specimens for the systemic test were subdivided into specimens or required size and amounts, washed with distilled water, and sterilized with ETO.

The materials were separately extracted with sodium chloride injection, a solution of sodium chloride injection and 5% ethanol (955 v/v), polyethylene glycol 400, and cottonseed oil at 37 "C for 9 h. Extract and correspond- ing blank controls (sodium chloride, 5% ethanol saline, polyethylene glycol 400, and cottonseed oil) were injected into groups of five albino Swiss Webster mice each in the amount and route set forth in table 5 of the USP.29 Test animals were observed immediately after the injection, again at 4 h, and then not earlier than 24,48, and 72 h after injection. Nine major organ systems were evaluated for demonstrated toxic symptoms: autonomic, behavioral, sensory, neuromuscular, cardiovascular, respiratory, oc- ular, gastrointestinal, and cutaneous. Each organ system

was scored as follows: 0, normal; +1, observable but modest; +2, marked; and +3, death.

For the intracutaneous test, the fur on the back of 16 healthy, thin skinned New Zealand albino rabbits was closely clipped. Two hundred microliters of each extract was injected intracutaneously at 10 sites on each of two rabbits (left side of spine). Similarly, at 10 other sites on each rabbit (right side of spine), 0.2 mL of the corresponding blank (sodium chloride, 5% ethanol saline, polyethylene glycol 400, or cottonseed oil) was injected. The injection sites were examined 24,48, and 72 h after injection for gross evidence of tissue reaction such as erythema, edema, and eschar. A nonirritating extract was one that did not show a greater response than control blanks.

USP Class XXI-Intramuscular Implantation. Intra- muscular acute toxicity was evaluated by the USP Class XXI, 7 Day Implant Test.29 Eight appropriately sized specimens of hot compression molded 60:40 and 90:lO POE were sterilized with ETO and were implanted into the paravertebral muscle on one side of the spine of each of two previously unused New Zealand rabbits by means of a trocar. In a similar fashion, two strips of USP Negative Control Plastic RS (USP, Rockville, MD) were implanted immediately in the opposite paravertebral muscle of each animal. After 7 days implantation the rabbits were euthanized and the tissue immediately surrounding the sample material was examined for hemorrhage, film, and/or encapsulation. Each parameter was scored from 0 to 7, indicating the magnitude of the toxic response.

In Vitro Degradation Studies

pH Exposure Study. To determine if pH had any ef- fect on the hydrolytic degradation rate of bulk POE, 12 specimens of molded 60:40 POE were placed into 15-cc polystyrene vials filled with Tris-buffered saline (20 mM Tris/l50 mM NaCl) at pH 7.4, sealed, and placed into a water bath at 37 "C. Another 12 60:40 POE specimens were placed in 15-cc polystyrene vials along with Tris- buffered saline, at pH 5.0, and maintained at 37 "C. Three specimens were removed at 1,3,6, and 12 weeks from each pH and tested to failure in three-point bending, as described later.

Statistical analysis was performed using BMDP analy- sis of variance and covariance (Programs P1D and P2V, BMDP Statistical Software, Inc. Los Angeles, CA).30 Statistical significance was assumed at ,y = 0.05.

After determination of flexural mechanical properties, specimens were placed in a vacuum oven and dried at 50 "C for 48 h. Specimen dry weight was then measured on an analytical balance and recorded. Specimen initial dry weight was measured and recorded before immersion in Tris-buffered saline.

POEs FOR USE IN SURGICAL IMPLANTS 55

Sterilization. To determine the effect of radiation ster- ilization on both initial mechanical properties and the degradation rate of mechanical properties, nine molded 60:40 POE specimens were sealed in 15-cc polystyrene vials with room air and irradiated with a cumulative dose of 2.52 mrads of gamma radiation at room temperature. Irradiation was accomplished using a Cobalt 60 source, and total exposure time was 171 min. Six of the irradiated polymer specimens were then placed in 15-cc polystyrene vials with Tris-buffered saline, at pH 7.4, sealed, and placed in a water bath at 37 "C. Three specimens were removed at 3 and 6 weeks and tested to failure in three- point bending to determine flexural mechanical properties.

Additionally, three molded 60:40 POE specimens were sterilized with ETO. The prehumidification cycle was 1 h at 45% humidity, and the exposure cycle was 2.5 h with a 12% ETO concentration at 8 psig followed by degassing at 55 "C for 48 h. Change in weight average molecular weight and polydispersity was determined before and after sterilization using gel permeation chromatography (GPC) calibrated with polystyrene standards. Residual ETO was determined using a Hewlett-Packard 5890 with a flame ionization detector against ETO standards prepared weight to weight in parts per million (ppm).

Exposure to Mechanical Loads. Experimental details and fixture design have been described p r e v i o ~ s l y . ~ ~ Briefly, 18 molded 60:40 specimens were exposed to intermittent cyclic loading in aerated Tris-buffered saline (pH 7.4) at 37 "C. The specimens were immersed continually but only loaded for 630 cycles at 1 Hz/day, to simulate the activity of a sedentary postsurgical patient. This value was calculated from the time spent walking for a sedentary adult multiplied by an average ~ a d e n c e . ~ ~ , ~ ~ The load was 8.0 N. This corresponded to a stress in the outer regions of the specimen of 10% of the initial flexural yield strength of the POE polymer. Specimens were turned over every day so that both sides of the specimens were exposed to both tensile and compressive loads, and to avoid adding the effects of creep to the experiment. Three specimens were removed at 2,4,8,16,32, and 40 days and tested in three-point bending to detect changes in stress-strain behavior.

To determine the effects of a constant load on degra- dation rate, 18 specimens were subjected to a static load of 3.6 N while exposed to Tris-buffered saline at body temperature. This corresponded to a stress in the outer regions of the specimen of 2% of the initial flexural yield strength of 60:40 POE. Three specimens were removed at 2,4,8,16,32, and 40 days and tested in the same manner as the above-mentioned specimens.

Finally, to determine to what extent loading augments the effect of the chemical environment on degradation rate, nine unloaded specimens were immersed in Tris- buffered saline at physiological pH and temperature. Three specimens were removed and tested at 7,21, and 42 days.

The relationship between flexural strength and time and the relationship between modulus of elasticity and time were determined by linear regression analysis for polymer specimens immersed in saline, and grouped according to type of load (no load, static load, and cyclic load). In addition, an analysis of variance to determine equality of regression lines among load groups was performed.

Statistical software was from the BMDP package (Pro- grams P1D and PlR, BMDP Statistical Software, Inc., Los Angeles, CA).30 Statistical significance was assumed at a = 0.05.

Effect of Molecular Weight on Degradation Rate. To determine what effect molecular weight would have on POE degradation rate, solvent case 60:40 POE films (n =

40) were placed in individual 7-mL vials filled with Tris- buffered distilled water (0.05 M and pH 7.4), sealed and held at 37 "C. Five specimens of each molecular weight (18,27,41, and 67 kDa) were removed at 10,18, and 31 days, and tensile mechanical properties were measured as described below. Specimen absorbed wet mass, dry mass change, and inherent viscosity were also monitored.

Absorbed wet mass was determined by rinsing each of the specimens with deionized water, blotting the surface to remove surface moisture, and then weighing immediately on an analytical balance. To determine dry mass change, specimens were then dried for 2 days at 200 torr and 40 to 50 "C before having their final mass measured. The per- cent absorbed wet mass was calculated as follows: percent absorbed mass = (wet mass - final dry mass)/final dry mass X 100. Finally, inherent viscosity was measured for 0.5% POE/chloroform solutions (w/w) at 30 "C.

Mechanical Property Determinations

Flexure Testing. The flexural mechanical properties of hot compression molded POE specimens were determined by three-point bending in accordance with ASTM Stan- dard D 790-81 Method l.34 The diameter of the load nose was 12.6 mm. The diameter of the supports was 6.4 mm, and the distance between the supports was 25.4 mm. Span-to-depth ratio was 16:l. An Instron Model 1125 materials testing machine was used to load the specimens. Crosshead speed was 2 mm/min.

Flexural yield strength and modulus of elasticity were calculated using the following equations:

S = 3PL/2bd2

where S = stress in the outer fibers at midspan, P = load, L = support width, b = width of specimen, d = depth of specimen, and

E = L3m/4bd3

where E = modulus of elasticity in bending, L = support width, b = width of specimen, d = depth of specimen, and m = slope of the load-deflection curve.

56 DANIELS ET AL.

Tensile Testing. The tensile mechanical properties of solvent cast 60:40 POE film specimens were determined in accordance with ASTM D-638-82.27 An Instron Model 1125 materials testing machine was used to load the film specimens. Pneumatic test fixture grips with serrated faces were set 25.4-mm apart when mounting the specimens, and 9.53 mm was recorded as the gage length for each specimen. Tests were performed at 105%/min strain rate in ambient air at room temperature.

Tensile yield strength was determined as S, = L,/A, where L , is the tensile load at yield from the strip chart recording, and A is the original cross-sectional area of the micro-dogbone film specimen. The modulus of elasticity was determined from the slope of the load/displacement curve as change in stress over change in strain before the tensile yield point. Percent elongation at yield was determined by dividing the change in gage length at the yield point by the original gage length and multiplying by 100.

RESULTS

Polymer Characterization

For the 90:lO POE, the DSC scans revealed a Tg of about 98 "C and revealed that the polymer began some degree of thermal degradation when heated above about 180 "C in air. The rheological scans indicated that the viscosity of the melt behaves normally between 110 and 180 "C. Within this temperature range, the melt viscosity of the polymer is rather high, lo4 and lo6 poise, at low shear rates. For the 60:40 POE, the DSC scans indicated a Tg of about 70 "C, and indicated that the polymer began to degrade at about 130 "C. The rheological scans for this polymer indicated that it behaves normally between 85 and 155 "C. Melt viscosity was also between 104 and 106 poise.

Acute Toxicity of Hot-Molded Specimens

Both 60:40 POE and 90:10 POE were rated nontoxic in cytotoxicity (Tissue Culture Agar Overlay, Table I), sys-

temic and intracutaneous toxicity (USP Class VI systemic and intracutaneous extract injection, Tables I1 and III), and intramuscular implantation (USP Class XXI, 7-day implant, Table IV). In all cases results were comparable to negative controls.

Effect of Saline pH on Hot-Molded Specimen Degradation

The effects of exposure to Tris-buffered saline, pH 5.0 and pH 7.4 at 37 "C are shown in Figure 4. Initial flexural yield strength of 60:40 POE was 65.2 ? 0.8 MPa (mean t SD) and the initial modulus of elasticity was 1.58 +- 0.02 GPa. POE retained approximately 80% of its initial flexural yield strength and 92% of its initial modulus of elasticity after 12 weeks in vitro exposure (7.4 pH). The final values for strength and stiffness were 50.9 2 0.5 MPa and 1.45 2 0.05 GPa, respectively.

Statistical analysis of variance and covariance for the two pH showed no significant difference ( p < 0.05) in the rate of mechanical properties degradation: strength, F(1, 26) = 3.7270, p = 0.056 and modulus, F ( l , 26) = 0.9112, p = 0.394.

The weight loss after 12 weeks in vitro exposure for pH 7.4 was 0.15 & 0.05% and pH 5.0 was 0.65 2 0.11% (Fig. 5).

Effect of Sterilization on Hot-Molded Specimens

Radiation sterilization (2.5 mrad) reduced initial flexural strength by 60% (27.2 2 1.5 MPa), and had a negligible effect on initial modulus (1.65 2 0.01 GPa) and markedly increased the degradation rate (Fig. 4). After 3 weeks exposure to Tris-buffered saline at pH 7.4 and 37 "C, irradiated 60:40 POE retained 13% of its poststerilization flexural yield strength (3.5 2 0.5 MPa) and 18% of its flexural modulus (0.3 2 0.04 GPa). Sterilization of the polymer with ethylene oxide had no measurable effect on the molecular weight or polydispersity of 60:40 POE. Residual ETO was measured at 3.8 ppm. The proposed limit for small (<lo g) implants is 250 ppm. No mechani- cal property tests of degradation studies were performed

TABLE 1. Results of Standard Tissue Culture Agar Overlay Response Index = Zone Index /Lysis Index for Direct Test

Sample Replicate #1 Replicate #2 Replicate #3 Average

Results: Noncytotoxic + Control 2.O/5.Oa 2.015.0 2.015 .O 2.015.0

o.o/o.o - Control 0.0/0.0 0.0/0.0 0.0/0.0 60:40 POE o.o/o.o 0.0/0.0 0 . 010 . 0 o.o/o.o

- Control o.o/o.o 0 . 010 . 0 o.o/o.o 0.0/0.0 0.0/0.0 90:10 POE o.o/o.o o.o/o.o 0.0/0.0

Results: Noncytotoxic + Control 2.01.5.0 2.015.0 2.01.5.0 2.015.0

aZone/lysis.

POEs FOR USE IN SURGICAL IMPLANTS

TABLE II. Results of USP VI Toxicity Testing Systemic Injection

Animals Showing Signs of Toxicity No. of Extract Mice 0 h 4 h 24 h 48 h 72 h

Controls Saline 5 0 0 0 0 0

5% EtOH 5 0 0 0 0 0 Oil 5 0 0 0 0 0

Peg 400 5 0 0 0 0 0 60:40 POE-No Toxic Response Noted Systemically

Saline 5 0 0 0 0 0 5% EtOH 5 0 0 0 0 0

Oil 5 0 0 0 0 0 Peg 400 5 0 0 0 0 0

90: 10 POE-No Toxic Response Noted Systemically Saline 5 0 0 0 0 0

5% EtOH 5 0 0 0 0 0 Oil 5 0 0 0 0 0

Peg 400 5 0 0 0 0 0

Effect of Mechanical Loads on Hot-Molded Specimen Degradation

Results of cyclic, static, and unloaded exposure POE specimens to Tris-buffered saline at body ture and pH are shown in Figure 6.

57

of 60:40 tempera-

Cyclic Loading Exposure. A combination of both ex- posure to Tris-buffered saline and intermittent cyclic load- ing decreased the initial flexural strength and modulus by 75% after 40 days. The final flexural yield strength was 12.6 2 9.3 MPa and the flexural modulus of elasticity was 0.37 2 0.2 GPa.

Static Loading Exposure. Strength decreased by 29% and stiffness by 20% after 40 days. The final flexural yield strength was 46.3 ? 0.6 MPa and the flexural modulus of elasticity was 1.51 ? 0.09 GPa.

on ETO sterilized material. However, little effect is an- ticipated because no change in polymer molecular weight was noted.

Unloaded Exposure. Both strength and stiffness de- creased by less than 10% after 42 days. The final flexural yield strength was 57.6 2 1.7 MPa and the flexural modulus of elasticity was 1.51 2 0.03 GPa.

TABLE 111. Results of USP Class VI Toxicity Testing lntracutaneous Injection

Average Score Material/ Extract Control Rabbit No. Sites 24 h 48 h 72 h

5% EtOH

Oil

Saline 60:40 POE Control

60:40 POE Control

90:lO POE Control

90:lO POE Control

60:40 POE Control

60:40 POE Control

90:lO POE Control

90:lO POE Control

60:40 POE Control

60:40 POE Control

90:lO POE Control

90:lO POE Control

Peg 400 60:40 POE Control

60:40 POE Control

1 1 2 2 3 3 4 4 5 5 6 6 7 7 8 8 9 9 10 10 11 11 12 12 13 13 14 14

10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10 10

58 DANIELS ET AL.

TABLE 111. Icontinued)

Average Score Material/ Extract Control Rabbit No. Sites 24 h 48 h 72 h

90: 10 POE 15 10 010 0/0 010 Control 15 10 010 010 o/o

90: 10 POE 16 10 o/o 010 010 Control 16 10 010 010 010

h0:40 and 90: 10 POE-No Toxic Response Noted Intracutaneous

A Erythremia/edema:

For data grouped according to type of load (no load, static, and cyclic), strength and modulus for all load conditions decreased in a linear fashion (Fig. 6). Flexural yield strength decreased significantly with exposure time for no load, static load, and cyclic load conditions. Modulus of elasticity decreased significantly with exposure time for all three load conditions (Table V).

Analysis of variance indicated that the change in slope of flexural strength as a function of time differed sig- nificantly between the three loading groups, F(4,48) =

26.611, p < 0.000001. In addition, the change in slope of the modulus of elasticity as a function of time also differed significantly between no load, static load, and cyclic loads, F(4, 48) = 22.186, p < 0.000001.

TABLE IV. Results of USP Class XXI Intramuscular Implantation Test

Hemrnorrhaging Encapsulation Other

60:40 POE-Nontoxic Rabbit 1 Control 1 0 Rabbit 1 Control 1 0 Rabbit 1 Test 1 0 Rabbit 1 Test 2 0 Rabbit 1 Tcst 3 0 Rabbit 1 Test 4 0

Rabbit 2 Control 1 0 Rabbit 2 Control 2 0 Rabbit 2 Test 1 0 Rabbit 2 Test 2 0 Rabbit 2 Test 3 0 Rabbit 2 Test 4 0

Rabbit I Control 1 0 Rabbit 1 Control 2 0 Rabbit 1 Test 1 0 Rabbit 1 Test 2 0 Rabbit 1 Test 3 0 Rabbit 1 Test 4 0

90:10 POE-Nontoxic

Rabbit 2 Control 1 0 Rabbit 2 Control 2 0 Rabbit 2 Test 1 0 Rabbit 2 Test 2 0 Rabbit 2 Test 3 0 Rabbit 2 Test 4 0

0 0 0 0 0 0

0 0 0 0 0 0

0 0 0 0 0 0

0 0 0 0 0 0

0 0 0 0 0 0

0 0 0 0 0 0

0 0 0 0 0 0

0 0 0 0 0 0

Effect of Molecular Weight on Cast Film Degradation

Tensile mechanical properties after in vitro exposure to Tris-buffered distilled water are shown in Figures 7,8, and 9. Films made from the lowest molecular weight 60:40 POE studied (18 kDa) were too brittle for me- chanical testing, therefore no data are included. Films fabricated from other molecular weight POEs (27,41, and 67 kDa) possessed similar initial tensile mechanical properties, (Table VI). No change in elastic modulus was observed for any of the three molecular weight POE films after 31 days in vitro exposure. However, the 27-kDa film showed a 21% decrease in yield strength (25 2 1.8 MPa)

FLEXURAL YIELD STRENGTH (MPa) 7 0 ,

4 0 1 * pH.7.4

* pH-5.0

-A- Irradlated. pH-7.4

0 2 4 6 8 10 12 14 TIME (WKS)

A

FLEXURAL MODULUS (GPa)

I t \ 0.5 t h * pH.7.4

* pH.5.0

I 4- Irradlalod. pH-7.4

0 2 4 8 8 10 12 14 TIME (WKS)

B Figure 4. Effects of saline pH at 37 "C in vitro and radiation steril- ization on mechanical properties of hot-molded 60:40 POE. (A) mean flexural yield strength (MPa) (mean t SD). (B) mean flexural modulus (GPa) (mean i SD).

POEs FOR USE IN SURGICAL IMPLANTS 59

TENSILE YIELD STRENGTH (MPa)

j 0 4

WEIGHT LOSS (4i)

* p H 7 4

0.4

4 0

30

20

10

0 1 31 0 10 18 31 0 10 18 31 0 10 18

EXPOSURE TIME (DAYS)

POE MOLECULAR WEIGHT

87,000 (D) 0 41,000 (D) 0 27,000 (D)

Figure 7. Effect of in vitro exposure to 7.4 pH Tris-buffered distilled water at 37 "C on mean tensile yield strength (MPa) of different molecular weight 60:40 POE solvent cast films (mean 5 SD).

-0.2 ' I 0 2 4 6 8 10 12 14

TIME fWKSl

Figure 5. Percent weight loss of hot-molded 60:40 POE after in vitro exposure to saline at pH 5.0 and 7.4 at 37 "C.

showed the largest decrease in dry mass ranging from 2.1 +- 1.8% (18 days) to 1.5 -+ 0.2% (31 days), but inherent viscosity changed little (Fig. 11). Absorbed wet mass showed no increase with increasing exposure time for all three molecular weight POE films studied (Figure 12).

and a 34% decrease in elongation to yield (7.6 2 0.2%) after 31 days exposure; the 41- and 67-kDa films showed no decrease in their mechanical properties.

Specimen mean dry mass dropped for all three mo- lecular weight films studied (Fig. 10). The 27-kDa film

FLEXURAL YIELD STRENGTH (MPa) DISCUSSION . -

The primary rationale for conducting this research was to provide preliminary fabrication, biocompatibility, and mechanical property data on poly(ortho esters) in order to determine their suitability for use in surgical implants that serve a temporary function.

7 I

30 t Polymer Characterization and Fabrication

Poly(ortho esters) are completely amorphous polymers and therefore viscous. POEs having molecular weights between 70 and 90 kDa were used for hot compression molding. The DSC scans of 60:40 and 90:lO POEs indicated a thermal degradation temperature averaging about 70 "C above each polymer's Tg, respectively, when

* No Load

-6- StatIc Load

-4- Dynamlc Load

0 5 10 15 20 25 30 35 40 45 TIME (DAYS)

A

FLEXURAL MODULUS (GPa)

2 1 - r ELONGATION TO YIELD (%)

2o I 1.5

T \ -6- Static Load

-6%- Dynamlc Load

8--

0 5 10 15 20 25 30 35 40 TIME (DAYS)

B

" 0 10 18 31 0 10 18 31 0 10 18 31

EXPOSURE TIME (DAYS)

POE MOLECULAR WEIQHT

67,000 (0) 41,000 (D) 0 27,000 (D)

Figure 8. Effect of in vitro exposure to 7.4 pH Tris-buffered distilled water at 37 "C on mean etongation to yield ("h) for different molecular weight 60:40 POE solvent cast films (mean 2 SD).

Figure 6. Comparison of in vitro mechanical loading conditions on mechanical properties of hot-molded POE specimens immersed in aerated 7.4 pH Tris-buffered saline at 37 "C. (A) mean flexural yield strength (MPa) (mean c SD). (6) mean flexural modulus (GPa) (mean 2 SD).

60 DANlELS ET AL.

TABLE V. Linear Regression Equations

Flexural yield strength versus time No load: Static load:

S b = -0.2084 ( T ) + 67.21, r2 = 0.7903, p < 0.0001 S b = -0.4919 ( T ) f 65.23, r2 = 0.9260, p < 0.0000~

Cyclic load: Flexural modulus of elasticity versus time No load:

S b = -1.1727 ( T ) + 69.91, r z = 0.8131, p < 0.00001

Eb = -0.0020 ( T ) + 1.614, r2 = 0.4694, p < 0.0140 Static load: E b = -0.0069 ( T ) + 1.567, Y 2 = 0.7522, JJ < 0.0000~ Cyclic load: Eb = -0.0264 ( T ) f 1.699, r2 = 0.7215, p < 0.00001

Sb, flexural yield strength; E b , flexural modulus of elasticity; T , time.

heated in air. This represents a rather narrow processing temperature range. In addition, dynamic rheology scans of the 60:40 and 90:lO POE indicated a rather high melt viscosity near the degradation temperature for each POE, suggesting melt-processing of these polymers by extrusion or injection molding may be difficult. However, a recent report on the physical and mechanical properties of absorbable polymers suggests degradation temperatures of POEs can be extended to over 300 "C when heated in an inert atmosphere (nitr~gen).~' Higher processing tem- peratures should lower POE melt viscosity, thereby aiding polymer processing in extrusion or injection molding.

Unlike the earlier specimen^,'^ solvent cast POE films produced as described here were free of microbubbles. Films made from the lowest molecular 60:40 POE studied (18 kDa) were brittle, and cracked or broke into pieces when handled. Presumably, the brittleness was due to the lack of chain entanglement necessary for strength and toughness in amorphous polymers.

Acute Toxicity

As mentioned earlier, recent clinical report^'^,'^ describ- ing a late stage noninfectious foreign body response for devices made from PGA and PLA may be related to the accumulation of large amounts of acidic degra- dation products, which may exceed local tissue clear- ance capabilities when end-stage polymer degradation is

For completely amorphous alpha-polyesters such as poly(DL-lactic acid) and poly(1actide-co-glycolide) in bulk form, their end-stage hydrolytic degradation prod- ucts are predominantly water soluble acidic monomers and oligomers.1'~16 For semicrystalline alpha-polyesters such as poly(L-lactic acid) and poly(glyco1ic acid) in bulk form, their end-stage hydrolytic degradation products are water soluble acidic monomers and oligomers as well as water insoluble spherulitic crystal^.'^^^^ Because alpha-

polyesters show measurable mass loss only at end-stage degradation because of their bulk hydrolyzing nature, the rapid accumulation of degradation products at the implanthissue interface that exceeds local tissue clearance capabilities appears to be a plausible hypothesis.

Presumably, the acutely nontoxic responses for POE noted in United States Pharmacopeia tests of cellular, systemic, intracutaneous, and muscle implant toxicity, are due to the initial surface hydrolyzing nature of POE, releasing a more balanced combination of nonacidic, alcoholic, and acidic degradation products gradually over time. Whether or not surface erosion is the predominant form of hydrolysis for bulk POE throughout the entire degradation process has not been determined, but if so, this is a major advantage over bulk hydrolyzing alpha- polyesters.

Although the preliminary data showing a lack of acute toxicity for poly(ortho esters) are very promising, the question of long-term toxicity cannot be answered until polymer implant studies in animals have been carried out to complete implant absorption. There is a preliminary re- port on the histological evaluation of POE reinforced with degradable calcium-sodium-metaphosphate microfibers in a rabbit model with promising results. Unfortunately, the study was terminated before complete absorption of the

This has been a failing of many animal implant studies of the past involving PLA and PGA implant^.",^"

In Vitro Degradation

The lack of difference in mechanical properties for 60:40 POE exposed to Tris-buffered saline at pH 7.4 and pH 5.0 seems extraordinary at first, because incorporation of as little as 0.1 wt% of an acidic excipient such as suberic acid into the polymer results in complete erosion of the polymer in a pH 7.4 buffer in as little as 1 day. Clearly, the insensitivity of the polymer to an external acidic

TABLE VI. Initial Tensile Mechanical Properties of Different Molecular Weight Solvent Cast 60:40 POE Films

Molecular Weight P a >

Elongation to Yield (%)

27 000 41 000 67 000

34.4 2 5.7 34.7 2 3.5 34.7 2 4.4

11.5 2 2.0 12.7 2 2.0 11.6 ? 1.8

~ ~

420 2 68 412 ? 31 411 2 31

Values are mean 5 SD, n = 5

POEs FOR USE IN SURGICAL IMPLANTS 61

MODULUS OF ELASTICITY (MPa)

500

400

300

200

100

0 0 10 18 31 0 10 18 31

EXPOSURE TIME (DAYS)

INHERENT VISCOSITY (mllg)

M 0 10 18 31

POE MOLECULAR WEIGHT

w 87,000 (D) 41,000 (D) 0 27,000 (D)

Figure 9. Effect of in vitro exposure to 7.4 pH Tris-buffered distilled water at 37 "C on mean modulus of elasticity (GPa) for different molecular weight 60:40 POE solvent cast films (mean 2 SD).

environment is due to the high initial hydrophobicity of the polymer. This makes the acid sensitive ortho ester linkages virtually inaccessible to the external acid. The lack of change of mechanical properties was further supported by weight loss data, showing virtually no weight loss at pH 7.4 and only minimal weight loss at pH 5.0 in 12 weeks.

Radiation sterilization reduced initial flexural yield strength by 60%, had a negligible effect on initial modu- lus, and markedly increased degradation rate. This latter effect is due presumably to reduction in molecular weight from radiation-induced chain scission. However, it may be possible to avoid this radiation effect. Preliminary experi- ments suggest that there is little chain scission if radiation exposure is at dry ice temperature (- 50 "C) under an inert atmosphere such as argon (Personal Communication, J. Heller, SRI International, Menlo Park, CA).

Data for the pH exposure study were obtained using specimens that were kept immersed in Tris-buffered saline in the absence of any mechanical load. Although no load degradation studies are typical for bioabsorbable polymers and composites, they are not representative of actual use in mechanical stabilization of tissues. A real device is exposed to either a quasi-static load or more likely, to a markedly dynamic load. The experi-

DRY MASS LOSS (%)

5 / - - - - - - - - - - ?

3 t

50

40

30

20

10

n " 0 31 0 31 0 31

EXPOSURE TIME (DAYS)

PO€ MOLECULAR WEIGHT = 87,000 (0) 41,000 (D) 0 27,000 (D)

Figure 11. Effect of in vitro exposure to 7.4 pH Tris-buffered distilled water at 37 'C on inherent viscosity for different molecular weight 60:40 POE solvent cast films.

ments reported here involved subjecting specimens to such loads during exposure. Loads were either static or dynamic and applied in four-point bending as previously reported.31

The data indicated that mechanical properties of specimens subjected to intermittent cyclic loading decreased at a significantly more rapid rate than specimens subjected to no load or static load conditions. The increase in degradation rate with mechanical loading is likely due to the opening up of microscopic cracks when the surface of a specimen is in tension. These microscopic cracks then provide a path for water to penetrate into the device that allows polymer hydrolysis to take place with consequent deterioration of mechanical properties.

The three higher molecular weight POE films (27,41, and 67 kDa) were generally as strong or stronger than other absorbable films reported in the literature. The 27- kDa molecular weight was much stronger and tougher than the 18 kDa, perhaps indicating that a sudden transi- tion in microstructure (i.e., chain entanglement) occurs between 18 and 27 kDa. However, unlike still higher molecular weight films, the 27-kDa film showed a de- crease in yield strength and elongation to yield after

WET MASS ABSORBED (%) 10

T 2 4

1 2

n n 10 18 31 10 18 31 10 18 31

EXPOSURE TIME (DAYS)

POE MOLECULAR WEIGHT = 87,000 (D) 41,000 (0) 0 27,000 (0)

Figure 10. Effect of in vitro exposure to 7.4 pH Tris-buffered distilled water at 37 "C on percent dry mass loss for different molecular weight 60:40 POE solvent cast films (mean & SD).

" 10 18 31 10 18 31 10 18 31

EXPOSURE TIME (DAYS)

POE MOLECULAR WEIGHT = 87.000 (0 ) 41,000 (D) 0 27,000 (D)

Figure 12. Effect of in vitro exposure to 7.4 pH Tris-buffered distilled water at 37 "C on percent wet mass absorbed for different molecular weight 60:40 solvent cast films (mean -C SD).

62 DANIELS ET AL.

31 days exposure, suggesting that degradation is faster at lower molecular weights.

During in vitro degradation, specimen mass dropped, but inherent viscosity changed little, supporting the hy- pothesis that POEs degrade initially by surface rather than bulk hydrolysis, in contrast to PGA and PLA. In addition, there was no increase in absorbed wet mass with time, again supporting the hypothesis of degradation of POE principally by surface erosion.

Related Work

As stated in the introduction, the rather high complication rate of noninfective inflammatory response reported in the repair of ankle fractures using PGA pins'3 may be in response to a high rate of release of acidic degradation products. Due to their hydrophobicity and consequent surface hydrolysis, POEs may degrade much more slowly releasing more pH neutral degradation products gradually over time, thus possibly reducing complications. This could be a real advantage in fracture fixation. However, the initial mechanical properties and in vitro dynamic loading studies suggest POE by itself is an inadequate material for fracture fixation devices for long bone fixa- tion of the femur or tibiae. Because successful fracture fixation has been reported for absorbable materials that have mechanical properties similar to cortical bone,' this suggests POEs may be useful in fiber-reinforced polymer composites as a matrix polymer for fracture fixation de- vices. Reinforcement of POE with absorbable fibers would increase initial mechanical properties and composites of POE would show superior load carrying capacity.

To date, related work has included development of an absorbable composite material of 90:lO POE rein- forced with randomly oriented, crystalline microfibers of calcium-sodium-metaphosphate, with flexural mechanical properties similar to cortical bone.42 This absorbable com- posite material may be useful in low-load fracture fixation devices. Also, hydrophobic films of 60:40 POE have been used as an outer coating on absorbable composites to retard the influx of fluids, slowing loss of strength of stiffness.43 In addition, in vitro lipid exposure studies of POE films suggest the material may be useful as an interpositional barrier in the prevention of postsurgical adhesion of soft tissues.44

CONCLUSIONS

Bioabsorbable POEs, based on two diols (trans- cyclohexane-dimethanol and 1,6-hexanediol) and a dike- tene acetal (3,9-bis(ethylene) 2,4,8,10-tetraoxaspiro [5,5] undecane), can be formed easily into solid shapes, both by hot compression molding and solvent casting. Forming, handling, and storage can be accomplished in dry air with little degradation. In acute toxicity tests, hot-molded

POE specimens are rated nontoxic, producing the same response as negative controls.

Hot-molded specimens degrade slowly in saline in the absence of applied loads, losing less than 10% of their stiffness and only about 20% of their strength in 12 weeks. The degradation is virtually unaffected by decreasing saline pH from 7.4 to 5.0. This demonstrates the rela- tive hydrophobicity of POE because incorporating small amounts of acid within the polymer markedly increases the degradation rate. Degradation rates are increased sub- stantially by adding dynamic mechanical loading to the saline exposure. This may be the case for other degradable polymers also, but no data of this type for other polymers were found in the literature. Presumably, the mechanical loading opens microcracks, allowing water to enter POE specimens more easily.

Solvent cast POE films are strong in tension (30+ MPa tensile yield) and reasonably tough (12-15% elongation to yield), except at low molecular weights. The low- est film studied (18 kDa) was brittle. The higher mo- lecular weight films (41 and 67 kDa) show no decrease in mechanical properties after 31 days in physiological buffer. A film with a somewhat lower molecular weight (27 kDa) offers similar initial strength and toughness but shows a decrease in mechanical properties at 31 days. This faster degradation may be advantageous in some implant applications. The films show a decrease in weight with exposure time but no change in either molecular weight or water absorption at 31 days. This combination supports the hypothesis that the relative hydrophobicity of POEs leads them to degrade initially by surface erosion.

POEs offer a promising combination of fabricability, acute biocompatibility, mechanical properties, and degra- dation rates for surgical implants intended for temporary use. Initial strength and stiffness of POEs is less than that reported for some PLAs and PGAs. However, the mechanical properties needed must be determined for each clinical application and POE mechanical properties appear to be sufficient for many cases. Also, POEs may well retain their properties longer after exposure due to predominate surface hydrolysis than is the case for other materials such as bulk hydrolyzing alpha polyesters. The slow rate of release could reduce complications reported for large fracture fixation devices of PGA. Further studies are warranted particularly to determine long-term (i.e., 6-12 months) degradation rates and biocompatibility of degradation products.

Funding was provided: by Technology & Ventures Division, Baxtcr Health Care, Ivrine, CA; by Osteotech, Inc., Shrewsbury, NJ; by SRI International, Menla Park, CA; and by the Orthopedic Bioengineering Laboratory at the University of Utah School of Medicine, Salt Lake City, UT. Also, the authors would like to thank Dr. Nathan L. Pace for his help with statistical analysis.

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Received September 3, 1991 Accepted October, 29, 1993